Development of cell-laden 3D scaffolds for efficient engineered skin substitutes by collagen gelation

Abstract

Conventional collagen scaffolds, which were fabricated like spongy types, have been used widely to promote wound repair since they can enhance various cellular activities including cell proliferation and migration, and even guidance of near cells to work as normal tissues functionally. Fabrication technology of 3 dimensional (3D) scaffolds including solid free-form fabrication and rapid prototyping methods is developing all the time in order to promote wound repair efficiently. In addition, researchers have been studying scaffolds containing other components, such as various cells, growth factors and related materials. In this study, we modified rapid prototyping methods and then set up a cell printing system, which is able to fabricate 3D cell-laden scaffolds for better skin tissue regeneration. These scaffolds have a layered structure and were manufactured using collagen, having optimal biocompatibility, and human primary skin cells including epidermal keratinocytes and dermal fibroblasts. Accordingly, these scaffolds are capable of proliferation and migration of keratinocytes and fibroblasts effectively. Therefore, we suggest that these cell-laden scaffolds can be used as engineered skin substitute (in other words, artificial skin), sufficiently.

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Introduction

Organ transplantation is the most efficient therapy to treat damaged tissue(s) and organs. Besides, development of autograft, allograft and xenograft of transplantable organs has progressed. However, organ transplantation has limitations such as immunosuppression following surgery as well as there being a shortage of donor tissues and organs.¹ Many artificial organs have been developed to overcome the issues with immunosuppression and demand. And development of these artificial organs was based on tissue engineering, which is a multidisciplinary study with a goal of manufacturing biological substitutes that replace, maintain or improve their own tissue and organs.^2,3

Recently, researchers have been focusing on the 3 dimensional (3D) scaffolds for generation of artificial organs using synthetic and natural polymers.^4,5 The structure and properties of 3D scaffolds are critical and important factors for the generation of organs. The general requirements for 3D scaffolds are as follows: (i) suitable mechanical properties to retain the structure and function for cell proliferation, (ii) interconnected pore structure with proper size to infiltrate cell and nutrient, (iii) appropriate surface chemistry to promote cell attachment and proliferation. Such scaffolds were used to regenerate bone, vessel, tendon, skin and so on.^6–12

Among various organs, the loss of integument as the largest organ in the human body, can occur by various wounds such as abrasion, bruising, stabbing, hacking, laceration, burning and so on. Especially, thermal trauma is the most common skin defect and scalding burns can result in a severe wound rapidly over a wide area and lead to death.¹³ Therefore, in the case of burn wound healing therapy, artificial dermis has been used to regenerate defected skin. However, these artificial dermis, including acellular dermal matrix and artificial collagen dermis, have limitations such as high cost, infection derived from donor pathogen, lack of donor tissue, and secondary autograft skin graft.¹⁴ To overcome these limitations, many researchers have studied artificial dermis through various fields including cell biology,^15,16 genetics,¹⁷ material engineering^18,19 and clinical medicine.^20,21 Furthermore, studies on the scaffolds including cells and cytokines are currently on the way to reduce reconstruction time of the wound bed.

These skin wounds can be classified into 4 types; epidermal wound, superficial partial-thickness wound, deep partial-thickness wound and full thickness wound, according to increasing depth of the injury.^13,22 Currently, dressings such as MatriDerm® (Germany), PelnacTM (UK) and Integra (USA) are used extensively to treat skin injury including deep wound or broad skin defects during surgery. These dressings, called artificial dermis, were fabricated by conventional methods such as freeze-drying techniques. These artificial dermis consisting of an average thickness of 1–1.2 mm collagens look like sponges and are used for efficient re-epithelialization and revascularization of the damaged skin via increase of cell proliferation and migration. However, cell proliferation and migration inside the scaffolds take a long time to regenerate defected skin. Recently, for this reason, artificial dermis containing cells and growth factors have been researched and applied clinically to reduce the wound healing times.^23–26

Herein, we modified a conventional rapid prototyping system and developed novel cell-laden 3D scaffolds including human primary epidermal keratinocytes and dermal fibroblasts by a collagen gelation method. These cell-laden scaffolds are composed of 4 layers collagen struts: 1 layer (surface area) as an epidermis containing keratinocytes and other 3 layers (bottom area) as a dermis containing fibroblasts. Besides, these scaffolds are made up of collagen, a major component of the extracellular matrix, having optimal biocompatibility. Consistently, these 3D scaffolds exhibited noticeably enhanced proliferation in vitro. Therefore, we suggest that these scaffolds can be used efficiently as engineered skin substitute.

