Bioprinting technologies are powerful additive biofabrication techniques to produce cellular constructs for skin tissue engineering owing to their unique ability to precisely pattern living and non-living materials in pre-defined spatial locations. This unique feature, combined with the computer controlled printing and medical imaging techniques, enable researchers and clinicians to generate patient specific constructs partly replicating the intricate compositional and architectural organization of the skin. Bioprinting has been used to automatically dispense hydrogels with skin cells located in prescribed sites that promote skin formation in vitro and in vivo. Current skin bioprinting approaches mostly rely on the sequential printing of fibroblasts and keratinocytes embedded within a homogeneous hydrogel. Although such approaches have already been translated to pre-clinical scenarios, they still present limitations in terms of fully replicating the cellular and extracellular matrix (ECM) heterogeneity in native skin. The success of bioprinting for skin repair strongly depends on the design of printable bioinks capable of supporting the function of printed cells and stimulating the production of new ECM components. To better mimic the human skin, novel developments in dedicated bioprinting technologies, in the design of bioinks, as well as in the printing of vascularised constructs are necessary. This paper presents an overview regarding the use of bioprinting for skin tissue engineering applications. The operating principles of bioprinting technologies are outlined along with requirements of printed skin constructs. Finally, pre-clinical results are summarized and future perspectives for the field are highlighted.
The integumentary system comprises the skin and respective appendages (hair, nails and sweat glands) and is responsible for many vital functions in the body, such as the homeostasis, protection, and transmission of sensorial information [52, 149, 201]. Skin is the most superficial and the largest organ of the human body, acting as an efficient barrier at the interface between the internal organs and the external environment. However, the permanent exposure to environmental stresses may result in a variety of skin injuries induced by acute trauma, thermal, chemical, mechanical, microbial and radiation effects. Additional causes of skin loss also include genetic disorders, chronic wounds and surgical interventions [52, 176]. Depending on the lesion type, extension and depth, both superficial and deep skin layers may be damaged and the innate healing ability of skin may be dramatically reduced or inhibited. In the most critical clinical scenarios, such as deep dermal or third degree burns and full-thickness wounds, hypodermal adipose tissue, fascia, in addition to internal tissues and organs, may also be injured, leading to significant morbidity, dysfunctionality and potentially death . Skin burns and chronic wounds represent one of the most debilitating, painful, and costly health conditions with a tremendous economic and social impact. Typically, these wounds require extensive hospitalization, labour intensive clinical procedures and costly wound care products, representing a major burden over total world healthcare expenditure [9, 153, 175].
Several skin therapies and wound care products have been developed and tested through pre-clinical and randomized controlled clinical trials with the ultimate goal of promoting the repair and regeneration of functional skin [44, 49, 84, 86, 88, 103, 124, 189, 195, 212]. Traditional treatments are based on the use of grafts (which includes auto-, allo- and xenograft varieties), herbal and animal-derived compounds, silver-containing and traditional dressings. These therapies have been used since ancient times mainly due to their clinical efficacy, simplicity and affordability, offering a broad range of therapeutic effects, such as anti-inflammatory, cell stimulatory, antimicrobial, and wound debridement [98, 99, 145, 159, 179, 181]. However, in certain types of skin wounds, such as deep burns and chronic wounds, more sophisticated and effective treatments are usually needed [20, 50, 87, 212]. Advanced therapies assume a prominent role in the management of such wounds, as they explore the basic principles and elements of tissue engineering to produce both cellular and acellular constructs for skin regeneration [52, 109, 149]. In the most classical approach, known as scaffold-based or top-down, scaffolds are fabricated through either conventional or additive manufacturing processes, and subsequently seeded with autologous or allogeneic cells. The cellular scaffold is then cultured in vitro to produce a tissue-engineered construct for subsequent implantation into the lesion site [66, 116, 146, 177]. Despite the fact that such an approach allows for a good control over the scaffold characteristics (when additive manufacturing processes are used) and cell-scaffold interactions, it fails in placing individual cells at specific locations throughout the scaffold, thereby not mimicking the intricate cellular organization of natural tissues at micro- and nanoscale. Other important drawbacks are related with the limited cell density and insufficient vascularization, which may lead to tissue malformation and heterogeneous tissue growth [121, 148]. To address these limitations, an alternative approach, termed scaffold-free or bottom-up approach, has been used to generate 3D structures resembling the biological and functional organization of human tissues [71, 102, 129, 139, 172]. In this approach, one strategy of biofabrication, termed bioassembly, is applied to generate hierarchical constructs with a prescribed two-dimensional (2D) or 3D organization. Larger and complex tissue constructs are obtained through the automated assembly of pre-formed cell-containing fabrication units (e.g., cell aggregates, cell fibres, cell sheets or microtissues) produced by different techniques, such as self-assembled aggregation, microfabrication or microfluidics [10, 78, 129]. Bottom-up approach addresses some of the limitations of the top-down approach by the creation of tissue constructs with multiple cell types placed in specific 3D locations, with a high cell density and possibility of printing vascular networks [16, 121, 166]. Bioprinting, one main strategy of biofabrication, is attracting significant interest from researchers working in the field of tissue engineering and regenerative medicine due to their unique ability to print single cells, bioactive molecules, biomaterials or cell-aggregates into structurally organized constructs in a layer-by-layer fashion with high resolution and accuracy [10, 78, 111, 136]. Bioprinting provides a powerful tool to arrange cells, biomaterials and soluble factors within a 3D environment on a length scale comparable to the complex heterogeneity found in natural tissue (10–100 μm), enabling new perspectives as well as unmatched possibilities in the design of biomimetic substitutes for tissue regeneration [10, 80, 125]. In the context of skin tissue engineering, bioprinting has been explored for the fabrication of 3D constructs containing skin cells positioned in distinct layers to resemble the anatomy of native skin [117, 184].
This review presents a detailed overview of the bioprinting processes as valuable technologies to produce cell-laden constructs for skin tissue engineering, outlining the operating principles, processable biomaterials and pre-clinical results. Current challenges in the field will be also discussed mainly focusing on the design of advanced bioinks and the vascularization of skin constructs.
Human skin properties and regeneration
The human skin is a multilayer organ that presents limited self-regenerative properties after injury. In contrast to other vertebrates, such as amphibians and mammalian foetuses through mid-gestation, skin scarring is the inevitable and permanent endpoint in mammals . To address the limited regenerative capability of adult skin and promote the restoration of skin barrier properties, there are diverse products and therapies available for clinical use (e.g., dressings, cell-based treatments and bioengineered skin substitutes) or currently under investigation (e.g., bioprinted cellular substitutes), showing promising outcomes.
