Authors: Jiahao Ye, Lei Ji, Liangle Liu, Keyu Zhou, Rui Zhu, Chuchu Sun, Lanjie Lei, Minghai Dai
Categories: Review Article, Biomaterials, Neuromodulation, Neuralization, Tissue engineering
Source: Materials Today Bio
Authors: Jiahao Ye, Lei Ji, Liangle Liu, Keyu Zhou, Rui Zhu, Chuchu Sun, Lanjie Lei, Minghai Dai
In recent years, a large number of biomaterial-based strategies have been developed for tissue repair and regeneration. Despite these advances, achieving functional recovery of regenerated tissues remains a significant challenge, primarily because of insufficient attention to neuromodulation. This review provides a comprehensive analysis of biomaterial-assisted neuralization approaches aimed at repairing damaged tissues and restoring physiological function. We elucidate the mechanisms underlying neuralized tissue repair through four key the neural response following tissue injury or lesion, neurogenic inflammation and immune modulation, neurovascular coupling effects, and the effects of neuromodulation on stem cell behavior. Notably, smart-responsive and electroactive biomaterials have facilitated neuralization, thereby improving functional integration. Furthermore, this review highlights novel advances in biomaterial-assisted neuralization strategies for tissue engineering. The discussion is organized around four key establishing structural and mechanical foundations conducive to neural regeneration, designing delivery systems for neural modulation, constructing microenvironments for electrophysiological regulation, and developing smart responsive biomaterials that facilitate neuralization. By examining current challenges and future directions, we provide innovative perspectives for the development of next-generation biomaterials to advance regenerative medicine.
Tissue and organ defects arising from traffic accidents, surgical complications, and diseases such as diabetes are a major source of distress for patients worldwide. To address such defects, tissue engineering has become an effective strategy for injured tissue repair [[1], [2], [3]]. Its goal is to construct biologically functional replacements in vitro or in vivo by rationally combining seed cells, scaffolding materials, and bioactive molecules, thereby addressing structural deficiencies and functional impairments resulting from injury or disease [[4], [5], [6]]. Previous research has shown that biomaterials with tunable compositions, structures, and physicochemical properties can mimic extracellular microenvironment, supporting cell differentiation, adhesion, and proliferation, thereby promoting tissue repair [7]. However, because the mechanisms underlying neural modulation remain incompletely understood, its effects on tissue repair are often overlooked. Notably, peripheral nerves are densely distributed in several tissues (such as skin, bone, and heart) and play critical roles in regulating their physiological functions [8,9]. Nerves regulate tissue-resident cell behavior and the immune microenvironment by secreting various neuropeptides and neurotrophic factors (NTFs). Conversely, denervation is often associated with delayed healing and poor tissue integration outcomes [10,11]. Here, we introduced the concept of neuralized tissue engineering. In tissue engineering, the significance of neuralization lies in leveraging the critical roles of the nervous system, particularly the neuromodulation and innervation, to facilitate tissue repair. Therefore, the neuralized tissue engineering we proposed herein specifically refers to an innovative technology that synergistically promotes the repair of damaged tissues, such as the nerves, skin, bone, and heart, through the integration of neurogenic cells, neuroinductive factors, and neuroconductive scaffolds. This approach facilitates recovery across multiple fronts, including nerve regeneration, neurogenic immune repair, neurovascular regeneration, and neural electrical signal transmission.
