Abstract
Melt-electrowriting (MEW) is a high-resolution additive manufacturing technique that has demonstrated significant progress in recent years. Owing to its precise control over fiber deposition, MEW is especially suitable for fabricating fine structures that mimic the natural extracellular matrix (ECM), thereby presenting considerable promise for applications in tissue engineering and regeneration. This review systematically examines the fundamental design principles and recent progress in MEW-based strategies for different tissue engineering and regeneration fields. Initially, the components of the MEW system, the underlying printing mechanisms, and the role of key process parameters are introduced, thereby providing a comprehensive framework for the rational design of scaffolds that replicate both the structural and functional characteristics of native ECM. Subsequently, the selection and performance of commonly employed biomaterials are discussed, with an emphasis on the versatility for diverse tissue engineering applications. The integration of MEW with bioactive materials is further highlighted as an effective approach to enhance the biological functionality of printed constructs and extend their therapeutic potential. Finally, current challenges and future perspectives are outlined, aiming to guide ongoing research and facilitate the clinical translation of MEW-based biofabrication technologies.
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1 Introduction
Constructing functional micro/nanoscale structures to mimic the extracellular matrix (ECM) is important for tissue regeneration in vitro and in vivo [1, 2]. As a complex architecture, the ECM exerts pivotal influences on various biological functions: (i) the ECM provides attachment cues for cells and organizes the individual cells into three-dimensional (3D) tissues [3]. (ii) ECM plays a dominant role in supporting the transport of nutrients to promote metabolic processes [4]. (iii) As an intricate network composed of multiple biomacromolecules, ECM can deliver topographical and biological signals, such as integrins, proteoglycans, and growth factors [5]. Therefore, an ideal artificial micro/nanoscale structure should possess a 3D architecture, nutrient channels, and topographical cues. In addition, therapeutic applications require appropriate structural dimensions and suitable mechanical properties for suturing [6, 7].
To accurately replicate the functional native ECM structure, researchers have extensively endeavored using various methods [8]. For instance, electrospinning is a typical technique employed to fabricate sub-microscale or nanoscale fibrous architectures, which can provide topographical cues for cells [9,10,11]. Furthermore, functional nanomaterials and bioactive factors are added to the electrospun fibrous films to regulate cellular behaviors, such as proliferation and differentiation [12, 13]. However, the random fiber orientation inherent to electrospinning limits its ability to replicate the organized 3D architecture of native ECM.
Additive manufacturing is a highly controllable manufacturing approach to fabricate structures with precise geometries, which has found widespread application [14, 15]. For example, two-photon polymerization and fused deposition modeling (FDM) printing are widely used to fabricate 3D architectures for tissue engineering applications. Two-photon polymerization enables to produce 3D structures with the smallest feature size down to nanometer scale, which provides an accurate tool to regulate cellular behaviors [16]. However, this method commonly requires expensive optical systems and specific photosensitive polymers which limits its biological applications. Moreover, FDM technique can produce sturdy and durable parts, thus is widely used in bone tissue substitutes. Teixeira et al. constructed a 3D-printed poly(lactic acid) (PLA) scaffold with poly-dopamine and collagen I coating by FDM, which exhibited an enhanced osteoinductivity. However, the FDM method unable to construct scaffolds with high-resolution and small pore diameters, which is not beneficial to cell-scaffold interaction [17].
Recently, electrohydrodynamic (EHD) printing has emerged as a high-resolution 3D printing technique to fabricate micro/nanoscale fibrous structures mimicking the ECM [18,19,20]. This printing technique precisely constructs micro/nanoscale fibrous scaffolds through fiber deposition, which combines elements of electrohydrodynamical fiber attraction and melt extrusion [21, 22]. EHD printing enables the creation of a biomimetic microenvironment with 3D structures to support cell growth and migration. The 3D porous structure has the ability to facilitate nutrient transfer, thereby regulating cell growth, proliferation, and activity [23]. As a specialized branch of EHD printing, Melt-electrowriting (MEW) can construct artificial tissue analogs using melt-based polymers as printing materials [24, 25]. Compared to solution-based EHD printing, MEW exhibits advantages, such as solvent-free processing and enhanced process stability. These benefits not only mitigate the risk of toxic residues but also improve controllability, which is facilitated by a lower surface charge density [26]. Moreover, surface modification of specific materials and the incorporation of related growth factors broaden the application scope of MEW in engineering various tissues. Hence, this unique printing method holds great promise for the precise production of micro/nanoscale fibrous scaffolds for tissue engineering.
This article reviews the latest progress in MEW, focusing on the fabrication of artificial tissue analogs as implantable materials for treating injuries and diseases. We summarize the principles, process parameters, and common biomaterials used in the MEW process. Furthermore, we highlight typical applications of MEW fibrous scaffolds in different tissues, involving cardiac tissues, bones, cartilages, vascular, ligaments and tendons, periodontal tissues, and nerves (Fig. 1). Finally, this review provides new insights into the design concept of the tissue engineering scaffolds with an emphasis on the latest designs and innovations of MEW scaffolds, especially in terms of their 3D structure, mechanical properties, and incorporation of bioactive materials. Finally, the main challenges and future perspectives are comprehensively summarized and discussed to inform future research directions.
Advances in the application of MEW in tissue engineering and regeneration. The schematic indicates the advantages of MEW on constructing functional scaffolds and replicating the intricate structure of ECM
2 Fundamentals and Mechanism of Melt-Electrowriting
A typical MEW system comprises a glass syringe equipped with a heating sleeve, a high-voltage source, a computer-aided translating system, and a collector (Fig. 2a) [27, 28]. During the MEW process, materials are melted within the syringe and then extruded through a metal nozzle under pressure (air pressure or mechanical extrusion). Meanwhile, a high voltage was applied to the nozzle, and the charged pendant droplet deformed to a conical shape (Taylor cone) (Fig. 2b). As the voltage intensifies, the cone further sharpens to achieve equilibrium between the electric field force and surface tension. When the charges overcome the surface tension, a charged jet is ejected toward the collector [29]. This process arises from the equilibrium of forces between gravity, surface tension, internal hydrostatic pressure, external gas pressure, and the electric force [21]. When the printing speed exceeds the translation speed, a phenomenon known as "jet lag" occurs, where the contact point of the polymer jet lags behind the nozzle [30]. Moreover, as the material droplets are deposited, charges are transferred to the conductive substrate and grounded [31]. With the assistance of the computer-aided translating system, the microscale fibers are deposited onto the collector substrate according to a specific path, thereby constructing the fibrous scaffolds. Furthermore, MEW integrates structural design with customized programs, making it possible to develop optimized biomimetic structures [32].