Materials and methods

Materials

The type I atelo-collagen (Matrixen™-PSP, Bioland Co., South Korea) was originated from porcine. The DMEM (Dulbecco’s Modified Eagle Medium) powder was purchased from Gibco (Cat. no. 12800-017, USA). Other cell culture materials were purchased from Invitrogen (USA).

For cultivation of keratinocyte and fibroblast, KGM (keratinocyte growth medium) and FGM (fibroblast growth medium) were purchased from Lonza (USA).

The 3D printing systems

The 3D printing system was composed of a 3D robotic system of two parts (DRM60, DRM130 series, Dongburobot, Korea) including a dispensing system (NEP-2000, EST, Korea) and a temperature-controllable stage to move easily at 3 axes. The stage consisted of 4 parts such as a circulating pump, a temperature controller, water chamber and a manufacturing plate. A sterilized 3rd deionized water was circulated into a plate by a circulating pump system to maintain the plate temperature at 25 °C to 60 °C. The 3D scaffolds specification was controlled by parameters of the 3D plotting system such as nozzle moving speed, nozzle tip, and pneumatic pressure. To fabricate cell-laden scaffolds, a neutralized collagen solution was placed in the barrel of the 3D plotting system. The moving speed of 29G (outer diameter: 340 μm) blunt-end nozzle was fixed at 2 cm s⁻¹ and the pneumatic pressure to extrude collagen was fixed at 150 ± 10 kPa.

The collagen solution for fabrication of 3D scaffolds

To fabricate cell-laden 3D scaffolds, porcine type I collagen was used. To make a collagen solution, the collagen powder was dissolved in 0.05 M acetic acid (pH 3.2) at final concentration of 10% (w/v). The DMEM powder was dissolved in 100 mL sterilized 3rd deionized water by enrichment up to 10 times to fabricate 10× enriched DMEM solution (10× DMEM). A 10% collagen solution and the 10× DMEM were mixed to maintain neutral pH (pH 7.0) at a ratio of 1 : 1 (v/v). Generally, it is required to neutralize collagen solutions since they are naturally acidic (pH 3.2). So, we neutralized the acidic collagen solution using 10× DMEM since 10× DMEM contained 3.7 g NaHCO₃, which is mildly alkaline in aqueous solution due to the OH⁻ group. It and 10× DMEM, which served as a buffer and nutrient, also provided an effect to protect cells from the external environment and to grow cells. These neutralized collagen solutions were mixed with cells gently to fabricate cell-printed scaffolds just on the verge of cell printing.²⁷

Mechanical properties of 3D scaffold

To confirm the mechanical properties, a compression test was performed. To test compression data, the scaffolds were cut into small strips (5 × 5 mm²) and stacked up in 4 layers. The test was conducted using a tensile instrument (Top-tech 2000; Chemilab, South Korea). The tensile test was performed in a ‘wet’ state. The stress–strain curves were recorded at a compression speed of 0.5 mm s⁻¹. All values are expressed as means ± standard deviation (n = 3).

