Natural 3D-Printed Bioinks for Skin Regeneration and Wound Healing: A Systematic Review

Three-dimensional bioprinting has rapidly paralleled many biomedical applications and assisted in advancing the printing of complex human organs for a better therapeutic practice. The objective of this systematic review is to highlight evidence from the existing studies and evaluate the effectiveness of using natural-based bioinks in skin regeneration and wound healing. A comprehensive search of all relevant original articles was performed based on prespecified eligibility criteria. The search was carried out using PubMed, Web of Science, Scopus, Medline Ovid, and ScienceDirect. Eighteen articles fulfilled the inclusion and exclusion criteria. The animal studies included a total of 151 animals with wound defects. A variety of natural bioinks and skin living cells were implanted in vitro to give insight into the technique through different assessments and findings. Collagen and gelatin hydrogels were most commonly used as bioinks. The follow-up period ranged between one day and six weeks. The majority of animal studies reported that full wound closure was achieved after 2–4 weeks. The results of both in vitro cell culture and in vivo animal studies showed the positive impact of natural bioinks in promoting wound healing. Future research should be focused more on direct the bioprinting of skin wound treatments on animal models to open doors for human clinical trials.


Introduction
Tissue damage or injury is a severe health problem that annually accounts for around half of the world's annual health care expenditure [1]. The wound healing mechanism is an immediate protective process that intervenes after the body suffers injury. During this process, damaged or destroyed tissues are disposed of, the vulnerabilities of skin tissues are managed, and skin integrity is restored [2,3]. This process, however, requires excellent patient care and suitable wound coverage. Although traditional wound dressings (i.e., gauze, lint, plaster, and bandages) shield the wound from contaminants, those dressings require frequent changing to avoid neighboring tissue maceration, in addition to their tendency to adhere to the injury, which makes it painful when replacing [4].

Skin Bioinks
The vast majority of the used wound healing bioinks were gelatin and collagen. Although gelatin hydrogel has high rheological properties, it showed zero viscosity at temperatures above 27±1°C [21], and all gelatin studies have examined the use of different crosslinking agents [21,23,26,28,32,33,35,36]. On the contrary, four of the six studies reported the ability to print collagen hydrogel without the need for chemical crosslinking agents [25,29,31,34]. The integration of alginate hydrogel with either gelatin [26,28,32] or honey [30] was also reported.

Collagen-Chitosan blends
Evaluating the rheological and printability of collagen-chitosan composite as a potential bioink.
In vitro NIH 3T3 cells NHS/EDC [19] CNF/GelMA Utilizing the use of deficient GelMA concentrations as supporting materials to CNF-based bioink In vitro Mouse 3T3 fibroblasts Ca +2 to crosslink CNF UV light to crosslink GelMA [23] Sulfated and Rhamnose-rich XRU Developing polysaccharide modification of 3D bioprinted XRU extract and evaluate its validity.
In vitro Human dermal fibroblasts (HDFs) Photo-crosslinking by UV light [20] dSIS slurry Studying the physicochemical and biological properties of dSIS bioink. In vitro Normal skin fibroblasts (NSFs) EDC [24] Viscoll Collagen Evaluating the impact of different collagen concentrations on viscoll to produce high fidelity constructs In vitro NIH 3T3 No crosslinking applied [25] Alginate/Gelatin Investigating the rheological behavior of alginate/gelatin as a complex construct.

In vitro AECs and WJMSCs
Two-steps gelation: a) Gelatin crosslinked by low temperature; b) Alginate crosslinked by Ca +2 [26] BCNFs+ SF/Gelatin Enhancing the resolution and the mechanical performance of SF/gelatin scaffolds.

In vitro & in vivo L929 cells & 12 mice
BCNFs work as a crosslinking agent [33] Fibrinogen and thrombin/Collagen I   (1) Cell proliferation assay on the 10% XRU hydrogels showed a 6.3-fold increase in HDFs cell number two weeks post-culture; (2) Coating XRU with collagen, further promoted cell proliferation with a 7.5-fold increase in cell number 14 days post-culture.