Principle and strategies for realizing the intricate structures of MEW technique. a Components of a MEW system; Reproduced with permission from Ref. [27], Copyright 2018, Wiley–VCH. b Schematic of the MEW working principle during the printing process. c Coiling (c1), sinusoidal (c2), and straight (c3) fibers printed at below and above CTS speeds; Reproduced with permission from Ref. [18], Copyright 2018, Wiley–VCH. d Principle of EHD printing of sinusoidal PCL fibers; Reproduced with permission from Ref. [43], Copyright 2024, American Chemical Society. e Schematics of multimaterial printing of piezoelectric cardiac microarchitectures combining melt-based, solution-based EHD printing and extrusion printing processes; Reproduced with permission from Ref. [44], Copyright 2024, Wiley–VCH. f Schematic of different collector substrates ranging from flat to wedge and curved dome-shaped structures; Reproduced with permission from Ref. [45], Copyright 2020, Elsevier. g Mandrel collector as printing substrate; Reproduced with permission from Ref. [47], Copyright 2022, Wiley–VCH. h Illustration of fiber deviation phenomena during printing process; reproduced with permission from Ref. [51], Copyright 2022, Elsevier. i Schematic of reducing fiber diameter to realize smaller fiber spacing; Reproduced with permission from Ref. [52], Copyright 2023, Wiley–VCH. j Overhang (j1) structure, and branches (j2) of fiber wall into two overhangs; Reproduced with permission from Ref. [53], Copyright 2020, Wiley–VCH
To attain a consistent fiber diameter and a stable scaffold structure, the optimization of process parameters has been thoroughly examined. Significant efforts have been directed toward refining key parameters, including the applied voltage, feed rate, nozzle-to-collector distance, collector velocity, and the type of collector [33,34,35,36,37]. Through optimization of the parameters, MEW can fabricate a full spectrum of various fibers with discrete diameters using a single nozzle [27, 38]. The voltage applied between the nozzle and collector directly influences the electrostatic forces and the maintenance of the Taylor cone, ultimately impacting printing accuracy. Therefore, during the layer-by-layer stacking process, maintaining a constant electrostatic force is crucial to minimize [39]. The feed rate (controlled by pressure) determines the diameter of the fiber, and the diameter of the fiber increases as the feed rate increases. However, the feed rate must be kept within an optimal range to prevent the material from extrusion difficulties and jet instability. The distance between the nozzle and the collector also plays a pivotal role in influencing the electric field strength, which in turn affects the formation of the Taylor cone and the jet. Collector velocity significantly impacts both the fiber diameter and the jet morphology. As the velocity increases, the jet exhibits three distinct morphologies. The minimum velocity that allows for the printing of linear fibers is termed as the critical translation speed (CTS) [40]. At this speed, the jet is perpendicular to the collector substrate and the collected fibers are straight. Below the CTS, the jet vibrates and causes the collected fibers to be coiled or sinusoidal (Fig. 2c) [18, 41, 42]. Above the CTS, the angle between the fiber and the collector substrate decreases, and the degree of jet lag increases, resulting in smaller fiber diameters [21]. In another study, sinusoidal fibrous scaffolds were fabricated based on a zig-zag movement of the collector, and their fiber morphology (such as period and amplitude) could be broadly adjusted (Fig. 2d) [43]. Taking advantage of this approach, multi-material serpentine scaffolds can be constructed using a multi-nozzle using 3D printing system (Fig. 2e) [44].
In addition, the interactions between the fibers and the substrate type can influence the fiber deposition accuracy. For example, Hou et al. found that using an aluminum foil with a 45° relative placement angle resulted in the smallest fiber deposition deviation ratio of less than 1.16% [36]. Furthermore, multi-biomaterial fibrous scaffolds on non-flat surfaces were explored by maintaining a constant distance between the nozzle and the collector to ensure a stable electrical field (Fig. 2f) [45]. Besides common flat substrates, mandrel collectors are commonly employed for constructing tubular scaffolds, which are extensively used in heart valves engineering and vascular grafts (Fig. 2g) [34, 46,47,48].
The printing accuracy of MEW is affected by various factors. The factors impacting the printing accuracy mainly include fiber deviation (including fiber overlapping and tilting) and residual charge densities [49]. Specifically, fiber deviation is influenced not only by processing parameters, such as material temperature, collector temperature, collector conductivity, applied voltage, and nozzle-to-collector distance, but also by scaffold design parameters and the inevitable presence of jet lag, including fiber diameter, layer number, and inter-fiber distance [50, 51]. Chang et al. introduced a post-process index (Ip), which compares the experimental proportion of void area in a top-view image of the fabricated scaffold sample (Pf) to the theoretical proportion in the scaffold design pattern (Pt), defined as Ip = Pf/Pt (Fig. 2h). It is concluded that when Ip is equal to 1, the printing precision can be improved by reducing the fiber diameter and layer number, and increasing inter-fiber distance [51]. This phenomenon has been previously investigated, leading to the conclusion that a uniform fiber arrangement can be achieved when fiber spacing is five times greater than the fiber diameter (Fig. 2i) [52].
Although the above stacking deviation is considered inaccuracy, researchers can achieve precise control over geometry and mechanical properties by leveraging microscale layer displacement. This advancement allows for the creation of innovative nonlinear geometries, such as overhangs (Fig. 2j1) and branches (Fig. 2j2), as well as seamless transitions and abrupt changes in print tracks. These capabilities showcase the unparalleled flexibility of MEW [53,54,55].
3 Polymer Materials for Melt-Electrowriting
Table 1 lists the biomaterials commonly used by MEW in the field of tissue engineering and regeneration in recent years. In general, due to the nascent stage of MEW development and the scarcity of printable polymers, polycaprolactone (PCL) and its derivatives continue to dominate as the primary polymers [29]. PCL boasts exceptional biocompatibility, a high molecular weight, a low melting temperature, rapid solidification, and superb processing capabilities [56]. Moreover, PCL maintains its Young's modulus, yield strength and crystallinity under heated conditions for hours or days [57, 58].
Researchers can use various methods to modify the performance of PCL to meet the requirements of different tissue engineering applications. First, MEW PCL scaffolds exhibit low hydrophilicity which affects the cell adhesion/proliferation [59]. To address this issue, researchers have explored various post-treatment methods, including oxygen plasma treatment [58] and surface modification [60]. These methods significantly enhance the hydrophilicity of PCL scaffolds and the adhesion ability of cells. Second, researchers have mixed molten PCL with nano-hydroxyapatite (nHA) to fabricate PCL/nHA composite scaffolds to promote bone formation [61]. In addition, PCL and other materials were dissolved in certain organic solvents. After the solvent evaporates, PCL-based printing materials with specific functions are obtained. For example, PCL, polyethylene glycol (PEG), and roxithromycin (ROX) were mixed in tetrahydrofuran and subsequently evaporated. The MEW composite scaffolds exhibited improved hydrophilicity, cell proliferation, and antibacterial properties [62]. Similarly, PCL/magnesium phosphate (MgP) composite scaffolds exhibit significant potential for periodontal regeneration, while poly(hydroxymethylglycolide-co-ε-caprolactone) (pHMGCL)/PCL composite scaffolds demonstrate tunable degradation profiles and mechanical adaptability for cardiac tissue engineering applications [63]. In addition, a porous scaffold fabricated by poly(ε-caprolactone)-block-poly (ethylene glycol)-block-poly(ε-caprolactone) (PCEC) has been integrated with gelatin methacryloyl (GelMA) hydrogels to synergistically enhance their mechanical properties. The biocompatible triblock copolymer PCEC exhibited improved hydrophilicity and degradability compared to pure PCL [64].
