Introduction

Tissue engineering and regenerative medicine are necessitated by the limited reparative capacity of post-natal tissues and organs. In the treatment of cutaneous defects, the use of tissue-engineered dermal substitutes, in their various forms, is increasingly routine, particularly as a life saving tool following acute thermal trauma. However, at present, even the most successful amongst these, cannot restore the full functionality and appearance of uninjured dermis.

Dermal substitute construction is variable and a range of materials, designs and cell sources have been investigated. Most commercially available products are based around polymer matrices, derived from both natural and synthetic sources. The majority of these, however, were developed during the 1990s and are designed to mimic the basic properties of the extracellular matrix (ECM).

Physiologically, on the simplest level, the ECM supplies the structural and organizational framework of developing and mature dermal tissue. This framework is created in the main through highly hydrated insoluble macromolecules such as fibrillar proteins (e.g. collagens), glycoproteins (e.g. elastin or fibronectin) and proteoglycans with associated glycosaminoglycans (e.g. versican). The tunability of this system is such that cells are able to create cell-specific micro-environments through bidirectional interaction with, and remodelling of, the ECM surrounding them.

Following damage to or destruction of dermal tissue, granulation tissue is formed. This loosely woven immature neomatrix is deposited by invading monocytes (mainly macrophages) and fibroblasts following platelet-mediated establishment of haemostasis. The early granulation tissue contains large quantities of fibrin, fibronectin, hyaluronan and also collagen types I and III [165]. Several growth factors have been identified as important for cell invasion of the site, proliferation and matrix deposition/remodelling including transforming growth factor (TGF)-β [164], platelet derived growth factor (PDGF) [85, 109, 142], fibroblast growth factors (FGFs) [85, 160] and matrikines released by matrix damage and remodelling. This early tissue provides the framework for the deposition of neodermis. Within a few weeks the fibronectin-rich matrix disappears. Hyaluronan is reduced and collagen type I/III fibres are slowly remodelled to contain less collagen III and reorganized into large bundles. The anisotropic architecture of the mature fibrillar matrix consists of both rigid supporting collagens, flexible elastic networks of elastin, and molecular connectors. This arrangement not only confers important mechanical properties but influences cell behaviour by the manner in which cells bind to, and therefore sense their microenvironment [134].

Cells adhere to the different ECM components either directly through cell surface receptors (e.g. integrins) or via intermediate factors with varying degrees of specificity and affinity. The density and number of specific cell surface receptors that bind the ligand initiates corresponding intracellular signalling events. These specific interactions, in part, control cell survival, cell phenotypes and drive cell fate decision [44, 66, 94]. Of additional importance is the spatial orientation of activated cell surface receptors. In some cases three-dimensional integrin activation is required to initiate a cellular response. An example of this is the regulation of MMP-13 production through 3D activation of α1β1 and α2β1 integrins in collagen substrates, which results in the coordinated stimulation of three MAPK classes (ERK1/2, JNK and p38). In this example, it is the balance between the p38 (activating) and ERK1/2 (suppressing) pathways through integrin-mediated ECM-cell signalling that affects the rate of enzymatic degradation of the matrix [134].

The extent and complexity of the inherent bioactivity of native dermal ECM has only recently been fully realized and is not yet understood fully. The ECM environment, which a transplanted cell is subjected to, influences cell survival and behaviour. The current gap in knowledge into what specific environmental cues are required for maximizing therapeutic efficacy of transplanted cells could explain the failure of early cellular dermal substitutes to behave as designed once in the wound bed [78, 79, 98]. It is probable that a more detailed knowledge of the complex interactions involved will allow improved substitute design. Recent advances in material synthesis and processing have facilitated a developmental move towards more sophisticated bio-inspired materials and away from educated trial and error bio-mimicry [43, 79, 176].

This review will discuss the current status of dermal substitutes and the biomaterial design that they are based on. The use of stem cells in regenerative medicine is reviewed in this context along with the characteristics and potential of various stem cell sources for application in this field. The article then moves on to focus on novel emerging techniques and materials that are likely to become important in future dermal substitutes, in particular, those containing stem cells.