Skin structure and constituents
The skin serves as a protective layer against the invasion of pathogens and external agents, playing a critical role in thermal regulation, homeostasis and transmission of sensorial information [142, 157, 201]. From an anatomical point, skin consists of three sequential layers of epidermis, dermis and hypodermis, each having distinct functions, biomechanical properties and cellular constituents (Fig. 1). Adnexal structures like nails, hair follicles, sweat, apocrine and sebaceous glands are responsible for cosmetic, protective and temperature regulation functions. Epidermis is the most superficial skin layer being composed of keratinized stratified squamous epithelium and organized in four or five main layers (Fig. 1) containing keratinocytes in different stages of proliferation, differentiation and keratinization [157, 178]. Although the epidermis is an avascular layer, it is under constant renewal due to the presence of epidermal basal cells and stem cells in the basal layer and hair follicles . The main cellular constituents include keratinocytes, melanocytes, Merkel cells and Langerhans cells [13, 149]. The underlying dermis, separated from the epidermis by a basement membrane, is mainly composed of ECM molecules arranged in a dense matrix of connective tissue rich in collagen and elastin fibres that provides structural and mechanical support to the skin . It is a highly dynamic, vascularized layer that supports the epidermis, provides flexibility and strength to the skin, and also contains blood vessels, sweat and sebaceous glands, hair follicles and sensory receptors [13, 115]. The predominant cells in dermis are heterogeneous subpopulations of fibroblasts, which are responsible for the production of connective tissue components . Other cell types like mast cells and macrophages are also present, playing important roles in immune and inflammatory responses [13, 79]. The hypodermis, which acts as an energy source and heat insulator, is located just below the dermis and is mainly composed of adipose tissue and collagen, attaching the dermis to underlying tissues [13, 115].
Wound healing process
From the onset of injury, the human body initiates a series of complex and coordinated signalling pathways, involving the interaction between several ECM components, cells, cytokines and growth factors, aimed at tissue repair and reconstruction of lost cutaneous tissue [110, 178]. The healing of an acute wound proceeds in a sequence of overlapping and well-orchestrated events throughout different phases of haemostasis, inflammation, migration, proliferation and maturation. The first response of the human body to an injury is haemostasis, which is mediated by platelets, and involves the formation of a fibrin clot that circumvents the bleeding, acts as a provisional ECM for cell migration, and loosely joins the wound edges [61, 188]. During haemostasis, platelets suffer degranulation and release a plethora of cytokines and growth factors (e.g. transforming growth factor-β (TGF-β), fibroblast growth factor (FGF), transforming growth factor-α (TGF-α), vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), and epidermal growth factor (EGF)) responsible for attracting inflammatory and vascular cells to the wound site [61, 110, 113]. In the inflammatory phase, neutrophils, monocytes (that differentiate into macrophages), lymphocytes and mesenchymal cells (that differentiate into fibroblasts) play a key role not only in the elimination of microbes and foreign particles, cellular debris and clearance of apoptotic cells, but also in the stimulation of keratinocytes, fibroblasts and endothelial cells [61, 110, 113]. Subsequently, the healing process continues with the formation of a scab and the migration and proliferation of epithelial cells. Keratinocytes, located at the wound periphery, are stimulated to migrate and proliferate into the lesion site to promote re-epithelialization by the synthesis of new epidermis . Fibroblasts and endothelial cells are also recruited to initiate the synthesis of ECM components, the formation of a granulation tissue, and the growth of new blood vessels from existing ones [51, 61, 178]. Some of the fibroblasts located within the granulation tissue are converted into myofibroblasts, which are involved in wound contraction and re-epithelialization [51, 110, 113]. At the final stage, wound healing involves a long maturation phase, in which the wound edges move towards the centre (wound contraction), and the new tissue is continuously remodelled in order to approximate its properties to those of the native skin. This phase can take several weeks to months, and is marked by changes in the ECM composition with re-organization of collagen fibres and collagen crosslinking [110, 180]. Despite the continuous remodelling process, adult wound healing predictably results in a scar tissue with a tensile strength that only achieves 80% of its native properties, impairing the flexibility, mobility and esthetical outcomes .
Therapeutic options for skin reconstruction
The most suitable therapeutic treatment for skin wounds depends on numerous factors including the wound type, number of affected skin layers, level of exudate, extension of injury, and the overall health status of the patient . The clinical ‘gold standard’ therapy remains the application of split thickness autologous skin grafts [52, 149, 176]. Liquid and semi-solid preparations, such as creams (e.g. silver sulphadiazine) and solutions (e.g. povidone-iodine, sodium hypochlorite), have also been used as supportive anti-infective wound care products in daily clinical practice. They are commonly used as cleaning and antimicrobial agents, allowing wound bed preparation and decontamination [8, 67]. To extend their efficacy, facilitate their handling and improve their impact on the wound, most of these therapeutic compounds have been embedded within gauzes or incorporated within dressing materials [2, 77]. More advanced products, like traditional and modern dressings, represent a cost-effective therapy to protect the wound from the external environment and to prevent wound contamination and dehydration [115, 150]. Due to their high water absorption rates, adhesion to the wound bed and increased pain during dressing changes, traditional dressings (e.g., gauze, cotton wool) are often employed as secondary dressings to allow the exudate drainage and to support the application of more effective products (e.g., skin substitutes) [77, 115]. Traditional dressings have been largely replaced by modern dressings, which are capable of creating and maintaining a moist environment in the wound bed, ideally suited for cellular migration and proliferation [81, 115, 124, 150, 151]. A great deal of interest has been focused on advanced dressings with the ability to release bioactive substances directly into the wound bed [6, 75, 161, 192]. Bioengineered skin substitutes represent the most advanced products available for skin wounds, allowing the regeneration of both epidermal and dermal skin layers [23, 50, 138, 176]. Major limitations are the permanent closure of the wound, inadequate vascularization, high manufacturing costs and lack of skin appendages [50, 149].
Bioprinting can be defined as the computer-aided patterning and assembling of cells, cell aggregates, biomaterials and biochemicals with a predefined spatial organization, to generate biologically relevant constructs with structural organization . Contrary to bioassembly, where the minimum fabrication units are pre-formed cell building blocks, bioprinting makes use of fabrication units down to molecular level . Bioprinting processes are capable of printing living and non-living materials in a computer-controlled and automated manner with high levels of resolution, accuracy and reproducibility, enabling the generation of biological substitutes with intricate architectures, precise geometrical configurations and biomechanical heterogeneity [16, 80, 91, 92]. Bioprinting technologies can be classified in three main categories: inkjet bioprinting, laser-assisted bioprinting and extrusion bioprinting (Table 1) [111, 125, 191]. In these processes, the material to be printed, known as bioink, is loaded in a reservoir and subsequently deposited onto a receiving substrate through the action of light, pressurized air, vibration, thermal or mechanical effects. Bioinks consist of hydrogel precursor solutions or decellularized extracellular matrix loadable with cells and/or bioactive factors that play a pivotal role on the overall reproducibility of the printing process and quality of the printed construct [100, 143, 169, 185]. Depending on the bioprinting technique, bioinks can be deposited as small droplets (inkjet and laser-assisted technologies) or continuous strands of material (extrusion-based technologies), resulting in different printing times and constructs with distinct levels of heterogeneity, resolution and accuracy [111, 125].
Inkjet is a high-throughput printing technology based on the dispensing of small drops (typically 1–100 picolitres) of a liquid bioink onto a collector substrate . Drops can be generated by either continuous inkjet or drop-on-demand (DOD) inkjet printing . DOD printing is the most popular method affording a rapid and fully controlled generation of droplets (20–50 μm) through the application of a pressure pulse to the bioink loaded into a reservoir placed just behind the printing nozzle [29, 38, 171]. DOD printing can be further categorized according to the mechanism used for the droplet formation and ejection as thermal inkjet (vapour bubble formation) or piezoelectric inkjet (mechanical actuation). Thermal inkjet printers use a heating element to vaporize a small volume of fluid, generating a bubble that expands and ejects a controlled volume of fluid as a single droplet. The main concern associated to thermal inkjet relies on the high temperatures generated close to the heating element during the printing process, which may affect the biomolecules and cells to be printed [170, 171, 213]. However, it has been demonstrated that, due to the short duration of the heating pulses (~2 µs), the increase in temperature of the bioink is only a few degrees, allowing the printing of viable cells with a low percentage of cell dead . In piezoelectric printers, a mechanical pulse generated by the deformation of a piezoelectric transducer under a controlled voltage is used to generate and eject the liquid droplet. The printing speed, drop size and stresses experienced by the cells are easily controlled by changing the voltage applied to the piezoelectric transducer [38, 171].
