Bioengineered smart trilayer skin tissue substitute for efficient deep wound healing

Materials Science & Engineering C 105 (2019) 110140

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Bioengineered smart trilayer skin tissue substitute for efficient deep wound healing

T

Swati Haldara,b,c, Akriti Sharmaa,b, Sumeet Guptad, Samrat Chauhand, Partha Roya,c, ⁎ Debrupa Lahiria,b, a

Tissue Engineering Lab, Centre of Nanotechnology, IIT Roorkee, India Biomaterials and Multiscale Mechanics Lab, Department of Metallurgical and Materials Engineering, IIT Roorkee, India c Molecular Endocrinology Lab, Department of Biotechnology, IIT Roorkee, Roorkee, Uttarakhand 247667, India d Department of Pharmacology, Maharishi Markandeshwar (Deemed to be University), Mullana, Ambala 133207, India b

A R T I C LE I N FO

A B S T R A C T

Keywords: Deep wound healing Trilayer bioengineered scaffold Skin regeneration Smart scaffold

Skin substitutes for deep wound healing require meticulous designing and fabrication to ensure proper structural and functional regeneration of the tissue. Range of physical and mechanical properties conducive for regeneration of different layers of skin is a prerequisite of an ideal scaffold. However, single or bilayer substitutes, lacking this feature, fail to heal full thickness wound. Complete scar free regeneration of skin is still a big challenge. This study reports fabrication of a trilayer scaffold, from biodegradable polymers that can provide the right ambience for simultaneous regeneration of all the three layers of skin. The scaffold was developed through optimization of different fabrication techniques, namely, casting, electrospinning and lyophilisation, for obtaining a tailored trilayer structure. It has mechanical strength similar to skin layers, can maintain a porositygradient and provides microenvironments suitable for simultaneous regeneration of epidermis, dermis and hypodermis. A co-culture model, of keratinocytes and dermal fibroblasts, confirms the efficiency of the scaffold in supporting proliferation and differentiation of different types of cells, into organized tissue. The scaffold showed improved and expedited wound healing in-vivo. Taken together, these compelling evidences successfully established the engineered trilayer scaffold as a promising template for skin tissue regeneration in case of deep wound.

1. Introduction Skin is indispensable when it comes to overall protection of the other underlying systems. Skin consists of three layers: outer epidermis, middle dermis and innermost hypodermis. Each of these layers varies in its architecture, cell composition, water content, physical properties and mechanical strength. New skin replaces dead and worn out old skin at regular intervals. While skin possesses limited self-healing capacity to mend minor bruises, the relatively superficial wounds can heal with topical medication. However, inherent regenerative property of skin is not sufficient for healing deep wounds, that involves a complex process of re-establishing functional tissue following disease or injury [1]. Commonly, deep wounds are treated with autologous skin grafting, a complete surgical process dependent on availability of donor site, which is a major bottleneck [2]. Moreover, post-surgical scars can restrict joint mobility, lead to complete/partial loss of organ function and skin enfeeblement due to misaligned architecture of the grafted skin

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[3,4]. Skin regeneration through engineered scaffolds can circumvent limited donor site availability. Skin layers differ in the compositions of their extra-cellular matrices (s. matrix) (ECM), structures, mechanical strengths and surface energies [5]. Since 1980s, polymeric scaffolds with collagen to cultivate autologous cells have been used for treating superficial wounds [6]. However, skin regenerative scaffolds need to perform beyond mere cell growth support, and provide adequate micro-niche for proliferation and differentiation of different types of cells to ensure the proper tissue regeneration. Single layer scaffolds cannot regenerate skin to its original form and quite often the regenerated tissue lacks functionality. Therefore, the idea of successful skin regeneration with bilayer scaffold has dawned upon tissue engineers and consequently, explored by several groups [7–12]. However, none of these attempts have been successful either due to lack of combination of materials to ensure proper water content of different layers of the skin or due to abrupt transition from hydrophobicity to hydrophilicity between two consecutive layers

Corresponding author at: Department of Metallurgical and Materials Engineering, IIT Roorkee, India. E-mail address: [email protected] (D. Lahiri).

https://doi.org/10.1016/j.msec.2019.110140 Received 3 April 2019; Received in revised form 11 August 2019; Accepted 26 August 2019 Available online 28 August 2019 0928-4931/ © 2019 Elsevier B.V. All rights reserved.

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Composite Implants B.V., The Netherlands) and TissueTech Autograft System (Fidia Advanced Biopolymer, AT, Italy) are the commercially available composite skin substitutes [33,54–74]. Since, these skin substitutes come with cultured cells on them, they take a long fabrication time, demand careful handling under aseptic conditions and have short shelf –lives, requiring multiple applications [15]. Thus, some of them are more appropriate as temporary dressing materials facilitating healing of partial thickness/epidermal wounds [33,58–66,70–72]. In addition, sometimes subsequent STSG is required for complete skin regeneration [58–60]. Besides, with cells from xenogenic sources some of these composite skin substitutes have been reported to have higher risk of inducing immunogenic reactions and concomitant graft rejection [55,58–60,66–68]. TissueTech Autograft System (Fidia Advanced Biopolymer, AT, Italy) is a combination of Hyalograft 3D (epidermis) and Laserskin (dermis), which are applied independently and consecutively [73,74]. Thus, holding them together in the wound site is a practical problem [75,76]. Skin substitutes up till now majorly focused on designing natural, artificial or mixed dermal and epidermal components. In case of full thickness wound healing, all the three layers of skin need to be regenerated, which can occasionally include muscular tissue as well. Current scaffolds cannot support simultaneous regeneration of different layers of skin. This is due to (1) lack of integrated dermal and epidermal components, and (2) lack of adherence between these components. Besides, there are other issues of subsequent surgical procedure (STSG), long fabrication time, short shelf life, aseptic handling, graft rejection and infection. Hence, the requirement of a simpler, easily synthesizable, less fussy, hassle-free regenerative skin substitute with a longer shelf life that will facilitate simultaneous regeneration of all the skin layers. The current study was undertaken to engineer a trilayer regenerative skin scaffold that will require a single step application at wound site, without the need for epidermal sealing following dermal regeneration. Through meticulous designing the separation of the layers upon application at the wound site will be prevented. The scaffold will be thick enough to fill up the depth of the wound providing protection from infection, besides promoting proliferation and spatial organization of different population of skin cells for natural tissue like regeneration. The design of the scaffold will be aimed at providing specific micro-niches for growth and maturation of different types of cells required to heal a deep wound. This will be achieved by modulating the physical and mechanical properties of each layer by varying their compositions and methods of fabrication. The topmost layer of the scaffold will be designed to have minimum porosity and maximum hydrophobicity like epidermis. The next layer will be fabricated in a way to enhance porosity and hydrophilicity to support dermal regeneration. Since the third layer is meant to support subcutaneous regeneration, it should be able to retain copious amount of moisture. Accordingly, the design and fabrication of the third layer will be optimized to make it the most hydrophilic out of the three layers. This will, in addition, enable it to support regeneration of tissues immediately underneath hypodermis. Uniqueness of the scaffold will be in its layerwise compositional variations intended to facilitate spatial organization of different types of cells and simultaneous regeneration, of all the layers of the skin. This design is expected to maintain a continuum between regenerated layers without affecting their individual integrity.