Biomaterials

Perhaps unsurprisingly, given its natural prevalence, collagen is the most widely utilized scaffold biopolymer available for use in skin substitutes. It is obtained from allogeneic or xenogeneic sources, with sufficient molecular homology existing to prevent a significantly detrimental immune response in most cases [16]. The collagen in these scaffolds is often reconstituted into porous, fibrous and hydrogel structures with physical properties that can be controlled through manufacture conditions, such as chemical cross-linking to prevent degradation. However, such treatments may compromise effectiveness, for example extensive cross-linking increases matrix rigidity and reduces cell attachment and viability [24, 143]. In its natural state, collagen has a superior degradation rate (up to 6 weeks) in comparison to reconstituted collagen (1 week) [19, 62, 72].

Numerous other natural materials have been used in scaffolds for engineered dermis including hyaluronic acid [25], fibrin [2], laminin [51] and elastin [60]. These materials are usually processed into porous scaffolds through lyophilization (freeze-drying) or phase separation processes. (Fig. 1) Naturally occurring ECM molecules are considered advantageous due to their cell interaction, adhesion and signalling properties. These interactions could include direct matrix-cell interactions, through the uptake of soluble factors or through the interaction of enzymatically degraded fragments, or matrikines, with cell surface receptors [17, 20, 154]. However, with the exception of elastin, the mechanical properties of these materials are often poor in comparison to the properties of synthetic materials, which can be tailored across a broad spectrum. Even elastin, with a half life of around 70 years [27], must be organized into its naturally occurring network to function properly and retain its stability [27]. When using natural materials disease transmission and immunogenicity remain a concern. In addition to these concerns, further complications include the availability of materials; especially allogeneic material, purification of selected materials, and batch-to-batch and source variability. These also serve to increase the cost of utilizing some raw materials, meaning substitutes developed from these materials are potentially more expensive. A possible way of side stepping these problems could be through the use of recombinant technology in their production. Due to these issues, there is a desire to develop synthetic materials that can be designed to best mimic the natural ECM.

Fig. 1
figure 1

Current strategies for the production of dermal substitutes. Current scaffolds tend to be in several main forms: a Hydrogels, b porous sponge (lyophilization/particulate leaching/phase separation), c non-woven fibre, d woven fibre, e honey comb mesh. These can be inserted as acellular scaffolds or can be seeded with autologous/allogeneic cells harvested from bioposy or cell line. Alternatively cell suspensions can be directly applied to skin defects as a liquid or spray

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Synthetic materials offer the possibility of improving material control, reducing batch variation, eliminating disease transmission and providing more cost-effective scalability. Synthetic materials investigated as scaffold materials include polyurethane (PU), polypropylene (PP) [138], poly(ethylene glycol), polyglycolide (PGA), polylactide (PLA) and polylactide-coglycolide (PLGA) [29, 101, 108], polytetra fluoroethylene (PTFE), polycaprolactone (PCL)[119], polyethylene terephthalate (PET), Poly(l-lactide) (PLLA)[169] and poly(ethyleneglycolterephthalate)-poly(butylenes terephthalate) (PEGT/PBT) [21, 39], [157] Despite their ease of manipulation and manufacture, the safety of these materials remains a concern [57]. Although they may not be directly toxic, the use of synthetic materials has in some cases been found to lead to a foreign body response and fibrous capsule formation surrounding the material [138, 155]. Even though, cell interaction with these materials can be limited; they can often be easily and effectively modified through the attachment of growth factors, attachment sites or coating with ECM molecules to improve cellular performance. Zisch et al. [176] covalently decorated PEG hydrogels with vascular endothelial growth factor (VEGF) which led to improved vascularization following grafting to chick chorioallontoic membrane and subcutaneous implantation in rats. A more simplistic approach was applied by Chen et al. [29] who investigated a hybridized PLGA/collagen mesh as a 3D culture system for tissue engineering skin and found that hybridization resulted in improved cell attachment and ECM deposition of dermal fibroblasts.

