L. Zhang, H. Ren, L. Wu, Z. Liu, A. Xie, X. Yao, J. Ju and M. Liu, J. Mater. Chem. A, , 12, DOI: 10./D4TAE
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The skin plays an important role in vitamin D synthesis, humoral balance, temperature regulation, and waste excretion. Due to the complexity of the skin, fluids loss, bacterial infection, and other life-threatening secondary complications caused by skin defects often lead to the damage of skin functions. 3D bioprinting technology, as a customized and precise biomanufacturing platform, can manufacture dressings and tissue engineering scaffolds that accurately simulate tissue structure, which is more conducive to wound healing. In recent years, with the development of emerging technologies, an increasing number of 3D-bioprinted wound dressings and skin tissue engineering scaffolds with multiple functions, such as antibacterial, antiinflammatory, antioxidant, hemostatic, and antitumor properties, have significantly improved wound healing and skin treatment. In this article, we review the process of wound healing and summarize the classification of 3D bioprinting technology. Following this, we shift our focus on the functional materials for wound dressing and skin tissue engineering, and also highlight the research progress and development direction of 3D-bioprinted multifunctional wound healing materials.
Keywords: Functional materials, Wound healing, 3D bioprinting, Dressing, Skin tissue engineering
166The skin is the largest integumentary organ in the human body, accounting for approximately 15% of the body weight[1]. The skin has a multilayered structure, divided into epidermis, dermis, and subcutaneous layer, which is an important barrier for167 the body to resist various damages from the external environment, such as mechanical interference, microbial invasion and ultraviolet (UV) radiation[2,3]. In addition, the skin is equipped with has basic functions such as thermoregulation, humoral balance, sensory detection, and immunological surveillance[4]. However, fragile skin is susceptible to external extremes or injuries, which usually lead to skin defects, functional impairment, fluid loss, and bacterial infections. Wound healing is a medical problem on a global scale, placing an enormous burden on human health and global healthcare system[5]. Statistics have shown that the global medical cost caused by incomplete chronic wounds in 6 million patients is as high as $20 billion[6-8]. Moreover, countless wounds cannot heal naturally due to the progressive degeneration and necrosis of tissue cells in the wounds of extensive injury and ulceration[9]. Therefore, establishing a reliable, safe, and simple treatment is an urgent problem to be solved[10].
The wound healing process is complex and dynamic, mainly including hemostasis, inflammation, cell proliferation, and maturation[11,12]. At present, for wounds that cannot be healed by the human body, such as large-area trauma and burns, traditional methods such as autograft[13], allograft[14], cell therapy[15], and skin substitute[16] are usually used to treat such wounds in clinic. However, these traditional approaches are often limited by insufficient donors, small scope of repair, immune rejection, and high costs[17,18]. Therefore, a large number of wound dressings and skin tissue engineering scaffolds have been developed to provide artificial substrates for wound repair and tissue regeneration[19]. In the process of wound repair, traditional dressings and skin tissue engineering scaffolds usually have problems such as inability to stop bleeding, susceptibility to wound infection and inflammation, and difficulty in achieving vascularization[20]. Among them, infection is the main obstacle in the wound healing process, which can cause the elevation of reactive oxygen species (ROS) and protease levels in the wound, excessive inflammation and other problems, and ultimately leads to incomplete wound repair and prolonged repair time[21]. At the same time, hemostasis is also particularly important for wound healing, which is related to the patient's life and subsequent wound healing[22]. Thus, functional materials with some or more characteristics have great potential in wound healing treatment. For example, the antibacterial materials (such as silver, zinc oxide, and chitosan) can inhibit or kill bacteria in wounds through multiple mechanisms: the anti-inflammatory materials (paeoniflorin, apigenin, and luteolin) can inhibit the production or release of anti-inflammatory factors to fight inflammation, and the hemostatic materials (such as chitosan, montmorillonite, and kaolin) can control wound bleeding either actively or passively; these properties are very important for wound healing[23,24]. Therefore, the addition of functional materials can endow wound healing materials with a variety of properties, thus promoting rapid and effective wound healing. In addition, in order to provide personalized treatment for different wound types, three-dimensional (3D) bioprinting technology, which has prominent advantages in wound healing and tissue regeneration, has been introduced into wound management[5].