Inkjet bioprinting is widely explored in the field of tissue engineering due to their unique characteristics of non-contact printing, high-throughput efficiency, automation, versatility, resolution, and possibility of parallel printing. This technology has been used to print a variety of compounds in prescribed 2D patterns, including growth factors , proteins , polymers , nanoparticles  and drugs . The capability of inkjet bioprinting to print mammalian cells with high accuracy, and little or even no reduction of cell viability, was also demonstrated using different cell types, such as embryonic rat motoneurons , human microvascular endothelial cells , mouse embryonic fibroblasts , and retinal ganglion cells . The formation of transient pores in the cell membrane has been observed during inkjet printing, which were reported as self-repaired in 2 h subsequent to printing . The formation of transient pores seems to be associated to temporary changes in the cell membrane permeability, and was further explored to transfect cells by co-printing plasmids encoding green fluorescent protein with porcine aortic endothelial cells . Despite these attractive capabilities, inkjet bioprinting is a nozzle-based technology, requiring the use of bioinks with low viscosity (typically below 20 mPa s−1) and low cell density (<106 cells ml−1) to prevent nozzle clogging [76, 111, 125]. Printing bioinks with those characteristics is usually associated with droplet spreading and diffusion onto the substrate, which significantly reduces the spot resolution and limits the biofabrication of true 3D constructs due to the lack of structural stability of printed layers to support subsequent layers. The surface tension of the bioink and the wettability of the receiving substrate are also two key properties affecting the bioink ejection from the reservoir and the resolution of the printed constructs [38, 140, 170]. Another important limitation of inkjet printing relies on the need for bioinks capable of preventing cell settling, sedimentation and aggregation during printing in order to prevent nozzle clogging and time-consuming intermediate re-suspending operations. To address these complications, alternative methods (e.g., addition of surfactants and ethylene diamine tetra-acetic acid) have been tested to improve the stability of bioinks, to maintain their cell-suspending ability, and to enhance the homogenization, without compromising printability and biocompatibility [29, 141].
The working principle of laser-assisted bioprinting is based on the conventional laser-induced forward transfer (LIFT) technology, involving the use of a focused laser pulse to induce the transfer of a bioink from a donor ribbon onto a collector slide . For tissue engineering applications, traditional LIFT apparatus has been modified mostly regarding the print ribbon and receiving substrate in order to improve the compatibility with biological materials. The bioink is coated onto the bottom side of the laser transparent substrate and consists in cells suspended in a liquid medium or in a viscous polymer solution [57, 58, 134]. A high-powered laser pulse (usually a near infra-red laser) is focused onto a thin metal layer (1–100 nm), placed above the laser transparent substrate, generating a high-pressure bubble that propels the material towards a parallel substrate [14, 57, 128]. The most widely used modified-LITF processes in biofabrication comprise matrix-assisted pulsed laser evaporation direct writing (MAPLE DW) and absorbing film-assisted laser-induced forward transfer with its variants such as the biological laser processing (BioLP). The distinctive feature between such processes is the inclusion of either an optical absorbing material (MAPLE DW) or a sacrificial absorbing layer (BioLP) at the interface between the laser-transparent ribbon and the material to be transferred [128, 163]. In BioLP, the laser absorption interlayer, usually a thin metal coating (Au, Ag, Ti, TiO2 ~100 nm), displays several functions, such as: (1) eliminate the interaction between the laser and the biological material, (2) protect cells from light exposure, (3) minimize the bioink heating, (4) promote a quick thermal expansion with a more efficient droplet ejection, and (5) increase the printing reproducibility [14, 57, 128, 165].
Laser-assisted bioprinting processes are attractive for cell printing applications due to their high-resolution (10–100 μm), reproducibility, nozzle-free nature and possibility of printing multiple biomaterials and cell types by incorporating those compounds in the same ribbon or having multiple parallel ribbons . As a result, these processes are promising as they allow the combination of biomaterials and cells in 3D environments, enabling the formation of close cell–cell and cell–matrix interactions. In contrast to inkjet, laser-assisted bioprinting enables the processing of bioinks with a broad range of viscosities (1–300 mPa s−1) and higher cell concentrations (108 cells mL−1) [59, 89, 125, 165], which is a prime requisite to mimic the native organization of human tissues. Although the absence of an orifice reduces the stresses to which cells are subjected, cell injury and dysfunctionality may occur due to thermal heating, optical irradiation and mechanical impact with the receiving substrate [57, 128, 165]. Previous research has shown that these effects can be significantly reduced, or even eliminated, by controlling the laser pulse characteristics, the bioink viscosity, the thickness of absorbing layer and the substrate properties [27, 53, 101]. These parameters can also be optimized in order to better control the printing process and improve the reproducibility and resolution . Laser-assisted bioprinting also enables a great control over the droplet size and number of printed cells per droplet via manipulation of the laser pulse energy, the laser spot size, the distance between the ribbon and the substrate, and the thickness of both energy absorbing layer and cell support layer [53, 57, 59, 89]. Several works have also shown the ability of laser-assisted bioprinting to print mammalian cells without affecting their viability or function, or inducing DNA damage [14, 53, 55, 58, 60, 89, 90, 200, 203].
The direct extrusion of bioinks onto a substrate is the most widely explored technique to produce complex and heterogeneous 3D constructs for tissue engineering. In this versatile technology, cell-laden polymeric solutions [28, 154], decellularized ECM components , cell suspensions , microcarriers  or tissue spheroids  are loaded into standard disposable syringes, and printed onto a building platform driven by pneumatic (pressurized air), mechanical (piston or screw) or solenoid (electrical pulses)-based dispensing systems [76, 111, 125, 137]. An important advantage of extrusion bioprinting is the ability to print highly viscous polymer solutions containing a wide range of cell densities (up 107 cells mL−1) [92, 182]. It also enables the biofabrication of clinically relevant 3D constructs by the deposition of continuous and larger hydrogel strands through a nozzle with variable diameters, usually in a range of 150–300 μm [92, 111, 125]. The extrusion bioprinting apparatus typically includes a computer-controlled robotic stage, multiple dispensing systems, and a construction platform that collects the printed material. To further improve its versatility and the range of processable biomaterials, additional modules can also be coupled and integrated to the traditional apparatus, serving multiple purposes such as controlling the temperature of reservoirs , enabling multimaterial printing , and allowing the in situ crosslinking through physical  or chemical  methods.