Aside from manufactured matrices, several products exist that consist of intact de-cellularized dermal matrices of allogenic or xenogeneic origin (e.g. Alloderm™, Oasis™, FortaFlex™, Repliform™). The hypothesis being that these matrices offer an ideal scaffold environment for the migration and proliferation of dermal cells through the retention of structure, attachment sites and matrix-bound growth factors. Reports demonstrate that 3D adhesions are formed rapidly in cells seeded to tissue-, or cell-derived acellular matrices [33]. This suggests that these matrices if processed competently may retain and present important information that promotes attachment, proliferation and migration to a further extent than other three-dimensional and two-dimensional materials [33]. However, problems can arise when using harsh de-cellularization protocols that may remove or denature growth factors and damage matrix architecture [30, 69]. Conversely, mild protocols may fail to remove cellular material efficiently, resulting in a subsequent immune reaction in response to retained foreign cell fragments [34, 52, 166]. Of the commercially available products mentioned above; Alloderm™ and Oasis™ have had relative success in their clinical application in the treatment of chronic ulcers and as dressings for split thickness graft donor sites. Accelerated healing and decreased pain were reported in some cases in comparison to conventional treatments [88, 111, 121, 136, 158, 159]. It is worth noting that intact de-cellularized ECM represents a mature end-point structure, both in terms of architecture and molecular composition. It is possible that this environment is not ideal for inducing a regenerative wound healing response from host cells. Finally, although a scaffold material can be designed as “permanent”, generally it is considered desirable that the transplanted scaffold can be safely assimilated into the body as new matrix is generated by the populating cells.

Stem cells

Currently, the transplantation of cells in regenerative medicine is limited by the poor survival rates and persistence of the transplanted material [83, 129]. The cellular constituents of current dermal substitutes only remain present for approximately 1 month after application with only a few percent of cells surviving initial engraftment [31, 147]. If cellular dermal substitutes are to succeed, further improvements need to be made. This could be achieved through the delivery of stem and progenitor cell populations within protective and bioinstructive environments [3].

Stem cells are defined by their ability to self-replicate and produce more specialized progeny [87, 141]. Their incorporation into dermal substitutes could lead to improved therapeutic activity by directly contributing to cellular content of the healing wound, the release of paracrine factors and activation of host stem and somatic cell populations [3]. Furthermore due to their plasticity, it may be possible to increase the non-fibroblastic cell content of substitutes without the need to incorporate several different cell populations. Previous work has sought to identify suitable stem or progenitor populations for tissue engineering applications. Several candidate populations have been identified drawing from embryonic stem cell (ESC), adult stem cell (ASC) sources and more recently through induced pluripotent stem cells (iPS cells) (Table 1). Each source has advantageous characteristics, though none are without disadvantages.

Table 1 Advantages/disadvantages of selected stem cell sources for use in dermal substitutes
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Embryonic stem cells are pluripotent stem cells derived from the inner cell mass of the blastocyst, which forms several days after fertilization [152]. Due to their pluripotency, and unique ability to maintain pluripotency in long-term culture, these cells make an attractive single source to generate cells of multiple, diverse lineages [18]. However, the use of ESCs has attracted widespread public controversy leading to complex and stringent regulations governing their use. Additional safety concerns exist with the use of these cells due to reports of teratoma formation [172]. However, since the generation of the first human ESC line in 1998, at least 225 subsequent human ESC lines have been generated [64, 152].

The physiological function of ASCs is the maintenance and repair of the tissues in which they reside. ASC populations reside within niches in tissue incorporating cell-cell, cell-matrix, soluble cues, mechanical properties and soluble factor gradients to maintain steady numbers of stem cells in a stable undifferentiated state within them [161] (Fig. 2). It was previously thought that ASCs were lineage restricted to their host-tissue. However, recent work has shown that some populations of these cells are multipotent and possibly even pluripotent [74], and, therefore, capable of differentiation into a wider range of cells than anticipated [84, 130]. Certain populations of these stem cells have characteristics such as abundance and ease of extraction, facilitating their use in regenerative medicine and tissue engineering [63]. Of these, bone marrow-derived mesenchymal stem cells (BM-MSCs) and adipose-derived stem cells (Ad-MSCs) are perhaps the most promising candidates for stem cell therapy. BM-MSCs are the most characterized ASC population [23, 54]. They are present in bone marrow in low density (~1 MSC per 5 × 103 mononuclear cells) [77]. Although no single BM-MSC specific marker has been identified positive expression of CD73, CD90 and CD105, when not expressing CD34, CD14 and CD45 has been used to identify MSCs from mixed populations [7, 153]. Bone marrow-derived stem cells have been used clinically in the treatment of cutaneous defects, showing improved healing outcomes in comparison to conventional treatments [13, 42].

