Recent Advances in the Design of Three-Dimensional and Bioprinted Scaffolds for Full-Thickness Wound Healing

Introduction

Cutaneous wound healing is a complex and dynamic process where the skin orchestrates a myriad of events to bring about tissue repair, regeneration, and restoration of tissue function.^1,2 To achieve desirable healing outcomes, it is imperative for wounds to be appropriately managed based on their location, size, and depth.^3,4 The treatment of full-thickness wounds (FTW), wounds where both the epidermal and dermal layers are lost, is of particular concern as the loss of an intact skin barrier can affect the whole body.^5,6

In FTW, key stem cell populations originating from the basement membrane and skin appendages (hair follicles and oil and sweat glands) are destroyed, compromising the skin's ability to regenerate itself.^7,8 In addition, complete reepithelialization and wound closure can only occur when a well-vascularized dermis is reinstated.^6,9 Even though eventual wound closure restores the skin's key barrier function, protecting against moisture loss and infection from pathogens, other important skin functions—thermoregulatory, excretory, metabolic, and sensory functions—are not restored in the absence of skin appendages.¹⁰

Further complications arise when full-thickness skin loss occur in burns that extend to greater than 20% total body surface area,^11,12 and/or in the presence of underlying comorbidities such as vascular disease, diabetes, and metabolic syndrome, which impair healing.¹³ In 2017, an estimated 8.9 million fire-/heat-related incidences were recorded globally.¹⁴ While burn injuries affect all populations, patient management remains suboptimal in many countries, resulting in considerable patient disability, disfigurement, and death.^14,15

Conversely, the prevalence of chronic wounds like neuro-ischemic ulcers (NIU), vascular ulcers (VU), and pressure ulcers (PU) is rising in many populations coincident with aging and metabolic syndrome.^16,17 In Singapore, the number of patients diagnosed with NIU and PU increased by 95% from 2013 to 2017.¹⁸ Over 97% of NIU patients also suffered from diabetes, while 86.5% of PU patients were older than 65 years. The incidence and impact of wounds that are hard to heal is increasing, resulting in even greater socioeconomic burden on patients, clinicians and societies.

Seeking to address these challenges, the demand to harness advanced medical technologies to improve wound management is escalating. In recent years, one technology in particular has gained steady traction: additive manufacturing, or three-dimensional (3D) printing, as it is more commonly known. In the biomedical field, 3D printing has already achieved considerable penetration, illustrated by the fabrication of pharmaceuticals, culminating in the FDA-approved SPRITAM^® for epilepsy,¹⁹ as well as the construction of patient-specific prostheses, devices, and instruments.²⁰

In the following, we discuss the potential of exploiting 3D printing as a tool to generate scaffolds for cutaneous wound healing, in particular FTW, and their application to supplement, or restore, the integrity and native functions to injured skin tissue. First, we look into the current wound management options to understand their limitations. Next, we delve into how these limitations may be addressed with 3D printing approaches. Finally, we explore how 3D printed scaffolds can be applied in FTW healing and the challenges in bringing such technology from bench to bedside.

Current Treatment Options and Their Limitations

Current clinical management of severe burns and chronic wounds often indicates surgical intervention. Commonly, skin tissue harvested from uninvolved donor site(s) is grafted onto sites of injury. While full-thickness skin grafts are successful, transplanted tissue is more commonly applied as split-thickness skin grafts (STSG), permitting larger areas to be grafted.^2,21,22

While skin grafting has beneficial outcomes,²² certain complications limit wider usage: (1) patients presenting with large surface area injuries and small stature (e.g., children) often have insufficient viable tissue suitable for harvest; (2) immune responses and disease transmission are not without risk when grafting tissues from unrelated (allogeneic) donors; (3) complications associated with donor site are not uncommon, especially in individuals subjected to severe trauma; in such situations, the creation of another wound can be counteractive; and (4) impaired or infected wound bed can lead to graft failure.²³ A number of commercial tissue-engineered skin substitutes have been developed to address some of these shortcomings. Over the past decade, the availability of tissue-engineered substitutes has reduced demand for autologous donor tissue, minimizing donor site morbidity and improving clinical outcomes for patients with large area skin injuries.

Current commercial skin substitutes can be broadly categorized as acellular and cellular porous constructs.²⁴ Scaffolds that are acellular, such as Integra^®, MatriDerm^®, and AlloDerm™, act as dermal templates intended to support infiltration of native cells, allowing skin regeneration, remodeling, and vascularization to take place. Some scaffolds contain live cells, sourced from allogeneic donors (e.g., Dermagraft^®, Apligraf^®, and Orcel™) or populated with autologous cells (e.g., TissueTech^®, PermaDerm™, and denovoSkin™). Proliferating cells within the scaffold provide an exogenous source of growth factors and cytokines that aid the wound healing and tissue repair process. Scaffold-free constructs, including cultured epithelial autografts (CEA) in sheet form (Epicel^®), or noncultured epithelial cell sprays (ReCell^®) are designed to augment wound healing, but are not suitable for FTW as they do not provide an intact extracellular matrix (ECM) to support attachment and survival. For a more detailed description of these skin substitutes and their applications, the reader is referred to comprehensive reviews by others.^25–28

While the application of tissue-engineered dermal substitutes has improved clinical outcomes, several issues remain to be adequately addressed. Graft failure in several products, for example Integra, Dermagraft, and Apligraf, has been attributed to infection.^29–31 Many dermal constructs consist of a single component like collagen, or hyaluronan, and do not recapitulate the compositional complexity evident in native ECM. Decellularized human-derived matrices like Alloderm may perform better in this respect, but these single-layered products often require concurrent, or secondary, surgical interventions with STSG or CEA to substitute for lost epidermis. In addition, the manufacturing of these commercial skin substitutes utilizes classical engineering methods such as layering, casting, and freeze-drying, all of which do not offer precise control to define the structural intricacies of native skin.^32,33

The realization that skin is not the simple tissue it appears from the outside—rather, skin is a complex tissue, populated with multiple cell types, multiple sensory, secretory, and protective structures, and structural complexity not found in any other mammalian organ—has driven interest in applying newer technologies, such as 3D printing, to regenerate the structural complexity and organization of the tissue.