High
The dSIS scaffold developed in the study can be a potential candidate for the application of skin defects with a high level of fidelity and rapid swelling ratio.

Low
The introduction of nanofibers from bacterial cellulose had a low impact on the printability of the composite bioinks.

High
In G8-G12 gelatin scaffolds, HDFs cell growth rates were approximately 14% higher than in the G6 gelatin scaffold. The mechanical properties were highly dependent on the pore size. [21]
One the 1st day, L929 cells exhibited a slightly slower growth on SS/GelMA scaffolds of 0.5, 0.33, and 0.2 GelMA in comparison to the control group. While on days 7 and 14 after culture, cell growth was delayed on both matrices and the control group. HaCaT and HSFs cell viabilities were exhibited higher on the scaffolds containing more SS.

High
The inclusion of silk sericin (SS) in the matrices was shown to promote HSFs cell growth. The study also suggested that SS/GelMA is suitable for HaCaT cell culture application as it showed high cell viabilities after seven days. [35] G-SF-SO 3 -FGF2 Scaffold porosity: The method explained, but no results presented On the first and third days, similar proliferation rates were noticed by CCK-8 assays with and without FGF2. On the 5th day, proliferation rates were enhanced significantly of almost 40% increase after treating with FGF-2.

High
Using 100 ng/mL of FGF2 led to ã 40% higher proliferation rate. Sulfonated SF coated scaffold promoted cell adhesion, proliferation, and growth. [36] Collagen The study found that FBs and KCs can be evenly printed layer-by-layer as a dermal-like layer and epidermal-like layer.
The 3D printing technique provides high dimensional control for engineering skin tissues.  BCNFs + SF/gelatin After seven days, cells could grow under the surface of the printed line at a range of 160-220µm. The hierarchical pore structure of the printed line allowed sufficient space for cell growth.

weeks
The findings showed that the arrangement of pore structure is beneficial for nutrient supply for the ingrowth of tissue post-implantation in vivo. [33] Fibrinogen and thrombin/Collagen I One-week post-surgery, the wound area was 66% of the original wound area in contrast to the control group wound area, which remained at 95% (n = 12). Two weeks post-surgery, the wound area was 15% of the original wound area, and the control group wound area was 40% (n = 8).

10-14 days
In situ 3D bioprinting of autologous cells accelerated the process of wound healing in approximately three weeks in comparison to other treatments. [31] S-dECM Three-weeks post-surgery, S-dECM bioink accelerated wound closure as it consists of different growth factors and cytokines capable of accelerating wound healing. Besides, cells encapsulated dECM accelerated wound re-epithelialization two weeks post-surgery.

SS/GelMA
The immuno-histochemical observation of IL-6 and TNF-α cytokines indicated acute inflammatory on the 7th day and decreased on the 14th day and hardly found on the 28th day.

weeks
Although further in vivo investigations are needed to validate the material, SS/GelMA hydrogel scaffolds represent possible candidates for the application of wound healing and tissue engineering. [35] Gel-SF-SO3-FGF2 Two-weeks post-surgery, the epithelial cells tended to migrate from the skin edges towards the wound center in the G-SF-SO3 group. Meanwhile, the dermis and epidermis layers were almost wholly repaired in the 3D G-SF-SO3-FGF group. On the 28th day post-surgery, the wound defect was completely closed in both G-SF-SO3 and G-SF-SO3-FGF2.