in the scaffold leading to layer separation during healing or due to improper biodegradation of the polymer/s used. More importantly, none of these reported studies attempted at mimicking the varying architecture of natural skin [7,8,13,14]. Nevertheless, quite a few substitutes are commercially available. They are either cellular (composed of regenerating cells), non-cellular (composed of biocompatible and/or biodegradable materials) or acellular [either allogenic (composed of de-cellularized ECM from genetically similar sources, e.g., human cadaver) or xenogenic (composed of de-cellularized ECM from genetically different sources, e.g., bovine and porcine cadavers) or composite (combination of cellular and non-cellular categories)] [15]. Cellular skin substitutes like, EPIBASE® (Genverier Lab, Sophia-Antipolis, France) and Recell® (Avita Medical Ltd. US) comprise of suspensions of keratinocytes, and keratinocytes and melanocytes from the patient, respectively, that is sprayed on the wound. Although, these substitutes do not have scarring or rejection issues, but they are not efficient in healing full depth wounds [16,17,18]. Non-cellular, biomimetic three dimensional (3D) scaffolds, like, Biobrane® (Bertek Pharmaceuticals, Morgantown, WV, US), Integra® (Integra Life Sciences, Plainsboro, NJ, US), Pelnac™ (Gunze Ltd., Medical Materials Center, Japan), Suprathel® (BioMed Sciences, Allentown, PA, USA), Terudermis® (Olympus Terumo Biomaterial Corporation, Japan) and Hyalomatrix PA® (Fidia Advanced Biopolymers, Italy) are all bilayer matrices that require split thickness skin grafting (STSG) following dermal regeneration to replace the frequently used non-bioresorbable/non-biodegradable, nylon (in case of Biobrane®) or silicone (in case of others) outer layers with epidermal sealings [19–33]. Different layers of skin maintain interfacial continuity. Since, the epidermal sealing is placed separately after dermal regeneration, the two layers fail to integrate at the interface. Thus, the regenerated skin is not similar to the natural one. Biobrane® and Integra® run the risk of infection due to accumulation of exudate underneath [19–23]. PELNAC™ also has a similar issue with an added problem of scarring [25]. Although, Suprathel® has strong anti-sepsis property and can stall bleeding efficiently, it is not very efficient in full thickness wound healing [27,28]. Hyalomatrix PA® is only good enough for partial thickness wounds and poses risk of infection [32,33]. Acellular allogenic skin substitutes, like, SureDerm® (HANS BIOMED Corporation, Seoul, Korea), Alloderm® (LifeCell Corporation, Branchburg, NJ, US) and GraftJacket® (Wright Medical Technology, Inc., USA) are variations of ECM obtained from decellularization of dermis from human cadavers. Therefore, these substitutes fairly stands up to the graft rejection issue. However, none of them can heal full thickness wound without STSG, rather they serve as the base for the subsequent surgical step [34–43]. To overcome the limited supply of allogenic sources of skin substitutes, xenogenic alternatives like, Permacol™ Surgical Implant (Tissue Science Laboratories plc, UK), OASIS® (Cook Biotech In, West Lafayette, US), Pri-matrix® (TEI Biosciences Inc., South Boston, MA, USA), Matriderm® (Medskin solutions, Dr. Suwelack AG, Billerbeck, Germany) are used. However, these skin substitutes, obtained from different species (mainly, bovine and porcine), frequently face graft rejection issues. Moreover, like their allogenic counterparts, these acellular skin substitutes are not stand-alone regenerative matrices and require subsequent STSG [44–53]. In addition, inefficient revascularization and full thickness wound healing are also reported [45–48]. Composite skin substitutes mostly comprise of a biocompatible/bioinert polymeric matrix that may or may not be biodegradable with one or more types of skin cells from allogenic or xenogenic sources growing on it. Epicel® (Genzyme Biosurgery, Cambridge, MA, USA), Epidex™ (Euroderm GmbH, Leipzig, Germany), Apligraf® (Graftskin) (Organogenesis/Novartis, Canton, MA, US), OrCel® (Ortec International, New York, USA), Laserskin®/Vivoderm (Fidia Advanced Biopolymers, Padua, Italy), Bioseed-S (BioTissue Technologies GmbH, Freiburg, Germany), CryoSkin (Altrika Ltd. UK), Hyalograft 3D® (Fidia Advanced Biopolymers, Italy), Transcyte® (DermagraftTC) (Advanced BioHealing, Inc., USA), Permaderm™ (Regenecin Inc. NJ, US), PolyActive® (Holland

2. Methods and materials 2.1. Materials Polycaprolactone (M.W. 80,000), N,N-(3-dimethylaminopropyl)-Nethyl carbodiimide (EDC) and N-Hydroxysuccinimide (NHS) were purchased from Sigma-Aldrich, India. Gelatin type A (175 bloom strength) was purchased from Alfa Aesar, India. Acetic acid (purity 99%) and dichloromethane (DCM) (purity 99.8%) were obtained from Thomas Baker, India. 3-(4, 5-dimethylthiozol-2-yl)-2,5 diphenyl 2