Other materials have gradually come into view in recent years. Notably, PLA and poly(L-lactic acid) (PLLA) have been employed in the fabrication of high-fidelity structures [65, 66]. PLA and PLLA were widely used in tissue engineering and bone fixation, because they possess better biodegradability and high stiffness compared to PCL [67, 68]. Furthermore, poly(L-lactide-co-ε-caprolactone-co-acryloyl carbonate) (PLACLAC), synthesized via open-loop polymerization, has demonstrated efficacy in fabricating fibrous scaffolds with robust mechanical properties, including rigidity, elasticity, and creep resistance, specifically engineered for tendon and ligament repair [42].
4 Melt-Electrowriting Scaffolds for Tissue Engineering Applications
The onset of pathologies is frequently linked to disruptions in the ECM, leading to pathological tissue remodeling and ultimately tissue dysfunction [92, 93]. Previous studies have demonstrated considerable efforts in the domains of tissue engineering and regenerative medicine, highlighting the significance of replicating the intricate architecture of native tissue. These biomimetic structures offer a promising avenue for the creation of functional tissue substitutes that can expedite the healing process [56]. Given that the collagenous fibers within each tissue exhibit distinct structures, all human tissues possess unique mechanical properties crucial for their functionality.
Advancements in MEW and biomaterials research have furnished the methodological and theoretical frameworks necessary for the implementation of biomimetic scaffolds. MEW enables the easy regulation of mechanical properties (such as the effective stiffness and anisotropy) of scaffolds. This capability allows for tailored designs that cater to tissues with varying sizes or structures. Hence, this strategy has been widely employed to fabricate scaffolds with microscale fibrous structure for tissue engineering applications (Table 2).
This section discusses the specific applications of MEW scaffolds in tissue engineering and systematically covers the main application fields, highlighting the unique challenges and demands of different tissues, etc. This section will successively explore the application progress and challenges of MEW scaffolds in the fields of cardiac tissue engineering, bone regeneration, cartilage tissue engineering, vascular grafts, ligament and tendon tissue engineering, periodontal regeneration and nerve repair.
4.1 Cardiac Tissue Engineering
Cardiovascular diseases (CVDs) have high morbidity and mortality rates, often leading to extensive myocardial cell loss in the left ventricle, which is subsequently replaced by fibrotic scar tissue [94, 95]. This scarred area loses its contractile ability, ultimately leading to heart failure [96]. Adult myocardial tissue has limited regenerative capacity, often leading to progressive pathological remodeling as the heart attempts to compensate for lost contractility [97].
Reconstructing cardiac tissue in vitro represents a potential strategy for treating myocardial tissue loss. This strategy lies in replicating the fibrous structure and mechanical environment of native cardiac tissue, which includes appropriate mechanical performance to facilitate contraction, the ability to guide the orientation of cell alignment, and outstanding conductivity to promote electrical signal propagation between myocardial cells [98]. Moreover, the conductive scaffold can promote synchronous contraction and the expression of proteins involved in muscle contraction and electrical coupling [99, 100].
Previous research has shown that the stiffness of heart muscle is approximately 10–20 kPa at the beginning of diastole, increasing to 50 kPa in a healthy heart by the end of diastole, and can reach 200–300 kPa in an infarcted heart; the passive strain of the myocardium can reach 15–22%, and the myocardial conductivity is approximately 0.57 S/m. By carefully integrating these prerequisites, researchers can develop innovative methods to address the potential loss of myocardial tissue, offering new hope for the treatment of CVDs and heart failure [101,102,103].
4.1.1 MEW Scaffolds for Cardiac Tissue Engineering
Various melt MEW scaffolds have been developed, and extensive research has focused on their mechanical properties. The most common design is grid structure, which exhibits isotropic mechanical properties. In contrast, the rectangular structure (Fig. 3a) exhibits anisotropic properties depending on its aspect ratio, providing flexibility in adjusting mechanical properties through precise manipulation of the aspect ratio [32, 91]. However, it is crucial to note that the long-term mechanical performance of these scaffolds as implants has not been thoroughly evaluated.
MEW scaffolds for cardiac tissue engineering. a Squared and rectangular MEW microfibrous scaffolds; Reproduced with permission from Ref. [91], Copyright 2017, Wiley–VCH. b Sinusoidal scaffolds printed at print speeds below CTS; Reproduced with permission from Ref. [69], Copyright 2023, Elsevier. c Sinusoidal PCL fibers with a period of 3 mm and an amplitude of 2 mm; Reproduced with permission from Ref. [43], Copyright 2024, American Chemical Society. d Printing path and structure morphology of auxetic structure; Reproduced with permission from Ref. [60], Copyright 2020, Wiley–VCH. e MEW scaffolds with hexagonal cells structure; Reproduced with permission from Ref. [70], Copyright 2018, Wiley–VCH. f Image of full scaffolds printed under final conditions: 6.5 kV, 800 mm/min, 2 bar; Reproduced with permission from Ref. [89], Copyright 2025, Wiley–VCH. g Morphology of PPy modified on fibrous scaffolds before and after; Reproduced with permission from Ref. [60], Copyright 2020, Wiley–VCH. h Schematic of the fabrication of multiscale conductive scaffolds; Reproduced with permission from Ref. [72], Copyright 2019, Royal society of chemistry. i Schematics of magnet-driven mechanical and piezoelectric stimulation for cardiac maturation; Reproduced with permission from Ref. [44], Copyright 2024, Wiley–VCH
To achieve synchronized deformations and long-term implantation, it is essential to design scaffolds that possess suitable mechanical properties. Chen et al. developed a fibrous scaffold with sinusoidal structures under specific printing conditions below the CTS (Fig. 3b). This structure has a smaller period than the serpentine structure (Fig. 3c) [44], so it can regulate cell behavior at the cellular scale. When stretched with constant frequency and strain, sinusoidal scaffolds exhibited a lower Young's modulus and a higher elastic limit strain compared to rectangular scaffolds, which is beneficial for synchronized contraction of beating cells. Moreover, MEW shows the flexibility of adjusting the anisotropy ratio of sinusoidal scaffolds [69]. Notably, the sinusoidal mesh could extend along its long axis by up to 23% before undergoing irreversible deformation, compared to 7% for straight fibers. The elastic strain energy of sinusoidal scaffolds was more than three times higher than that of rectangular grids. Analogously, the serpentine architectures can mimic the curved collagenous fibers in the cardiac ECM, and the scaffolds show outstanding flexibility and elasticity compared with straight ones [43]. This structure is more path-dependent, and therefore, more controllable than sinusoidal structures constructed below the CTS.