Major concerns with extrusion bioprinting are related to (1) the low printing resolution (~200 μm), (2) the compromise between the shape fidelity and cell compatibility, and (3) the shear stresses generated during the passage of cells through the needle [111, 125]. The printing resolution mainly depends on the device characteristics (e.g., nozzle diameter, flow rate, and printing speed), the material properties (e.g., rheological behaviour, wettability, surface tension) and the cell density. Environmental conditions and crosslinking methods also have a critical role on the achievable resolution [17, 31, 32, 111]. Since extrusion bioprinting involves the deposition of cellular bioinks through a small nozzle, printed cells are subjected to mechanical stresses during the process, which may lead to a loss of cell viability and even to alterations in the phenotype . Several works showed that stresses transmitted to the cells are a function of the dispensing nozzle characteristics (geometry, diameter and length), applied pressure, deposition speed, and bioink rheological properties. These studies also demonstrated the possibility to determine and to reduce the shear stresses transmitted to the cells through experimental and numerical studies [17, 19, 30, 46, 127]. To address the unique requirements of extrusion bioprinting regarding the print fidelity and biological characteristics, research efforts have been focused on the development of bioinks exhibiting appropriate rheological, mechanical and biological properties [28, 32, 74, 82, 100, 152, 169, 174]. A multitude of crosslinking mechanisms, including thermal gelation, ionic and photocrosslinking, have also been explored to induce the in situ gelation of printed materials with the ultimate goal of improving the mechanical properties, the shape fidelity and the formation of interconnected 3D pores throughout the construct [4, 28, 107], which still remains a major challenge.
Bioprinted cellularized skin hydrogels: from in vitro to in situ printing
The interest in bioprinting technologies to produce cell-laden constructs for skin applications has been significantly increasing mainly due to their ability to pattern heterogeneous cell populations, soluble factors and biomaterials in predesigned 3D locations with high degree of automation and reproducibility. These features provide new possibilities in the development of biomimetic constructs resembling the anatomical and cellular organization of skin tissue (Fig. 2) [90, 95], with potential to address the limitations of current clinical therapies, including the vascularization, regeneration of skin appendages and integration with the host tissue [52, 149]. Bioprinting processes also enable the integration with medical imaging techniques (e.g., micro-computed tomography, magnetic resonance imaging) and path-planning devices, allowing the design of patient specific implants for direct printing either in vitro or in situ [184, 200].
Requirements of printed skin
Printed skin constructs are currently explored for several purposes including wound healing and in vitro models for pharmaceutical testing and disease model (e.g., inflammatory skin disease). Advances in microfluidic and bioprinting technologies have also allowed the miniaturization of skin models through the development of skin-on-a-chip devices, which offer a valuable alternative to animal experimentation . In order to closely reproduce the main properties of native skin and provide reliable models for clinical translation, printed skin must fulfil key compositional and functional characteristics:
Cellular composition: native skin is composed of multiple layers containing distinct cellular populations. The epidermal layer contains keratinocytes, melanocytes, Merkel and Langerhans cells. In the dermis, fibroblasts are the predominant cells, while the hypodermis is composed of adipocytes. To replicate as close as possible the heterogeneous cellular composition of native skin, each cell type should be precisely dispensed at specific 3D locations in their respective layer. The density and ratio between cell populations within the construct must be controlled in order to promote cell–cell interactions that regulate autocrine and paracrine signalling.
Biochemical composition: in the dermal region, cells are embedded within a rich ECM that provides a multitude of cell-adhesion and cell-remodelling sites, allowing cells to remodel and interact with the surrounding matrix. The basement membrane physically separates skin layers, providing motifs for cell anchorage, which is essential for stabilization and diffusion of molecules. In printed skin, the biochemical cues, their function, location and density must be precisely controlled. The presence of cell-adhesion sites and cell-proteolytic domains is fundamental to replicate the dynamic and interactive nature of ECM. Also the ability to sequester and release bioactive molecules, for instance via growth factor-binding domains, is of paramount importance for the sustained release of cytokines and growth factors during the wound healing.
Zonal organization and material composition: native skin is a highly complex tissue composed of a multitude of ECM components, which composition and ratio varies according to the region. The dermal layer is rich in collagens (I, III, IV, VII), fibronectin and glycosaminoglycans. Dermis is also organized into an upper ‘papillary’ (high ratio of collagen type III to type I) and a lower ‘reticular’ (low ratio of collagen type III to type I) region . At the epidermal–dermal junction, a basement membrane mostly composed of laminin and collagen IV provides a physical barrier between two layers . Materials for skin printing should consider the fibrillar and hydrated nature of ECM. Although is difficult to develop a formulation containing all the skin components, it’s fundamental to engineer advanced materials which stiffness and composition could be readily altered according to the location throughout the construct. The design of materials able to mimic the basement membrane constitutes an important requisite in printed skin.
Tissue architecture and shape: skin has thickness in a range of 1.0–2.5 mm, depending on the location in the body, with epidermis measuring approximately 0.2 mm . The thin papillary region contains higher cellular density along randomly orientated collagen fibre bundles, while the thicker reticular dermis consists of highly orientated fibre bundles. The heterogeneity of skin throughout the body requires precise control over the layer thickness in order to print constructs matching the wound depth. In addition, the thickness and fibre orientation in both papillary and reticular dermis should be considered to improve biomimicry. The shape of the printed construct should also be defined according to the specific application.
Barrier function: epidermis contains keratinocytes at different stages of differentiation with fully stratified keratinocyte layers providing an efficient barrier to the penetration of external agents across the skin. Lamellar granules, produced by keratinocytes, secret several key compounds (lipids, proteins, hydrolytic enzymes) for skin barrier properties . To recreate the barrier properties of skin, printed constructs must contain an epidermal region consisting of densely packed keratinocytes. Constructs need to be cultured at air–liquid interface (ALI) to promote stratification and the barrier properties carefully evaluated. Late or incomplete stratification and keratinization of printed constructs must be evaluated, in particular, when immortalized keratinocyte cells are used.
Appendages and pigmentation: skin contains adnexal structures such as nails, hair follicles, sweat, apocrine and sebaceous glands. Melanocytes produce melanin, a pigment that confers coloration to the skin and protects from UV light. Merkel and Langerhans cells participate in sensing and immunological defence, respectively. The development of skin appendages in printed skin is a major challenge in the field. Skin pigmentation can be addressed by printing melanocytes in the epidermis, while sweat glands can be engineered by printing epidermal-derived stem/progenitor cells.
Vascularization: dermis contains numerous blood vessels that provide nutritional support to the skin. Vasculature includes larger blood vessels at the interface between the reticular dermis and hypodermis, which branch into small vessels to supply superficial plexus at the junction between the papillary and reticular dermis, as well as the epidermis by diffusion. The ability to engineer functional vasculature in 3D is critical for nutrient supply throughout thick skin constructs. Printing branched, perfusable channels could be a viable solution to promote efficient mass transfer of the skin. Complementary strategies could also involve the co-printing of stromal and endothelial cells as well as the tethering and release of angiogenic factors.
Hydrogels are the golden standard materials for skin regeneration as they provide a hydrated and highly permeable microenvironment in which, depending on the composition (e.g., cell-adhesion and proteolytically sensitive motifs) and properties (e.g., stiffness), encapsulated cells might migrate and proliferate in a similar way to what happens in the native ECM . Hydrogels have been widely used to assist the regenerative and reparative processes of many tissues [41, 198, 199] and several studies were performed assessing their potential as wound healing adjuvant, mainly due to their porosity, permeability, viscoelastic properties and low immunogenicity (depending on the source) [21, 39, 64, 114]. Hydrogels are defined as crosslinked 3D network structures obtained from a range of synthetic or natural polymers that can absorb and retain large amounts of water [7, 168]. They are usually classified according to the nature of the crosslinks established between the polymer chains into chemical or physical hydrogels. Chemical hydrogels involve the formation of covalent bonds, while physical hydrogels are obtained when physical interactions are established between the polymeric chains (molecular entanglement, ionic interaction and hydrogen bonding) [168, 210]. Hydrogels display many of the desired characteristics of an ideal skin substitute, as they promote wound debridement, while providing a suitable moist environment to stimulate the healing process. In addition, they can absorb wound exudates, allow gaseous exchanges and prevent bacterial invasion. Hydrogels also provide thermal isolation and display low adherence to the wound bed, which facilitates dressing change without trauma to the patient. Furthermore, the fine-tuning of hydrogel properties is easy to achieve, allowing control over stiffness, biodegradability rates and permeability to either cells or soluble factors. Transparent hydrogels allow to follow-up the wound healing process without removal. When carrying cells, either fibroblasts or keratinocytes separately  or together , these hydrogels have shown better performance in wound healing and host tissue integration.