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shear test using the same universal low load tensile tester used for tensile testing mentioned above. Scaffold for this experiment was prepared by electrospinning layer 2 on a measured width of semi-dried layer 1 to create a zone of overlap measuring 2 mm in length. 20 × 5 mm strips of this sample were tested for overlap shear keeping the gauge length 2 mm and grip separation rate 0.1 mm/s. Thickness of the overlapping zone that varied between 1.2 mm to 1.5 mm was used to calculate the stress. Three independent experiments were done to calculate the average maximum adhesion strength. Data is also represented as stress versus displacement graph. The third layer being highly porous and hygroscopic was suitable for neither of the above techniques.

tetrazolium bromide (MTT), 4′, 6-diamidino-2-phenyindole (DAPI) and dimethyl sulfoxide (DMSO) were purchased from HiMedia, India. Growth media, FBS, antibiotics and Trysin-EDTA for cell culture were procured from Gibco, Life Technologies, India. 2.2. Fabrication of trilayer scaffold The topmost layer of trilayer scaffold (layer 1), mimicking epidermis, was synthesized using casting method. A 20 ml solution of 10% PCL was prepared in DCM. The solution was stirred at 60 °C and 450 rpm for 30 min before pouring into a glass mould. The mould was kept at 70 °C for 1 h followed by overnight incubation at room temperature. The middle layer of the scaffold (layer 2), mimicking the dermal structure of human skin was electrospun. A 20% PCL solution was prepared in glacial acetic acid. The solvent was heated up to 90 °C, then PCL was added in small batches and stirred at 400 rpm till the solution became clear and homogenous. The solution was then loaded in a 2 ml syringe and electrospun at 1.2 ml/h flow rate, under 15 kV voltage over a 10 cm tip-target distance. The middle layer was electrospun on a semi-dried top layer to ensure adherence between both the layers. The third layer (layer 3), resembling hypodermis, is a lyophilized hydrogel made from gelatin. 15% gelatin solution was prepared in deionized water by stirring at 35 °C and 400 rpm. The ensemble of the top and middle layers was spread out on gelatin solution, frozen first at −20 °C for 4 h and then at −80 °C, overnight. The assembly was then lyophilized for 36 h. The entire scaffold was finally crosslinked with EDC-NHS to prevent the highly hygroscopic gelatin from readily absorbing water or other aqueous medium and thus ensuring adequate durability of the scaffold during wound healing.

2.5. In-vitro biodegradation assay Rate of degradation of scaffold was evaluated in vitro by following the loss in weight after incubating the scaffold in phosphate buffer saline (PBS) containing collagenase (to simulate the in vivo wound environment) over a month. Initial dry weights of pieces of freeze-dried scaffolds of equal dimensions were taken. The samples were then incubated in PBS solution at 37 °C for 4 weeks. Finally the samples were removed at an interval of five days, dried and weighed to determine the percentage weight loss. 2.6. Cell viability assay The human keratinocyte (HaCaT) and human dermal fibroblast (PCS-201, HDF) cell lines were obtained from American Type Culture Collection in frozen conditions. After revival of the cells in Dulbecco's Modified Eagle's Medium (DMEM) cell viability was assessed by MTT assay. Briefly, the cells were seeded at a density of 1 × 104 cells per well in 24-well plates on UV irradiated scaffolds in triplicates. The cells were incubated in CO2 incubator for 7 days. The cell viability was assessed by adding 0.5 mg/ml MTT followed by incubation at 37 °C for 4 h. The media was then removed, and 500 μl of DMSO was added to dissolve the formazone crystals. Then 200 μl of this solution was aliquoted into a 96-well plate, and optical density measured at 570 nm using a plate reader (FLUOstar OPTIMA, BMG Labtech, Germany) to calculate the percentage cell survival using following formula:

2.3. Microstructure and surface analysis The fabricated scaffold was sputter coated with gold to make it conductive and then imaged using field emission scanning electron microscope (SEM, EVO 18 special edition, Zeiss, Germany) at an acceleration voltage of 15 kV. 2.4. Mechanical studies Tensile properties of individual top and middle layers and complete scaffold were evaluated in dry conditions at room temperature using a low-load universal tensile testing machine (Bose ElectroForce® 3200 series III test instrument). The 20 × 5 mm strips of the samples were tested with a gauge length of 2 mm and rate of grip separation of 0.1 mm/s. The sample thickness varied between 0.7 mm to 1 mm. The samples were tested in triplicates to calculate Young's modulus (E) yield strength (YS) and ultimate tensile strength (UTS) of scaffold and individual layers. Adhesion strength between first (L1) and second (L2) layers was measured through pull-off test using a portable pull-off adhesion tester (PosiTest AT, DeFelsko Inspection Instruments, USA). A composite scaffold with layers 1 and 2 was prepared as mentioned above. 10 mm dolly was affixed to layer 1 of 2 × 3 cm scaffold mat with a non-aggressive two-component solvent-less epoxide adhesive. The adhesive contained resin and hardener mixed in 1:1 ratio and was cured for 24 h before conducting the test. The glue was chosen according to ISO 4624 guidance and ASTM D 4541 standard. The dollies were thoroughly degreased, subjected to abrasion by rubbing on sand paper and subsequently cleaned before affixing to prevent glue failure during the test. The actuator of the pull-off tester was attached to the dolly adhering to layer 1 of the scaffold and load was perpendicularly applied manually. The pull-off tester was calibrated according to the manufacturer's manual prior to the test. The maximum force withstood by the dolly was recorded from three independent experiments and used to calculate the average adhesion strength. Shear stress between these layers was measured through overlap

%Cell Survival =

OD from cells grown on implant × 100 OD from cells growm on regular tissue culture well

2.7. In-vitro co-culture of keratinocytes and fibroblasts The scaffold with the casted (top) and electrospun (middle) layers was tested for its ability to support co-culture of two different populations of cells: keratinocytes (HaCaT) representing epidermis and human dermal fibroblast (PSC-201, HDF), representing dermis. Manually tailored snap well plate (Sigma Aldrich, India) was used to set up a co-culture. Briefly, the membrane of the snap well ring was replaced with the bilayer scaffold. Keratinocytes were seeded on the casted layer at a density of 1 × 105 cells and grown for 24 h at 37 °C. The structure was flipped and HDF cells were seeded at a density of 1 × 105 cells on the electrospun layer. The cells were co-cultured on the construct for 7 days following which they were fixed in 4% formaldehyde and immunostained with cell type specific antibodies. The keratinocytes were stained with anti-ZO-1 antibody (Invitrogen, India) and HDF cells with Phalloidin AlexaFluor555 (Life Technologies, India). Anti-ZO-1 antibody stained the tight junction membrane protein, Zonula Occludens (ZO-1) in keratinocytes and Phalloidin stained factin in HDF cells. Images were captured with Nikon CLSM A1R (Japan). Finally z-stacking was used to ensure that the scaffold could prevent mixing of the two cell populations during co-culture. 3