Fig. 2
figure 2

Schematic representation of adult stem cell niche and factors involved in controlling stem cell behaviour within it. The undifferentiated stem cell (blue) is situated within the niche surrounded by supporting cells (red), specific extracellular matrix molecules and in proximity to microvasculature. The fate of the stem cell is controlled within the niche through direct interactions with the ECM and supporting somatic cells. Further to this, the release of paracrine and autocrine factors alters the behaviour of the resident stem cells. The release of these factors can be influenced by cytokine/growth factor release and/or neuronal signalling (green cell) to tune stem cell behaviour to the tissue requirements, e.g. increased division and migration during wound healing

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Adipose-derived stem cells offer an attractive alternative to BM-MSCs. Adipose tissue has the highest abundance of stem cells of any tissue in the body, which can be easily harvested through liposuction surgery and demonstrate multipotency in vitro [55, 179]. Flow cytometric analyses of Ad-MSCs have revealed that they share similar surface receptors with BM-MSCs. This importantly includes the activated lymphocyte adhesion molecule CD166, which has been shown to identify BM-MSC populations with multipotent differentiation potential, and integrin β1 which is also associated with the cells in the epidermal stem cell compartment [53, 75, 89]. Ad-MSCs have yet to be used in the clinic to treat cutaneous injury but have shown promise in animal wound healing models [5].

Hair follicle-derived stem cell (HFSC) populations offer a further option for application in stem cell therapy due to their natural position and roles within the skin, their ease of extraction and multipotency [4, 32, 70, 135]. Of the discussed ASC populations HFSC have been least investigated for tissue engineering skin. However, they may represent an important future resource in the development of dermal substitutes.

Another candidate population of cells for use in the development of cell instructive dermal substitutes would be the newly promising iPS cells. Since the discovery by Takahashi and Yamanaka [150] that pluripotency could be induced in somatic cells through the delivery of several growth factors, the field of iPS cells has rapidly expanded. Their research elegantly demonstrated the direct reprogramming of mouse embryonic fibroblasts through the enforced expression of four transcription factor genes (Oct3/4, Sox2, Klf4 and c-Myc) [112, 150]. Since this initial work, several groups have confirmed and improved this technique and generated human iPS cells that are epigenetically and developmentally indistinguishable from ES cells [92, 97, 123, 124, 162]. The c-Myc gene has since been shown to be dispensible for reprogramming which has resulted in reduced incidence of malignant transformation in iPS derivatives [112]. Recently, virus-free generation of iPS cells was achieved by Zhou et al. [174] and Kim et al. [80] using recombinant proteins. Zhou et al. [174] used polyarginine protein transduction domain tagged proteins repeatedly added to culture media containing valproic acid (VPA) to reprogramme murine fibroblasts. In a similar way, Kim et al. [80] used proteins tagged with a highly basic peptide sequence derived from human immunodeficiency virus-TAT protein to reprogramme human fibroblasts. These tagged proteins were released from HEK293 (human embryonic kidney) cells engineered to over express them [80]. The virus free reprogramming of cells, removes the major safety concern of retroviral gene transfer-induced mutagenesis. This technique potentially means that a tailor made population of pluripotent stem cells can be constructed rapidly from a small sample of the patient’s own somatic cells. Several safety concerns, such as tumor formation potential, remain to be addressed prior to clinical application of stem cells from this source. Although this field is progressing rapidly, akin to other breakthroughs, it is likely to be some time before the transition from laboratory to clinic is made possible [12, 28, 112].