3D bioprinting is an important branch of 3D printing technology applied to life science and medicine. 3D printing is a rapid prototyping technology that constructs 3D geometric shapes through computer-aided design and layer-by-layer deposition of materials. 3D bioprinting is based on the principle of additive manufacturing to accurately deposit bioinks containing biomaterials, growth factors and even living cells in a controllable space to create complex tissue structures to simulate natural tissues or organs[25-29]. In skin repair, this technology can precisely match the geometric shape of wound healing materials and tissue defects, so as to achieve rapid and effective wound healing[30]. So far, the combination of 3D bioprinting technology and a variety of functional materials can produce the reproducible and personalized 3D constructions with multiple functions, such as antibacterial, anti-inflammatory, antioxidant, hemostasis, and antitumor properties. Liu et al.[31] fabricated a Gel/PCL/ PDA cores/shell fiber scaffold for controlled anticancer drug release by depositing polydopamine (PDA) and polycaprolactone (PCL) on the surface of 3D-printed drug-loaded alginate-gelatin hydrogel scaffolds. The scaffold can be implanted at the resection site of patients with malignant tumors for local cancer treatment through drug release (doxorubicin) and photothermal therapy. In addition, it can repair surgically resected defect tissue and promote wound repair.
In this article, we describe the principles, advantages, and disadvantages of different 3D bioprinting technologies, and review the fundamentals of the wound healing process. In addition, we focus on the classification and characteristics of different functional materials, as well as the important application of 3D-bioprinted functional materials for wound healing, aiming to provide new ideas and useful references for the preparation and further development of multifunctional wound healing materials using 3D bioprinting technologies in the future.
The most significant organ of our body, the skin, has numerous important functions such as secretion, regulation, and protection[32]. However, the structure and168 function of this organ are susceptible to burns, cuts, surgical incisions or illnesses, such as diabetes[33]. Skin injuries caused by physiopathologic, physical, and chemical factors usually trigger complex, highly integrated and overlapping self-healing process, involving hemostasis, inflammation, migration, proliferation, and tissue remodeling[34]. Immediately after an injury, the hemostatic response begins and blood vessels temporarily constrict for 5'10 minutes, helping to slow down the blood flow. During hemostasis, platelets aggregate at the site of injury, while fibrin forms a clot to prevent blood loss and microbial contamination[35] (Figure 1a).
The complete process of wound healing. Stages of wound healing include (a) hemostasis stage, (b) inflammatory stage, (c) proliferative stage, and (d) tissue remodeling stage.
The inflammatory and hemostatic phases occurred almost simultaneously. At this stage, under the complex interaction of cytokines, inflammatory cells such as neutrophils and monocytes recruited to the wound site differentiate into macrophages and produce ROS and proteases to destroy and remove foreign particles, bacteria, and tissue debris at the wound site[36,37]. In addition, macrophages release various growth factors and cytokines to induce proliferation and migration of fibroblasts. The inflammatory phase usually lasts for 2'5 days[35,38] (Figure 1b).
The proliferative phase generally begins around the third day after injury and will last about 2'4 weeks. This stage mainly includes granulation formation, epithelialization, and angiogenesis. Fibroblasts migrate from the surrounding tissue to the wound site to produce extracellular matrix (ECM) components such as collagen and proteoglycans, thereby forming pale pink granulation tissue[39]. Granulation tissue provides a matrix for epithelial cells to cover the wound during epithelialization, and re-epithelialization is completed when the epithelial cells have completely filled the defect wound. On the other hand, endothelial cells in the blood vessel wall promote the formation of new blood vessels, and also create new capillaries in the existing blood vessels. In addition, fibroblasts differentiate into myofibroblasts, which contract and close the wound[2,40] (Figure 1c).
Remodeling is the final and most clinically important stage of wound healing. The remodeling phase begins in the third week after the injury and may last from 1 to 3 years. During this stage, inflammatory cells, fibroblasts, and endothelial cells migrate from the wound site or die. Various growth factors induce collagen deposition and orderly arrangement, thereby enhancing the strength of new tissue. The ECM gradually transforms into scar tissue or functional skin[41] (Figure 1d). During the various stages of wound repair, the interference of any factor (such as wound infection, oxidant, and excessive inflammation) may affect the successive stages of wound repair, which169 may result in the formation of chronic wounds[42]. Therefore, application of functional materials for external interventions on these adverse factors in the process of wound repair is critical to avoid the occurrence of chronic ulcers and wound healing[38,43].
In recent decades, the materials used for wound healing are divided into naturally derived materials and synthetic materials[44]. Natural materials mainly include collagen, chitosan, fibrin, hyaluronic acid, gelatin, and sodium alginate. Synthetic materials include poly(lactic-co-glycolic acid) (PLGA), poly(ether-ether-ketone) (PEEK), poly(lactic acid) (PLA), polycaprolactone (PCL), and poly(glycolic acid) (PGA)[44]. So far, researchers have developed a large number of wound dressings and tissue-engineered skin substitutes based on the above materials. However, there is still much more room to improve the materials for skin wound healing[45]. The existing wound healing materials and various functional materials can be combined according to the depth, scope, and pathological state of different types of skin wounds, thereby meeting the ever-evolving needs of patients[46]. Therefore, a large number of functional materials are used as modern wound dressings and skin tissue scaffolds, including antibacterial materials, antiinflammatory material, conductive material, antioxidant, hemostatic materials, flexible material, and antitumor material. Representatives of various functional materials and their related mechanisms are shown in Figure 2 and Table 1.