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Fig. 1. Microstructure of the multilayer scaffold resembles that of skin. [A-C] Schematic [B] shows the three layers of the scaffold in a backdrop of haematoxylin-eosin (HE) stained section of skin [Adapted from reference [82]]. The panel right next to the schematic is an optical image of the trilayer scaffold, each layer being clearly visible and marked as L1, L2 and L3. [C] Corresponding SEM images of the layers are shown in the right panel (under the heading SEM images of scaffold layers). [A] The left panel shows the SEM images of corresponding layers of skin [Adapted from references [83–85]]. [D] Cross-sectional view of the scaffold showing three layers as L1, L2 and L3 in a magnified view in [E]. [F] Graph showing percent porosity of each layer of the scaffold. [Adaptation from reference [84] reproduced with permission; references [82, 83, 85], and are under Creative Common License].

Mountant (Sigma, India). The stained sections were imaged under a light microscope (Carl Zeiss Axiovert, Germany).

2.8. Animal studies All animal experiments were done according to protocols approved by Animal Ethical Committee of MMU, Mullana (Approval number: MMCP-IAEC-17-02 dated 23/02/2017). The female adult 12–14 months old Sprague Dawley rats weighing 170–230 g, were used to evaluate the efficacy of the scaffold in healing in vivo wounds. The animals were housed in ventilated cages with sawdust beddings and food and water provided ad libitum. The temperature of the animal house was maintained at 22 ± 1 °C with 12 hour light/dark cycle. All the experiments including surgery were performed during the light phase. The study was conducted according to excision model of wound healing as reported earlier [77]. The animals were anesthetized, coat around mid-dorsal area of each animal shaven and paired wounds of 8 mm diameter each were created with a sterile biopsy punch (Tejco Vision, India). The animals were randomly divided into four groups: Normal, Positive, Standard and Experimental. Circular pieces of the scaffold measuring 8 mm in diameter were used to cover the inflicted wounds in the ‘experimental’ group of animals. The animals in the ‘standard’ group were treated with a known wound-healing drug, Neosporin while those in the ‘positive’ group were left untreated after the wounds were inflicted. The fourth group of normal animals were included to compare the histologies of the scaffold regenerated, Neosporin healed and self-healed tissues. Each group had six animals. Wound healing was evaluated by measuring the area of the wounds at intervals of 1, 3, 5, 10 and 15 days, post-surgery. After 15 days, regenerated/healed tissues were harvested, fixed in 10% Neutral Buffered Formalin [100 ml buffer containing 10 ml 37% formaldehyde, 0.8 g Sodium chloride (NaCl), 0.4 g Dipotassium hydrogen phosphate (K2HPO4) and 0.65 g Potassium dihydrogen phosphate (KH2PO4)], embedded in paraffin and 8 μm tissue sections were cut using microtome (Senior Rotary Microtome, Model: RMT-30, Radical Instruments). The sections were then stained with haematoxylin and eosin (HE). The sections were deparaffinised with xylene, rehydrated by dipping in a gradient of diluted ethanol (100%, 90%, 70%, 50%) for 2 min each, rinsed in tap water followed by staining in haematoxylin for 3 s. Excess haematoxylin was washed out for 5 min under running tap water and the sections were stained with eosin for 5 s. After a thorough wash with water, the sections were dried and mounted with coverslips in DPX

3. Results and discussion 3.1. Microarchitecture Resembling the native skin architecture as closely as possible is an important aspect for successful scaffold fabrication, because anatomical architecture defines the pattern of cell growth and adaptation that would eventually ensure formation of functional tissue [78]. A trilayer scaffold was designed for this purpose, where each layer mimicked different layers of the skin. The epidermal layer of the skin forms the outermost protective covering for the internal body organs followed by dermis. Due to high tensile strength, epidermis provides protection to the underlying layers from physical abrasions. Being made of keratinocytes, epidermis is highly hydrophobic and prevents loss of moisture from underlying layers. Therefore, to simulate the tensile strength and hydrophobicity, PCL was used to prepare first layer (layer 1) of the scaffold. Dermis is relatively more porous than epidermis to support growth of fibroblast cells. To retain the strength yet offer a porous structure, again PCL was used to synthesize the second layer (layer 2) of the scaffold. While, layer 1 was casted into a thin smooth surface with minimal porosity for growth of the polarised keratinocytes as found in the epidermal layer of skin, layer 2 was electrospun to create a nanofibrous structure to mimic the dermis, which is mainly composed of collagen fibres. Thus, the techniques used to fabricate first and second layers determined their microstructures. The third layer (layer 3) of the scaffold served a dual purpose of mimicking the lowermost hypodermal layer of the skin as well as providing a self-adhesive foundation for adherence of scaffold to the wound bed. Gelatin was thus, chosen for layer 3 as it forms a gel with adhesive properties. The consolidated architecture of the scaffold is a lyophilised structure with large interconnected pores. Layer 2 was electrospun on casted, semi-dried layer 1 to ensure adherence between these layers. This bilayer was spread out on gelatin solution in a chamber before lyophilisation. This created a suction pressure from bottom of the chamber upwards, which along with the adhesiveness of gelatin enabled physical adherence between all the three layers of the scaffold. 4