Another interesting structure is the auxetic structure (Fig. 3d), which can be considered as a macroscopic version of the sinusoidal structure [60]. Unlike the microscale sinusoidal structure, auxetic structures exhibit bidirectional simultaneous deformation based on a negative Poisson's ratio. This leads to expansion in multiple directions simultaneously and enhances unique properties, such as shear resistance, indentation resistance, and synclastic curvature, making it an attractive option for cardiac patches [103]. Inspired by beehives, Castilho et al. generated fibrous scaffolds with hexagonal 3D microstructures that can achieve large and anisotropic reversible deformations (Fig. 3e). These scaffolds demonstrated potential for minimally invasive applications and were able to extend along the x-direction by 35%–40% before permanent deformation, compared to 2%–3.5% for rectangular scaffolds. Notably, the energy absorption and release capacity of hexagonal scaffolds during elastic deformation was 20–40 times greater than that of rectangular structures [70]. Furthermore, some latest research focuses on the combination of hydrogels and MEW scaffolds. For example, Vega et al. fabricated a bio-inspired scaffold with diamond-shaped pores by MEW, this geometry exhibited a large deformation beyond 50% in the short direction (Fig. 3f). Moreover, the flexible scaffold was encapsulated within a fibrin-cell hydrogel suspension, which exhibited a high cell density and thereby enabled large-scale synergistically deformation of the scaffold [89]. Qiu et al. utilized a consecutive hybrid bioprinting strategy by combining MEW and extrusion-based bioprinting to construct oriented tissues on the fibrous scaffold. The resultant aligned myoblast constructs could be differentiated into multinucleated myotubes with enhanced muscle-specific protein and gene expression [90].
In summary, diamond, hexagonal, auxetic, and sinusoidal structures demonstrate superior mechanical behavior compared to rectangular scaffolds. Their enhanced elastic response and higher failure strain facilitate synchronized cardiac tissue contraction. Furthermore, these studies emphasize the significant potential of MEW in constructing fibrous scaffolds with tailored structure and mechanical properties to meet the requirements of cardiac tissue.
4.1.2 Electrically Conductive Functionalization of MEW Scaffolds
Endowing MEW scaffolds with suitable conductivity can significantly enhance signal propagation for cardiomyocytes (CMs) [98]. Conductivity is a crucial evaluation criterion for cardiac fibrous scaffolds. Over the past decades, advancements in the modification of conductive materials on the MEW scaffolds have primarily centered around three methods: sputter coating [69], in-situ chemical oxidative polymerization [60], and hybrid printing [72].
During sputter coating, a thin layer of gold nanoparticles is deposited onto PCL scaffolds. When subjected to high-intensity electrical fields, these scaffolds demonstrated a significant increase in their ability to coordinate synchronous contractions between cardiac cells after 120 h of electrical stimulation [69]. However, compared to sputter coating, scaffolds modified with polypyrrole (PPy) through in-situ chemical oxidative polymerization exhibit superior conductivity and enhanced chemical stability (Fig. 3g) [60].
Moreover, hybrid EHD printing presents an alternative approach to achieving conductive functionality. This method allows for the construction of a poly(3,4-ethylenedioxythiophene)–poly(styrenesulfonate)–polyethylene oxide (PEDOT:PSS–PEO) conductive layer that mimics the multiscale, layer-specifically oriented architectures of native myocardium in a controlled manner (Fig. 3h) [72].
Beyond conductive materials, piezoelectric materials have gradually gained significant attention with the advancement of electroactive materials [104]. A recent study demonstrated multimaterial-printed serpentine microarchitectures comprising PCL fibers, polyvinylidene fluoride (PVDF) fibers, and PCL/Fe3O4 magnetic microfibers (Fig. 3i). PCL microfibers were used as mechanical support and PVDF microfibers were used for piezoelectric stimulation. The magnetic deformation of the multi-material structure was controlled by an external magnet through PCL/Fe3O4 fibers. This engineered piezoelectric cardiac scaffold shows significant structural and functional repair of infarcted myocardium, demonstrating the therapeutic potential of piezoelectric cardiac structures in the treatment of myocardial infarction [44].
4.2 Bone Regeneration
Bone tissue engineering scaffolds can provide effective strategies for bone injury repair and replacement, such as osteoporosis, osteonecrosis, and osteomalacia [105, 106]. These scaffolds can replicate the natural bone ECM, promoting the adhesion, proliferation, and differentiation of bone cells [107,108,109]. In recent years, progress has been made in the creation of bone tissue engineering scaffolds capable of providing essential osteoconductive and osteoinductive cues that support robust bone regeneration [110, 111]. Various methods have been employed to fabricate these scaffolds, including lyophilization [112], solution electrospinning [113], and additive manufacturing techniques [114]. Among these methods, MEW has the ability to fabricate high-resolution biomimetic fibrous structures in a controlled manner, and its microarchitecture plays a pivotal role in regulating bone healing response.
4.2.1 MEW Scaffolds for Bone Tissue Engineering
During the process of bone healing, the formation of new blood vessels is a key connection in the repair. They promote bone tissue reconstruction by providing oxygen, nutrients and cellular support. MEW scaffolds can support the growth of smaller vascular networks during bone repair, primarily composed of vessels with diameters at or below 20 μm [115]. Moreover, FDM scaffolds have been introduced to promote bone regeneration. These scaffolds feature a 0°/90° lay-down pattern and possess exceptional biomechanical properties, and the large pores of FDM scaffolds can facilitate the development of larger vessels [116, 117]. Kelly et al. investigated how scaffolds’ microstructures in FDM and MEW affect the healing of large bone defects. The results revealed that FDM PCL scaffolds facilitated the formation of large bone spicules through their pores, whereas MEW PCL scaffolds contributed to the development of a rounder bone front during the healing process. Furthermore, both FDM and MEW scaffolds demonstrated the ability to support vascularization, FDM scaffolds supported greater total vessel formation, including the development of larger vessels (Fig. 4a) [71]. Compared with FDM scaffolds, MEW scaffolds supported higher levels of new bone formation and encouraged the formation of a dense microvascular network with minimal infiltration of larger vessels into the defect area.
MEW scaffolds for bone regeneration. a FDM and MEW scaffolds after 21 days of human MSCs; Reproduced with permission from Ref. [71], Copyright 2022, IOP Publishing Ltd. b Different pore-structured PCL scaffolds fabricated by MEW; Reproduced with permission from Ref. [73], Copyright 2019, American Chemical Society. c Composite scaffolds composed of MEW PLLA lattice and gelatin/genipin/bioglass hydrogel; Reproduced with permission from Ref. [66], Copyright 2022, Elsevier. d PCL fibers modified with mineral nanostructure; Reproduced with permission from Ref. [126], Copyright 2020, Wiley–VCH. e PCL fibers modified with CaP; Reproduced with permission from Ref. [73], Copyright 2019, American Chemical Society
Recreating the biomimetic structure of bones holds immense potential for treating bone injuries and diseases. Researchers have demonstrated that pore sizes of 100 μm and larger are conducive to bone formation. This is attributed to their high global stiffness, local fiber stiffness, superior seeding efficiency, ability to maintain a spread cellular morphology, and significant enhancement of human mesenchymal stromal cells (MSC) collagen and mineral deposition [75, 118]. However, pore sizes exceeding 500 μm also enhance vascularization and bone formation through their large surface areas and excellent interconnectivity [119]. High specific surface areas offer numerous benefits, including optimal cell attachment, increased protein adsorption sites, potential enhancement of the ligand density available for cell binding, and improved oxygen delivery [120,121,122]. Notably, offset and 250 μm homogeneous scaffolds can provide these capabilities (Fig. 4b). In addition, porous scaffolds with a porosity-graded architecture serve as an excellent template for mimicking the natural bone ECM [73].