Through 3D printing, precise control over scaffold microarchitecture can guide and promote cellular colonization, while spatial integration of heterogeneous materials and minority cell populations can replicate the diverse and varied functions intrinsic to full-thickness native skin. Three-dimensional printing further provides the ability to fine-tune printing resolution and design, creating flexibility, and capability to personalize scaffolds to individual and frequently irregular wound dimensions. Although currently facing several regulatory and logistical challenges, the potential for point-of-care tissue therapeutics should also be recognized. The use of 3D printing to achieve desirable features of an ideal wound scaffold is illustrated in Table 1.^33,34

Overview of 3D Printing Techniques for Dermal Scaffolds

Three-dimensional printing is an additive manufacturing process that uses digital information to fabricate user-defined objects layer by layer. Its accuracy, cost-effectiveness, and ability to produce geometrically complex shapes is recognized in many industries such as architecture, medicine, and manufacturing. In accordance with the ISO/ASTM 52900 standard, a total of seven different 3D printing processes are established^35,36; material jetting, material extrusion, and vat polymerization are the most widely used processes and will be briefly outlined below. Three-dimensional bioprinting is a relatively newer approach, which synergizes with existing technologies by introducing living cells as another additive to the printing “inks.” Printed live tissue constructs are intended for “biomanufacturing” (e.g., biotherapeutics), tissue-specific testing (e.g., cosmetics and toxicity), transplantation, and biomedical R&D applications.

Material extrusion

Extrusion-based 3D printing is the most common printing technique, first developed by S. Scott Crump in 1988, trademarked as “fused deposition modeling.” The basic principle of extrusion-based printing involves pneumatic, or mechanical, forced ejection of liquefied materials from a nozzle onto a build platform. Moving the print head in an XY plane facilitates 2D layer-by-layer addition according to the computer-aided design (CAD) instructions, producing the desired 3D model.³⁷ In comparison to other 3D printing techniques, extrusion-based printing is relatively simple and cost-effective and has excellent compatibility with a wide variety of materials.³⁸

Material jetting

Inkjet printing is an alternative 3D printing technique derived from 2D inkjet document printing technology. Droplets are ejected from drop-on-demand printheads that generate pressure pulses using thermal, piezoelectric, or electromagnetic methods.³⁹ The technique distributes individual ink droplets to specific locations on the substrate and initiates polymerization using UV light after each layer. The addition of successive layers builds up to create a 3D structure. Surface quality and dimensional accuracy of the 3D models printed by this method are affected by process parameters such as part positioning, printing orientation, nozzle cleanliness, and machine preparation.⁴⁰ However, light-based cross-linking is unsuitable for living constructs, potentially causing off-target genetic defects. Evidence also indicates the quality of finished products tends to be inhomogeneous and deteriorates rapidly, due to inefficient cross-linking. Consequently, inkjet printing is utilized for applications requiring high-speed printing, rather than printing accuracy.⁴¹

Vat polymerization

Laser-/light-assisted 3D printing, also known as stereolithography (SLA), is a freeform fabrication technique known for its high resolution, accuracy (≤20 μm), and compatibility with a wide range of materials. SLA devices comprise a pulsed laser source, focusing system, energy-absorbing layer, and a vat of liquid photopolymer resin. Three-dimensional models are generated from the bottom-up: a computer-controlled laser beam triggers cross-linking of photosensitive polymers in the vat of bulk ink, layer by layer. Cross-linking kinetics in SLA are determined by factors such as light intensity, scanning speed, and exposure time.⁴²

Three-dimensional bioprinting

The flexibility, speed, and ability to create complex multicomponent constructs have attracted attention for its potential to address unmet medical needs. Three-dimensional bioprinting, a subcategory of 3D printing, is an innovative approach purposed with creating biomimetic tissues by manipulating bioactive molecules, biomaterials, and living cells. Biologically functional constructs generated through bioprinting are mechanically stable, resemble native tissue microenvironments, and support the metabolic and physiological needs of complex living tissues.⁴³ The three main 3D bioprinting modalities of inkjet, extrusion, and laser-/light-assisted bioprinting address multiple aspects of biomanufacturing. However, each of these methods has strengths and weaknesses evident from variations in cell viability, resolution, and fidelity of printed constructs.

The challenge for bioprinting using inkjet technology is shear stress. Shear stress is a phenomenon caused by tangential forces and experienced by the fluid ‘ink’ as it passes through the jet heads. It is common to all flowing liquids; however, it has greater significance for fluid flow in small diameter tubes. Friction with tube walls increases the temperature of flowing fluids. Compression and decompression also occur as the ink passes through the nozzle, adversely affecting cell physiology, organelle integrity, and cell viability. Strategies intended to minimize these physical parameters, such as modifying fluid viscosity, optimizing cell densities, and flow parameters, complicate the fabrication of biologically relevant constructs. This is especially challenging when fabricating tissues with high cell densities, as additional difficulties, for example, nozzle clogging, are encountered.⁴⁴ Despite these drawbacks, the low cost, efficient printing speed, high resolution, and capacity to generate concentration gradients ensure inkjet bioprinting remains popular.⁴⁵

Extrusion bioprinting is the most frequently used bioprinting modality due to its compatibility with a wide diversity of materials such as hydrogels, synthetic polymers, and cell spheroids.^46,47 Compared to inkjet-based bioprinting, extrusion bioprinting can print constructs at high cell densities. However, high shear stress is associated with high viscosity inks formulated for extrusion printing, reducing the viability of printed cells (40–86%),⁴⁸ which is less than the viability of cells typically printed with inkjet-based bioprinting (>90%).⁴⁴ Extrusion pressure and nozzle aperture can be altered to improve cell viability, but this also affects printing speed and resolution.