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their diverse structural morphologies, owing to the difference in their fabrication techniques [90]. Casting forms a compact structure with minimal porosity, whereas, electrospinning produces nanofibrous structures that individually have extremely low tensile strength. The casted PCL material, therefore, has high Young's modulus depicting high mechanical strength. This suffices well for this layer to serve as a barrier to environment stresses. Inner layers, made of electrospun PCL and gelatin, offer a more gentle, porous environment that is essential for cell attachment. Higher Young's modulus plays an important role in controlling the cellular behavior due to the process of durotaxis occurring in human skin cells [91]. Not only the mechanical strength of the intact scaffold is important, it is also essential to maintain a balance between the decrease in strength due to scaffold degradation and the increase in strength due to deposition of new ECM. Otherwise, it may result in abnormal phenotypic changes occurring within the tissue. Maintenance of adhesion between different layers, particularly the first two ones, of the scaffold after its implantation at the wound site is critical for successful healing. These layers are designed for gradual biodegradation and prolonged presence to support through the entire healing process. In contrast, layer 3 is meant for offering instant adhesion of the scaffold to the wound bed with rapid integration with the wound environment to provide the ECM for expedited healing through sped up cell proliferation. Therefore, being fast degrading, strength of layer 3 will not influence the cellular behavior and organization. Moreover, being a hydrogel layer, it could not be tested for tensile and shear strengths as the sample cannot be clamped or glued for load application. Therefore, the adhesion strength between first and second layers of the scaffold was measured through pull-off and overlap shear tests. The adhesion strengths thus obtained from pull-off and overlap adhesion tests were 5.15 ± 0.144 MPa and 10.43 ± 0.745 MPa, respectively (Fig. 2B, Table 2), which corroborated well with each other.

The microstructure of different layers of the scaffold was evaluated through SEM imaging. For clarity of understanding, a schematic diagram is provided to denote the locations of different layers of the scaffold corresponding to different layers of the skin. A compilation of SEM images of different layers of skin (borrowed from the duly mentioned sources) has been included in the figure for a comparative analysis. Fig. 1 represents this schematic with separate panels showing the SEM images of the surfaces of layers 1, 2 and 3 (Right: top, middle and bottom panels, respectively). SEM images of epidermis, dermis and hypodermis of natural skin are provided as references (Left: top, middle and bottom panels, respectively). An optical image of the cross section of the scaffold with all the three layers is also shown in Fig. 1. SEM images of the three layers of the scaffold show striking resemblance to respective layers of natural skin (Fig. 1, compare top, middle and bottom panels on left and right). Epidermis is non-porous with a scaly appearance (Fig. 1, left top panel). This microtopology was successfully achieved in layer 1 of the scaffold (Fig. 1, right top panel). Likewise, the porous architecture of dermis (Fig. 1, left middle panel) was nicely mimicked in layer 2 (Fig. 1, right middle panel). The highly spongy hypodermis with large interconnected pores (Fig. 1, left bottom panel) was achieved in the layer 3 of the scaffold (Fig. 1, right bottom panel). The dense structure of layer 1 is crucial in preventing entry of any foreign particle as well as to withstand external environmental stresses if the scaffold is to be implanted without any sutures [79]. Also, keratinocytes, the cells in epidermis, grow in a compact film like pattern. Thus, it is essential to have a template that supports the growth of a continuous film, at the same time maintaining some level of interconnectivity with the underlying layers. This is essential for the diffusion of nutrients, wastes and intercellular communication through signalling molecules, required for keratinocyte differentiation to maintain conducive epidermal homeostasis [80]. Hence, the porous structure of layer 2. Layer 3 requires to be highly absorbant and retentive to efficiently sponge the wound exudates while catering to the insistent moisture requirement for hypodermal regeneration. Therefore, layer 3 was made highly porous. Lyophilized gelatin being hygroscopic and absorbant was used in layer 3 [81]. The scaffold was crosslinked with EDC-NHS to maintain the structural integrity of the hygroscopic gelatin containing layer 3 which was confirmed by FTIR. The crosslinked gelatin showed sharp peaks at 3433 cm−1 and 1648 cm−1 indicating formation of NeH and C]O bonds, respectively (Fig. S1). These peaks are characteristics of amide bonding eCOHNe taking place between amine and carboxylic groups of amino acids in EDC-NHS crosslinking, thus confirming the occurrence of cross linking in gelatin [86]. Cross-linking prevented unwanted water absorption and retention by gelatin layer and thus, increased the durability of the scaffold.

3.3. Evaluation of hydrophobicity/hydrophilicity of the scaffold The hydrophilic/hydrophobic characteristic of the scaffold influences cell adhesion and cell migration during tissue regeneration to a large extent. The surface wettability of the three layers was characterized by measuring the water contact angles at room temperature. Images of water contact angles for different layers of the scaffold are shown in Fig. S2A. Higher the contact angle, higher is the hydrophobicity [92]. Layer 1 has an average contact angle of 117.91 ( ± 2.26 SE), which marks it as hydrophobic. Average contact angles of layers 2 and 3 are 79.29 ( ± 1.64 SE) and 56.3 ( ± 0.85 SE), respectively, suggesting reduced hydrophobicity. Hydrophobicity of the scaffold varies unidirectionally from top to bottom: top layer being most hydrophobic, middle layer moderately hydrophobic, while bottom layer being highly hydrophilic. This arrangement is expected to establish a gradient of moisture content in the scaffold, similar to one present in skin. The topmost layer of our skin, epidermis, is extremely hydrophobic in nature, which is critical in preventing redundant adsorption of external fluid as well as entry of foreign particles inside the body. In order to serve similar purpose, the topmost layer of the scaffold is also engineered to be hydrophobic, to prevent infection of the wound by any foreign entity. Dermis is less hydrophobic and more porous in nature. Therefore, the middle layer was designed to be relatively less hydrophobic. This moderate hydrophobicity along with a porous architecture is expected to maintain a balanced gradient of wettability across the scaffold layers, and also regulate the outflow of exudates, simultaneously. The lowermost layer was made hydrophilic in nature to serve the purpose of efficient attachment and integration on the wound site. Additionally, this layer is expected to maintain water in a three-dimensional network so as to facilitate macromolecular movement and provide efficient exchange of nutrients, gases and waste, an important requirement for tissue regeneration. The composition and method of fabrication, eventually, determine the structure of individual layer of the scaffold. This, in turn, helps in mimicking the overall trend of