Classification and representative materials of functional materials. Abbreviations: CS, chitosan; PANI, polyaniline; PPy, polypyrrole; CAT, catalase; PLA, poly(lactic acid); MMT, montmorillonite; SA, sodium alginate; PDMS, polydimethylsiloxane; SE, silicone elastomers; PTX, paclitaxel; Cur, curcumin.
Impose long-term interference to bacterial growth through electrostatic interaction
Increase ROS level to destroy DNA, RNA, polysaccharides, lipids and proteins of bacterial cells
Promote the transformation of macrophages from a pro-inflammatory M1 phenotype to a pro-healing and anti-inflammatory M2 phenotype at the wound site
Inhibit cyclooxygenase to block the synthesis of inflammatory mediators such as prostaglandins and thromboxane, thereby exerting anti-inflammatory effects
Scavenge ROS that plays an eventful role in the inflammatory process
Provide electrical stimulation at the wound site by increasing electrical conductivity to activate ion channels and transduce signals downstream to guide the migration and proliferation of skin cells
Scavenge the free radicals (ROS) and inhibit the generation of ROS, and block free radical chain transfer
Activate the enzymatic antioxidant system in the body and stimulate the formation of non-enzymatic antioxidants in the body[54]
Activate platelets and promote red blood cell aggregation to rapidly form blood clots
Bind to plasma and activate coagulation factors of the internal coagulation cascade
Swell after absorbing fluid to form a physical barrier, causing blood to aggregate and coagulate[57]
Insert between polymer molecular chains, weaken the inter-molecular chain stress and increase its mobility
Interfering with DNA, RNA or protein synthesis
Generate ROS
Downregulate migration and proliferation of cancer cells by regulating several signaling pathways
Antibacterial materials are able to inhibit or kill bacteria (or fungi). Antibacterial materials can be divided into four categories according to their chemical structure and composition: (i) inorganic compounds (silver ions and copper ions)[64]; (ii) organic compounds (guanidine salt, quaternary ammonium salt, and quaternary phosphorus salt)[65]; (iii) natural antibacterial agents (antibacterial peptides, chitin, and chitosan)[64]; (iv) composite antibacterial agents (inorganic/organic, inorganic/ inorganic, organic/organic, and composite materials)[66].
170The mechanism of antibacterial materials inhibiting or killing bacteria (or fungi) includes various aspects. For example, chitosan, quaternary ammonium salt, metal cations, and metal oxide nanoparticles can interact with the bacterial membranes directly, and the positively charged antibacterial agents are adsorbed and permanently retained on the negatively charged bacterial membrane through electrostatic interaction, thereby causing long-term interference with bacterial growth via preventing glucose metabolism, cellular respiration, and oxygen uptake[67]. It has been reported that chitosan'fibrin composite (CF) scaffolds impregnated with quercetin (Q-CF) as wound dressing exhibited good bactericidal performance against Escherichia coli and Staphylococcus aureus. At the same time, the wound healing experiment in albino rats in vivo showed that Q-CF scaffold could accelerate wound healing.
In addition, antibacterial agents such as metal cations and metal oxide nanoparticles can also damage bacterial DNA, RNA, polysaccharides, lipids, and proteins by increasing ROS levels to achieve bacterial killing action[48]. Recently, Guo et al.[68] added PDA as an antibacterial component to a matrix of magnesium ions (Mg2+) and polyacrylamide (PAM) to prepare an excellent composite antibacterial hydrogel PDA-PAM/Mg2+. This composite hydrogel exhibited excellent tissue adhesion and synergistic photothermal antibacterial activity, and was effective against S. aureus and E. coli after near-infrared (NIR) light irradiation. A wound infection rat model revealed that PDA-PAM/Mg2+ hydrogel wound dressing could promote collagen deposition and tissue regeneration, which could accelerate wound healing.
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Inflammation is the body's immunobiological response to infection. Inflammation can be chronic or acute, longer or shorter in duration, and the main symptoms are heat, redness, pain, swelling, and even loss of function[69].171 Inflammatory response usually exists in the process of wound healing, and persistent inflammatory response is one of the major reasons for delayed wound healing. Anti-inflammatory materials can inhibit the production or release of anti-inflammatory factors, thereby promoting wound healing process.