2-4 weeks
FGF2 growth factor enhanced the wound healing, re-epithelization as well as promoting blood vessel formation, and expression of various corresponding markers. [36] Gelatin-alginate Post-surgery, the scaffold treatment group showed a significant decline in the wound area. The wound diameter decreased from 0.8 cm on the 1st day to 0.2 cm on the 14th day. The whole wound was nearly healed with almost no crust. On the 14th day, the control group seemed to be covered with hard black crusts, and the mean wound diameter was 0.7 cm. In comparison to the control group, the treatment group formed granulation tissue with uniform and layered wound thickness, which indicates that the scaffold support cell migration and proliferation.
14 ± 1 day The use of gelatin-alginate was found to decrease wound bleeding and perfusion post-implantation. The scaffold also found to facilitate wound maturation and healing.
For 3D bioprinting in animal studies, the studies included around 151 animal subjects. Each study included 12-40 animals, but one study [34] did not disclose the number of animals used. Four studies reported the use of mice [31][32][33][34], two studies reported the use of rats [35,36], and one study reported the use of porcine [31].

Skin Bioinks
The vast majority of the used wound healing bioinks were gelatin and collagen. Although gelatin hydrogel has high rheological properties, it showed zero viscosity at temperatures above 27 ± 1 • C [21], and all gelatin studies have examined the use of different crosslinking agents [21,23,26,28,32,33,35,36]. On the contrary, four of the six studies reported the ability to print collagen hydrogel without the need for chemical crosslinking agents [25,29,31,34]. The integration of alginate hydrogel with either gelatin [26,28,32] or honey [30] was also reported.

Biocompatibility Measures
Most of the natural-based bioinks were reported to have excellent biological properties. Thirteen of sixteen in vitro studies reported high cell proliferation rates. Even though significant changes in proliferation rate were not evident in three studies [19,22,35], they reported high cell viability. Seven studies reported good cell viability [20,22,24,25,29,30,33], five reported a minimum of 85.07-98% cell viabilities [22,25,26,29,34], and one reported some dead cells indicating low cell viability [24].
Furthermore, fourteen studies reported high cell growth, and only dSIS slurry [24] and SS/GelMA [35] bioinks were found not to facilitate cell growth. All in vivo studies results showed excellent matching with in vitro studies results except for SS/GelMA [35], which showed unique wound healing property after two weeks post-treatment.

Quality Evaluation
The risk of bias of the included studies was conducted using a modified version of the OHAT. In general, the experimental conditions of all reported bioinks were duly mentioned, and almost all studies have low reporting and performance risk of bias. Five of the six in vivo studies have a low risk of bias due to reporting outcome details and fulfilling the selection criteria. Four of twelve in vitro studies showed a low risk of bias as well. In contrast, eight studies have a moderate risk of bias due to the lack of skin cell representation and short follow-up periods, and only one study was found to have a high risk of bias due to high reporting and selection biases (i.e., finding was not clear, adverse events and probability values were not reported; follow-up period, statistical analysis, and outcomes measures were not suitable). The results of the risk assessment are summarized in Table 4. Checklist

Overview of the Included Studies
This systematic review shows that natural 3D bioprinted skin substitutes can promote full wound closure based on the pooled results from 18 in vitro cell culture and in vivo animal studies. Most of the 3D bioprinted skin substitutes facilitated cell proliferation, adhesion, and differentiation, and most in vitro studies reported high cell viabilities. Moreover, all animal studies declared total wound area reduction on animals wounded dorsal two weeks post-surgery. However, beyond the limits and practical concerns of evaluating in vitro cell culture and in vivo animal studies and comparing these results to human needs, it must be accepted that animal studies encompass the first level of evidence.
The primary objective of using 3D bioprinting in wound healing is to apply the rapid treatment directly to the injured tissues. Albanna et al. have successfully printed fibrinogen and thrombin/collagen I incorporated HFBs and HKCs directly on the dorsal of mice and porcine models (Figure 2). This study resulted in accelerating the process of wound healing in approximately three weeks in comparison to other treatments. The immunohistochemistry study revealed that HFBs and HKCs were found, together with endogenous cells, within the dermis and epidermis layers of the wound 3-6 weeks post-surgery [31]. strength and low viscosity above 27±1°C, and that limits gelatin usage in 3D bioprinting. It is often mixed with other natural biomaterials, such as alginate [26,28,32] and silk-fibroin [33], to overcome the low formability. Moreover, gelatin methacrylate (GelMA) is also a potential wound healing bioink due to its high thermal sensitivity and photo-crosslinking ability. GelMA is also known to have good biocompatibility, and of promoting cell to cell interaction and cell migration. Furthermore, the advantageous mechanical stability of GelMA after UV crosslinking was used to induce a high shape fidelity of natural-based bioinks, such as cellulose nanofibrils [23] and silk sericin [20]. Geometric information obtained via scanning is then inputted in the form of an STL file to orient the scanned images to the standard coordinate system; (d) The scanned data with its coordinate system is used to generate the fill volume, and the path points for nozzle head to travel to print the fill volume; (e, f). Output code is then provided to the custom bioprinter control interface for generation of nozzle path needed to print fill volume. Figure and caption reused from Albanna et al. [31]. Used under the Creative Commons License (http://creativecommons.org/licenses/by/4.0/).