The transplantation of stem cells, notwithstanding the source of the cells, has encountered several obstacles, in particular, controlling the fate of the engrafted cells once applied to the wounded dermis. Several schemes have been suggested such as those by Discher et al. in which materials could be engineered to contain stem cells within niche environments designed to maintain the ‘stemness’ of the cells. As progenitor cells move out from this niche environment, differentiation could be directed through engineered ECM design, as in natural tissue [9, 106, 156]. An alternative approach could be to attract circulating native stem cell populations once a scaffold is transplanted. This could be achieved through incorporation of some homing signal and/or specific binding sites to mobilize resident stem cell populations and encourage residence within the substitute [35]. Sasaki et al. [139] investigated whether systemically delivered MSCs are able to differentiate into multiple skin types and contribute to wound healing in murine models. They found that wound-site intradermal injection of the chemokine SLC/CCL21 increased MSC homing to the defect and resulted in accelerated wound repair. A related factor, stromal cell-derived factor-1 (SDF-1), is rapidly overexpressed following tissue injury [10] and is known to recruit endothelial progenitor cells to wound sites [14, 15, 46]. Recently, Rabbany et al. [132] investigated the effects of a SDF-1-releasing alginate scaffold on cutaneous wound healing in a porcine model [132]. The authors reported that wounds treated with SDF-1 fully closed without scar formation. Interestingly, they go on to link this finding to the sonic hedgehog (Shh)-Gli pathway, which is responsible for regulating foetal organ development and implicated in scarless wound healing [36]. Sonic hedgehog has been shown to upregulate SDF-1 and also enhance SDF-1’s recruitment of progenitor cells to wound sites [8, 86]. The authors suggest that Shh-mediated SDF-1 upregulation may be the mechanism by which decreased scarring is achieved in foetal wound healing and could be exploited to achieve similar results in adult tissues [132]. Materials designed to deliver these chemokines and other factors in conjunction with stem cell transplantation could result in increased efficacy of stem cell treatments.

Aside from the delivery of soluble factors, substrate stiffness is now recognized as an important factor affecting the lineage commitment of stem cells [41, 137]. Engler et al. [41] and Saha et al. [137] demonstrated that in 2D culture. MSCs cultured on soft substrates commit to neurogenic fates, on moderately soft substrates commit to myogenic fates and on rigid substrates commit to osteogenic fates. Recently, Pek et al. [125] demonstrated that substrate stiffness has a similar effect on MSCs cultured in 3D systems. By utilizing an inert thixotropic polyethylene glycol–silica nanocomposite gel, Pek et al. [125] were able to demonstrate the direct effect of matrix stiffness whilst controlling the biological cues presented to the cells within the system, in particular, the effect of integrin-ligand binding mediated cell responses [125].

Current strategies for stem cell-assisted cutaneous repair

It is hypothesized that stem cells can improve healing through three basic mechanisms; (1) creation of an environment that enhances the regenerative capacity of endogenous cells, (2) transdifferentiation, and (3) cell fusion [99, 131, 145, 146]. Numerous studies have demonstrated the potential of transplanted ASCs to improve wound healing rates in animal models and clinical trials [5, 42, 47, 48, 114, 133]. These early trials have shown that cell preparation and delivery methods alter the therapeutic efficacy of the transplanted stem cells leading to contrasting reports in the literature [5]. For example, a murine model treating dermal wounds in diabetic mice with human adipose-derived stem cells, found that when the cells were delivered as a suspension, no advantageous healing effect was observed [5]. However, when cells were delivered as three-dimensional aggregates, wound healing was significantly accelerated [5]. Other studies deliver Ad-MSCs [3, 117, 120] and BMSCs [59, 93, 99, 107, 114, 173] in dermal substitutes. One reason for the observed benefit of cells delivered in this way could be that the harsh, necrotic environment of the wound may lack the appropriate milieu to support the transplanted stem cells. It is clear that the environment in which the cells are delivered must confer some level of protection to maintain viability whilst allowing direct contact with, and migration into, the wound environment.