Cell isolation and cultivation

To fabricate cell-laden 3D scaffolds, the primary cells, human epidermal keratinocytes (HEK) and human dermal fibroblasts (HDF), were obtained from MCTT (Modern Cell and Tissue Technologies, Seoul, Korea). Normal HEK and HDF, isolated from foreskins obtained through routine circumcisions, were cultured as described previously.^15,28,29 Briefly, keratinocytes and fibroblasts were isolated from neonatal foreskin obtained through clinical circumcisions. The washed foreskin biopsies were cut into the small pieces and incubated in HBSS containing 10 mg mL⁻¹ dispase (Dispase II, Gibco, USA) for 1 hour at 37 °C. The epidermis was separated from dermis, and then perturbed by pipetting after being incubated in basal medium containing 0.05% trypsin solution for 30 minutes at 37 °C to release keratinocytes. The collected keratinocyte by centrifugation were cultivated in a 100 mm culture dish with keratinocyte growth medium containing the supplementary growth factors. To collect fibroblasts, the dermis was incubated in PBS containing type II collagenase (Sigma, USA) for 1 hour at 37 °C, and fibroblasts were harvested by centrifugation. The harvested fibroblasts were cultured in the growth medium containing the supplements.

Cell isolation and cultivation for in vivo test

To test efficiency of cell-laden 3D scaffolds in an in vivo mouse model, keratinocytes and fibroblasts were isolated from ICR mouse (HSD:ICR (CD-1®), 1 day, Koatech, Kyunggi, Korea). After peeling off full skin from mouse under the anesthesia with 20 μL anesthetic drugs (zoletil (Virbac, France) : rompun (Bayer, Germany) = 1 : 2), the skin was soaked in the Defined K-SFM (Gibco, USA) including 10% FBS and 1% antibiotics. Then the skin was immersed in PBS including Dispase II (5 mg mL⁻¹) and 10% FBS for 1 hour at 37 °C to divide epidermis (keratinocyte) from dermis (fibroblast). To isolate the keratinocytes, the epidermis was soaked in PBS including 0.25% trypsin/EDTA for 12 minutes at 37 °C. These separated keratinocytes were filtered by 70 μm cell strainer to dissociate single keratinocyte after neutralizing by PBS including 20% FBS. To isolate the fibroblasts, the dermis was dipped in DMEM including 500 units collagenase type IV for 1 hour at 37 °C after being chopped into small pieces. To neutralize TE buffer, DMEM including 20% FBS was added to the medium including collagenase. Then, the separated fibroblasts were filtered by 70 μm cell strainer.

Scaffolds fabrication condition

First of all, the cell-laden 3D scaffolds were fabricated at a clean bench to maintain aseptic conditions for cell printing techniques. The cell-laden 3D scaffolds were fabricated at two different conditions to confirm the optimal manufacturing process. The pH value of the collagen solution was fixed at pH 3, 5, 6, and 7 by 10× DMEM. And plate temperature was set up at 27, 30, 33, 36, 39, 42, 45, and 46 °C.

Analysis of characteristics

The morphology of cell-laden 3D scaffolds including HDF and HEK was observed under an optical stereomicroscope (SZ 61, Olympus, Japan) and the laden cell in the scaffolds was confirmed under an inverted microscope (IX70, Olympus, Japan). And stained scaffolds sections were observed under an optical microscope (DM 750, Leica, Japan).

The mechanical properties of cell-laden 3D scaffolds including HDF and HEK were evaluated by measuring the tensile properties. The scaffolds were cut into small pieces (5 × 5 × 1.2 mm) and stacked up in 4 layers. The uniaxial analysis was carried out using a tensile machine (top-tech 2000, Chemilab, South Korea). To analyze stress–strain curves of cell-laden 3D scaffolds, the moving speed of the stretching jig was fixed at 0.5 mm s⁻¹.

Live/dead staining

To confirm the viability of printed HDF in the cell-laden 3D scaffolds after cell printing, the cells in the cell-laden 3D scaffolds were exposed to 2 μM calcein AM and 4 μM ethidium homodimer-1 (LIVE/DEAD® Fixable Stains, Life Technologies, USA) for 10 minutes in an incubator in dark conditions. The stained scaffolds were observed under a microscope and the microscopic images were captured by digital camera. A green and red color in the images indicated live and dead cells, respectively.