Alginate
Alginate has been used in different 3D bioprinting applications because of its high shearthinning and rapid gelation post-printing. However, alginate has many limitations as crosslinking delay may reduce the shape fidelity of the bioprinted constructs, low cell viability as rapid crosslinking limit cell-to-materials interaction. An attempt was conducted by Datta et al. to overcome those limitations by decreasing alginate viscosity using honey to increase cell viability without altering alginate printability. While alginate is qualified for most of the physicochemical properties needed for 3D bioprinting, it suffers poor cell adhesion properties, requiring efforts to enhance the cell adhesion without sacrificing the physicochemical properties [30]. Printing simple alginate solutions were found to have low shape fidelity, although researchers attempted to increase alginate viscosity or extrude it with chemical crosslinkers such as Ca +2 [26]. (c) Geometric information obtained via scanning is then inputted in the form of an STL file to orient the scanned images to the standard coordinate system; (d) The scanned data with its coordinate system is used to generate the fill volume, and the path points for nozzle head to travel to print the fill volume; (e,f). Output code is then provided to the custom bioprinter control interface for generation of nozzle path needed to print fill volume.

Bioinks Materials & Combinations
Many types of natural-based bioinks, composite or stand-alone materials, have been proposed to restore the skin integrity and accelerate the wound healing process due to their desirable properties, such as resembling skin ECM, high printability, and excellent biocompatibility as hydrogels are the most commonly used biomaterials [14].

Collagen
Collagen, as a hydrogel, exhibited desirable biodegradability, high shape consistency at 37 • C, and excellent microstructure of micro-and macropores that promote cellular attachment and proliferation [29]. However, collagen direct 3D bioprinting is still limited as collagen solutions have poor printability, especially when incorporated with cells or tissue spheroids [25]. Notably, despite the limited collagen printability, no chemical crosslinking was applied over most of the studies. Instead, this property was overcome by either admixing with other materials such as fibrinogen and thrombin [31], chitosan [19], by using fibrillar collagen [29], by using low concentrations of collagen (2-4%) [25], or by controlling cell suspensions and densities [22]. In the same context, proteins gelation of matrices such as collagen is usually initiated by pH or temperature control or by both. Although this approach is valid for thin structures, it showed diffusion or thermal transference limitations in thick structures (1 to 3 mm), which may lead to the appearance of gelled and non-gelled regions. High levels of pH or temperature may also lead to severe harm to cells [22].