Designing 3D biomaterials to control stem cell-assisted dermal regeneration

The challenge, thus, lies in the identification, design and manufacture of dermal scaffolds which maximize the healing potential of transplanted stem cells. Effectively manipulating the enormously complex cell interactions and regulators of cell fate in three dimensions still represents a significant problem even in vitro. The majority of information that we have about cell signalling networks and interactions has been derived from 2D in vitro culture systems. It is a challenge, therefore, to micro-engineer a 3D environment when mechanisms of intra- and inter-cellular signalling networks in this setting are not well known [95]. This situation is exacerbated by the tissue-dependent variability of the niche itself down to the molecular level and the dissimilar responses of stem cells from different sources to similar materials and microenvironments [118]. The materials present may as a result have a differential effect depending on cell source. This, and the lack of understanding of how dimensionality affects signal transduction pathways could be addressed with the development of quantitative models [95, 96]. Using high throughput techniques to test biologically driven assumptions into cell behaviour through accurate creation of specific cell-matrix interactions is likely to be an efficient method for the development of materials with the complexity needed to accurately direct cellular processes [6, 45, 100, 115].

Several materials and fabrication techniques have been developed or recently applied to tissue engineering that have the potential to make the manufacture of such complex materials possible (Table 2; Fig. 3). The presentation and release of growth factors as a means of instructing cell behaviour is an area which has received considerable interest. Micro- and nano-spheres incorporated into matrices of various forms can be designed with release profiles to deliver factors in a concentration and time-specific manner [148, 163, 170]. This approach also protects the growth factors from degradation. Alternative approaches tether growth factors to scaffold polymers to be released and activated through cell-mediated enzymatic action. One example of this cell-demanded release, pioneered by Hubbell, involves VEGF covalently bound to poly (ethylene glycol) hydrogel matrices [38, 177, 178]. This couples growth factor concentration directly to cell remodelling and was shown to induce controlled, natural-like blood vessel growth and formation in chick chorioallontoic membrane and rats [38].

Table 2 Current advantages/disadvantages of some novel fabrication techniques
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Fig. 3
figure 3

Flow chart representing some of the design considerations for next generation dermal substitutes. The design of dermal substitutes starts with the choice of substrate material which can be from a biological source or a synthetic material. The grey box represents selected instructive cues and promising methods of incorporation into substrates. This instructive scaffold may then be inserted as an acellular substitute or cells may be introduced (purple box). The incorporation of viable cells into the designed substrate is also an important design consideration with several populations to select from. The 3D spatial positioning of these cells within scaffolds should improve the speed and functionality of substitute assisted healing. The efficacy of transplanted cells is dependent on delivery method and scaffold material design

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Matrix tethering and micro/nano-sphere incorporation could also be used to create spatial and temporal concentration gradients in three dimensions [144]. Gradients of bioinstructive molecules presented in these ways, effectively reduce the amount of the particular molecular factors needed to produce the desired cellular response [144]. Aside from chemical gradients, the incorporation of other gradients such as (but not exclusive to) pore size, substrate stiffness and cell attachment site distribution or type are also important. Spatial patterning of some or all of these physical and biological cues mimics the natural composition of dermis and other tissues. For example, the recreation of the dermal–epidermal bilayer could be facilitated by gradients of keratinocyte/fibroblast-specific attachment sites, growth factors, differentiation factors and scaffold properties. These factors could work synergistically to induce correct cell fate decisions from transplanted and endogenous stem and progenitor populations for rapid, successful wound healing.

One of the promising novel techniques to emerge in recent times is bioprinting. Bioprinting can be defined as the automated, computer-aided deposition of biological materials and cells in a predesigned pattern [102]. It is possible through this relatively new technique to construct 3D-engineered tissues via layer-by-layer deposition. The layered structure of the skin may mean that this technique is particularly applicable to the creation of dermal substitutes.