Cell proliferation assay

To evaluate cell proliferation of HDF in the cell-laden 3D scaffolds, each scaffolds was analyzed using a MTT cell proliferation assay (sigma, USA). The cells in the scaffolds were merged into 5 mg mL⁻¹ MTT solution for 4 hours at 37 °C. Then, absorbance was detected at 570 nm using a multimode reader (DTX 880, Beckman Coulter, USA) after the reaction was stopped using DMSO.

Culture of 3D skin collagen scaffolds

HDF (5 × 10⁵ cells per 3 layer) and HEK (3.2 × 10⁶ cells per 1 layer) were used for fabricating the 3D skin collagen scaffolds. And then an air–liquid culture method was performed to evaluate maintenance of the cells within the 3D collagen scaffolds. Briefly, the scaffolds were cut into circles of 8 mm diameter using a biopsy punch, lifted onto an insert grid (140656, Thermo Scientific, Denmark), added to the level of the grid using E-medium (consisting of Dulbecco’s modified Eagle’s medium and F12 medium in a 3 : 1 ratio plus 10% fetal bovine serum, 5 μg mL⁻¹ insulin, 5 μg mL⁻¹ transferrin, 2 × 10⁻⁸ M T3, 0.1% gentamicin, 10⁻¹⁰ M cholera toxin, and 0.4 μg mL⁻¹ hydrocortisone) and incubated in a 5% CO₂ incubator for 7 days. To obtain immunohistochemical data, the scaffolds were cryosectioned to 40 μm-thick and the sections were stained with primary monoclonal antibodies against CK-10 (MA1-06319, Thermo Scientific, Netherlands), vimentin (M0725, Dako, Denmark) overnight at 4 °C. The slides were then washed in phosphate-buffered saline and incubated with biotinylated horseradish peroxidase-conjugated secondary antibody for 2 hours. And then the slides were washed in phosphate-buffered saline and incubated with streptavidin for 1.5 hours. The sections were incubated in diaminobenzidine (Dako) until the desired staining intensity was reached. And the slides were counterstained with Fast Red. The sections were also stained with hematoxylin and eosin (H&E) for detection of cell distribution.

Full-thickness excision and grafting in mouse skin

For the in vivo animal study, ICR mice (female mice, body weight 20–25 g, 6 weeks old) were kept in the local animal care facility according to the institution guidelines. Twenty mice were included in these experiments. The mice were kept in separate cages in the animal laboratory equipped with controlled conditions to optimize animal care. And the mice had free access to rodent feed and water ad libitum under the standard laboratory guidelines.

Mice were anaesthetized with zoletil/rompun and 1 × 1 cm² full-thickness excision of skin was made on the mid back. After injury by excision, the mice were randomly divided into five groups of five mice each. Group 1 was a non-excision (normal) group as a control. Group 2 was a non-treatment group which received only vaseline gauze dressing after excision by 1 × 1 cm². Group 3 and 4 were experimental groups receiving the 3D scaffolds and vaseline gauze. Group 3 was treated with 3D collagen scaffolds, which did not include HDF and HEK. Group 4 was treated with cell-laden 3D collagen scaffolds, which included HDF and HEK.

To obtain the experimented skin samples, the mice were sacrificed in the CO₂ chamber individually. Then the skin was excised including a 2 mm margin from the experimented area. The excised skin was washed 2 times with PBS to remove blood and dirty particles and then soaked in 4% paraformaldehyde for 2 days to ensure sufficient permeation. And then, the fixed skin was dipped in 30% sucrose solution including antibiotics for 2 days. To obtain immunohistochemical data, the skin was cryosectioned to 8 μm-thick and stained using haematoxylin and eosin.

This animal study was conducted in accordance with guidelines and approval of the Institutional Animal Care and Use Committees (IACUC) of Hallym University (Hallym-2010-78).