Gelatin
Gelatin is another commonly used bioink that presented high degradability, biocompatibility, and suitable rheological properties. Nevertheless, pure gelatin solutions have weak mechanical strength and low viscosity above 27 ± 1 • C, and that limits gelatin usage in 3D bioprinting. It is often mixed with other natural biomaterials, such as alginate [26,28,32] and silk-fibroin [33], to overcome the low formability. Moreover, gelatin methacrylate (GelMA) is also a potential wound healing bioink due to its high thermal sensitivity and photo-crosslinking ability. GelMA is also known to have good biocompatibility, and of promoting cell to cell interaction and cell migration. Furthermore, the advantageous mechanical stability of GelMA after UV crosslinking was used to induce a high shape fidelity of natural-based bioinks, such as cellulose nanofibrils [23] and silk sericin [20].

Alginate
Alginate has been used in different 3D bioprinting applications because of its high shear-thinning and rapid gelation post-printing. However, alginate has many limitations as crosslinking delay may reduce the shape fidelity of the bioprinted constructs, low cell viability as rapid crosslinking limit cell-to-materials interaction. An attempt was conducted by Datta et al. to overcome those limitations by decreasing alginate viscosity using honey to increase cell viability without altering alginate printability. While alginate is qualified for most of the physicochemical properties needed for 3D bioprinting, it suffers poor cell adhesion properties, requiring efforts to enhance the cell adhesion without sacrificing the physicochemical properties [30]. Printing simple alginate solutions were found to have low shape fidelity, although researchers attempted to increase alginate viscosity or extrude it with chemical crosslinkers such as Ca +2 [26].

Skin-decellularized Extracellular Matrix (S-dECM)
Extracellular matrix (ECM) represents the non-cellular part of a tissue or an organ, and it mainly assembles the microenvironment network for the cell to perform specific functions. Each tissue has its well-constructed ECM, which consists of several components and proteins that maintain the native structure and support cell migration. Interestingly, the ECM can be derived by using an appropriate protocol and reused as a scaffold for tissue regeneration [37]. Kim et al. successfully decellularized porcine skin-tissue and formed a printable dECM bioink. They found that, in comparison to collagen bioink, the 3D bioprinted skin equivalent using derived ECM bioink promoted dermal compartment stabilization, enhanced epidermal organization, and provided more physiological relevant skin functions in vitro. Moreover, dECM-based 3D skin encapsulated EPCs, and ASCs promoted neovascularization and re-epithelialization as well as wound closure in vivo [34].

Bioink Biocompatibility & Cellular Behavior
Bioinks biocompatibility was duly investigated, and some of the possible reasons that may affect cell viability, adhesion, proliferation, migration, and differentiation were reported. In general, cytotoxicity should be evaluated when proposing a potential material for medical use. Most of the included studies performed MTT assay to ensure no cytotoxicity or inflammation caused by the cell-to-materials chemical interaction. Notably, only silk sericin/GelMA bioink was found to cause acute inflammation on the 7th day, which disappeared at the end of the follow-up period [35].
Bioink pore size should also be considered when choosing a bioink as small pore sizes cause a lack of nutrition and oxygen supply, which led to low cell viability and slower cell migration. Choi et al. studied the effect of gelatin pore size on cell behavior and found that the proliferation rate of HDFs increased by 14% in pore size of 580 µm compared to 435 µm after 14 days [21]. However, using natural bioinks is favorable because of their suitable inter-molecular network. For example, fibrillar collagen is well-known to have a suitable micro-and macropores structure, which was found to highly intervene in increasing cell viability and promoting high cell attachment and proliferation [29].
In the same context, bioink concentration crucially affects cell viability as high concentrations lead to compacted cells. An evaluation of the impact of using different collagen concentration into viscoll on cell viability, found that decreasing collagen concentration from 4% to 2% resulted in increasing the cell viability from 87.2% ± 2.1% to 97.2% ± 1.2% (p < 0.05) [25]. Nocera et al. studied the effect of using smaller collagen extract on NIH 3T3 cell viability and found that decreasing the concentration from 100-extract to 25-extract promoted cell viability from 85.07 ± 6.73% to 111.31 ± 3.65% (p < 0.05) [29]. Xu et al. also studied the effect of admixing small concentrations of GelMA with cellulose nanofibrils (CNFs) on cell proliferation. They found that three days after culture, there was twice the number of cells on CNF/GelMA bioink compared with CNF bioink alone [23].
On the other hand, growth factors are essential morphogenetic proteins that influence cell activity and direct tissue repair and regeneration [38]. Xiong et al. studied the effect of using a fibroblast growth factor (FGF2) on cell proliferation. They found that adding 100 ng/mL of FGF2 growth factor to the scaffold significantly enhanced the proliferation rate (~40% to~75%), tissue morphology, and the assembly of the collagen fibril ( Figure 3) [36].