The term bioprinting encompasses a number of techniques including inkjet [116, 167], laser guidance direct write [122], laser-induced forward transfer (LIFT) [65], extrusion-based printing [168] and electrostatic-based jetting [58]. These techniques have several characteristics that can be harnessed for the creation of next generation dermal substitutes. The computer-aided nature of the technique means that precision deposition of cells, ECM and bioactive agents in defined amounts is possible with relative ease [116]. Importantly, through the different ink dispensers of printers, it is possible to use multiple ECM components simultaneously. Non-contact processes, where a jet is formed for microsecond time-scales, also mean that the deposition of bio-ink can be made on a variety of substrates and even liquid [37]. Ink-jet delivery systems have received particular attention but several limitations currently exist such as high shear stress on extrusion and high impact on deposition of the bio-ink droplet [104]. In bio-printing cells, maintaining viability and gene expression profiles represents a challenge, with some success reported [56, 113, 116]. The viscosities, surface tensions and densities of the bio-inks used in these systems are limited by the process requirements, restricting the concentrations of some ECM molecules that may be used [104]. This becomes important in the construction of 3D bio-printed substitutes where deposited material must be capable of retaining shape and offering structural support. At present 3D construction relies on either layer-by-layer deposition of bio-inks which can be gelified or through the use of a 2D ‘bio-paper’ material which is then stacked sequentially to give a 3D structure [105, 116]. It has been suggested that the latter method may be the only feasible solution to mechanical problems arising when using direct 3D gel-based approaches [56]. This approach could also improve general handling properties for clinical application of dermal bio-printed constructs. In addition, the form of bio-paper utilized for 3D construction may allow the combination of bio-printing with another emerging processing technique, electrospinning. Ultimately it may be possible to develop in situ bio-printing clinical tools. This technology could feasibly allow the direct regeneration of cutaneous injury [103].

Electrospinning supplies a simple, cost-effective method to construct porous scaffolds with uniform fibres in the nano-scale [22]. The process involves charging a polymer solution or melt to a high potential and injecting it through a blunt needle towards a grounded collector, typically 10–30 cm away. The electrical potential of the resulting fibre drives it across the air gap between the needle and collector, with the solvent evaporating during this transit. This results on the formation of dry polymer fibres on the collector [11]. The scale of the fibres can be produced in these scaffolds corresponds to the natural ECM, with reports demonstrating that they promote normal cell-cell and cell-matrix interactions [61, 91]. The majority of the natural and synthetic biomaterials used to construct nanofibrous scaffolds are those previously used to construct scaffolds by other methods. Electrospun scaffolds have been shown to promote fibroblast proliferation in vitro [175]. Working on full thickness wounds in mice, superior ‘take’ and levels of contraction were observed in collagen electrospun scaffolds in comparison to freeze-dried scaffolds [128].

The electrospinning technique lends itself to the creation of sequentially layered scaffolds [11]. Through altering the spin parameters, physical properties such as fibre diameter, porosity and fibre alignment (anisotropy) can be adjusted. As an example, fibre diameter, alignment and density could be adjusted to create a scaffold that structurally mimic the reticular and papillary dermis. Yang et al. [171] electrospun a collagen-PCL multi-layered scaffold containing viable human dermal fibroblasts and keratinocytes deposited directly through the electrospinning process. Although viability of the cells was maintained it has been suggested that care needs to be taken when spinning cells directly due to exposure to high electrical field and organic solvents often used in the process [11]. To create biologically instructive scaffolds growth factors may be directly added to the polymer solution prior to spinning with studies indicating that biological activity is maintained in vitro and in vivo [49].

Conclusion

Over recent years, biomaterials science has developed novel techniques which can be applied to tissue engineering and regenerative medicine. Amongst these newly emerging techniques, bio-printing and electrospinning have the potential to produce scaffolds which incorporate instructive factors and viable cells. A number of studies have demonstrated the principles of viable cell positioning and bioactive molecule patterning [73, 116] and managed to control stem cell fate [76, 127]. The translation from these improved scaffolds to instructive three-dimensional dermal substitutes is not a simple task. Mimicking the complexity of the natural matrix may not be a practical option when producing commercial dermal substitutes. Which factors and properties are essential to create a niche for stem cells conducive for the regeneration of dermal tissue requires further investigation. The precise positioning of molecules and cells may allow the construction of differentiated structures currently lacking from dermal substitutes like nerves, sweat glands and pilosebaceous units. The precise and automated nature of bio-printing makes it a particularly attractive choice for the fabrication of dermal substitutes containing stem cells. In the future, it may even be possible to bio-print bespoke dermal substitutes in situ.