A major challenge in skin bioprinting, is designing suitable bioinks to produce 3D cellular constructs with intricate geometries, shape fidelity, and high resolution in the placement of cells. In the biofabrication field, traditional approaches to generate such constructs often involve (1) the sequential printing of cell-laden hydrogels or melt extruded thermoplastic fibres, (2) the formulation of viscous bioinks by adding high molecular weight polymers (e.g., hyaluronic acid, HA), or (3) the use of increased polymer concentrations and crosslinking densities [152, 173, 174, 193]. The combination of hydrogel bioprinting with melt extrusion has been successfully explored to design well-defined 3D constructs with improved mechanical properties, which is of special interest for load bearing tissues , but of limited application in soft tissues. The printing process of viscous bioinks might affect the distribution of cells throughout the construct ultimately leading to tissue heterogeneity, while highly crosslinked hydrogels may lead to excessively stiff matrices that impose severe restrictions to cell spreading, migration and proliferation [11, 24, 130]. Thus, it is crucial to ensure that the bioink has the appropriate rheological properties to be printed and, at the same time, is able to maintain the shape upon deposition. Alternative strategies have been tested to improve the printing fidelity without compromising the bioink cytocompability, including the printing of partially crosslinked hydrogels, the crosslinking of hydrogel solutions during bioprinting, and the deposition of biocompatible sacrificial materials that provide initial support for the printed hydrogel pattern [43, 135].
The development of hydrogels to be applied as a bioink is a complex and multidisciplinary process that requires knowledge about the principles of printing mechanisms and the fluid properties required for a given liquid to be printable. The interactions between the bioink and the receiving substrate as well as its rheological properties are important parameters determining the resolution and accuracy of printed constructs [38, 57, 76, 111, 123]. In order to adapt a material for bioprinting technologies, it is important to take into account whether the technical characteristics of the printing device afford the design of biologically relevant constructs for a given application. By placing suitable cell types in appropriate positions, the tissue construct may then mature into a tissue/organ and late achieve functionality either inside a bioreactor or in vivo . The ideal properties of hydrogels for bioprinting include stability, sterilization, biodegradability, adequate mechanical properties and swelling characteristics, but it is also required that both hydrogel chemistry and crosslinking mechanisms promote cell function . Cells must not only survive to the printing process and remain viable, but should also be able to migrate within and from the implant, proliferate, establish cell–cell/cell–matrix interactions, and also secrete growth factors and other healing mediatiors. In addition to these important general considerations, the properties of bioinks should adapt to the printing mechanism used to produce tissue constructs. For extrusion bioprinting, in which the bioink flows through a nozzle towards the substrate, rheological properties are particularly relevant. The rheological behaviour of bioink determines how the viscosity is affected during the printing as a response to the shear stress. Printing cells embedded in a hydrogel precursor solution also requires careful selection of the viscosity in order to ensure homogenous distribution of cells and prevent nozzle clogging. This is particularly critical for inkjet printing owing the limited range of admissible viscosities . In addition to viscosity, surface energy, tension and charge constitute other relevant properties of biomaterials for bioprinting, as they exert great influence on droplet formation and ejection [57, 123]. Moreover, the bioink should also undergo a fast and cytocompatible gelation after deposition in order to maintain the prescribed shape and spatial location of cells at high resolution.
Despite recent advances in the design of printable bioinks for various applications (Table 2), the number of bioinks for bioprinting cellular skin substitutes is still limited to a restricted number of natural hydrogels including alginate, collagen, gelatin, fibrin and hyaluronic acid. The interest in these materials mainly relies on the fact that their composition and structure resembles the ECM properties of native tissues . These polymers can also be easily converted into either physical or chemical hydrogels under cell-compatible conditions through a myriad of crosslinking pathways . In addition, certain natural polymers (e.g., collagen) also present cell-proteolytic domains and/or adhesion motifs, providing recognition sites for embedded cells and allowing the cell-mediated degradation of the hydrogel network in a similar way to the remodelling of native ECM. Natural hydrogels, in particular protein-derived hydrogels, usually present concerns regarding weak mechanical properties, high immunogenicity and virus/disease transmission, though they retain their biological activity being the most widely used ECM-mimicking materials [47, 147]. Synthetic hydrogels constitute an important alternative due to their controlled chemical composition and tuneable mechanical properties. The drawback of most synthetic materials is the lack of cell-adhesion moieties, which can be solved by the incorporation of cell instructive cues into the polymer backbone .
Printed skin constructs for wound healing
A major application of printed skin constructs is to promote the wound repair and regeneration. Functional skin is an important need in medical field to treat patients with a variety of injuries, including (1) burn wounds, which lack of the skin barrier and are more susceptible to infection, (2) chronic ulcers, which are difficult to heal and associated to high healthcare costs, and (3) scars, which cause severe anaesthetic limitations and movement disabilities. In response to these needs, bioprinting has been explored to create sophisticated bi-layered skin substitutes containing dermal and epidermal components, which mimic the native organization of skin tissue.
Michael et al.  used the laser-assisted bioprinting to produce multi-layered cellular skin constructs and evaluate their ability to promote the in vivo regeneration of full-thickness skin wounds in nude mice (Fig. 3). Skin substitutes were created by printing 20 layers of NIH3T3 fibroblasts followed by 20 layers of HaCaT keratinocytes suspended in collagen onto a sheet of Matriderm™ and subsequently incubated under submerged conditions overnight. In vitro studies showed that fibroblasts started to produce collagen and keratinocytes formed a dense epidermis, though no differentiation was observed. After 11 days of in vivo implantation, skin substitutes showed a good integration into the native tissue. Small blood vessels were also detected in the printed skin substitutes, growing from both the wound bed and wound edges into the Matriderm® towards the direction of printed cells. Immunohistochemistry assays confirmed the presence of E-cadherin between keratinocytes, indicating the establishment of cell–cell contacts. This work demonstrated the capability of laser-assisted bioprinting to produce 3D biological skin substitutes capable of developing functional tissues in vitro and promote skin regeneration in vivo, which represents an exciting advance in the field of 3D skin bioprinting.