Results and discussion

Overall structure of cell-laden 3D scaffolds

First of all, we confirmed that the structure of the scaffolds is suitable for a 4 layer strut, consisting of 1 keratinocytes layer as a epidermis and 3 fibroblasts layers as a dermis to mimic normal skin tissue (Fig. 1A). Therefore, these cell-laden 3D scaffolds (1 keratinocytes layer and 3 fibroblasts layers) were completed by a 3D robotic system as shown Fig. 1B. The cell-laden 3D scaffolds dimension is 10 mm × 10 mm × 2 mm (W × D × H). The strut and pore size of these scaffolds were measured to be 300.65 ± 29.35 μm and 294.47 ± 47.08 μm, respectively. And one cell-laden 3D scaffold contained 5 × 10⁵ cells (fibroblasts) and 3.6 × 10⁶ cells (keratinocytes).

Fig. 1 Overall structure of cell-laden 3D scaffolds. (A) A scheme of cell-laden 3D scaffolds, consisting of 1 layer (surface area) with keratinocytes and the other 3 layers (bottom area) with keratinocytes by cell printing techniques. (B) Morphology of fabricated collagen cell-laden 3D scaffolds.

Temperature optimization of cell-laden 3D scaffolds

Next, cell-laden 3D scaffolds were manufactured under various conditions, such as different temperature and pH to confirm optimal conditions for fabrication. First of all, we tried to optimize the temperature for fabrication of cell-laden 3D scaffolds. The morphology of the temperature-dependent scaffolds is as shown Fig. 2A. To analyze the morphology of the scaffolds, a strut of the scaffolds was fabricated as one layer. As a result, the optimal temperature for scaffolds fabrication is between 36 and 39 °C. In the case of the 37 °C plate temperature, the strut size and pore structure inside the scaffolds were uniform compared to other temperature conditions. Whereas, the collagen scaffolds were not fabricated totally below 36 °C or over 46 °C. In addition, when the plate was not at optimal temperature (27–33 °C, 42–45 °C), a spreading phenomenon of the collagen strut occurred. Therefore, the pore size of the collagen scaffold was nonhomogeneous and inappropriate to induce cell proliferation. In a previous study, neutralized collagen solution formed fiber at 35–40 °C quickly. On the other hand, neutralized collagen solution was not able to form a strut below ∼35 °C, because collagen fibrillation is processed slowly below ∼35 °C.^30–34 Also in our study, the neutralized collagen solution turned into water and cannot form a strut above 42 °C. This phenomenon surmised that neutralized collagen was denatured under high temperature conditions.

Fig. 2 Temperature-dependent cell-laden 3D scaffolds. (A) Morphology of one-layer cell-laden 3D scaffolds, which were fabricated at different temperature conditions without cell. (B) Strut size distribution of cell-laden 3D scaffolds at various temperature conditions. (C) Diameter of cell-laden 3D scaffolds struts, fabricated at different plate temperatures. (D) Table of relationship between temperature and diameter.

Uniform diameter distribution of the strut is also a considerable factor when the struts in the scaffolds were printed on the temperature-controllable stage at a variety of temperatures to confirm optimal gelation temperature conditions of neutralized collagen solutions. As the plate temperature increased from 27 to 39 °C, the strut size decreased to 300 μm. And also when the range of the plate temperature was between 27 to 33 °C, the strut size distribution of the scaffold was very wide. This was because of the inappropriate temperature for collagen gelation. Finally, the strut of the scaffolds was spread out on the plate and the scaffolds were not manufactured to an adequate structure. On the other hand, at 45 °C the strut was not spread adequately compared with temperatures below 36 °C although the strut size was about 300 μm (Fig. 2B).

Moreover, the strut size was noticeably smallest in fabrication conditions at 39 °C (Fig. 2C and D). These results revealed that the stage temperature of collagen gelation is optimal to 39 °C to fabricate cell-laden 3D scaffolds.

pH optimization of cell-laden 3D scaffolds

Collagen gelation depends on pH conditions as well as temperature. Therefore, we tried to fabricate cell-laden 3D scaffolds in a pH-dependent manner to find out the optimal pH. Fig. 3 shows the morphology of collagen scaffolds, which were fabricated at different pH levels, i.e. 3, 5, 6, and 7. In the case of pH 3, a strut of collagen scaffolds was not maintained uniformly and retained pore structure and adequate 3D structure for proper cell proliferation and migration. At pH 5 and 6, the scaffolds cannot be used for skin regeneration due to their swollen strut since primary human skin cells such as keratinocyte and fibroblast cannot live in acidic environment. As anticipated, the scaffolds were fabricated adequately to regenerate damaged skin when the collagen solution was at pH 7. And these conditions provided a cell viable environment.