Significant progress has been made in recent years in the design and biomaterials employed for bone regeneration. Various biomaterials, such as PCL, PLLA, and poly(lactic–co-glycolic acid) (PLGA), have been dedicated to fabricate bone tissue engineering scaffolds. The application of PLLA scaffolds is limited due to their lack of flexibility and relatively low biological activity [123]. To overcome these limitations and create natural bone ECM environment, Xu et al. fabricated a MEW PLLA lattice/gelatin/bioglass reinforced scaffold with a controllable interconnected porous architecture (Fig. 4c). This composite scaffold exhibited an enhanced storage modulus and an exceptional ability to form apatite crystals in vitro [66].
4.2.2 The Combination of MEW Scaffolds with Bioactive Materials to Promote Bone Repair
Recreating the mineral environment of bone holds immense potential for treating bone injuries and diseases. Studies have shown that the mineral structure of bone consists of needle-shaped mineral units that aggregate to form an irregular 3D nanoscale structure. These unique topographies endow bone with exceptional biological and mechanical properties [124].
To achieve this nanoscale structure and promote bone healing, nHA was coated onto various scaffolds (Fig. 4d). The resultant scaffolds promote osteogenic differentiation and polarize human macrophages toward an M2 phenotype, thereby enhancing the production of anti-inflammatory cytokines, such as IL-10 [125,126,127]. In addition, calcium phosphate (CaP), another common biomaterial that resembles the inorganic mineral phase crystals found in bone, can modulate osteoclast and osteoblast activity by releasing calcium and phosphorus ions (Fig. 4e) [73]. The modification of these biomaterials created an environment conducive to bone growth which improves the infiltration, proliferation, and attachment of osteoblasts on the scaffolds.
The chemical precipitation method is a prevalent technique for modifying the mineral composition on the surface of MEW scaffolds by integrating PCL fibers and functional materials. For instance, to circumvent surgery-induced bone infection, which may hinder bone repair and regeneration, ROX was incorporated into fibrous scaffolds. The results demonstrated that PCL/PEG/ROX scaffolds exhibited an initial burst drug release followed by sustained long-term release, which is beneficial for preventing and treating bone infections [62].
Furthermore, composite scaffolds can enable effective biomimetic functionalization [128]. To enhance the flexibility and relatively low biological activity of PLLA, Xu et al. combined the MEW PLLA scaffold and gelatin/bioglass underwent chemical changes, with the amino (NH3+) groups of the gelatin molecular chain serving as active sites for the formation and induction of nHA crystal growth. As the COO−/Ca2+/COO− ionic bonding extended, phosphoric ions (PO43−) were incorporated to promote the nucleation of nHA [66].
4.3 Cartilage Tissue Engineering
Articular cartilage is capable of bearing high and regular loads, which can effectively protect the ends of articulating joints for decades under normal physiological conditions [129, 130]. However, damage to articular cartilage accelerates joint degeneration and leads to osteoarthritis, but its regenerative capacity is extremely limited [131, 132]. The construction of engineered artificial cartilage analogues primarily requires two essential prerequisites: simulating the regional structure of natural cartilage (the superficial layer, transition layer, and deep layer), and the incorporation of cartilage growth factors [133]. The former endows the tissue with load-bearing capacity through collagen fiber arrangement and proteoglycan distribution, while the latter promotes the differentiation of chondrocytes and tissue formation [134,135,136].
Currently, cell-based strategies are the most common methods for treating articular cartilage defects. However, they often lack suitable structural support and result in poor long-term implantation performance [137]. MEW has been widely used to construct a network that supports cell proliferation and differentiation, effectively recapitulating the structure of natural cartilage and enabling it to support regular load-bearing activities [138].
4.3.1 MEW Fibrous Scaffolds for Cartilage Implants
To improve the structural organization of engineered cartilage tissues, researchers have combined the benefits of cellular self-assembly with MEW printed fibrous scaffolds [139]. These scaffolds can be used to guide the deposition of a biomimetic collagen network within engineered tissue, mimicking the structure and function of native articular cartilage. Furthermore, several key metrics are considered to evaluate the mechanical properties of these scaffolds: the equilibrium modulus, dynamic modulus, compression modulus, and Young's modulus [140].
Guidance on these mechanical properties provides a guideline for researchers in the field of engineered cartilage analogues. One such innovation involves the use of MEW scaffolds to enhance the mechanical properties of hydrogels (Fig. 5a). This innovative strategy synergistically integrates the hydrogel that contribute to regional regulation of cell fate and MEW fibrous scaffolds-facilitated modulation of adhesion, proliferation, and differentiation. Notably, this approach demonstrates considerable potential for alternative implants and achieves native cartilage-like mechanical properties [64]. The advantage of MEW can also be used to fabricate sophisticated bi-layered microfibrous architecture which can duplicate the zonal mechanical properties of articular cartilage (Fig. 5b) [141]. Moreover, hybrid bioprinting provides a versatile and scalable approach for biological joint resurfacing, in which inkjet printing was used to deposit numerous cells into MEW scaffolds that can self-assemble into an ordered array of spheroids, finally generating a hybrid tissue (Fig. 5c). By employing this method, the scaffold exhibited impressive equilibrium moduli, reaching 180 kPa and 214 kPa at 20% and 30% strain, respectively. These values are indicative of the compressive stiffness of the tissue's solid matrix and are comparable to those of native tissue (0.2–2 MPa). The tensile Young's modulus of the engineered cartilage was 1.4 MPa, approaching the range of native cartilage (5–12 MPa) [76]. Furthermore, the porous structure of these scaffolds can serve as a cell collector, facilitating the construction of cartilaginous scaffolds and enhancing gene expression associated with chondrogenesis [77].
MEW scaffolds for cartilage tissue engineering. a Integrated tri-layered composite scaffold was constructed using the UV-assisted, stepwise infiltration and crosslinking procedures. As shown, with the assistance of UV-crosslinking, the cell and growth factor-laden GelMA hydrogel precursor solution of the respective layer was successively infused into the fiber networks for the B, D and S layers to construct the tri-layered fiber-reinforced hydrogel composite; Reproduced with permission from Ref. [64], Copyright 2020, Elsevier. b Design approach and sequential MEW of the bi-layered fibrous scaffolds; Reproduced with permission from Ref. [141], Copyright 2019, Elsevier. c MEW and inkjet bioprinting were combined into a sequential biofabrication framework: a defined number of MSCs were ink-jetted into box-like MEW scaffolds, which supported spontaneous cellular aggregation within each microwell; Reproduced with permission from Ref. [76], Copyright 2022, Elsevier. d Schematic illustration showing covalent immobilization of TGF on microfibrous scaffolds; Reproduced with permission from Ref. [74], Copyright 2022, Wiley–VCH
4.3.2 Functionalization of MEW Scaffolds with Chondrogenic Factors
Cartilage growth factors hold immense potential in promoting cartilage formation and development. The incorporation of chondrogenic growth factors within bio-scaffolds architectures enables promote the formation of cartilage while recapitulating the multifaceted functional characteristic of native cartilage.
Cytokines, such as transforming growth factor-β (TGF-β), exert a profound influence on chondrogenic differentiation throughout developmental stages, maintaining chondrocyte quiescence and preventing hypertrophy. Thus, the strategy of immobilizing TGF-β onto MEW PCL polymers has shown promise in supporting chondrocyte differentiation and facilitating cartilage tissue formation. In a previous study, a novel atmospheric-pressure plasma (APPJ) surface treatment method was explored, which allows for the covalent immobilization of TGF-β onto 3D-printed polymeric meshes. This surface modification technique enhances protein loading efficiency and bioactivity while preserving mesh strength and fostering chondrogenic differentiation (Fig. 5d) [74, 142].