Laser-/light-assisted bioprinting (LaBP) is a printing modality that does not subject cell-laden bioinks to nozzle-associated shear stress and compression/decompression. Thus, LaBP is able to print high-viscosity biomaterials with minimal concerns of clogging.⁴⁹ LaBP also offers acceptable printing resolution (i.e., precision) and cell viability postprinting; however, LaBP is not caveat free. To achieve high precision, LaBP printing is slow. When printing live tissue, additional measures must be invoked to prevent printed constructs from dehydrating.^46,47 The technology is also relatively wasteful; a surplus of bioinks makes it less economical than competing technologies. Thus, optimization of the printing parameters is critical in ensuring printed constructs maintain the viability of printed cells, and is capable of supporting the physiology and functions of printed structures.^46,47

Three-Dimensional Printed Acellular Dermal Scaffolds

Printed scaffolds aid in restoring the integrity and native functions to injured skin tissue. Material structure and ECM components, in particular, are the principal cues that provide form and function for solid tissues. Thus, one emergent strategy is to fabricate cell-free tissue scaffolds. Biochemical and structural cues are incorporated into the scaffold to create a prohealing environment guiding endogenous tissue repair processes to repopulate and “repair” the tissue defect. Compared to cell-populated scaffolds, this approach circumvents issues associated with histocompatibility (host versus graft disease) and permits a wider selection of materials to be considered when designing and fabricating acellular scaffolds. For example, FDA-approved thermoplastics with desirable physicochemical properties, such as poly(lactic acid) (PLA), can be used.⁵⁰ In this section, we review how scaffold materials and additives can aid in wound healing.

Scaffold material

Materials employed to support 3D printed tissue scaffolds include natural proteins (e.g., collagen, gelatin, and fibrin) and polysaccharides (e.g., hyaluronan, alginate, chitosan, and cellulose). Synthetic alternatives include polymers (e.g., PLA, poly(caprolactone) (PCL) and poly(lactic-co-glycolic acid) (PLGA)).⁵⁰ Scaffold composition is traditionally limited by its “printability” and “biocompatibility” properties within the “biofabrication window.”⁵¹ However, by combining different materials (i.e., blending), composite scaffolds with desirable properties may be generated with the capability to support cell adhesion and proliferation.

Blending scaffold materials that provide structural support and cellular adhesion is exemplified by Liu et al. who mixed alginate and gelatin.⁵² Cross-linked alginate conferred mechanical strength to the scaffold, while gelatin contributes arginine-glycine-aspartate cell adhesion motifs, supporting cell attachment.⁵³ Alginate/gelatin scaffolds implanted onto FTW on the dorsum of mice promoted cell migration and proliferation, supporting improved formation of granulation tissue compared with control mice that were left untreated.⁵² In a separate study, a scaffold blend formulated from silk-derived sericin (providing adhesion motifs) and gelatin methacrylate (GelMA, providing mechanical strength) also improved cell viability and rheological properties that allowed the formulation to be printed.⁵⁴

In contrast to material mixtures, composite scaffolds can also be constructed in distinct, connected layers. To achieve an optimal balance in scaffold porosity, Wang et al. employed extrusion printing to fabricate a bilayer structure of alginate and PLGA (Fig. 1A).⁵⁵ When implanted onto rat FTW, highly porous alginate promoted cellular ingrowth, vascularization, and collagen deposition, while a semipermeable PLGA outer layer protected the underlying alginate from moisture loss and passively enhanced the scaffold's mechanical characteristics. This construct, modeled after the dermal and epidermal layers of human skin, was found to be more effective in promoting collagen deposition and blood vessel formation compared to wounds dressed with simpler constructs of either alginate or PLGA alone (Fig. 1B).

In addition to facilitating cell adhesion and migration, the native surface property of scaffold materials may also provide beneficial effects for wound healing, for example, inhibition of bacteria growth. The cationic surface of chitosan is well documented to confer antibacterial properties.⁵⁶ Antibacterial activity was especially highlighted by Intini et al. who found that printed chitosan scaffolds prevented wound infection in diabetic rats compared to inert carboxymethylcellulose-/alginate-based scaffolds.⁵⁷ Nonetheless, in some situations, high charge density may hinder cell adhesion. This can be balanced by incorporating negatively charged materials like gelatin, forming a gelatin/chitosan composite exhibiting improved fibroblast attachment and scaffold printability.⁵⁸

Scaffold with bioadditives

The inclusion of bioadditives with foundational scaffold is another critical aspect of 3D printed acellular constructs. Bioactive compounds can be embedded into scaffold structures to tailor biological activities and create a more holistic wound healing environment.

Various additives have been incorporated into 3D printed scaffolds to inhibit the growth of bacteria. Natural polysaccharides such as chitosan (produced by crustaceans) and pectin (produced by terrestrial plants) possess inherent antibacterial properties. Upon inclusion into scaffolds, these compounds retain their inherent properties and add novel functionality to the scaffold.^59,60 Lactoferrin, an antimicrobial peptide secreted by mammals, has been homogenously mixed with PCL-formulated inks and the extruded scaffold demonstrated to retain antimicrobial activity.⁶¹ Antibiotics, nafcillin and tetracycline, have been incorporated into polymer scaffolds to provide controlled drug release.^62,63

Alternatively, bactericidal activity can also be affected through inorganic metal ions. Nanocrystalline silver, in particular, has well-established efficacy and broad-spectrum antimicrobial properties, and is especially useful for controlling multidrug-resistant pathogens.^64,65 Therefore, several investigators have attempted to impregnate silver nitrate (AgNO³⁺)^66,67 or silver nanoparticles (AgNPs) into dermal scaffolds by 3D printing.^68,69

The ability of AgNO³⁺-loaded PCL scaffolds to inhibit growth of Staphylococcus aureus was highlighted by Muwaffak et al.⁶⁷ As an additional benefit, scaffolds were printed according to the complex contours of individual wounds, providing better conformation to uneven wound boundaries and more effectively delivering silver ions directly into the wound.