Fig. 3 pH-dependent cell-laden 3D scaffolds morphology of one-layer cell-laden 3D scaffolds dependent on different pH of collagen solutions.

Mechanical properties of cell-laden 3D scaffolds

The mechanical properties of scaffolds are a very important element to organ transplantation.³⁵ To adapt cell-laden 3D scaffolds into the clinical field, these scaffolds should have porous structures not only with good biocompatibility but also with high mechanical strength.³⁶ Therefore, we carried out an analysis of mechanical properties using a universal tensile machine in order to confirm mechanical properties of cell-laden 3D scaffolds, which was fabricated including a keratinocyte and fibroblast in 4 layers at 37 °C. As a result, the value of Young’s moduli of scaffolds is 0.01 ± 0.001 kPa as Young’s moduli of scaffolds were measured by a tensile machine (Fig. 4). Generally, the Young’s modulus of scaffolds is low since the scaffolds were not cross-linked. However, it is necessary to increase the strength for clinical applications. Therefore, the following procedure is to enhance the strength using a cross-linking reagent in the scaffolds.

Fig. 4 Strength of cell-laden 3D scaffolds Young’s moduli of 3D cell-laden scaffolds under wet conditions.

Characteristics of cells inside cell-laden 3D scaffolds

In general, cell viability decreased during the cell printing process as cells were extruded from nozzle due to a shear stress.³⁷ First of all, we carried out live/dead cell staining to confirm that cells inside scaffolds retained their properties after 3D cell printing through collagen gelation methods. In Fig. 5A, these scaffolds were fabricated using 1 layer including fibroblast and then were immunostained with calcein AM (green, live cells) and ethidium homodimer-1 (red, dead cells). As a result, the ratio of the live and dead cells was 84.9% and 15.1%, respectively. Therefore, these conditions (from 36 to 39 °C) is a suitable temperature for cells to survive, whereas cell viability was significantly decreased above 39 °C.³⁸

Fig. 5 Activity of cells inside cell-laden 3D scaffolds. (A) The images of a live/dead cell staining after one-layer cell printing with HDF (×200). Cell-laden scaffolds are soaked in PBS including 2 mM calcein AM and 4 mM ethidium homodimer-1. (B) MTT assay of cell-laden scaffolds using a fibroblast after 1, 3, 7 days of cell printing. * is P < 0.005 and ** is P < 0.001. (C) The images of cell released and migrated from scaffolds after 5, 10, 15, 36 days of fabrication (×100). (D) The immunohistochemical images of 3D cell-laden scaffolds through air liquid culture. Each section was stained with H&E, CK-10 and vimentin for location of fibroblast and keratinocytes.

We then tested a proliferation assay using MTT solution to confirm that cells inside cell-laden 3D scaffolds can proliferate properly (Fig. 5B). The scaffolds were fabricated to three layers including only 5 × 10⁵ cell of fibroblasts. And a control group was used to culture the cells on the TCP (tissue culture plate). After 1, 3 and 7 days, each of the scaffolds were soaked in the MTT assay solution for 4 hours and then DMSO was added to stop the reaction according to directions for the MTT assay. As a result, in the 3D cell-laden scaffolds, cells are marginally increased in a time-dependent manner. But, the cells of the control group were cultured in a 6 well plate. After 3 days, the proliferation rate had not increased. This was likely due to the lack of area for cell proliferation in the case of 2D cell culture plate. However, the cell-laden 3D scaffold provides enough area for skin regeneration.