In addition, the combination of MEW fibrous scaffolds with hydrogels can significantly enhance the mechanical properties of hydrogels. Burdick et al. increased the modulus of soft norbornene-modified hyaluronic acid (NorHA) hydrogels by approximately 50 times using MEW scaffolds with a spacing of 400 µm while retaining the chondrogenic potential of the hydrogels. Moreover, NorHA hydrogels demonstrated the capacity to enhance the cartilage formation and maturation [79].
4.4 Vascular Grafts
CVDs remain the leading cause of death worldwide, and the therapeutic schedule of CVDs mainly relies on the implantation of vascular grafts. The fabrication of vascular grafts should replicate the native vascular structure to facilitate cell penetration and possess suitable mechanical performance. Mimicking the mechanical properties of native blood vessels is crucial for maintaining graft patency and preventing hemodynamic failure [46]. A mismatch in radial mechanical compliance, particularly at the anastomosis site, can lead to intimal hyperplasia, ultimately resulting in graft failure. Furthermore, in the design of double-layer scaffolds, the inner layer can serve as an antithrombotic barrier, and the outer layer promotes the directional migration of cells and signal transmission through controllable fiber arrangement.
MEW is extensively utilized in the biofabrication of vascular grafts, particularly those with small diameters [143]. The fabrication of tubular grafts depends on cylindrical mandrels, deposited fibers are collected on a rotating mandrel with linear slide translation (Fig. 6a), tubular structures of various diameters can be fabricated using different-sized mandrels. The ability to fabricate tubular grafts with a low pore size and a high number of pivot points by controlling process parameters has been previously demonstrated (Fig. 6b) [34]. Moreover, achieving suitable printing parameters is essential to ensure the highest possible accuracy in pattern fabrication, such as minimizing jet lag by adjusting the printing speed to the CTS [144].
MEW scaffolds for vascular grafts. a MEW vascular grafts fabricated by rotating mandrel; Reproduced with permission from Ref. [34], Copyright 2018, Elsevier. b Tube with a winding angle of 20° and 30 pivot points, with an internal diameter of 1.5 mm and wall thickness of approximately 200 μm; Reproduced with permission from Ref. [34], Copyright 2018, Elsevier. c Micrographs of printed tubular scaffolds for each of the six patterns. Scale bar = 2 mm; Reproduced with permission from Ref. [48], Copyright 2022, Wiley–VCH. d Morphology and geometric dimension of tubular PCL scaffolds (4.0 mm inner diameter) produced via MEW with varying architectural arrangements; Reproduced with permission from Ref. [46], Copyright 2023, IOP Publishing Ltd
The MEW vascular grafts exhibit high reproducibility, minimal variation in fiber diameter, and negligible pore spacing deviation [46, 81]. Moreover, the MEW technique can adjust the mechanical properties of scaffolds by optimizing the microstructure of the scaffold. Previous research has established the impact of microstructure on the overall mechanical attributes of these grafts, such as the diamond, auxetic, closed cell (straight and wavy), and open cell (straight and wavy) (Fig. 6c) [48].
To evaluate the mechanical properties of vascular grafts, longitudinal and circumferential tensile tests are employed to assess uniaxial mechanical properties, facilitating the derivation of graft strength and elastic modulus based on the stress–strain response. These indirect testing methods can be used to estimate equivalent burst pressure, as the high porosity of MEW scaffolds renders them unsuitable for direct bursting pressure methods [145]. In addition, human arterial vessel burst pressures range from 1200 to 4200 mmHg, while arterial compliance ranges from 8.0% to 17.0% per 100 mmHg, these indexes can be used as reference standards for the design of vascular grafts [146, 147].
Grafts exhibiting better compliance tend to possess superior dynamic performance. The diamond and auxetic structures are typical structures with the above properties. The auxetic structure demonstrates a negative Poisson's ratio with bidirectional stretching, whereas the diamond structure undergoes radial contraction. In addition, significant differences are observed between open-cell (straight and wavy) and closed-cell structures in terms of elastic modulus. The closed-cell structure boasts higher stiffness with an elastic modulus of several MPa, compared to the open-cell structure, which falls within several kPa. Furthermore, the open-cell structure has a better deformability at the direction of longitudinal (up to 150%) compared to closed-cell structure. This exceptional deformation capability of the open-cell (straight and wavy) structure stems from the offset design of fibers parallel to the loading axis, with wavy fibers playing a less significant role compared to straight fibers.
Furthermore, pore spacing influences the overall mechanical performance of the grafts. Klein et al. investigated the production of highly porous, biocompatible microfibrous scaffolds featuring a sinusoidal microstructure (Fig. 6d). Their findings revealed that scaffolds with small-diameter sinusoidal structures and 250 μm spacing exhibited mechanical properties akin to native blood vessels and encouraging initial cell culture results. A similar trend was observed for strength and failure, suggesting enhanced mechanical performance in scaffolds with reduced pore spacing [46].
4.5 Ligament and Tendon Tissue Engineering
Tendon and ligament injuries are common worldwide, yet treatment remains challenging due to limited donor availability, healing complications, and frequent reinjury risks [148]. This drives the need for biomaterial-based replacements that can effectively restore damaged tissues. These connective tissues exhibit unique mechanical features including nonlinear stretching properties, highly aligned cell patterns, and wavy collagen fiber characteristics that are difficult to reproduce [82]. Collagen fibers serve dual roles by supporting tissue regeneration while maintaining mechanical stability, enabling the tissues to endure cyclic stretching during movement.
Recently, Groll et al. fabricated a robust, elastic, and creep-resistant sinusoidal scaffolds utilizing MEW, which exhibits enhanced mechanical properties after UV cross-linking (Fig. 7a) [42]. In addition, the scaffold has demonstrated cytocompatibility and the ability to mimic natural ligaments and tendons, the sinusoidal structure significantly enhanced the flexibility of this scaffold, enabling it to manage the loading and unloading of collagen fibers during movement (Fig. 7b) [82]. Meanwhile, the geometric morphology of the fibers can guide cells to align along the fiber path [149]. To further construct a biomimetic tendon structure, researchers have developed a multiscale scaffold with a hierarchical structure, which consists of a shell and a core (Fig. 7c). This biomimetic tendon exhibited strong tensile strength and the ability to accelerate the regeneration of tendon tissue [150].
MEW scaffolds for ligament and tendon tissue engineering. a Tensile testing of straight and sinusoidal scaffolds with and without cyclic preloading; Reproduced with permission from Ref. [42], Copyright 2018, Elsevier. b Scaffolds with crimped fibers before and after rolling; reproduced with permission from Ref. [149], Copyright 2017, Elsevier. c Multiscale structures of biomimetic tendon scaffolds before and after rolling; Reproduced with permission from Ref. [150], Copyright 2024, Wiley–VCH
4.6 Periodontal Regeneration
Periodontal disease refers to inflammatory pathologies affecting the periodontium (the specialized tissues supporting teeth), which arise from dysregulated host immune responses triggered by pathogenic microbial biofilms adhering to dental surfaces. Further loss of periodontal ligament (PDL) fibers may lead to loss of alveolar bone support after the destruction of connective tissue [151, 152]. Due to the complexity of periodontal tissue structure, tooth-supporting tissue regeneration is a complex clinical challenge [153]. In recent decades, significant efforts have been devoted to mimic the 3D structure of ligament-to-bone interfaces. This structure offers the advantage of hindering the migration of epithelial and connective tissues while providing an optimal environment for cell adhesion, proliferation, and osteogenic differentiation [154].