Shi et al. demonstrated the bactericidal efficacy of AgNP-loaded scaffolds using a mouse FTW model inoculated with Escherichia coli and Staphylococcus aureus.⁶⁸ Polydimethylsiloxane (PDMS) scaffolds impregnated with AgNPs clearly reduced the number of recoverable, viable bacteria (colony-forming units) in the infected wounds. Interestingly, AgNP-loaded PDMS scaffolds infused with silicon oil were found to reduce the presence of live bacteria in infected wounds even further. Wounds treated with PDMS scaffolds without AgNPs, and PDMS scaffolds loaded only with silicon oil, healed more slowly than uninfected wounds. Where patients may be allergic to silver, incorporation of gallium, or copper, has been suggested as possible alternatives.^67,70

In addition to providing a porous substrate for anchorage and controlling the growth of pathogenic microorganisms, other studies have evaluated the value of introducing mitogens and growth factors into the wound environment.⁷¹ Major growth factor families involved in the cutaneous healing cascade are platelet-derived growth factors (PDGF), epidermal growth factors (EGF), insulin-like growth factors, fibroblast growth factors (FGF), and transforming growth factor beta (TGF-β), with recombinant versions of PDGF, EGF, and FGF commercially available for wound management.^72,73

For smaller wounds, treatment with mitogens and factors targeted to accelerate wound closure are unlikely to deliver measurable benefit; however, for large surface area wounds (e.g., burns and chemical injuries), substantial clinical benefits become self-evident when growth factors reduce time to wound closure.⁷² Using a rat excisional model, with a relatively large 16 mm splinted wound, Gao et al. demonstrated the benefits of EGF on wound closure.⁷⁴ EGF immobilized onto synthetic PLGA scaffolds with mussel-derived adhesion protein supported up to 99% wound closure after 21 days. Other growth factors, for example, FGF, incorporated into printed gelatin/silk-fibroin scaffolds, are also effective.⁷⁵ Using a 20 mm FTW rat model, Xiong et al. found that, while granulation tissue was evident in all wounds, only scaffolds loaded with FGF exhibited full neo-epidermis coverage over wounds. Notably, injuries dressed with FGF-loaded scaffolds also exhibited mature collagen organization.

One hypothesis for the poor healing observed in chronic wounds is the presence of elevated proteolytic activity in the wound bed, which is speculated to degrade endogenous growth factors and mitogens; thus, the wound environment is deprived of appropriate signals and may even initiate a nonhealing environment in the first place.^71,76 To address this possibility, Wan et al. designed a bilayer scaffold comprising AgNP-containing gelatin hydrogel overlaid on extruded GelMA loaded with PDGF isoform that is composed of two B subunits (PDGF-BB) (Fig. 2A).⁶⁹ This double-layered design was intended to protect wounds from infection, through the antibacterial upper layer, and simultaneously deliver mitogenic growth factors to granulating tissue in wound bed, through the basal layer. In diabetic mice, PDGF-BB-laden scaffolds accelerated granulation tissue formation with increased vascularization and collagen deposition (Fig. 2B). With the dermal component reinstated, subsequent reepithelialization and wound closure were achieved more rapidly than wounds dressed with homogenous gelatin scaffolds.

It is recognized that prolonged exposure to mitogenic growth factors may have adverse consequences, for example, hyperproliferation, fibrosis, and dysplasia.⁷⁷ This may account for the increased incidence of neoplasms and dysplasias reported in patients with diabetic foot ulcers treated with high doses of the FDA-approved recombinant PDGF-BB—Becaplermin.⁷⁸ Similar data have been reported for TGF-β1, a cytokine applied therapeutically for its anti-inflammatory activity, also stimulates fibrosis during the remodeling phase.⁷⁹ To minimize the formation of scars (fibrosis) in patients with severe burns, Navarro et al. evaluated the incorporation of halofuginone into 3D printed keratin scaffolds.⁸⁰ Halofuginone, a semisynthetic drug, interrupts downstream phosphorylation in the TGF-β signaling cascade, attenuating collagen synthesis and fibrogenesis.⁷⁹ Using a porcine cutaneous burn model, wounds treated with halofuginone-impregnated scaffolds were observed to heal more slowly.⁸⁰ However, healed wounds displayed improved collagen organization and qualitatively better healing outcomes.

Additives that do not contribute directly to wound healing have also been incorporated into scaffold formulations to improve and monitor the wound healing process. Anti-inflammatory drugs like diclofenac sodium,⁸¹ and local anesthetic lidocaine,⁸² help to alleviate pain during the healing continuum, benefitting tissue articulation and pliability. Monitoring wounds for infection is critical to the prevention of clinical complications of wound healing. Mirani et al. incorporated pH-sensitive dyes into alginate scaffolds as a way to visualize infection state.⁸³ The inclusion of pH indicator with antibiotic-eluting scaffolds has the potential to provide bimodal diagnostic and therapeutic functionalities for application in infected wounds.

Bioprinted Dermal Scaffolds

Incorporating living cells with 3D printing of scaffolds realizes the promise of reproducibly fabricating living biomimetic human tissues. By accurately controlling the organization and location of specific cell populations and structures to replicate the complexity of mature skin, bioprinting offers a potential solution to the limited pool of donors for clinical-grade STSGs that are needed for the management of FTW. Considering the variety of skin equivalent constructs suitable for clinical applications, we will discuss the commonly used bioink materials and cells, as well as the practicalities of fabrication: tissue constructs fabricated on the bench and allowed to mature before transplantation (in vitro bioprinting), or fabricated in situ, directly at the site of injury.

