Treatment of extraordinarily large deep skin defects remains a great clinical challenge. Today’s “gold standard” approaches to cover such skin defects are primarily split- and full-thickness skin autografts, as well as skin flaps, skin expansion techniques, and dermal substitutes [62
]. Laboratory-grown skin substitutes offer a novel promising treatment option for patients suffering from severe, full-thickness skin injuries [20
]. Those patients need an artificial skin substitute due to the shortage of healthy donor skin sites for autografts.
Therefore, our laboratory-Tissue Biology Research Unit (TBRU) at the University of Zurich, Switzerland, and several laboratories worldwide have developed dermo-epidermal skin substitutes (DESSs) containing both—dermal and epidermal skin layers [12
]. Our hypothesis is that a skin substitute closely resembling native human skin can yield more satisfactory clinical results both functionally and cosmetically. Of note, the addition of the missing dermal component significantly enhanced the mechanical stability and mesenchymal-epithelial interaction of those skin substitutes.
Moreover, our laboratory also investigated the mechanical modifications of the dermal compartment for better surgical handling [71
]. Consequently, we developed a plastically compressed hydrogel based on a collagen type I matrix serving as a dermal template for DESSs [71
]. Uncompressed collagen hydrogels are fragile and fold when manipulated with forceps, but after plastic compression, they are significantly more stable, which improves handling. Importantly, this dermal template has been successfully used for the establishment of large DESSs (7.5 × 7.5 cm) and was eventually tested in a pig animal model [36
Recently, these autologous human DESSs (DenovoSkin) were successfully used in phase I clinical trial on 10 children at the University Children’s Hospital, Zurich [75
]. We recently started a phase II clinical trial in various burn and reconstructive surgery centers in Switzerland (University Children’s Hospital Zurich, University Hospital Zurich) and Europe using those autologous skin grafts [75
However, the prevalent challenge facing this innovative approach is the lack of sufficient vascularization to support the survival and viability of the above-mentioned DESSs after transplantation. This aspect is of utmost importance for the treatment of non-healing chronic wounds.
In a pilot study, we overcame these hurdles by the in vitro generation of capillary networks in DESSs using endothelial and mesenchymal progenitors derived from the stromal vascular fraction (SVF) of human adipose tissue (Figure 2
]. This approach holds great promise for future clinical applications as adipose tissue represents a convenient, abundant, and easily accessible cell source [78
]. Moreover, this concept was previously successfully used in the engineering of precisely sized osteogenic constructs, increasing the efficiency and uniformity of bone tissue formation [79
Importantly, the rapid onset of blood infiltration in the skin substitutes had efficient effects on promoting epithelial and dermal tissue repair in vivo by (1) increased collagen type I expression, (2) elevated cell proliferation rate of both dermis and epidermis, (3) improved graft take rate, (4) reduced expression of wound healing markers such as cytokeratin 16 (CK16) and cytokeratin 17 (CK17), and (5) rapid achievement of epidermal homeostasis. These results confirm that such a co-culture-based pre-vascularization strategy of tissue-engineered skin grafts emerges as an efficient method to noticeably improve wound healing, cell engraftment, and skin function after transplantation [20
3.2.1. Cell-Laden Hydrogels as Wound Dressings
More recently, attention has been attracted to engineering hydrogels as skin templates to encapsulate cells and bio-macromolecules to support cell–cell and cell–microenvironment interactions.