3.2. Mechanical characterization of the scaffold The mechanical behavior of layers 1, 2 and complete scaffold were evaluated using uniaxial tensile loading under dry condition and plotted as stress vs strain curve (Fig. 2A). Ultimate tensile strengths (UTS) of layers 1, 2 and entire scaffold were found to be 20.45 ± 0.64 MPa, 1.40 ± 0.07 MPa and 20.71 ± 0.40 MPa, respectively and that of natural skin is 27.2 ± 9.3 MPa (Table 1) [87]. The Young's moduli of layers 1, 2 and whole scaffold were found to be 212.20 ± 3.22 MPa, 6.80 ± 0.11 MPa and 153.54 ± 17.78 MPa, respectively (Table 1). Interestingly, the Young's modulus of natural skin varies from 0.005 to 140 MPa [88]. Thus, the mechanical properties of the scaffold resemble those of natural skin. Ideally, the mechanical strength of a scaffold should be similar to that of the anatomical site of implantation. However, at the same time, it must also be strong enough to allow any kind of surgical handling and maintain its integrity from the time of implant to the completion of wound healing process [89]. The top and middle layers, though composed of similar material, showed great variation in their tensile properties. This is attributed to 5

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Fig. 2. Mechanical properties of the scaffold. [A] Stress vs strain curve of layers 1 (solid black line) and 2 (broken grey line), individually and the complete scaffold (broken dotted grey line). [B] Stress vs displacement curve showing adhesion between layers 1 and 2.

mechanical properties and moisture distribution of natural skin in the scaffold.

Table 2 Adhesion strength between first (L1) and second (L2) layers of the scaffold.

3.4. Scaffold degradability The scaffold, proposed in this study, is made of all biodegradable polymers. This was to make sure that the scaffold gets automatically removed, without any secondary/removal surgery, after regeneration is completed. To study degradation behavior, the fabricated scaffold was incubated in phosphate buffer saline (PBS) containing collagenase, at 37 °C over a period of one month and the loss in weight of the scaffold at regular intervals was recorded and represented as percentage (%) weight loss as a function of time (Fig. 3). Approximately, 75% of weight loss was observed within initial 5 days, while at the end of the experiment, about 10% of initial mass remained. This is consistent with the essential temporary nature of scaffold that needs to be eventually replaced by the native tissue [93]. Basically, biodegradation occurs when a biomaterial interacts with tissue fluids under physiological conditions. As evident from the curve, the scaffold shows rapid degradation in the initial few days, following which degradation slows down. This pattern of degradation could be attributed to high volume of gelatin present in the scaffold structure. Gelatin is a natural polymer and degrades faster than PCL, which is a synthetic biopolymer [94]. The degradation of gelatin helps its incorporation in the extracellular matrix, thus facilitating wound healing [95]. Additionally, the adhesive properties of gelatin help in cell attachment and provide a nutrient rich

Name of the test

Adhesion strength (MPa ± SE)

Pull-off test Overlap shear test

5.15 ± 0.144 10.43 ± 0.745

Table 2 with the values of adhesion strength between layers 1, 2 measured through pull-off and overlap shear tests. Data is represented as mean ± standard error (SE) of values obtained from three independent experiments.

environment for cell growth. An average of three weeks are required for deep skin wound to heal. Complete disintegration of the scaffold before three weeks would not suffice its purpose as a supportive template. However, the relatively slow degradation of PCL is expected to help maintain the structural integrity of the scaffold throughout the process.

3.5. In vitro biocompatibility evaluation Cell viability on scaffold was determined using MTT assay. Layers 1 and 2 were seeded with HaCaT (keratinocytes) and PCS-201 (human dermal fibroblast, HDF) cells, respectively. Same cells grown on tissue culture dish were taken as control. Actively growing cells convert MTT to formazan crystals that when dissolved in DMSO produces a purple solution. The intensity of purple colour and thus, the corresponding optical density at 570 nm (OD570) are directly proportional to the

Table 1 Mechanical properties of the scaffold in comparison to human skin. Mechanical properties

Layer 1

Layer 2

Trilayer

Human skin

Ultimate tensile strength (UTS) (MPa ± SE) Young's modulus (E) (MPa ± SE)

20.45 ± 0.64

1.40 ± 0.07

20.71 ± 0.40

27.2 ± 9.3I

212.20 ± 3.22

6.80 ± 0.11

153.54 ± 17.78

0.005 − 140II

Table 1 with the values of ultimate tensile strength (UTS) and Young's Modulus (E) of layers 1, 2 and complete scaffold. Data is represented as mean ± standard error (SE) of values obtained from three independent experiments. I, IIValues reported from references [87,88], respectively. 6

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either sides showed cell population specific morphologies: keratinocytes spread as a sheet on the non-porous layer 1 while fibroblasts colonized the fibrous layer 2 (Fig. 4C, panels b & c). This strongly suggests that the scaffold would be efficient in restoring the native architecture of the damaged skin during tissue regeneration. Both, keratinocytes and HDF cells showed improved proliferation in the coculture setup as compared to control cells growing on tissue culture flasks, since the number of cells growing on the scaffold was higher than the control cells (Fig. 4B, compare panels b with a and e with d). Probably, through paracrine signalling two cells types facilitate and support each other's growth and modulate their functions. The fibroblasts, on interaction with its ECM, secrete various kinds of signalling molecules that support keratinocyte proliferation. These signalling molecules are also responsible for the characteristic arrangement of keratinocytes that occur after differentiation to form a polarised epithelium exhibiting tight junctions. The porous fibrous structure of the electrospun layer provides fibroblasts a microenvironment similar to native ECM, building the secretome required for keratinocyte differentiation. Percent cell survival of keratinocytes and dermal fibroblasts when determined through MTT assay was found to be comparable to respective controls and not better. For MTT assay, the keratinocytes and fibroblasts were grown independently on layers 1 and 2, respectively. Since, the cells were not co-cultured, the effect of paracrine signalling in the form of improved cell growth was not observed.