Bioink materials

Bioinks are the critical enablers of 3D bioprinting, providing tissue structure, organization, and facilitating biological activities. An ideal bioink should possess certain physicochemical, rheological, and biological properties.⁸⁴ Hydrogels based on natural polymers currently represent the most common bioink formulations in skin bioprinting due to their compatibility, hydrophilicity, degradability, and capacity to support cell adhesion. Naturally derived hydrogels such as collagen, gelatin, alginate, and decellularized skin tissue matrix will be discussed together with applications in skin bioprinting in the section below. For comprehensive reviews of bioinks, readers are referred to the excellent articles by Gopinathan and Noh,⁸⁴ Morgan et al.⁸⁵ and Ashammakhi et al.⁸⁶

Collagen is among the most abundant protein in the human body. It consists of three polypeptide chains intertwined in a helical structure. Apart from being the main structural protein present in ECM, collagen possesses tissue-like physicochemical properties, biocompatibility, and acts as a reservoir of growth factors and critical biological response modifiers.⁸⁷ Type I collagen was used to bioprint simple skin substitutes using LaBP in two stages: embedding murine fibroblasts in collagen as a dermal equivalent and overlaid with human keratinocytes as an epidermal equivalent.⁸⁸ This construct recapitulated the two distinct layers of native skin populated by the dominant cell types, keratinocytes and fibroblasts, which upon cultivation, generated epidermis and dermis-like structures. In a follow-up study, the bilayered skin constructs were transplanted onto FTW created on the back of nude mice.⁸⁹ Grafts were analyzed 11 days after implantation and shown to comprise partially differentiated keratinocytes, with new blood vessels growing from the wound bed into the collagen-rich dermis. The constructs integrated and regenerated skin tissue similar to native skin.

Gelatin, the denatured form of collagen, is derived through partial hydrolysis of collagen. Possessing cell-adhesive capabilities, high biocompatibility, biodegradability, and low antigenicity, gelatin is a popular biomaterial with versatile applications in the biomedical field.⁹⁰ When used in bioprinting, gelatin is typically combined with other biomaterials to enhance and optimize mechanical and rheological properties. For instance, human amniotic epithelial cells and mesenchymal stem cells (MSCs) derived from Wharton's jelly remained viable in a skin substitute fabricated using a blend of gelatin-alginate.⁹¹ A second approach used extrusion printing to overlay a silk fibroin-gelatin composite bioink laden with keratinocytes on top of fibroblasts to support the viability and differentiation of epidermal keratinocytes.⁹² This bioprinted skin construct was found to be dimensionally stable and yielded transcriptomic and proteomic profiles similar to profiles determined from native human skin.

Another commonly used natural polymer, alginate, is an anionic polysaccharide derived from brown algae.⁹³ Alginate is an attractive material for biomedical applications having desirable properties of biocompatibility, low cytotoxicity, and ubiquity. Polymerization is easily initiated in the presence of divalent calcium cations [Ca²⁺].^94,47 In contrast to gelatin, which is derived from mammals, alginate does not possess mammalian cell binding motifs. To overcome this absence, other materials (including gelatin) are blended with alginate to provide cell adhesion motifs. To illustrate, Pourchet et al. evaluated a unique bioink formulated with alginate, gelatin, and fibrinogen, for skin bioprinting.⁹⁵ The construct was assembled using sequential printing of human dermal fibroblast (HDF)-containing bioink, overlaid with human epidermal keratinocytes, and subsequently cultivated at the air-liquid interface to initiate maturation. The structure and organization of the mature skin construct were authenticated with histopathology and found to express epidermal and dermal markers characteristic of native human skin.

To more closely recapitulate the native human skin tissue microenvironment, decellularized extracellular matrix (dECM), derived from native donor skin tissue, has been employed as biocompatible scaffolds for skin regeneration. dECM is produced by physicochemical manipulation of (usually healthy) native tissue, intended to devitalize the tissue.^96,97 Through careful selection of the physicochemical method, dECM retains the full molecular inventory of native ECM. Jorgensen et al. added decellularized and solubilized ECM from human skin tissue to increase ECM complexity of their bioprinted fibrin-based constructs.⁹⁸ With dECM supplementation, cell viability of the encapsulated primary fibroblast improved compared to using fibrin-based hydrogels. In addition, the mixture of dECM and fibrinogen bioink demonstrated improved printability and physical microenvironment of the composite construct was found to be similar to human skin.

Due to the low availability of human skin, other investigators have opted to utilize decellularized xenografts, which can be readily sourced in larger quantities, for example, pig skin^99–101 or pig small intestinal submucosa.¹⁰² Won et al. utilized porcine skin-derived dECM to formulate bioinks populated with HDFs.⁹⁹ The bioprinted construct is reported to support >90% HDF viability, and retain the expression of genes associated with skin development. The authors concluded that the porcine-derived dECM contributed essential growth factors and bioactive species needed to support fibroblast survival, proliferation, and metabolic function. Notably, when used to bioprint full-thickness 3D skin tissue constructs, bioinks incorporating porcine dECM exhibited superior mechanical stability to homogenous constructs assembled using type-I collagen bioinks.¹⁰¹ Bioink formulations incorporating porcine dECM successfully supported human stem cell survival, proliferation, differentiation, and vascularization, resulting in enhanced wound healing, reepithelialization, and neovascularization in mice FTW model.

Cell selection

Bioprinting enables the specific delivery of cells to targeted defects where the cells secrete various cytokines, chemokines, and growth factors that can provide therapeutic effects. Autologous, allogeneic, or xenogeneic primary and stem cells are some of the cell types used in 3D bioprinting.¹⁰³ Besides providing structural support, paracrine communication between the dermis and epidermis is critical to maintain homeostasis and respond to vital threats, infection, trauma, and injury.¹⁰⁴ The physiology and activity of skin cells in situ (i.e., intact native 3D tissue) are subject to complex and exquisite regulatory mechanisms at both local tissue levels, and also by systemic (e.g., autonomic) mechanisms. Ex vivo and in vitro tissue models and substitutes have yet to reproduce this level of polyfunctional complexity, ignoring the comprehensive modes of regulation. While many investigators have used immortalized cell lines to emulate mammalian tissues, these systems unfortunately fail to recapitulate the full complement of tissue functionality and regulation evident in vivo.