Rather than printing 3D skin constructs in vitro for subsequent implantation, efforts have also been focused on the development of integrated bioprinting strategies for the in situ printing of cellular skin substitutes. In situ skin printing is challenging as it requires the development of dedicated bioprinting devices, the integration with imaging systems to obtain data from the wound site, and the design of bioinks capable of instructing printed cells to perform their native functions. Binder et al. [18, 184] developed a device for the in situ skin printing composed of a cartridge delivery system containing a series of inkjet nozzles and a laser scanning system, both mounted on a portable XYZ plotting system. The data obtained from the laser was used to reconstruct a 3D model of the wound, which was subsequently employed to determine the skin area that was missing from the wound. Afterwards, the printheads filled the wound site with a bioink containing skin cells in a layer-by-layer manner. The potential of the system to induce the skin regeneration was firstly evaluated through the printing of human keratinocytes and fibroblasts suspended in a fibrinogen-collagen precursor solution into a full thickness skin lesions (3.0 × 2.5 cm2) created on nu/nu mice . After printing each layer, a thrombin solution was sprayed on top of the deposited layer to induce hydrogel formation. Results showed the complete closure of the wound in 3 weeks and the formation of new skin with similar properties to the native tissue. Histological analyses also revealed that the newly formed skin contained organized dermal collagen and a fully developed epidermis. The same research group used a similar procedure and bioprinting technique to print amniotic fluid-derived stem cells (AFSC) and bone marrow-derived mesenchymal stem cells (MSC) suspended in a fibrinogen-collagen gel into full-thickness skin wounds in mice . The cell-biomaterial solution was dispensed to the skin wounds (2.0 × 2.0 cm2) with uniform cell distribution. The closure and re-epithelialization of wounds treated with printed AFSC and MSC were significantly better than those of wounds treated only with the fibrin-collagen hydrogel. Histological sections of skin samples harvested 2 weeks after printing/implantation showed well-defined and organized epidermal layers of the regenerated skin in MSC and AFSC-treated wounds, which was accompanied by increased neovascularization and blood vessel maturation, particularly in the case of AFSC-treated wounds. However, findings also demonstrated that cells only remained transiently at the lesion site and did not permanently integrate with the host tissue. Another limitation of this bioprinting approach relies on the slow crosslinking time of the bioink system, which increased the overall procedure.
In addition to suitable mechanical, swelling and biodegradation characteristics, bioinks for in situ skin printing should exhibit fast crosslinking to reduce the printing time, support cell activity and eventually act as a delivery vehicle of biological factors. In order to meet such requirements, Skardal et al.  developed a photocrosslinkable heparin-conjugated hyaluronic acid (HA-HP) hydrogel for in situ printing capable of sequestering and releasing cell secreted growth factors. The bioink system, composed of thiolated HA with conjugated heparin groups, thiolated gelatin, PEGDA cross-linker and I2959 as photoinitiator, was loaded with AFS cells and printed directly onto full thickness skin wounds (2.0 × 2.0 cm2) in a nu/nu murine model (Fig. 4). After printing, the bioink was crosslinked in situ via thiol-ene photopolymerization upon exposure to UV light. In vitro tests indicated the effectiveness of heparin-mediated growth factor binding in modulating the release of pro-angiogenic AFS-secreted FGF and VEGF from HA-HP hydrogels. Wounds treated with cell-laden HA-HP hydrogels showed slightly higher closure rate and re-epithelialization compared to both HA-treated and non-treated wounds, though the differences between groups were not statistically significant after 14 days post-printing. Histological characterization of the harvested skin tissue indicated that HA-HP hydrogels promoted higher microvessel density which may be due to the sequestration of and release of AFS cell-secreted growth factors as well as increased deposition of elastin. Although this work demonstrates the potential of photopolymerizable heparin-containing HA-based bioinks combined with AFS cells for in situ skin printing, further characterization regarding the effectiveness of this strategy on the wound healing and quality of new skin is necessary.
Printed skin constructs as in vitro models
The interest in printing in vitro human skin models as platforms for toxicity and efficacy testing of topical medications and cosmetics has been significantly increasing due to the need for cost-effective alternatives to animal experimentation. Although there are several works demonstrating the printing of bi-layered skin constructs and evaluating their properties [156, 164], there is a lack of studies using and testing their applicability for pharmaceutical or cosmetic testing. The majority of these studies employ skin equivalents exclusively created by manual methods [160, 205] or fabricated using components obtained through 3D printing technologies combined with manual deposition of cell-laden solutions .
Lee et al.  used a robotic cell dispensing technology composed of a set of electromechanical microvalves to fabricate 3D cell-laden constructs in vitro mimicking the native skin anatomy (Fig. 5). Their approach involved the printing of collagen hydrogel precursor solutions, keratinocytes and fibroblasts suspension, followed by the crosslinking of each printed layer using a nebulized sodium bicarbonate solution as a gelation agent. Results indicated no significant differences on the viability of printed cells compared to manually plated cells, while confocal microscope images of printed 3D constructs after 4 days of in vitro culture showed distinct layers of fibroblasts and keratinocytes. This strategy was further explored and optimized in another work from the same group, where particular attention was given to the optimization of printing parameters, cell densities and histological characterization of skin constructs . Printed multi-layered constructs were cultured under submerged conditions followed by exposure to the air–liquid interface to stimulate proliferation of keratinocytes. At day 14, printed constructs exhibited a dense epidermal layer and a sparsely populated dermal layer without cell invasion and migration across the two regions. Histological analysis revealed epidermis compaction along the period of ALI culture, accompanied by the formation of terminally differentiated stratum corneum and thigh junctions between keratinocytes. Printed constructs maintained the shape and dimensions without shrinkage during the culture period, while conventional constructs obtained by manual deposition presented continuous shrinkage and shape modification. Further improvements on the density of fibroblasts in the dermis, as well as on the stratification and keratinization, are required to better mimic the native skin tissue. In another work, laser-assisted bioprinting was explored to produce 3D skin grafts consisting of a collagen matrix containing fibroblasts and keratinocytes in two main regions, aiming at mimicking the anatomic organization of native skin . The bi-layered 3D skin substitute was created by printing 20 layers of fibroblasts followed by 20 layers of keratinocytes. Both cell types were embedded in a collagen matrix and transferred onto a receiving slide coated with Matriderm™ (composed of collagen and elastin). After 10 days of in vitro culture, cells remained viable and located at the printed positions, without mixing between fibroblasts and keratinocytes. Immunohistochemical analyses revealed the formation of a laminin layer (main protein in basement membrane) between the fibroblasts and keratinocytes layers.
In another study, Cubo et al.  used an extrusion-based bioprinter to create a human plasma-derived bilayered skin using fibroblasts and keratinocytes. The dermal component was created by printing fibroblasts embedded within a plasma-derived fibrin hydrogel, which was allowed to polymerize for 30 min before the seeding of keratinocytes to generate the epidermal component. For in vitro testing, skin constructs were printed on transwell inserts, allowed to differentiate at the air–liquid interface for 17 days and evaluated for their resemblance to the native tissue. Histological analysis showed the formation of skin equivalents with structural resemblance to the human skin, characterized by the presence of spread fibroblasts within the dermal compartment and the terminal differentiation of keratinocytes on the epidermal component. To evaluate the in vivo differentiation capability of printed constructs, the substitutes were printed, incubated overnight and immediately grafted on full-thickness wounds created on the backs of immunodeficient athymic mice. Immunofluorescence results using different differentiation markers indicated the presence of stratum basal, stratum spinosum, stratum granulosum and stratum corneum on the epidermal component, as well as the formation of the dermo-epidermal junction at the interface with the dermal component. Recently, Kim et al.  developed a hybrid bioprinting approach to fabricate a human skin model with a functional transwell system in a single-step process, as a low-cost alternative to commercial transwell inserts. In their strategy, an extrusion module capable of printing hydrogel bioinks and thermoplastics was used to print the dermal compartment and a polycaprolactone (PCL) mesh to reduce collagen shrinkage during the construct maturation, respectively. The fibroblast-loaded collagen bioink was printed onto the PCL mesh to create the dermal region, while an inkjet module was subsequently applied to print the epidermal region through the deposition of keratinocytes on the top of the dermis-like layer. The dermal region exhibited smaller shrinkage compared to the control group (fibroblasts in collagen gel), while inkjet printed keratinocytes displayed a more uniform distribution and higher proliferation rate when compared to manually deposited cells. After 14 days of ALI culture, fibroblasts were well-spread throughout the dermal region expressing collagen type I, while stratified epidermis were formed on the epidermal region with expression of both early (keratin 10) and late (involucrin) differentiation markers.