Fig. 3. Biodegradation of the scaffold. Graph showing the percentage (%) weight loss of the scaffold in the presence of PBS over a period of thirty days. Each data point represents the average % weight loss ± standard error (SE). Samples were evaluated in triplicates for each time point.

number of surviving cells and represented as percentage survival relative to the control. The proportion of viable cells grown on the scaffold were similar to that of control (tissue culture flask), for both the cell lines, indicating that the scaffold material was not toxic to the cells and that it provided a substratum almost similar to tissue culture dish (Fig. 4A). This data confirmed that the degradation products released from the scaffold were thus non-toxic and hence, did not jeopardize metabolic activities of the cells growing on them. The ultimate goal of the fabricated scaffold is to sustain cell growth, maintain different populations of cells in segregated layers, particularly, preventing the highly proliferative fibroblasts from infiltrating the epidermis, thus, ensuring original tissue structure and composition in the regenerated skin. For this purpose, a co-culture model was set up using snap well plate. The well membrane was replaced with the scaffold, consisting of layers 1 and 2. The layer 1 was seeded with keratinocytes followed by seeding of layer 2 with fibroblast cells. In an independent experiment, these layers with cells growing on them were subjected to SEM imaging (Fig. 4B, panels c & f). The cells were also differentially immunostained for exclusively cell specific proteins: Zona Occludens (ZO-1) for keratinocytes and f-actin for human dermal fibroblasts (Fig. 4B). Keratinocytes, seeded on casted PCL (layer1), when stained with anti-ZO1 antibody, were found in a layer across the surface the scaffold (Fig. 4B panel b), showed a typical keratinocyte specific growth pattern as shown in control cells (Fig. 4B panel a). The SEM image revealed that these cells colonized the surface of layer 1 as a sheet (Fig. 4B panel c). Similarly, the electrospun fibrous layer 2 of the scaffold efficiently supported the typical fibroblastic growth (Fig. 4B panel d) of the HDF cells as evident from the Phalloidin AlexaFluor 555 stained f-actin fibres (Fig. 4B panel e). HDF cells also infiltrated the cross-section of the electrospun layer 2, adhered to the nanofibres and proliferated in three dimensions. SEM image clearly displayed these cells to be intertwined and dispersed within the fibres of layer 2 with actin fibres protruding from them (Fig. 4B panel f). Immunostaining also indicated the growth of healthy cells on the scaffold. Surface localization of ZO-1 in keratinocytes grown on layer 1 was improved and f-actin mesh was more prominent in fibroblasts grown on layer 2, relative to those grown on tissue culture dish (Fig. 4B, compare panels a & b, c & d). Immunostaining further confirmed the integrity of this coculture of cell types within the scaffold. Z-stacking of the scaffold with keratinocytes and fibroblasts, growing on each side of the scaffold, showed no mixing of the cell populations (Fig. 4C, panel a), during coculturing. In a supporting experiment, the SEM image of the transverse section of the scaffold with keratinocytes and fibroblasts grown on its

3.6. In-vivo study In the next part of the study, the efficiency of the scaffold in healing wounds was evaluated in vivo on excision model. The animals were divided into three groups: Positive (with untreated wounds), Standard (where the wound was treated with standard drug, Neosporin), and Scaffold (where the trilayer scaffold was implanted on the wound site). The digital images of the gross appearance of wounds for all categories are shown in Fig. 5A for different time intervals, post-surgery. The rate of wound closure/reduction in wound diameter was used as a parameter to check the healing efficiency of the scaffold. After 15 days of surgery, < 20% of the wound was closed in the animals belonging to the untreated group (Positive group). Almost 60% wound closure was observed in case of the animals treated with Neosporin (Standard group) while, this percentage was nearly 90 for the animals treated with the scaffold (Scaffold group) (Fig. 5B). Further, insight into the regenerative capacity of the scaffold was obtained by analysing the histological sections of the regenerated tissues. The HE stained sections of wound tissue from positive control group showed presence of inflammatory and granular cells (labelled as G in panel II of Fig. 5C), indicating the wound to be in its inflammatory phase. It was marked by clear absence of differentiated layers of skin. Occurrence of epidermal hyperplasia was prominent, indicating absence of directional tissue regeneration. Absence of required framework for the infiltered fibroblast inhibited the same from developing focal adhesion points that regulate cell-signalling process. (Fig. 5C, panel II). In the Neosporin treated section from standard group, the wound healing process was marginally enhanced as compared to positive control, as the number of inflammatory and granular cells had reduced indicating that healing process has moved beyond the inflammatory phase. The formation of the dermal and hypodermal layers have also been initiated (Fig. 5C panel III). However, regenerated epidermis is lacking which is also evident from less efficient wound closure in the Neosporin treated group as compared to the scaffold treated one. The cross section of scaffold implanted tissue showed all the three regenerated layers of skin with no traces of inflammation (absence of granular cells) (Fig. 5C panel IV). Although collagenation was apparent in both drug and scaffold treated sections, however, the drug treated section lacked a well-defined epidermis. Presence of prominent epidermal layer in scaffold treated section illustrates paracrine activity taking place amongst the cells leading to differentiation of 7

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Fig. 4. Biocompatibility of the scaffold. [A] Histogram percent (%) survival of keratinocytes and fibroblasts on layers 1 and 2, respectively of the scaffold by MTT assay. Each data point represents mean ± standard error (SE) of values from three independent experiments. [B] Healthier proliferation of keratinocytes and fibroblasts co-cultured on the scaffold is shown through representative immunostaining of Zona Occludens (ZO-1) (green) (panels a & b) and f-actin (red) (panels d & e). F-actin microfilament shown with yellow arrow (panel e: inset). Morphologies of the co-cultured keratinocytes and fibroblasts cells are shown through SEM images (c & f, respectively). Fibroblast adhered to electrospun layer 2 and surrounding fibres of the scaffold are shown with yellow and red arrows, respectively (panel f: inset). [C] Representative zstacking of scaffold with co-cultured keratinocytes (stained ZO-1) and fibroblasts (DAPI stained nuclei) showing separate layers of these cells in the scaffold (a); SEM image showing co-cultured distinct keratinocyte sheet and fibroblast mesh (b). Yellow arrows showing the jagged edges of the sectioned layer 1 and red arrows depicting fibroblasts that have infiltrated layer 2. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

performance skin regenerative biocompatible material. This vascularization was found to be completely absent in positive control (Fig. 5C, compare panels II & IV). Deep wound healing involved skin grafting until recently, before

keratinocytes thus, ruling out the probability of epidermal hyperplasia, while minimizing the occurrence of scar development. Efficient vascularisation in scaffold treated samples was yet another hallmark of wound healing, confirming the efficacy of the scaffold as a high-