MEW scaffolds have been widely explored to replicate the 3D nanofibrous ECM of the PDL and bone compartment. The matrix of the PDL and bone compartment is characterized by a parallel fibrous structure and pore size gradients (Fig. 8a) [83, 155]. The fibrous scaffolds facilitate cell and collagen alignment, relying on topographical cues and the fibrous structure to provide the necessary conditions for accelerated periodontal regeneration [80]. To reconstruct the interfaces of PDL and bone compartment, Malda et al. designed a novel graded scaffold comprising three distinct zones: a bone zone, an interfacial zone, and a PDL-to-bone zone (Fig. 8b). These graded composite scaffolds hold great application prospects in periodontal regeneration [63].
MEW scaffolds for periodontal regeneration. a MEW fibrous scaffolds with distinct fiber configuration and highly ordered architectures; Reproduced with permission from Ref. [83], Copyright 2023, KeAi. b MEW scaffolds of the crosshatch random structure, highly aligned PCL fibers for the PDL zone, 20MgP fibers for the bone zone, and interfacial crosshatch and PDL zones from top and side views, respectively; reproduced with permission from Ref. [63], Copyright 2023, American Chemical Society. c Surface morphology of PCL scaffolds coated with CaP layer; Reproduced with permission from Ref. [84], Copyright 2023, KeAi. d MEW scaffolds after F/CaP coating; reproduced with permission from Ref. [85], Copyright 2021, Wiley–VCH. e Single fiber of PCL and 20MgP (MgP particles identified with white arrows); Reproduced with permission from Ref. [63], Copyright 2023, American Chemical Society
Different functional materials were coated on the scaffolds to enhance the osteogenic capacity. For example, CaP is widely employed to stimulate rapid bone regeneration (Fig. 8c). Given that tooth tissue comprises cementum and alveolar bone, materials that facilitate bone repair are equally applicable in dental contexts [84]. Notably, fluorinated calcium phosphate (F/CaP)-coated MEW PCL fibrous scaffolds not only expedite bone regeneration and foster concurrent regeneration of tooth–ligament–bone interfaces but also exert significant impact on Porphyromonas gingivalis (Fig. 8d) [83, 85]. Furthermore, the osteoconductive materials with the chemical composition can enhance bone regeneration. Among them, MgP bioceramics can promote bone regeneration ability (Fig. 8e) [63].
Given the microbially abundant oral environment and the scarcity of antibacterial biomaterials in clinical practice, Zhang et al. introduced an innovative method to encapsulate copper (Cu) within mesoporous silica nanoparticles (MSNs) to form a Cu-loaded MSNs (Cu@MSNs) bilayer scaffold. The resultant scaffold prevented interference from non-osteoblastic cells and offered the dual advantages of supporting bone ingrowth and inhibiting bacterial infections [154].
4.7 MEW Scaffolds for Nerve Repair
Peripheral nerve injury has always been difficult and challenging in clinic because of its complex physiological structure and uncontrollable pathophysiological process [156]. Nerve tissue damage, particularly when exceeding 5 mm in severity, can impair this conduction ability and hinder self-regeneration [157]. Autologous nerve transplantation has drawbacks, such as limited donors, donor function loss, neuroma formation, nerve deformation or dislocation, and mismatch in nerve diameter. In addition, acellular allograft requires long-term immunosuppressive therapy, which may increase the risk of infection [158]. Consequently, the implantation of engineered nerve graft substitutes has emerged as a promising alternative for peripheral nerve repair.
Nerve tissue comprises highly oriented and densely packed cellular fiber bundles [159]. The Schwann cells (SCs) are arranged in strings along the axons and envelop them, which facilitates the rapid conduction of nerve impulses due to their unique distribution [160, 161]. Consequently, the current research focuses on establishing connections between damaged nerve tissues to facilitate repair. Yao et al. fabricated a microfibrous scaffold with cell-scale spacing (5–20 μm) by MEW. The biomimetic design can significantly promote the cell migration speed and the growth length of neurites (increasing by 40%). This optimized micron-scale fiber arrangement provides a new strategy for the design of neural tissue engineering scaffolds by offering biomimetic ECM topological cues, reducing the mechanical resistance of cell movement, and guiding the directional extension of neurites simultaneously (Fig. 9a) [52].
MEW scaffolds for nerve repair. a Morphology of dorsal root ganglions cultured for 7 days on PCL microfibers with different fiber spacings at day 7; reproduced with permission from Ref. [52], Copyright 2023, Wiley–VCH. b Schematics of the preparation of injectable Au–PCL scaffolds with high conductivity and high-throughput stimulation device for nerve cells; Reproduced with permission from Ref. [86], Copyright 2020, Elsevier
Furthermore, as neural cells are excitable, their functions are significantly influenced by external electrical stimulation, which serves as a crucial factor in enhancing neural cell differentiation and nerve regeneration [162]. Chen et al. coated a PCL fibrous scaffold with a conductive gold nanolayer, significantly enhancing scaffold’s conductivity. The conductive scaffolds showed the potential to promote the formation of the ECM and stimulate neurite growth (Fig. 9b) [86]. In addition, O’ Brien et al. fabricated a hybrid scaffold consisting of neurotrophic ECM and PPY coated PCL scaffolds, which can direct axonal regrowth across the injury lesion [163]. Notably, MEW not only provides a versatile platform for fabricating biomimetic scaffolds that support the survival and functional behavior of electroactive nerve cells but also enables new opportunities to study neural network behavior and nerve repair mechanisms in vitro [87, 88].
5 Challenges and Future Perspectives
In this review, we have summarized current design strategies and fabrication processes for producing tissue engineering scaffolds using the MEW technique. Owing to its unique ability to replicate native tissue microarchitecture with high fidelity, MEW has emerged as a valuable biofabrication tool in regenerative medicine. Continued development of MEW now focuses on overcoming application-specific challenges and expanding its functional potential across diverse tissue types.
5.1 Cardiac Tissue Engineering
In cardiac tissue engineering, the design of MEW scaffolds should possess suitable mechanical properties, multilayer aligned structure, and suitable conductivity. However, the designs of MEW scaffolds in recent years have been unable to meet all these requirements simultaneously. In future research, these requirements need to be taken into consideration. For instance, oriented fibers can be printed on the surface of the anisotropic fibrous scaffolds to promote cell orientation and structure-driven processes [164]. While researchers have developed scaffolds with diverse microstructures to enhance deformability, these designs remain limited to two-dimensional deformation. Achieving scaffolds that can accommodate complex, synergistic deformation with the beating heart remains a significant challenge. Furthermore, PCL remains the primary biomaterial used in MEW, but its high stiffness can hinder mechanical signal transmission from beating CMs to the scaffold. Thus, developing new, compliant materials suitable for MEW is a critical direction. In addition, effective strategies for modifying scaffolds with bioactive, conductive materials are needed to replicate the electroactivity of native cardiac tissue. Low cell seeding densities also impair the functional integration of CMs with the scaffold, further limiting synchronous contraction. Thus, future research should focus on the integration of MEW scaffolds with cell-laden hydrogels. Addressing these challenges will be essential for advancing MEW in cardiac applications.