One strategy to minimize this disconnect from native tissues and ex vivo, or in vitro models, is to utilize freshly isolated, or primary tissues/cells. The rationale is that freshly isolated, or primary tissues/cells retain physiological “memory” and maintain their in vivo physiology and function for at least several population doublings. Bilayered human skin constructs fabricated by Cubo et al. propagated primary human fibroblasts and keratinocytes on hydrogel scaffolds of plasma-fibrin.¹⁰⁵ These constructs remained viable, proliferated, differentiated, and produced well-developed stratum corneum and basale lamina/basement membrane, indicative of autonomous skin tissue repair. Moreover, rete pegs/ridges and neoangiogenesis were evident in these synthetic constructs following in vivo transplantation. In another study performed using collagen as the bioink, primary cell-laden bioprinted scaffolds transplanted into FTW on the dorsum of mice were observed to support effective cutaneous wound healing.¹⁰⁶

Apart from applications in wound healing, bioprinting can incorporate cells such as melanocytes to restore pigmentation and deliver cosmetic benefit. Pigmented full-thickness skin models suitable for grafting can be assembled using autologous melanocytes and keratinocytes and applied using bioprinting. Even in the absence of chemical, or ultraviolet light stimulation, melanogenesis and the transfer of melanosomes to adjacent keratinocytes are reported to occur within in vitro skin constructs.¹⁰⁷ Similarly, Ng et al. successfully bioprinted a 3D skin construct that matched with the native human skin. In contrast to skin constructs fabricated by traditional casting approach, the bioprinted skin exhibited well-stratified epidermal layers with uniform distribution of melanosomes.¹⁰⁸

The intrinsic capacity of stem cells for self-renewal and multilineage differentiation is shifting the paradigm for regenerative medicine and potential applications for bioprinting. Skin stem cells have been isolated from a number of niches, and utilized for clinical applications, including skin repair.^109,110 Koch et al. utilized the laser-induced forward transfer (LIFT) technique to fabricate a 3D skin construct using keratinocytes, fibroblasts, and MSCs, suspended in blood plasma and alginate hydrogel.¹¹¹ All cells within the construct exhibited cell viabilities of >90% and proliferated normally, with minimal evidence of apoptosis (as measured by DNA fragmentation). Cell surface immunophenotyping indicated that the phenotype of LIFT-bioprinted MSCs was unchanged.

Other stem cell populations, such as amniotic fluid-derived stem cells (AFSCs) and bone marrow-derived MSCs, have also been assessed for their wound healing capabilities.¹¹² Deposition of these cells suspended in fibrin-collagen bioinks onto FTW on nude mice demonstrated improved wound closure and reepithelialization, compared to FTW treated with cell-free, bioink-only control interventions. Seeking to prolong paracrine activity and support wound healing, Skardal et al. evaluated a heparin-conjugated hyaluronic acid hydrogel, able to sequester cytokines and growth factors secreted by AFSCs.¹¹³ When this formulation was directly printed onto FTW created in mice, samples loaded with AFSCs were found to reepithelialize, vascularize, and heal before FTW treated with unloaded control formulations. The authors interpreted this to indicate the hydrogel facilitated the sustained release of AFSC-secreted cytokines and bioactives.

Bioprinting approaches

The application of bioprinting to fabricate scaffolds and living tissue substitutes for the treatment of FTW may be broadly divided into two main approaches: in vitro fabrication and in situ grafting. In vitro bioprinting is an approach where skin constructs are fabricated in vitro using appropriate cell-laden bioinks, cultivated and permitted to mature before transplantation/grafting onto the wound site (Fig. 3A).^114,115 This approach requires the wound bed to have an adequate blood supply to ensure the survival of the in vitro fabricated graft. This unmet need has driven considerable research of in vitro bioprinting thus far, to develop vascularized dermal constructs. To this end, Baltazar et al. incorporated human endothelial cells derived from cord blood human endothelial colony-forming cells and human placental pericytes into conventional skin constructs fabricated with HDFs and keratinocytes.¹¹⁶ They observed host microvasculature invasion and the formation of epidermal rete pegs/ridges in wounds on the dorsum of immunodeficient mice, 2–4 weeks post-transplantation. Similarly, bioprinted skin constructs containing human dermal microvascular endothelial cells demonstrated neovascularization in an FTW model in vivo.¹¹⁷ Accelerated healing and reduced wound contraction were noted in mice treated with bioprinted skin constructs, compared to mice with wounds treated with commercial dressings. Comparable neovascularization results were also observed by Zhou et al., who applied scaffolds containing human umbilical vein endothelial cells (HUVECs) and human skin fibroblasts to both mice and porcine FTW.¹¹⁸ The contribution of delivered cells to aid in dermal regeneration was further validated in the porcine model as wounds recovered with more mature collagen organization and human nuclei were found incorporated into new hair follicles and blood vessels. An alternative strategy was adopted by Jorgensen et al., reporting their attempt to recapitulate collagen remodeling in wounds by using a trilayer bioprinted skin structure of epidermis, dermis, and hypodermis fabricated using distinct cell populations suspended in a fibrinogen composite bioink.¹¹⁹ Keratinocytes and melanocytes formed the epidermis and fibroblasts, follicular and papillary dermal fibroblast and microvascular endothelial cells formed the dermis, while preadipocytes formed the hypodermis. This full-thickness construct exhibited superior wound healing through rapid reepithelialization and reduced wound contraction when compared to untreated wounds and wounds treated with hydrogels.