Structural Design & Mechanical Properties
An effective bioink should possess excellent mechanical properties and should not breakdown post-printing. A bioink should also have high swelling ratios to maintain moisture wound area to exchange nutrients and facilitate cell proliferation. In the literature, human skin was found to have a young's modulus average of 100 to 1100 kPa [34]. The swelling ratio has an inversible relationship with young's modulus values, whereas increasing the dSIS filament distances from 500 to 700 μm increased the swelling ratio from 69% to 79% and decreased the Young's modulus from 26.6 ± 3.8 to 9.7 ± 3.1 kPa (p < 0.05) [24]. The same results were reported with CNF crosslinked BDDE [27] and alginate/gelatin [28].
Bioinks should maintain their shape once they leave the tip of the printing nozzle. Overall, proper bioink viscosity ensures high shape fidelity and minimizes the possibility of structural collapse after printing [32]. Shear-thinning is another critical parameter as bioinks should have excellent shear-thinning properties to avoid clogging during the printing process and to regain immediate structural consistency post-printing to be ready to support the next layer [19,26,27]. For example, a period of 1 min was required to ensure the transformation of collagen to gel-state to preserve a solid base for the printing of the next layer [22]. Additionally, the rigidity of the printed scaffolds appeared to affect cell proliferation profoundly. As the rigidity of CNF increases within a tunable range of 3-8 kPa, cell proliferation was promoted [27].

Animal Models & Wound Healing
Early treatment of wounds is critical to avoid wound aggravation and tissue damage over time, due to the hypertrophic scarring. Patients undergoing late treatment often experience severe scarring. Cell suspension densities is another critical factor, as using high densities cause low cell viability. Lee et al. reported the use of an inkjet bioprinting system and studied the effect of using different cell suspension densities and droplet size on cell viability. The study found that cell viability varied proportionally with cell suspension density and inversely with the space between droplets for both keratinocytes (KCs) and fibroblasts (FBs) skin cells. At very low cell suspension density (0.5 million cells/mL) and large droplet spacing (400 mm), FBs cell viability was moderate (84%). Similarly, at a high cell suspension density (5 million cells/mL) and small droplet spacing (400 µm), KCs cell viability was lower (94%) [22]. Moreover, cell adhesion is profoundly affected by the matrix thickness, whereas a higher percentage of cell attachment was observed with 3 mm samples than with 2 mm thick samples. A large thickness scaffold promoted cells to adhere [27].

Structural Design & Mechanical Properties
An effective bioink should possess excellent mechanical properties and should not breakdown post-printing. A bioink should also have high swelling ratios to maintain moisture wound area to exchange nutrients and facilitate cell proliferation. In the literature, human skin was found to have a young's modulus average of 100 to 1100 kPa [34]. The swelling ratio has an inversible relationship with young's modulus values, whereas increasing the dSIS filament distances from 500 to 700 µm increased the swelling ratio from 69% to 79% and decreased the Young's modulus from 26.6 ± 3.8 to 9.7 ± 3.1 kPa (p < 0.05) [24]. The same results were reported with CNF crosslinked BDDE [27] and alginate/gelatin [28].
Bioinks should maintain their shape once they leave the tip of the printing nozzle. Overall, proper bioink viscosity ensures high shape fidelity and minimizes the possibility of structural collapse after printing [32]. Shear-thinning is another critical parameter as bioinks should have excellent shear-thinning properties to avoid clogging during the printing process and to regain immediate structural consistency post-printing to be ready to support the next layer [19,26,27]. For example, a period of 1 min was required to ensure the transformation of collagen to gel-state to preserve a solid base for the printing of the next layer [22]. Additionally, the rigidity of the printed scaffolds appeared to affect cell proliferation profoundly. As the rigidity of CNF increases within a tunable range of 3-8 kPa, cell proliferation was promoted [27].