Currently, a variety of anti-inflammatory materials have been employed to counteract inflammation; for example, paeoniflorin and PDA can inhibit inflammation by promoting the transformation of macrophages at the wound site from the pro-inflammatory M1 phenotype to the pro-healing and anti-inflammatory M2 phenotype[49]. Aspirin, ibuprofen, and asiaticoside block the synthesis of inflammatory mediators (such as prostaglandins and thromboxane) by inhibiting cyclooxygenase (COX), thereby exerting anti-inflammatory effects[70]. It has been reported that asiaticoside (AS) not only has antiinflammatory activity but also has favorable effects on fibroblast proliferation and collagen synthesis[71]. Seon et al.[50] used AS to prepare a collagen-AS/εPLL doublelayer scaffold, in which the upper layer was loaded with εPLL with antibacterial effect, and the lower layer was composed of collagen with AS nanofibers. This scaffold exhibits anti-inflammatory and bactericidal effects by adjusting the TLR4/MAPK/NF-kB signaling pathway. Furthermore, a Sprague Dawley (SD) rat model of fullthickness inflammation demonstrated that the collagen-AS/εPLL scaffold could accelerate inflamed full-thickness wound closure and re-epithelialization to promote wound repair. Therefore, the collagen-AS/εPLL bilayer scaffolds have great application potential in the field of tissue engineering. In addition, a study has shown that cerium oxide nanoparticles can eliminate ROS, which plays an important role in the inflammatory process, to achieve anti-inflammatory effect[51].
Conductive materials refer to carbon nanomaterials, conductive polymers, and metal nanoparticles with electrical conductivity and electrical conductivity above 10'6 S/m. Conductive polymer materials such as polyaniline (PANI), silver nanowires (AgNW), graphene oxide (GO), polypyrrole (PPy), polythiophene (PTh), and their derivatives (mainly aniline oligomers and poly(3,4-ethylenedioxythiophene) [PEDOT]) have been widely used in biomedical fields such as flexible sensors, health monitoring, wearable devices, drug delivery systems, and tissue engineering[37]. Studies have confirmed the role of conductive materials in skin repair. Liang et al.[52] developed an injectable antibacterial conductive QCSG/ GM/GO hydrogel by using the conductive material GO functionalized with glycidyl methacrylate-modified quaternary ammonium chitosan (QCSG) and crosslinked gelatin methacrylate (GM). In addition to good antibacterial properties in vivo/in vitro, the full-thickness defect repair model of mice infected with methicillin-resistant S. aureus (MRSA) has proven that the conductive hydrogel can promote wound healing in the repair of infectious skin tissue.
Moreover, in wound care and tissue engineering, conductive polymer materials provide electrical stimulation to activate ion channels by increasing the conductivity of the wound site, thereby transmitting downstream signals that guide the proliferation and migration of skin cells, such as keratinocytes and fibroblasts[72,73]. Zhou et al.[74] developed a kind of conductive multifunctional PGFP scaffold cross-linked by branched polypyrrole@ polydopamine (PPy@PDA) nanoparticles, aldehyde F127, and poly(glycerol-amino acid) (PGA) (F127-Phe-CHO). PPy@PDA endowed the PGFP scaffold with skin adhesion behavior, controllable electrical conductivity, and photothermochemical tumor therapy. In addition, a fullthickness MRSA-infected wound model showed that this PGFP scaffold could promote collagen deposition, vascular endothelial differentiation, granulation tissue formation, and accelerate skin regeneration. This multifunctional scaffold has great potential in multimodal therapy of tumor/infection-damaged skin.
Based on the definition, oxidative stress represents a disproportion between the production and scavenging of ROS. ROS act as signaling mediators, which are involved in the regulation of growth, differentiation, proliferation, autophagy, and apoptosis of many cells. During wound repair, the controlled level of ROS can moderate the oxidative damage in the wound site and promote epithelial cell proliferation (proliferative phase), angiogenesis, and tissue repair[75,76]. An overproduction of ROS will disrupt the redox balance of cells, leading to a cascade of inflammatory responses that increase tissue damage and hinder wound healing[77,78]. Antioxidants can convert ROS into more stable molecules, such as water and oxygen, through complex catalysis; this explains why antioxidants are also known as reactive oxygen scavenger[79]. To date, a number of antioxidants have been used to manage ROS levels.