Due to their attractive properties, hydrogels are the most common materials used as a scaffold to culture cells for skin repair applications. Hydrogels can provide an appropriate scaffold for cell encapsulation due to their 3D matrix, which is replete with water and their biodegradability properties. Additionally, the majority of them are biocompatible, which means the hydrogel does not provoke any adverse reaction, rejection, or immune response upon implantation [80
]. Furthermore, mechanical properties of hydrogels including stiffness, viscoelastic behavior, and initial state recovery (self-healing) are tunable by copolymerization, nanoparticle incorporation, and changing the polymer concentrations and ratios [3
]. Moreover, hydrogels are biomimetic structures, mimicking the natural microenvironment enabling cell–cell interactions and interactions with the surrounding tissue [85
There are some commercial cell-laden hydrogels available on the market. TransCyte (Advanced BioHealing, Inc., New York, NY, USA and La Jolla, CA, USA) is a bio-engineered skin wound dressing, formerly marketed as Dermagraft-Transitional Covering [87
]. This dermal substitute is composed of human newborn fibroblasts, which are then seeded on the nylon mesh of Biobrane with a thin silicone layer regulating moisture vapor from the wound. TransCyte can be applied to the wound site and is protected by adhesive strips; in some cases, surgical staples are used [88
In general, Figure 3
represents a concept of a cell-laden hydrogel for skin tissue engineering. First, cells are isolated from a patient and cultured in a hydrogel matrix. The hydrogel supports the growth of cells and skin formation, which can be grafted back to treat the skin defect of the patient. Cell-laden hydrogels are usually prepared by mixing the isolated cells in a pre-polymer solution followed by a crosslinking using an ionic or chemical crosslinking such as a thermal or photo-crosslinking mechanism. However, the type of photoinitiator, light, and temperature need to be controlled to protect the cells and increase their viability after crosslinking [3
There are different kinds of cell-laden hydrogels, which can be divided based on their scaffold structure such as porous and stimuli-responsive hydrogels. In porous hydrogels, a porous bioscaffold containing cells forms a foam or crosslinked hydrogel, which can be applied as a skin substitute onto a wound site [91
]. In particular, porosity plays a critical role to allow host cell infiltration into the 3D network and to improve protein transport and diffusion to mimic native tissue structure and function. For instance, if the pores are too small, they might be blocked by cellular penetration, ECM formation, and vascularization of the inner areas of the scaffold [81
]. When designing a hydrogel scaffold for tissue engineering purposes, pore-related parameters including morphology, volume, size, distribution, throat size, wall roughness, and the interconnectivity of pores are important [93
It has been demonstrated that the optimum pore size is 5–15 μm for ingrowth of fibroblast, 20 μm for hepatocyte ingrowth, and 20–125 μm for the regeneration of adult mammalian skin [94
]. Thus, pore size and distribution should be taken into consideration while engineering hydrogels for skin tissue engineering and to mimic native skin structure.
On the other hand, several scaffolds have been developed using stimuli-responsive materials that can release the encapsulated cells and biomolecules into the host tissue when triggered by different internal or external stimuli. To synthesize such a scaffold, first, a stimuli-responsive polymer is mixed with cells; second, the pre-gel solution is applied via injecting/spraying on the wound site [96
]. Then, due to the external stimuli like thermal or photo stimulators, the polymer forms a physical gel leading to cell encapsulation within a 3D scaffold.
The in situ formation of cell/hydrogel scaffold structure facilitates the delivery of encapsulated cells, growth factors, and necessary nutrients at the wound site via using negligibly invasive techniques [97
]. Various studies have been conducted using stimuli-responsive materials. For instance, in the study of Eke et al. [98
], a UV-crosslinked biodegradable hydrogel was employed as a scaffold containing adipose-derived stem cells (ADSCs) to stimulate vascularization in difficult-to-heal wounds. In this study, methacrylated gelatin (GelMA) and methacrylated hyaluronic acid (HAMA) were used to synthesize the hydrogel network. Afterward, a photoinitiator, to induce photo-crosslinking, and cells were added simultaneously into the pre-hydrogel solution (Figure 4
a). The mechanical stability of composite hydrogels was engineered by varying the GelMA and HAMA concentrations and ratios. Although the modulus of this hydrogel (6 kPa) is lower than that of native skin, it proved to be easy to handle and manipulate, which is important in the laboratory and in the clinics. Further, in vitro results showed that those hydrogels provide an appropriate microenvironment for the proliferation of ADSCs. Additionally, in vivo studies demonstrated that stem-cell-loaded hydrogel scaffolds significantly improved vascularization at the wound site compared to their cell-free counterpart groups (Figure 4
3.2.2. 3D Bioprinting of Cell-Laden Hydrogels for Wound Dressings
There are various approaches to produce cell-laden hydrogels for skin engineering [99
]. Recently, 3D printing of cell-laden hydrogels emerged as a novel fabrication technique. This method involves the printing of hydrogel with cells in a layer-by-layer manner to fabricate a complex bioscaffold [100
]. The main advantage of this method in skin engineering is the ability to develop clinically relevant skin constructs that closely mimic the native skin architecture and heterogeneity. However, the success of bioprinting for skin regeneration is strictly dependent on the engineering of appropriate printable bioinks to support the function of cells and stimulate the fabrication of new ECM after printing. Hydrogel-based materials are one of the most promising bioinks for skin regeneration applications, as they have some unique properties such as their tunable rheological behavior, which is vitally important for delivering cells by printing [89
Several review papers have extensively discussed the main aspects of 3D bioprinting and its different techniques [101
]. Here, we mention some recent studies regarding cell-laden hydrogel-based skin substitutes developed via 3D bioprinting.