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Fig. 5. The scaffold can heal wounds efficiently in vivo. [A] Optical images showing the inflicted wounds and their closure over fifteen days in animals belonging to positive, standard and scaffold groups. [B] Wound diameters noted over fifteen days in different groups of animals. [C] Haematoxylin-Eosin (HE) stained transverse sections of skin tissues from normal (a), positive (b), standard (c) and scaffold implanted (d) groups of animals. E: epidermis, D:dermis, H: hypodermis, B: blood vessels (shown with black arrows), G: granular cells, collagenation shown with green arrows. * Denotes statistical significance wrt positive group for wound closure. ** and *** represent p-value < 0.05 and < 0.001, respectively.

regeneration of all layers of the skin and promote vascularization (Fig. 5). For this the fabrication technique was suitably tuned to ensure successful assemblage of the different layers making sure that they do not fall apart when applied in vivo. The scaffold architecture provided cues for directed layer specific skin tissue regeneration. The final product is a thick mat that can fill up the wound providing protection with ensuing scar free tissue regeneration. The first layer, made of casted PCL is hydrophobic and mimics the architecture of epidermis. The electrospun second layer being porous supports the dermal regeneration. The third layer made of lyophilized gelatin is highly hydrophilic and regenerates hypodermis. The continuum of increasing hydrophobicity from layer 3 upwards matches well with the moisture content profile of the skin. These layers, although adhered to each other on-site, due to their microstructural differences prevented infiltration of cells from one layer into other in vitro. As was evident from the data that the fibroblasts were restricted to the electrospun fibrous layer 2, while the keratinocytes spread as cobbleshaped sheet on casted non-porous layer 1. In vivo, this structure of the scaffold ensured that no fibroblast mediated epidermal hyperplasia are

different types of biocompatible/biodegradable scaffolds became more preferable. Some of these skin scaffolds are also available commercially for clinical use in patients. However, none of these scaffolds is capable of simultaneous regeneration of different layers of skin let alone, other damaged tissue in a deep wound. Efficient healing of deep wound, with its successful regeneration, requires not only cell proliferation but their organization also. Thus, an ideal regenerative scaffold for deep wounds should (1) serve as an instant cover-up for the gash, (2) prevent infection, (3) expedite wound closure followed by healing, (4) ensure proliferation of different types of cells, and (5) facilitate proper arrangement of these cells during regeneration to ensure scar free healing. The trilayer regenerative scaffold, discussed in the current study, not only facilitated wound closure, but also ensured restoration of original architecture and composition of the skin. Various compositions of the layers of the scaffold leading to different microstructures, helped to meet the specific requirements of each regenerating tissue type. The three distinct layers of the present scaffold have been synthesized through different methods like, casting, electrospinning and lyophilisation to generate micro-environments suitable for supporting

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developed. Moreover, restricting fibroblasts within layer 2 generated focal adhesion points for keratinocytes. Crosstalk between focal adhesion points and hemi-desmosomes in keratinocytes induces paracrine cell signalling, which in turn, promotes keratinocyte migration, maturation and eventual wound closure [96]. However, the fibrous nature of layer 2 facilitated adhesion, proliferation and required migration of fibroblasts within the constraints of the layer and development of blood vessels in vivo [97]. Moreover, the fibrous nature of layer 2 provides an option of modulated drug release to prevent infection and inflammation [98]. Due to least hydrophobicity and high moisture absorbability, gelatinous layer 3 ensures rapid exudate sequestration and adherence of the scaffold to the wound bed, besides, providing the physical and chemical support for hypodermal regeneration.

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4. Conclusion To summarize, the present study chronicles the development of a tri-layer PCL-gelatin scaffold, with similarity to actual skin layers in terms of architecture, physical and mechanical properties. The in vitro co-culture model, established using keratinocytes and fibroblast cells, indicated that the multilayered scaffold was able to maintain each cell population segregated while allowing their proliferation and ECM deposition. In vivo results confirmed wound-healing efficacy of the scaffold and showed that the regenerated tissue exhibits morphology similar to that of native skin. To conclude, the high porosity, good mechanical properties of the construct, together with the biocompatibility, biodegradability and enhanced, wound healing properties make the construct highly suitable for application in skin tissue engineering. The current data thus warrants further detailed preclinical/clinical studies to appreciate the unique features of this scaffold that can be used in clinical applications in near future. Declaration of competing interest There are no conflicts of interest associated with this publication. Acknowledgments The authors appreciate the support from Dr. Murali Kumarasamy, Center of Nanotechnology, Indian Institute of Technology Roorkee. The authors also wish to thank the laboratory staff of the Centre for Nanotechnology and the Department of Metallurgical and Materials Engineering, IIT Roorkee, for maintaining the experimental facilities. Appendix A. Supplementary data Supplementary data to this article can be found online at https:// doi.org/10.1016/j.msec.2019.110140. References [1] R.F. Diegelmann, M.C. Evans, Wound healing: an overview of acute, fibrotic and delayed healing, Front. Biosci. 9 (2004) 283–289. [2] R. Guo, S. Xu, L. Ma, A. Huang, C. Gao, The healing of full-thickness burns treated by using plasmid DNA encoding VEGF-165 activated collagen-chitosan dermal equivalents, Biomaterials 32 (2011) 1019–1031. [3] P.P. van Zuijlen, J.J. Ruurda, H.A. van Veen, J. van Marle, A.J. van Trier, F. Groenevelt, R.W. Kreis, E. Middelkoop, Collagen morphology in human skin and scar tissue: no adaptations in response to mechanical loading at joints, Burns 29 (2003) 423–431. [4] T.A. Wilgus, A.M. Ferreira, T.M. Oberyszyn, V.K. Bergdall, L.A. Dipietro, Regulation of scar formation by vascular endothelial growth factor, Lab. Investig. 88 (2008) 579–590. [5] S. Deepthi, K. Jeevitha, M. Nivedhitha Sundaram, K.P. Chennazhi, R. Jayakumar, Chitosan-hyaluronic acid hydrogel coated poly(caprolactone) multiscale bilayer scaffold for ligament regeneration, Chem. Eng. J. 260 (2015) 478–485. [6] C.J. Doillon, Skin replacement using collagen extracted from bovine hide, Clin. Mater. 9 (1992) 189–193. [7] J. Pu, F. Yuan, S. Li, K. Komvopoulos, Electrospun bilayer fibrous scaffolds for enhanced cell infiltration and vascularization in vivo, Acta Biomater. 13 (2015) 131–141.

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