Antioxidants are mainly divided into enzymatic antioxidants (low molecular compounds, endogenous molecules, including catalase, superoxide dismutase, and glutathione peroxidase) and nonenzymatic antioxidants (with many exogenous and endogenous molecules, such as PDA, curcumin, polyphenols, and flavonoids)[80]. In the process of skin tissue repair, these antioxidant materials can accelerate wound healing by controlling172 oxidative stress, enhancing the effect of growth factors, and improving the wound microenvironment. Therefore, some researchers combined antioxidants with other materials to treat wound healing[56]. Tang et al.[55] prepared a pGO-CS/SF scaffold composed of chitosan (CS) and silk fibroin (SF) combined with PDA-reduced GO (pGO) with good electroactivity and antioxidant properties. pGO endowed the pGO-CS/SF scaffold with multiple functions. Due to the presence of reducing catechol groups on pGO, the scaffolds could scavenge ROS to reduce cellular oxidation. Moreover, the pGO-CS/SF scaffolds had good electrical conductivity, which could regulate cell behaviors. The fullthickness skin repair model in rats showed that pGO-CS/ SF scaffolds could accelerate tissue regeneration. Therefore, the results suggested that the pGO-CS/SF scaffold might be a promising wound dressing.
The first stage of the wound healing process is hemostasis, and effective hemostasis is very important for subsequent wound healing[81-83]. Since the inherent hemostatic mechanism cannot effectively control bleeding, the timely use of hemostatic materials can reduce morbidity and mortality[22,24]. Therefore, the development of materials with excellent hemostatic activity is of great interest for controlling hemostasis and preventing of blood loss in the early stage of emergency trauma. So far, various materials used as hemostatic agents can be classified into natural hemostatic materials (fibrin, gelatin, chitosan, and sodium alginate), inorganic hemostatic materials (zeolite, montmorillonite, and kaolin), and synthetic hemostatic materials (cyanoacrylate, acrylic, and polylactic acid)[84,85]. An ideal hemostatic material should be biodegradable, biocompatible, and low-cost, as well as can achieve rapid hemostasis within 2 min[86].
The hemostatic mechanism of hemostatic materials is usually divided into active and passive pathways. The active pathway is to initiate the blood coagulation process by specifically initiating the coagulation cascade. For example, chitin and chitosan can promote the aggregation of red blood cells and rapidly form blood clots by activating platelets[58]. Kaolin can combine with plasma and activate coagulation factors of internal coagulation cascade to promote hemostasis[59]. The passive approach requires specific surface properties (antithrombotic, anti-infective, and hemocompatibility) of hemostatic materials to achieve hemostasis[24,57]. Resistant starch and cellulose can quickly form a physical barrier through rapid water absorption and expansion, leading to blood aggregation and coagulation[60]. The development of composite hemostatic materials can improve the hemostatic efficiency and shorten the hemostatic time through a variety of hemostatic mechanisms. In recent years, composite hemostatic materials have been developed to improve hemostasis efficiency and reduce hemostasis time. Recently, Zheng et al.[87] developed a novel W-8HAP-2PVA hemostatic aerogel based on ultralong hydroxylapatite (HAP) nanowires that could release Ca2+ to trigger the coagulation cascade and promote platelet adhesion. The porous structure of this aerogel could aggregate platelets and blood cells by rapidly absorbing water, further promoting thrombosis and accelerating hemostasis. In addition, this aerogel could accelerate the healing of diabetic mouse wound healing model. These results demonstrated that the W-8HAP-2PVA aerogel was an excellent hemostatic material for future clinical and emergency applications.
Flexible materials generally refer to polymer materials that have certain flexibility, stretch, bend, twist, and deform without losing performance[88]. Common flexible materials include silicone elastomers, polycaprolactone (PCL), poly(lactide-glycolide) (PLGA), polydimethylsiloxane (PDMS), polyester (PET), polyimide (PI), polyethylene naphthalene glycol (PEN), and the flexible component material PLA, which are commonly used in flexible electronics[89], soft robotics[90], and biomedical engineering[91]. The application of flexible materials in biomedical engineering is usually to integrate various electronic components on flexible substrates to form flexible circuit boards with high flexibility and elasticity like skin.
In wound healing research, bioengineered materials with high mechanical properties are widely used, while flexible products are relatively rare; however, the stiffness of materials may have an impact on cell behavior[92,93]. Flexible materials can be inserted between polymer molecular chains to weaken the stress between molecular chains and increase their fluidity, thus giving wound healing materials similar softness to natural skin, which is conducive to rapid tissue repair. Gao et al.[61] reported the preparation of flexible bilayer poly(lactide-glycolide) (PLGA) skin scaffolds using a solvent exchange deposition model (SEDM) phase separation combined with a rapid in situ formation system of electrospinning technology. The addition of this flexible biodegradable polyester makes the scaffold flexible, which is conducive to cell growth, and effectively promotes wound healing in rats.