Animal Models & Wound Healing
Early treatment of wounds is critical to avoid wound aggravation and tissue damage over time, due to the hypertrophic scarring. Patients undergoing late treatment often experience severe scarring. Before proposing the capacity of using an effective wound treatment, it must demonstrate high biocompatibility and non-cytotoxicity in vitro. It should also stimulate wound healing and tissue re-epithelization in vivo. Bioprinting human cells resulted in rapid epithelialization represented by 4-5 weeks of acceleration time of wound re-epithelialization [31].
Layering skin constructs in regular pore size and structure significantly influenced nutrition supply and cell ingrowth in the wound area [33]. To ensure scarless wounds, the treatment should be placed evenly in an organized manner on the wound. Xiong et al. reported that the application of the gelatin-sulfonic acid-FGF scaffold on rats' wounded dorsal helped to smoothen the wound post-surgery, and the cross-sectional results showed complete wound closure in addition to the existence of more blood vessels [36]. Furthermore, the cross-sectional area treated by SS/GelMA showed the formation of new collagen with high fibroblast proliferation similar to the healthy tissue seven days post-surgery followed by complete wound closure on the 4th week, thus proving excellent wound healing properties [35].
For the survival or integration of the new tissue or organ into the surrounding tissue, suitable vascularization is required. Many attempts have been made to build vascularized skin scaffold by using natural-based biopolymers [36] or by printing with interconnective pores sizes between 50 and 500 µm and micropores with diameters lower than 10 µm [29], or by decellularizing skin ECM [34].

Limitations of the Present Review
This systematic review has several limitations. No specific risk of bias checklist was found to assess in vitro studies. Instead, the OHAT tool was adapted to evaluate both in vitro and in vivo studies. Furthermore, using 3D bioprinting for wound healing is still undergoing animal studies, and no human randomized clinical trials were identified. Another limitation is that the observation time and measurements vary among studies, which causes high heterogeneity in the results. Hence, a meta-analysis was impossible to be performed.

Conclusions and Future Perspectives
This systematic review identified the potential in vitro cell culture and in vivo animal studies reporting 3D bioprinting skin substitutes. First of all, this review confirms the significant benefits of using 3D printed natural-based bioinks for skin repair and regeneration. Natural bioinks showed excellent ability to mimic the three-dimensional microenvironment structure of native skin tissue and to promote cell adhesion, proliferation, migration, and mobility. Furthermore, in vivo visualization showed full wound closure four weeks post-surgery with well-organized dermal and epidermal layers. This review reported the importance of many bioink properties that should be found to accelerate the wound healing process for a better therapeutic approach. We recommend the use of natural bioinks for wound healing and suggest performing more in vitro studies with the use of a variety of skin cell representations other than dermal fibroblasts, which is known to survive the harsh environment.
Despite the limited number of conducted studies, in situ bioprinting is one of the most promising advances in skin tissue engineering, which can be used by surgeons to print complex organs efficiently and rapidly. Yet, the main challenge is the ability to build tissue details more precisely which required the integration of different fields, including engineering, biology, and medical science. In addition, some new cross-linking techniques, such as two-photon cross-linking and directed on tip UV light, might promote structural control over the existing bioinks. Self-healing hydrogels constitute another interesting direction as they can be printed, retain their pre-vascularized microstructure, and can be used as self-healing scaffolds for wound healing.