Moreover, cell-laden 3D scaffolds have the capability of release and migration of the cells inside the scaffold strut. Therefore, we observed release and migration of cells inside the scaffold strut in a time-dependent manner. Cells inside the strut released initially after 5 days of scaffold fabrication. And then the released cells were attached, proliferated and migrated into the cell culture plate. According to analysis of image J, these cells were confluent over 95% after 36 days. Therefore, these data indicated that cells inside scaffolds are able to leave the scaffolds and proliferate.

Distribution of cells including keratinocytes and fibroblasts is a key factor in order that cell-laden 3D scaffolds are used as engineered skin substitute efficiently. Therefore, the cell-laden 3D scaffolds were maintained through air–liquid culture methods and stained with CK-10 and vimentin, which is a specific marker for keratinocytes and fibroblasts, respectively. The keratinocytes inside the scaffold were mostly distributed on the surface of the scaffold similar to the epidermis, and the fibroblasts inside the scaffold were well dispersed below the keratinocytes similar to the dermis. Accordingly, these data indicated that each cell (keratinocyte and fibroblast) forming the skin maintained their position regardless of time. In other words, cell location can be controlled to regenerate damaged skin during fabrication of cell-laden 3D scaffolds. Besides, these cells in the scaffolds were not only found in the dermal or epidermal layers, but they were also released out of the scaffold and proliferated into the wound bed.

Efficiency of cell-laden 3D scaffolds in in vivo animal model

Fig. 6 shows an operation process (a) and immunohistochemistry data (b). Cell-laden 3D scaffolds were located on the wound bed, which was an excision site on the mouse mid back (Fig. 6A). Then the wound dressing was covered on the scaffolds using vaseline gauze. After 1 week, damaged skin samples are immunostained with haematoxylin and eosin. As a result, groups 3 and 4 resulted in quick and almost perfect repair, compared with group 2. The damaged skin was regenerated almost completely and clearly. Also hair follicles were regenerated almost perfectly on the wound bed. However, groups 3 and 4 were not dramatically different. We speculate that scaffolds promoted skin regeneration since the scaffolds contained collagen, but the time of cell release from scaffolds was not enough. Thus two weeks or longer might be required to confirm a satisfactory effect.

Fig. 6 Efficiency of cell-laden 3D scaffolds in full-thickness excision mouse model. (A) The process of experiment for full-thickness excision and cell-laden 3D scaffolds grafting. (B) H&E staining images of (a) normal skin (b) only full-thickness excised skin (c) full-thickness excised and only 3D scaffolds without cells grafted skin (d) full-thickness excised and only 3D scaffolds with cells grafted skin one week after operation.

Conclusions

Herein, we tried to develop cell-laden 3D scaffolds, which can be used for better regeneration of damaged skin. Therefore, we presented collagen scaffolds containing collagen and human primary skin cells including keratinocytes and fibroblasts by 3D cell printing systems. Our cell-laden 3D scaffolds are well designed for skin regeneration since collagen, a major component of the extracellular matrix, has outstanding biocompatibility and cells inside the scaffolds were constructed as a cell-laden layered structure to mimic human normal skin. Besides, skin cells inside scaffolds are certificated since these cells-containing cell therapy products are sold commercially as a cultured epithelial autograph (CEA). Moreover, the keratinocytes and fibroblasts are dispersed properly within the scaffolds for efficient skin regeneration. However, our cell-laden 3D scaffolds are not used as an engineered skin substitute (ESS) commercially since scaffolds are difficult to handle due to lower stiffness factors in spite of sufficient efficiency for wound healing. Currently we are trying to increase strength and stiffness by crosslinking cell-laden 3D scaffolds using non-toxic chemical reagents such as genipin.^39–42 Therefore, the developed cell-laden 3D scaffolds using non-toxic chemical reagents, enhance their strength and then carry out in vivo study again to be used as ESS effectively in the future.

Acknowledgements

This study was supported by a grant of the Korean Health Technology R&D Project, Ministry of Health & Welfare, Republic of Korea (A 120942).

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Footnote

† These authors contributed equally to this study.

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