Antitumor materials are indispensable materials for suppressing residual or recurring cancer cells in patients with malignant tumors whose tissues are surgically removed. It is divided into natural drugs (anthocyanin and curcumin) and chemically synthesized drugs173 (quinoline derivatives, doxorubicin, and paclitaxel)[94,95]. The mechanism of antitumor materials killing or inhibiting cancer cells can be divided into three aspects. Chemotherapy drugs such as doxorubicin and paclitaxel inhibit tumor growth by interfering with DNA, RNA, or protein synthesis of tumor cells. Some photosensitizers such as indocyanine green and berberine can induce apoptosis of cancer cells by producing ROS or singlet oxygen[32]. In addition, polyphenols, such as anthocyanin, curcumin, and quercetin, can increase the content of active oxygen and downregulate cancer cell migration and proliferation by regulating several signaling pathways, such as EGFR/ MAPK signaling pathway[63].
Long-term controlled release of either natural anticancer drugs or chemotherapy drugs is very important for tumor treatment. 3D porous scaffolds have been widely used in cancer therapy and tissue engineering due to their good capabilities in drug controlled release[96-99]. Zhao et al.[32] designed and developed a multifunctional biomimetic cellulose nanofiber (CNF) in situ liquid wound dressing (CNF-ILWD). CNF-ILWD was simultaneously loaded with photothermal agent (indocyanine green) and chemotherapeutic drugs (doxorubicin) during the preparation process. NIR, temperature, and pH multiple response switches could efficiently control the drug release of CNF-ILWD to kill residual tumor cells in wounds and deep layers of skin, and eliminate bacterial biofilms and harmful bacteria. Therefore, drug-loaded CNF-based wound dressings can be used for postoperative tumor therapy and to promote the repair of infected wounds. The functional material products recently used for skin wound repair are presented in Table 1.
Despite the significant advancements in the field of tissue engineering, a large number of functional or multifunctional wound healing materials are still afflicted with problems such as morphological inconsistencies with wounds, difficulty in generating natural vascular networks and skin appendages, and difficulty in nutrient and oxygen exchange between tissue cells[100,101]. Also, it is hard to meet the diverse needs of wounds in complex situations. In recent years, 3D bioprinting technology has emerged as an ideal strategy to replace traditional low-precision cell spraying and seeding techniques to deposit cells, biomaterials, and bioactive molecules into precise 3D geometric patterns. Computer control provides tools for the development of vascular and adnexal regeneration, thereby replicating the anisotropy of natural skin[102,103].
3D bioprinting is an advanced additive manufacturing technology, which can distribute bioink containing biological materials, cells, or other active substances in a controllable space, so as to repeatedly manufacture 3D functional structures of various shapes and sizes with high flexibility[104]. According to the molding principles and printing materials, current bioprinting technologies mainly include extrusion-based bioprinting, laser-assisted bioprinting, digital light processing-based bioprinting, inkjet bioprinting, and microfluidics-assisted bioprinting[105,106].
Extrusion-based bioprinting is the most popular form of bioprinting that applies mechanical actuation or pneumatic pressure to extrude a bioink from a nozzle continuously, and deposit it layer-by-layer to form a 3D structure[107,108] (Figure 3a). Extrusion bioprinting systems can be classified into screw, piston, and pneumatic type according to their working principles[109]. Compared with other bioprinting technologies, extrusion-based bioprinting is relatively simple and low-cost, can handle high-viscosity bioinks, and has excellent compatibility with multiple materials (decellularized extracellular matrix [dECM], microcarriers, polymers, hydrogels, and cell aggregates)[110]. However, this system suffers from lower print resolution (50'400 microns) and longer production times due to the small nozzle diameter. Furthermore, when the cell density in the ink is too high, the high shear stress during extrusion reduces the number of viable cells[101,110].
Bioprinting technology. (a) Extrusion bioprinter is a continuous extrusion of cell-containing liquid bioink using manual or pneumatic force. (b) Schematic diagram of the laser bioprinting device. (c) Schematic illustration of the DLP-based bioprinting device. (d) Inkjet bioprinter sequentially ejects small droplets of hydrogels and cells to construct tissue. (From ref.[125] licensed under Creative Commons Attribution 4.0 International license.) (e) Four typical 3D bioprinting techniques correspond to four ways of cutting potatoes. (Reprinted with permission from Gu Z, Fu J, Lin H, et al., , Development of 3D bioprinting: From printing methods to biomedical applications. Asian J Pharm Sci, 15(5):529'557[123]. Copyright © Shenyang Pharmaceutical University.) (f1) Rendered image of the handheld skin bioprinter. (f2) Picture of the 3D-bioprinted microfluidic box. (Reprinted with permission from Hakimi N, Cheng R, Leng L, et al., , Handheld skin printer: In situ formation of planar biomaterials and tissues. Lab Chip, 18(10):'[126]. Copyright © The Royal Society of Chemistry .)