Another approach gaining traction for the clinical management of wound healing is bioprinting in situ. In this approach, cell-laden bioinks are applied directly to the wound, usually after debridement, using a portable hand-held, or robotic, automated printing device. This strategy offers several inherent advantages over in vitro bioprinting, including controlled spatial deposition of various cell types on site and the bypass of laborious in vitro differentiation steps. When applied to large surface area wounds, such as burns, it can potentially initiate early reepithelialization and reduce time to wound closure. It can also be coupled with a hand-held laser scanner to acquire precise spatial information regarding the wound topology for defined cell delivery. For instance, Albanna et al. bioprinted fibrin/type I collagen bioinks containing allogeneic/autologous skin cells directly onto FTW using a novel mobile skin bioprinting system with scanning technology (Fig. 3B).¹²⁰ In both murine and porcine models, the cell-treated wounds closed faster and with minimal wound contraction compared to the control animals. The presence of proliferating keratinocytes, highly organized dermis, and mature microcapillary bed in the regenerated tissue closely resembles native skin. To eliminate the time-consuming wound scans and CAD modeling, Hakimi et al. developed a hand-held bioprinter suitable for direct printing of biomaterials onto excisional wound models in situ (Fig. 3C).¹²¹ A hemostatic bioink (cell free) printed directly onto a porcine FTW stopped bleeding within 5 min of application, while the untreated wound required tens of minutes to achieve hemostasis. More than half of the treated group progressed to complete reepithelialization (wound closure) compared to the control group. This technology has obvious applications for acute trauma and emergency uses. Stem cells delivered through in situ methods were also evaluated for their wound healing capabilities. A hand-held bioprinter was utilized by Cheng et al. to bioprint MSC-laden fibrin-hyaluronic acid sheets on large full-thickness burn wounds.¹²² Macroscopic visualization and histological analysis revealed that the MSC-treated wounds provided an excellent healing profile with reduced scarring and inflammation, contraction, and rapid skin restoration with a physiologically thick epidermis and dense collagen.

Future Directions for 3D Printed Scaffold Designs

Modifications to existing materials, new materials, and combinations of materials and bioadditives are constantly evaluated to address currently unmet physical and biochemical complexities of wound healing. In the studies discussed above, we have seen evidence that acellular and cellular scaffolds are able to aid healing of FTW, through the use of biomimetic skin constructs, bioactive components, and representative cell populations. Parallel technological improvements in bioprinter design also pave the way for point-of-care manufacturing such that printed materials may be delivered directly for rapid personalized treatment of cutaneous wounds. Table 2 summarizes the studies discussed, in particular, studies with in vivo validation of wound healing outcomes. Moving forward, physical and biochemical complexity in 3D printing will be continuously exploited to include new materials and smart designs to create a more personalized and dynamic wound healing environment.

Smart materials

To capture the dynamic nature of skin, novel, smart materials are being developed, creating new classes of printable “intelligent” materials; this is termed “4D printing.”¹²³ Scaffolds for 4D printing should possess one or more of the following functions: (1) respond to an external stimulus through change of shape, self-assembly, or self-actualization; (2) detect and possibly quantify the external stimulus; and (3) repair damage to its own structure without any additional material (self-heal).^124,125

Self-healing scaffolds for wound healing were recently explored by Zhao et al. using a diabetic rat model.¹²⁶ Printed hydrogels formulated with polyaniline and cross-linked with polyvinyl alcohol were loaded with AgNPs to provide antibacterial activity. The scaffolds' ability to self-heal was demonstrated when severed scaffolds were able to reconnect. Self-healing scaffolds should ensure the physical barrier over wounds remains intact and, ideally, help to extend the scaffold's lifetime postgrafting. Notably, these “smart” scaffolds were also observed to attenuate inflammation and necrosis in infected FTW, allowing the wounds to progress to the proliferative stage earlier than control untreated FTW.

Self-healing scaffolds with stimuli-responsive behavior were investigated by Deng et al.¹²⁷ Polymer blends of nanoclay, N-isopropyl-acrylamide (NIPAM), and carbon nanotubes displayed self-healing ability and initiated deswelling responses when exposed to heat from thermal and photothermal sources. In the context of wound healing, swelling/deswelling behavior allows the scaffold to initiate automated release of bioactive compounds in response to internal stimuli from the wound environment, or on-demand release in response to external stimuli.¹²⁸ The printability and cytocompatibility of NIPAM were further investigated by Zhang et al.¹²⁹ A cross-linked network of NIPAM, acrylic acid, and fibrin, loaded with HUVECs, fibroblasts, and keratinocytes was printed in layers onto polyvinyl alcohol scaffolds. When these living constructs were lifted to the air-liquid interface and matured for 2 weeks, the multilayered construct retained tissue organization, high cell viability, evidence for squamous differentiation, and cornification, as well as spreading of subcutaneous endothelial cells.

Complex scaffold design

We predict that future developments in 3D printing will translate into increasingly complex designs providing improved temporal control over scaffold behavior. On-demand and temporal delivery of embedded bioactive compounds can prevent overexposure to growth factors or antibiotics, thus creating a more physiologically relevant wound healing environment. Parallel progress in design and development in the field of sensors and bioelectronics add exciting possibilities of integrating real-time monitoring into tissue constructs, providing clinically relevant, wound-specific data to support more effective management of skin health.

Complex scaffold designs integrating therapies with diagnostic capabilities in the same platform are an emerging strategy that could be a game changer. Khademhosseini et al. have screen-printed a pH sensor and microheater on a 3D printed flexible thermoresponsive poly-NIPAM substrate, to monitor changes in pH associated with wound infections and release antibiotics when required.^130,131 Adopting a similar strategy, Pang et al. have designed an on-demand drug-release platform to respond to changes in wound temperature, a classical clinical indicator of infection.¹³² The concept was validated in a porcine FTW model where embedded sensors and wireless communication remotely triggered the release of antibiotics into the wound bed. Although visual methods for monitoring wound status have been explored,^54,83 real-time monitoring increases the sensitivity and accuracy of collected data. With continued advancement in printable conductive materials^126,127 and circuitry designs,¹³³ fully or partially 3D printed multicomponent theranostic platforms promise a “new epoch” in the clinical management of wound healing.