Although there are various hydrogels used for bioprinting, they are mainly restricted to natural polymers such as alginate, collagen, gelatin, fibrin, and hyaluronic acid [104
]. In a study by Cubo et al. [106
], the authors used a 3D bioprinter to produce a human-plasma-derived bilayered skin for the treatment of burn injuries and traumatic and surgical wounds. The authors used primary human fibroblasts and keratinocytes that were obtained from skin biopsies [106
]. The dermal part was formed by printing human fibroblasts embedded within a plasma-derived fibrin hydrogel. Their results showed the production of skin equivalents with structural resemblance to the human skin, confirmed by the presence of fibroblasts spread within the dermal compartment and the terminal differentiation of keratinocytes.
In another study, Yanez et al. [107
] employed the 3D bioprinting technology to integrate capillary-like endothelial networks into a dermo-epidermal skin graft including neonatal human epidermal keratinocytes (NHEKs) and neonatal human dermal fibroblasts (NHDFs), both embedded in a fibrin–collagen hydrogel matrix. In this work, human dermal microvascular endothelial cells) were mixed with thrombin and printed on top of a manually plated layer of collagen-NHDF cells containing fibrinogen. After synthesis of the fibrin hydrogel, a layer of collagen-containing NHEK cells was pipetted on top of it to create a bilayered network. In order to take into account in vivo considerations, printed structures were implanted into skin full-thickness wounds on the back of athymic nude mice to examine the healing process. Wound treating behavior was compared with control (no treatment) and Apligraf (discussed previously) groups. Wounds healed with printed substitutes needed 14–16 days to heal, contrasting with 21 days in the control group and 28 days in the group implanted with Apligraf [107
]. Moreover, histological characterization demonstrated the formation of dermal and epidermal skin layers comparable to the native skin, which is accompanied by the presence of new microvessels in the mouse tissue. The authors concluded that the neoangiogenesis was triggered mainly by the presence of endothelial cells that were seeded in the skin graft. In addition, human keratinocytes which are known to secrete various angiogenic growth factors stimulated vascularization [107
Further, Hakimi et al. [108
] developed a handheld skin printer that allowed in situ formation of skin tissue sheets of different homogeneous and architected compositions (Figure 5
). They demonstrated that this system is compatible with dermal and epidermal cells incorporated with ionic crosslinkable alginate, enzymatically crosslinkable proteins (e.g., fibrin), and their mixtures with collagen type I and hyaluronic acid. Additionally, in a study by Liu et al. [109
], a cell-laden alginate/gelatin temperature-dependent hydrogel was used as a bioink for an extrusion-based 3D bioprinting using amniotic epithelial cells and Wharton’s-jelly-derived mesenchymal stem cells. In this study, the hydrogel bioink was engineered by changing the alginate/gelatin concentration ratio to achieve optimum rheological properties for 3D printing.
There are other recent research examples on developing cell-laden hydrogel bioinks to print skin layers or substitutes, focusing on natural hydrogels [110
]. However, it is vitally important to design specific physical, mechanical, and biological properties of hydrogel bioinks by modification of the gel composition, concentration, and ratios to provide promising biomaterials for future skin tissue engineering applications. Generally, an appropriate hydrogel bioink should be cell compatible and able to incorporate/encapsulate cells before and after crosslinking. Moreover, hydrogels should have an optimum viscosity and shear-thinning behavior to maintain the steady state of the gel and protect cells during printing, especially for extrusion-based 3D printers. Moreover, the bioink hydrogel should mimic the physical and mechanical features of native skin after printing to produce appropriate cues for cells to differentiate and proliferate.