Laser-assisted bioprinting uses an energy source (continuous monochromatic laser energy or pulses) to irradiate a light-absorbing layer, thereby causing the bioinks to be deposited as droplets on the printing platform by light[111] (Figure 3b). Depending on the laser source, laser-assisted bioprinting can be subdivided into laser direct writing (LDW), laser-induced forward transfer (LIFT), and matrix-assisted pulsed laser evaporation (MAPLE)[112]. Laser-assisted bioprinting has a high system resolution and open nozzle structure, which can precisely arrange small volume of cell droplets in 3D spatial positions, eliminating the problem of nozzle blockage. In addition, as a noncontact printing technology, it can prevent cell and bioink contamination to a certain degree. However, this technology can only select photosensitive polymers for printing, and photopolymerization requires additional chemical modification of materials, which limits the extensive use of various biological materials. In addition, this technology has high maintenance cost and long production time, which leads to low printing efficiency and difficulty in printing large tissues and organs[113,114].
Digital light processing-based (DLP) 3D bioprinting uses a digital micromirror device (DMD) to project a designed optical pattern onto an ink container, by174 manipulating light to induce the bioink in the exposed area to polymerize and cure a complete layer[115]. As the platform is raised and lowered, each new cured layer is bonded to the previous one, resulting in a complex and smooth structure[116] (Figure 3c). DLP bioprinting technology has high printing speed (printing time of seconds to minutes) and high resolution (200 nm'6 μm) with shorter printing time. Furthermore, it enables the use of bioinks with high cell concentrations (>106 cells/mL) without causing clogging of the nozzles[116]. Because of these advantages, this technology can simulate the precise structure and cell viability of natural tissues, leading to breakthroughs in the printing of functional living organ structures. However, DLP printing can only use photocurable bioinks, and the UV light used during polymerization may have an impact on cell viability.
Inkjet bioprinting is a noncontact printing process in which bioinks loaded into nozzles are stacked into structures in the form of droplets[106,117-120] (Figure 3d). This bioprinting techniques can generally be divided into two types: thermal inkjet bioprinting and piezoelectric inkjet bioprinting[104,121]. A major advantage of inkjet bioprinting is high resolution (50 μm), which enables the fabrication of complex scaffolds by printing multiple materials with high fidelity into relevant dimensional structures[116]. In addition, it has the advantages of high printing speed (10,000 drops per second), simultaneous printing of multiple ink cartridges, and low technology cost[122]. At the same time, inkjet bioprinting also has some limitations. For example, its small nozzle diameter and easy clogging limit its ability to print bioinks with high cell concentration and high viscosity[116]. Additionally, exposure of cells to high temperature of the nozzle and shear stress also reduces cell viability[122]. These four typical 3D bioprinting processes correspond to the inverse processes of potato slicing, shredding, dicing, and mashing, respectively[123] (Figure 3e).
175As science and technology continue to advance, current bioprinting techniques are also improved. For example, microfluidics-assisted extrusion bioprinting is a micro-device printing technology based on microfluidics, which enables precisely controlled deposition of multiple materials to obtain 3D structures in a relatively short period of time[124] (Figure 3f1 and 3f2). As an additive bio-manufacturing technique, 3D bioprinting can offer an essential strategy for wound dressings or skin tissue engineering to manufacture personalized construct precisely and dexterously in a short time, which would shorten the waiting time and reduce the suffering of the patients as well as accelerate regeneration of skin function.
This paper introduces various bioprinting methods, functional materials, and their applications in wound dressing and skin tissue engineering. 3D bioprinting emerges as an additive bio-manufacturing technique possessing the advantages of high resolution, flexible operation, repeatable fabrication, and high-throughput output for printing the intricate 3D structures that match the geometric shape of skin wound[117,158], thus it has been widely used in wound dressings and skin tissue engineering scaffolds in recent years[19,43]. As one of the development trends of advanced materials, multifunctional materials have become an attractive option for wound dressings and skin tissue engineering scaffolds. However, the cytotoxicity that may occur when the dosage of these multifunctional materials exceeds the cytotoxicity threshold is not negligible. Moreover, unlike traditional bandages, current 3D-bioprinted hydrogel dressings usually suffer from poor mechanical strength and stability although possess multiple functions, and do not function on knees and joints for long periods due to poor adhesion. Also, current 3D-bioprinted dressings required more research in overcoming the challenges of scars, nonoxygen permeable and damaged skin cells[159].