Printed skin constructs with appendages and pigmentation
To create biomimetic skin constructs, the complexity of the tissue must be addressed behind generating the two major macro-structures of skin, i.e., dermal and epidermal layers. Thus, a great deal of interest has recently been focused on the integration of appendage structures into skin constructs, including sweat glands and pigmentation.
The regeneration of sweat glands using 3D printed ECM-mimic environments was addressed by Huang et al.  using a gelatin-based bioink. The bioink composed of gelatin, alginate, mouse plantar dermis components (collection of protein molecules with bone morphogenic protein (BMP-4) the predominant growth-factor protein) and EGF was designed to provide a 3D niche to promote differentiation of epidermal progenitors into sweat gland cells. Several bioink compositions (with or without EGF and with plantar or dorsal dermis components) were printed using an extrusion system, crosslinked with calcium chloride, and their properties evaluated regarding cell viability and differentiation. Cellular hydrogels were deposited onto a cold glass chip (gelatin at gel state) and physically crosslinked with calcium chloride after printing, supporting cell spreading and growth. Results also showed that embedded cells within hydrogel containing EGF and plantar dermis components differentiated into sweat gland cells, as demonstrated by expression of luminal epithelial markers. After the in vitro characterization, 3D constructs were implanted into mice burned paws to access their performance. Printed hydrogels containing both EGF and plantar dermis components restored the sweat gland function upon 14 days of implantation, performing better than the other tested groups. Histological analysis also revealed the formation of sweat glands in regenerated skin and degradation of the hydrogel network. Although this study did not address the mechanisms of sweat gland restoration, the suggested bioprinting approach could be a useful tool for further research on the restoration of skin appendages. In a later study, the same research group demonstrated that the 3D architecture of printed constructs determines the sweat gland morphogenesis . 3D porous constructs (20 × 20 × 5 mm) were printed using different bioinks (with and without plantar dermis components) and printing nozzles (300 and 400 µm) and their effect on sweat gland cell differentiation and morphogenesis evaluated. After 14 days of culture, epithelial progenitors in constructs produced with the 300 µm nozzle showed higher expression of luminal epithelial markers, which is characteristic of differentiation into sweat gland cells. At later culture period (28 days), specific morphology of sweat gland tissue was only present in constructs printed with the 300 µm nozzle and the plantar dermis-containing bioink. The authors suggest that the larger pore size of constructs produced with the 300 µm nozzle together with the release of plantar dermis components (inductive factors) from the constructs might influence and guide the differentiation of printed cells. However, the mechanisms and cues that regulate cell response to the 3D architecture and drive cell differentiation and sweat gland formation still remain unknown.
Skin pigmentation is associated to the production of melanin from melanocytes and its transfer to surrounding keratinocytes, being an important requirement of skin constructs. Min et al.  used a micro-valve based bioprinter that deposits small liquid droplets to create full-thickness skin constructs containing pigmentation. The dermal component was created by printing five layers of collagen precursor solution, which was crosslinked through neutralization with sodium bicarbonate. During the fabrication, a suspension of fibroblasts was printed and embedded in the layers 2–4. After 1 day of culture, a suspension of melanocytes was printed on the top of the dermal layer in two different configurations (Fig. 6). After 4 days of culture, a suspension of primary keratinocytes was then printed on the top of the constructs, followed by submerged culture (1 day) and ALI culture for 4 days. Histological characterization showed the formation of distinct dermal and epidermal layers with terminal differentiation of keratinocytes, as shown by loricrin staining. Printed melanocytes in the epidermal layer formed multiple clusters of dark pigmentation, as observed by visual inspection and bright-field microscopic images. H&E staining revealed the presence of pigmentation and melanin granules, while Melan-A IHC staining showed that pigmented cluster consisted of melanocytes and keratinocytes, though additional characterization is required to investigate the location of both cell types. Despite promising outcomes, constructs exhibited non-uniform pigmentation, being necessary further optimization to improve the homogeneity in skin pigmentation.
Printed vascularized skin constructs
The formation of biologically relevant constructs containing a functional, perfusable vascular network is a major hurdle towards the full implementation and clinical translation of skin bioprinting. Vascularization of bioprinted skin constructs is essential to recapitulate the vasculature of dermis and to provide nutrients and oxygen to the embedded cells. It also plays a key role in the integration of implanted constructs with the host tissue, which remains a limitation of current skin substitutes. In the absence of a functional vascular network, nutrients, oxygen and soluble factors are supplied to the cells by diffusion, which is a slow and inefficient process that may lead to suboptimal cell nutrition, cell necrosis or delayed new tissue formation . In vivo, vascularized tissues contain an organized system of larger blood vessels that branch into smaller vessels, which ultimately branch into capillaries distributed throughout the tissue. Capillaries are in the close proximity (maximum ~200 μm) to the cells allowing nutrition, oxygenation and removal of waste products . The natural formation of a vascular network in vivo is regulated by two underlying processes of vasculogenesis and angiogenesis. The first occurs during early embryonic development to form a primitive capillary network, while the second takes place in adults aiming at the revascularization of damaged or diseased tissues from existing blood vessels [133, 167]. Recapitulating these biological processes in printed tissues is challenging and requires a proper spatiotemporal interplay between vascular cells, resident cells and surrounding ECM. Although several strategies have been explored to promote vascularization in skin constructs [62, 88], engineering branched vascularization in 3D hierarchical architectures remains a great challenge. Bioprinting technologies are attractive to engineer a vascular tree within thick constructs by the precise layer by layer deposition of multiple cell types and ECM-like bioinks into prescribed spatial locations at high resolution . Promising approaches to generate vascular networks have been reported using inkjet [34, 209], laser-assisted , and extrusion bioprinting [16, 92, 132]. In general, such approaches are based on four main strategies: (1) direct patterning vascular cells onto a receiving substrate , (2) continuous printing of polymeric bioink loaded with endothelial cells followed by polymer removal , (3) printing perfusable channels in a 3D construct for subsequent injection of a cell suspension into the empty channel , and (4) printing of multicellular spheroids . Vascular cells currently explored in bioprinting include endothelial cells such as human umbilical vein endothelial cells HUVECs or human dermal microvascular endothelial cells (HMVECs), and smooth muscle cells such as human umbilical vein smooth muscle cells.