Complex scaffolds can be designed without bioelectronics to also deliver distinct compounds at specified junctures during wound healing. Since wound healing is a complex process, temporal changes (upregulation or downregulation) of cytokines, chemokines, growth factors, and noncoding RNAs can have significant effects on wound healing outcomes.^1,71,134 Taking advantage of unique degradation rates and spatial deposition of materials, composite scaffolds can be designed to achieve sequential release of bioactives targeting specific wound healing events.¹³⁵ Spatiotemporal control in such scaffolds allows for sophisticated preprogrammed drug or growth factor release. However, the manufacture of multilayered and multicompound encapsulating scaffolds can be tedious. In a separate proof-of-concept study, spatial deposition of different materials was easily achieved with 3D printing and different fluorophore dyes, which represented model drugs, were released in a time-dependent manner.¹³⁶ This has also been demonstrated in the pharmaceutical field, where 3D printing is used to construct multicompartment tablets and devices for controlled drug release with complex dosing regimen.^137,138

Personalization

Three-dimensional printing and bioprinting have enabled the design and fabrication of custom (personalized) scaffolds. Patient-/wound-specific scanning and spatial control over material deposition allow for both in vitro⁶⁷ and in situ¹²⁰ construction of customized scaffolds that conform with the unique topology of an individual wound. At the same time, bioactive additives can be adjusted according to clinical diagnosis, wound type, comorbidities, and presence of infection(s).^68,69

Further customization is possible, allowing regenerated skin to match patient's skin tone, hair density, and sweat gland patterns through tailored incorporation of different cell populations. Pigmentation similar to native human skin is currently achieved by bioprinting of human fibroblast, keratinocytes, and melanocytes.^107,108 Functional sweat glands have been successfully differentiated in vitro from bioprinted MSCs, but this has only been recapitulated in mice wound models.¹³⁹ Although Abaci et al. successfully differentiated human hair follicles from dermal papilla cells in structure-specific collagen microwells, they highlighted that the process could be improved with bioprinting.¹⁴⁰ Therefore, continued investigation into bioprinted human skin appendages is necessary to achieve skin regeneration. Personalized regenerated skin can help to complement the original esthetics of patient's skin, for example, higher hair density in scalp skin or increased pigmentation for darker skin. More importantly, it allows for recovery of important functions specific to area of skin loss, for example, thermoregulatory function by sweat glands in palm of hands and soles of feet.

Regulation

Before we see a new generation of 3D printed skin substitutes being translated to the clinical and commercial setting, gaps in the current regulatory framework need to be addressed. Over the past decade, FDA's Center for Device and Radiological Health has reviewed and approved several 3D printed medical devices for clinical use,²⁰ while in 2015, the FDA's Center for Drug Evaluation and Research approved the 3D printed drug SPRITAM for treatment of epilepsy.¹⁹ Although the FDA's Center for Biologics Evaluation and Research has been exposed to various projects within the bioprinting field, there are currently no approved 3D printed tissue substitutes due to the investigative nature and regulatory consideration of such products.¹⁹

Part of these regulatory considerations includes safety and efficacy. As with any other commercially available wound healing products, validation of 3D printed or bioprinted scaffolds for FTW healing application should meet recommended biocompatibility evaluation endpoints such as testing for cytotoxicity, sensitization, and intracutaneous reactivity.¹⁴¹ Scaffold characteristics such as in vivo degradation should be thoroughly assessed. Although degradation products of synthetic polymers like PLA and PLGA have been determined to be noncytotoxic, crystalline particles of PLA L-enantiomer can invoke inflammatory responses¹⁴²; and in cases where scaffolds degrade too rapidly, acid build-up of degradation products can lead to undesirable pH changes in local tissue environment.¹⁴³

To determine biocompatibility and efficacy accurately, appropriate preclinical models that reflect the immunity and complexity of human FTW are essential. While it may not be feasible to replicate the healing environment of large burn wounds or nonhealing wounds perfectly,¹⁴⁴ investigators should be aware of the limitations of animal models used and work toward circumventing them, for example, using splinted wound models in mice to prevent healing through muscle contraction, testing with porcine wound models as their skin physiology and healing mechanism is similar to humans, or using immunocompetent animals to investigate for scaffold biocompatbility.¹⁴⁵ Through careful interpretation of efficacy results, we can then distinguish those investigative products ready for clinical validation.

Recognizing that 3D printing is becoming an increasingly common manufacturing technique for medical products, the FDA recently released (2017) documented guidelines for additive manufacturing of medical devices.¹⁴⁶ However, these technical guidelines do not address “biological, cellular or tissue-based products manufactured using additive manufacturing technology.” Bioprinted wound healing scaffolds are inherently different from other skin substitutes or printed implants on the market. Therefore, it is unclear if 3D printed tissue substitutes should be regulated as an advanced drug delivery method, medical device, biological cellular therapy, or combination product.¹⁴⁷

Although complexity is a distinct advantage in 3D printed scaffolds, it may also contribute to regulatory complications. Different combinations of bioprinter, printing technique, and raw material can lead to subtle yet important variations in the final product. In addition, personalization of the wound healing scaffold may create differences in the designing and manufacturing process, making it even more challenging to standardize a defined level of efficacy.¹⁴⁸ Compared to traditionally manufactured products, 3D printed scaffolds may require additional oversight to ensure quality control is maintained at each processing phase.¹⁴⁷

Moving forward, regulatory frameworks are essential to advise how current laws and product classifications apply to 3D printed tissue constructs.¹⁴⁸ Encouragingly, the FDA and several international authorities, including Singapore's Health Sciences Authority, recognize such regulatory considerations and are continuously engaged in dialog with stakeholders to enable safe and effective translation of 3D bioprinting into wider application addressing unmet clinical needs.^19,149

Conclusion

The relentless innovation drive behind printing techniques, biomaterial formulations, scaffold designs, and skin research beckons an upward trend of advancement in 3D printed wound healing scaffolds. However, for the existing and future innovations to achieve clinical impact, printed scaffolds must not merely fulfill the basic requirement of supporting wound closure, but it must aspire to enable scarless and functional skin repair. To establish the significance of novel constructs, continued rigorous in vitro testing and in vivo validation are required. Last, but not least, regulatory procedures need to be in place to guide the translation into clinically relevant products.

Disclosure Statement

No competing financial interests exist.

Funding Information

This research is supported by the Agency for Science, Technology and Research (A*STAR) REI2020 Advanced Manufacturing and Engineering (AME) Programmatic grant (A18A8b0059) Additive Manufacturing for Biological Materials (AMBM) project SP1.4.