The main distinctions of skin tissue engineering compared to the wound dressing are the loaded cells and bioactive factors. Whether it is to print the bionic skin structures with cell-encapsulating bioink, or to inoculate cells on the noncellular-printed scaffolds, the requirements of skin tissue engineering scaffolds for printing materials and conditions are stricter than that of the printed dressings, such as biocompatibility and viscosity of the bioinks, suitable temperature and pH, and sterile microenvironment for cell survival[114,160]. Although significant progress has been achieved in tissue engineering over the years, only a limited number of bioinks have the tissue matching characteristics and the ability to promote tissue generation[161]. At present, it is still a major challenge for skin tissue engineering to configure multifunctional bioink with printability, biocompatibility, and excellent mechanical integrity under individual condition[162]. Therefore, the design of mixed bioink should integrate the advantages of natural bioink and synthetic bioink to prepare bioink that is conducive to cell growth and can support cell survival in the printing process[160]. In addition, cell encapsulation bioink can use various types of cells, such as fibroblasts, keratinocytes, mesenchymal stem cells, and induced pluripotent stem cells, as cell sources[117]. Stem cells, such as induced pluripotent stem cells, can differentiate into various types of skin cells, but they are sensitive to the shear stress imposed on the cells during printing and are difficult to survive[161]. Autologous cells from patients are the source of gold-standard cells in skin bioprinting. While reproducing all functions of tissues and organs, they have no rejection reaction to patients, and can survive with sufficient vitality and maintain functions during the printing process. However, the normalization and standardization of human clinical trials related to 3D bioprinting cell-encapsulated bioinks before skin bioprinting can be translated to clinical application is another challenge, and it will take several years to develop a dedicated regulatory framework or dedicated regulatory guidance to make 3D bioprinting sustainable[163].
Another challenging problem that prevents skin regeneration is angiogenesis during skin repair[107]. The skin structure needs highly developed vascular network to supply nutrients and oxygen[164]. In addition, the bioprinting of complete skin with multilayer complex structure is still a difficult problem in tissue engineering. The thickness and texture of the epidermis, dermis, and subcutaneous adipose layer of the bioprinted skin should match the patient's natural skin, while the recovery of multiple functional skin appendages, such as sweat glands, hair follicles, and sebaceous glands, should be consistent with the normal skin anatomical structure and function[117,165]. At present, for most wound healing materials, the exploration of their mechanism and the evaluation of their therapeutic effects are carried out in animal models, such as mice. The phenomena and effects observed in animal models may not be fully applicable to humans[166]. For example, there are some significant differences between mice and humans in inflammatory reaction and cell behavior, and the complex microenvironment effects in vivo also make the experimental results uncertain, which is a very important limitation for translation[166]. Therefore, clinical validation should be carried out in larger skin defect models or chronic skin wound models, so as to enable their direct application in the future[167].
It can be predicted that combining the most advanced tissue engineering strategies and the achievements of current and ongoing research; it is very promising to develop fully functional bioprinted skin. Recent in situ bioprinting research has shed light on the concept of biological manufacturing of tissue directly in the living body[46]. Advanced in situ 3D bioprinting technology to combine multiple functional materials and bioactive factors to create fully functional bioprinted skin is a rapid skin construction technology with lower rejection rate. In addition, it can create specific organs from patients' cells in lesser time and lower cost, thus making the research and development process simpler, faster, and better. Moreover, in situ bioprinting should be integrated with other functions, such as realtime monitoring, higher degrees of freedom, equipment miniaturization, and dynamic surface printing[44,46]. In short, the structural complexity of the bioprinted skin structure requires further enhancements through the collective efforts of various technologies, in a bid to create a fully functional skin with lesser time and lower cost.
3D-bioprinted wound dressings and skin tissue engineering scaffolds have been widely used for skin wound repair. They are made of natural or synthetic polymers and can promote wound repair and tissue regeneration. At present, the main183 challenges facing wound healing materials are the further development of multifunctional materials, the progress of biological printing technology, and the construction of skin's functional structure. In the future, we believe that continuous advances in skin research, healing product design, material formulation, and printing technology can not only ease the preparation of new multifunctional wound healing materials but also lay a foundation for the clinical application of functional bionic skin. In the coming years, multifunctional, multimaterial, and multiscale manufacturing will be the focus in the research on 3D bioprinting of wound healing materials.
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This work was supported by the National Natural Science Foundation of China (/), the Fok Ying Tung Education Foundation (), the State Key Laboratory of Fine Chemicals (KF), and the Fundamental Research Funds for the Central Universities (DUT21YG113/DUT22YG213/DUT22YG116).
The authors declare no conflict of interest.
Conceptualization: Kedong Song, Huan Fang, Jie Xu Project administration: Kedong Song, Hong Wang Supervision: Kedong Song, Yi Nie, Hong Wang, Bo Pan Visualization: Huan Fang, Jie Xu, Hailin Ma, Kedong Song Writing ' original draft: Huan Fang, Jie Xu, Hailin Ma, Jiaqi
Liu, Erpai Xing
Writing ' review & editing: Kedong Song, Yuen Yee Cheng, Huan Fang, Jie Xu, Jiaqi Liu
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