5.2 Bone Regeneration
Future work in bone tissue engineering should focus on designing scaffolds that mimic the hierarchical and anisotropic architecture of native bone, which is crucial for promoting regeneration and vascularization [165]. Moreover, the extensive existing vascular networks can promote nutrient transportation and further enhance bone repair and regeneration. Thus, the integration of vascularization into MEW scaffolds can further enhance the biomimetic function and provide a more efficient regenerative capability for the bone defect area. Another interesting research area is about the formation mechanism of nHA during the bone repair process. Future research can utilize the fibrous scaffold constructed by MEW to study this process. Among them, the incorporation of piezoelectric materials is a promising research topic, as piezoelectric nanoparticles can effectively promote the differentiation, migration, and bone regeneration of osteoblasts [166].
Notably, the clinical translation still remains the current focus of research. Although experiments in vitro can demonstrate the biocompatibility of MEW scaffolds, further research on bone tissue engineering scaffolds still needs to pay attention to the interaction between native tissue and the MEW scaffolds. Specifically, when MEW scaffolds are implanted, they can undergo deformation within the defect and compression from surrounding muscle tissue. MEW scaffolds should mitigate this deformation and support new bone formation.
The investigation of surface topography represents another area for future exploration. Appropriate surface topography cues can enhance hydrophilicity, which profoundly influences the biological interactions between the surrounding environment and biomaterials. This includes the adhesion and adsorption of proteins, cell behavior, and ultimately the process of osteointegration. Furthermore, hydrogels possess excellent drug loading and controlled release capabilities. Therefore, the combination of fibrous scaffolds and hydrogels is a promising candidate material for treating bone defects [167].
5.3 Cartilage Tissue Engineering
Current MEW scaffold designs for cartilage repair mainly focus on mechanical properties and growth factor delivery, with insufficient consideration given to the curved surface structure inherent to native cartilage—a critical feature for load-bearing function [168]. Furthermore, conventional polymeric fibers often yield stiff and non-compliant constructs that fail to mimic the viscoelastic behavior of cartilage, potentially causing damage to adjacent joint tissues. While PCL offers good biocompatibility and printability, its slow degradation rate and limited biological activity restrict its effectiveness. Despite these promising directions, there remains a pressing need for more refined large-animal studies that closely replicate human physiological conditions to validate the translational potential of such innovative scaffold designs.
5.4 Vascular Scaffolds
In vascular tissue engineering, variability in scaffolds design and implantation techniques remains a major challenge. Future efforts should aim to replicate native vascular architecture by incorporating layer-specific, directional, and multi-cellular designs [169]. Moreover, further research is necessary to optimize the reendothelialization of MEW scaffolds to create a smooth and continuous surface akin to a natural lumen. Strategies such as incorporating ECM components can enhance mechanical performance, but ensuring sufficient sealing of scaffolds remains a persistent challenge. In terms of mechanical performance evaluation, simulating pulsating flow mechanical stimulation through dynamic culture is valuable for assessing the performance of the scaffold before conducting in vivo tests for long-term patency.
5.5 Ligament and Tendon Tissue Engineering
MEW has shown unique potential in the fabrication of ligament and tendon scaffolds. However, the clinical translation remains difficult due to the need for hierarchical fibrous structures that replicate both nanoscale and macroscale features, including the complex ligament–bone interface. Further studies may focus on the combination of MEW and electrospinning to construct the multiscale structures [170]. In addition, ligament and tendon regeneration require scaffolds that can mimic the in vivo mechanical loading environment. Static culture conditions are often insufficient to guide the directional differentiation of cells. Integrating 4D printing with stimuli-responsive materials (e.g., pH- or temperature-sensitive polymers) may offer a route to dynamic, biomechanically responsive scaffold systems.
5.6 Periodontal Regeneration
The construction of the ligament-to-bone interface in periodontal tissue is a critical challenge. The challenges of multi-tissue regeneration necessitate careful consideration of tissue structure, cell type, and soft-to-hard tissue interfaces. Another significant concern is the risk of bacterial infection during surgical treatment, which can lead to oral diseases caused by bacteria in the periodontal tissues. Therefore, future studies should explore the incorporation of antibacterial materials into the biomimetic scaffolds to address potential bacterial infections during surgery. Moreover, modifying the wetting properties of the surface of the fibrous scaffold through surface treatment can affect the antibacterial ability of the scaffold [171]. Large animal models simulating chronic periodontal inflammation are also necessary to evaluate the scaffold performance under clinically relevant conditions.
5.7 Nerve Repair
Although MEW technology has broad application prospects for neural repair, it still faces the following challenges. One challenge is to construct neural structures that are close to native neural structures. The nerve conduit requires an internal hollow structure to guide the growth of axons [172]. Therefore, future research needs to construct a hollow 3D structure using a rotating mandrel collector. Another challenge is the implantation of nerve scaffolds which are difficult to achieve interface and signal connectivity with natural tissues in vivo. Hence, the future research should focus on optimizing the design of neural interfaces to boost the efficiency and precision of signal transmission. Meanwhile, integrating interdisciplinary research findings to drive technological innovation is crucial. Artificial intelligence or machine learning technologies can be incorporated to improve their efficacy and accuracy in neural repair. In addition, it is imperative to strengthen the bridge between basic research and clinical application, ensuring that the latest scientific advancements are swiftly translated into clinical practice.
Moving forward, the advancement of MEW in tissue engineering will rely on the convergence of material science, biofabrication technologies, and systems biology. Rather than solving isolated engineering challenges, future progress will depend on designing integrated, adaptable platforms that respond to biological complexity and therapeutic demands. Collaboration across disciplines—including regenerative biology, data science, and clinical medicine—will be essential to translate MEW from precision manufacturing into functional, patient-specific solutions. In the future, MEW is positioned not just as a fabrication tool, but as a strategic enabler in the development of next-generation regenerative therapies.
Data availability
Data availability is not applicable to this article as no new data were created or analyzed in this study.
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Acknowledgements
This work was supported by the National Natural Science Foundation of China (Grant Numbers: 12402377, 12272253), the Fundamental Research Program of Shanxi Province, China (Grant Numbers: 202403021223002 202203021212215, 20210302123132). J.C. acknowledges support from the UCL GEF (574696). Furthermore, the authors also acknowledge the use of BioRender.com for creating part of the figures.
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Jia, J., Lei, Q., Liu, Z. et al. Advances in Melt-Electrowriting of Fibrous Bioscaffolds: A High-Resolution Manufacturing Strategy for Tissue Regeneration. Adv. Fiber Mater. 8, 499–525 (2026). https://doi.org/10.1007/s42765-025-00658-y
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DOI: https://doi.org/10.1007/s42765-025-00658-y