To address the need for vascularization in skin substitutes, Yanez et al.  used the inkjet technology to fabricate capillary-like endothelial networks into a dermo-epidermal skin graft consisting of neonatal human dermal fibroblasts (NHDFs) and neonatal human epidermal keratinocytes (NHEKs) embedded in a fibrin-collagen matrix. HMVEC were mixed with thrombin and inkjet printed on top of a manually plated layer of collagen-NHDF cells containing fibrinogen. After formation of the fibrin hydrogel, a layer of collagen containing NHEK cells was pipetted on top of it to produce bi-layered structures. These structures were then implanted into skin full-thickness wounds on the back of athymic nude mice to investigate healing properties and integration with the native tissue. Wounds treated with printed substitutes required 14–16 days to heal, contrasting with 21 days in the control group (no skin graft) and 28 days in the group implanted with Apligraf® (commercial tissue-engineered skin substitute based on collagen with embedded fibroblasts and further cultured with keratinocytes). Histological characterization of the wounds treated with the printed substitutes on post-operative day 14 showed the formation of dermal and epidermal skin layers similar to the native skin, accompanied by the presence of new micro-vessels in the mouse tissue. It was suggested that new blood vessel formation was stimulated by the presence of endothelial cells, indicating partial integration of the skin graft into the host tissue, which was further confirmed by the presence of human cells in the mouse tissue. More recently, Abaci et al.  developed an integrated strategy to generate perfusable vascular networks in human skin equivalents using primary and induced pluripotent stem cell (iPSC)-derived endothelial cells (iECs). 3D printing technology was explored to create moulds with the desired vasculature pattern and functional components for the fabrication of embedded vasculature. Sacrificial channels composed of crosslinked alginate were printed and embedded with dermal fibroblasts embedded with a collagen matrix. After 7 days of culture, keratinocytes were seeded on the top of the dermal compartment, cultured under submerged conditions for 7 days and subsequently maintained at ALI interface for additional 7 days. After this period of time, perfusable channels were created by removing the alginate using a sodium citrate solution, followed by the injection of either HUVECs or iECs, which established proper endothelial barrier function. Vascularized skin constructs grafted on the back of immunodeficient mice for 14 days were perfused with blood, while both non-vascularized constructs and constructs with acellular channels remained non-perfused, which suggest that engineered vasculature promotes and guides neovascularization of skin constructs. Although the skin construct was created by manual deposition of cell-laden solutions rather than using bioprinting technologies, this study reports an integrated approach to generate vascularized skin equivalents for tissue engineering applications. A similar approach to fabricate perfusable vascular channels within skin equivalents was recently reported by Mori et al. , who’s demonstrated the applicability of the vascularized skin equivalent for dermatological studies and for drug testing.
Conclusions and future perspectives
The field of skin bioprinting has experienced exciting advances in the direct printing of cellularized hydrogel constructs as effective alternatives to current therapies for skin repair and regeneration. Bioprinting technologies have been applied to create 3D constructs that recapitulate, to some extent, the architecture and cellular composition of native skin. Although those processes allow the precise patterning of skin cells and ECM components in the form of 3D constructs capable of promoting in vitro and in vivo skin formation, the field of skin bioprinting is still in an early stage, facing numerous challenges in diverse areas. A critical issue is the availability of bioprinters capable of producing constructs that replicate the intricate ECM composition of skin. Since most of the bioprinters currently used in laboratories result from technical modifications of available technologies and/or equipment, they usually have limited resolution and accuracy, which imposes serious restrictions to the processing conditions, range of printable biomaterials and characteristics of the printed construct. Despite the existence of dedicated bioprinting technologies capable of processing biologically relevant biomaterials and living cells already available for commercialization (e.g., 3D Bioplotter®, 3DDiscovery®), the printing speed and resolution should be increased in order to better reproduce the heterogeneity found in natural tissues and produce constructs with clinically relevant dimensions in acceptable periods of time. Another important aspect relies on the inclusion of physical and chemical crosslinking methods in the bioprinting process to induce rapid gelation during biofabrication in order to maintain shape fidelity and prevent loss of resolution. The development of hybrid bioprinting approaches for skin tissue engineering, through the combination of bioprinting and other biofabrication techniques with different scales of resolution, is also an important challenge that could enhance the functional outcomes of constructs and overcome technical limitations. Another important limiting factor in skin bioprinting is the lack of printable bioinks exhibiting appropriate rheological, mechanical and biological properties. As a result of the distinctive requisites of each bioprinting strategy, most of the natural and synthetic hydrogels that have been applied for skin wound healing need to be re-engineered to allow adequate processing through bioprinting. Besides printability, such materials must exhibit a plethora of characteristics, including biocompatibility, biodegradability, ability to preserve cell viability and function during/upon printing, appropriate immunological properties and the ability to promote integration with host tissue. Bioinks should also allow the introduction of cell adhesion sites and cleavable bonds to control cell behaviour and biomaterial degradation rate, respectively. Current bioinks for skin bioprinting are essentially limited to the use of alginate, collagen, hyaluronic acid and fibrin, which restricts the range of mechanical and biological properties of the fabricated constructs. Although the ability of bioprinting approaches to generate 3D constructs with complex architectures, while allowing good control over the spatial location of biomaterials, cells and biologically active compounds, has already been demonstrated, replicating the complexity of native skin tissues has not yet been successfully achieved. To improve the level of biomimicry, efforts should be made to consider additional aspects regarding the skin appendages, gradients of ECM components, and functional vascular networks. Skin appendages are responsible for several important functions, including protection, sensing and temperature regulation, while the gradients of ECM components play a primordial role in both cell behaviour and biomechanical properties of the skin. Novel bioinks allowing the reproduction of ECM gradients present in different skin layers should be designed and their effects on cell viability and function well characterized and documented towards regulatory approval. Although vascularization is essential for the long-term viability of the construct, it remains a huge challenge in almost all tissue-engineering strategies. Recent developments in bioprinting of functional vascular networks can be explored to design 3D hierarchical skin constructs with embedded vascular cells, which will represent an important advance in the field. Finally, efforts should be concentrated in the translation of bioprinting therapies into clinical practice. This is a long and difficult issue that depends on several factors, including the scale-up of the process, and many regulatory and commercialization aspects. Since novel bioprinting approaches for skin repair and regeneration are complex and multidisciplinary, aspects related to the bioprinting technology, bioinks, cell sources, integration with existing medical devices (e.g., robotic surgical arms, medical imaging systems, real-time imaging technology), automation, reproducibility and affordability must be considered right from an early stage to reduce the time to reach the market. Collaborative work should also be focused on the establishment of good manufacturing practices, standardization of assays to control the properties of manufactured constructs, and methods for the monitoring and inspection of the printing process.
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This work was financially supported by European Regional Development Fund (ERDF) through the COMPETE 2020—Operational Programme for Competitiveness and Internationalization (POCI), Norte Portugal Regional Operational Programme (NORTE 2020), under the PORTUGAL 2020 Partnership Agreement, and by Portuguese funds through Portuguese Foundation for Science and Technology (FCT) in the framework of the project Ref. PTDC/BBB-ECT/2145/2014. Ruben F. Pereira, Aureliana Sousa and Cristina C. Barrias are grateful to FCT for the doctoral grant SFRH/BD/91151/2012, post-doctoral grant SFRH/BPD/90047/2012 and FCT Investigator research position IF/00296/2015 (FCT and POPH/ESF), respectively.
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Pereira, R.F., Sousa, A., Barrias, C.C. et al. Advances in bioprinted cell-laden hydrogels for skin tissue engineering. Biomanuf Rev 2, 1 (2017). https://doi.org/10.1007/s40898-017-0003-8
- 3D bioprinting
- Extracellular matrix
- Tissue engineering