Neurogenic seed cells, the core component of neuralized tissue engineering, are typically cultured on scaffold materials to form functional units for tissue construction [[12], [13], [14]]. Among neurogenic seed cells, Schwann cells (SCs) play a particularly significant role in neuralization-based tissue engineering [15,16]. They accelerate axonal growth and myelin formation as well as secrete NTFs, thereby facilitating nerve regeneration and improving the microenvironment of damaged or diseased tissues [17]. In addition, stem cells have become a major focus of research owing to their differentiation potential. In particular, adult stem cells have attracted significant attention for their promising clinical applicability [18,19]. Biological scaffold materials serve as three-dimensional (3D) platforms that support cell adhesion, migration, differentiation, and proliferation. Moreover, they simultaneously provide a favorable microenvironment for tissue repair [7]. For neuro-based tissue engineering, key characteristics, such as biocompatibility, degradability, and appropriate pore structure should be present in ideal scaffold materials. In addition, the surface topology and other characteristics of these scaffolds must closely match those of the target tissue, whereas the efficient delivery of bioactive signals remains essential [20,21]. Recent studies have paid more attention to the construction of electrophysiological microenvironments as a creative strategy for promoting neuralization. In this context, scaffolds with electroactive properties are particularly advantageous, as they can more effectively exert neuromodulatory functions [22,23]. To precisely modulate the neuromodulatory behavior of seed cells within a scaffold, both biochemical and physical signals are required. Bioactive molecules, such as NTFs and neuropeptides, can be preloaded into scaffolds for sustained release, thereby supporting neuralization [24]. Furthermore, advances in smart-responsive biomaterials have enabled the incorporation of exogenous physical stimuli, including electrical stimulation, light of specific wavelengths, and ultrasound, all of which play critical roles in neural signaling [25]. Collectively, these strategies contribute to guiding tissue repair and functional reconstruction [26].
Beginning with the key mechanisms of tissue repair in neuralization, this review provides an overview of some novel advances in biomaterials for neuromodulated tissue regeneration. First, we summarize the biomaterials suitable for supporting neuralization, with particular emphasis on electroactive composite biomaterials. Second, we outline strategies for biomaterial-assisted neuralized tissue engineering, including the provision of structural and mechanical support for neural regeneration, the development of pro-neuromodulatory delivery systems, as well as the construction of electrophysiological microenvironments. We also highlight the applications of these biomaterials in nerve, skin, bone, and cardiac repair through their neuromodulatory effects. Finally, we discuss the current obstacles and future prospects of biomaterial-based neuralized tissue regeneration. Although previous reviews have separately explored the applications and significance of biomaterials in promoting nerve regeneration, designing bioelectrical scaffolds, and constructing neuroimmune microenvironments, most have been confined to single domains. Meanwhile, in this review, we integrate these three interdisciplinary areas, synthesizing multiple perspectives on how biomaterials promote tissue repair through neuromodulation. This provides a novel theoretical framework and therapeutic approach for designing scaffolds for functional tissue regeneration.
The repair mechanisms underlying neuralization involve a multidimensional synergy. Following tissue injury, the regenerative response of nerves is critical, relying on axonal regeneration, synaptic reconstruction, and the release of neurogenic factors (including neurotransmitters, neuropeptides, and NTFs) to promote neuronal survival and restore neurological connectivity [21]. Subsequently, neurogenic inflammation and immunomodulation contribute to repair, as neuromodulation improves the inflammatory microenvironment [27]. In parallel, neurovascular coupling, mediated by vasoactive substances released from neurons, regulates local blood flow, ensures the delivery of nutrients and repair factors, and promotes neovascularization [28]. Finally, neural signals modulate cell proliferation, differentiation, and directional migration by activating specific receptor pathways, thereby orchestrating the spatiotemporal balance between tissue regeneration and functional remodeling [29].
Nerve repair after tissue injury or lesions is a complex biological process, with neuronal survival and axonal regeneration serving as key components of functional recovery. In general, repair in the peripheral nervous system (PNS) is relatively efficient, whereas that in the central nervous system (CNS) is limited because of the formation of glial scars by astrocytes, weak intrinsic growth capacity of mature neurons, restricted regenerative potential of oligodendrocytes, and presence of inhibitory molecules [[30], [31], [32]].
SCs, the myelinating glial cells, represent a cornerstone of peripheral nerve regeneration owing to their remarkable plasticity [33]. Following injury, SCs rapidly dedifferentiate and become activated, clear degenerated axons and myelin fragments through phagocytosis, and secrete NTFs and chemokines to promote neuronal survival and attract regenerating axons. They subsequently align to form Büngner bands, which guide the extension of axons, while synthesizing extracellular matrix components that establish a pro-regenerative microenvironment. Ultimately, SCs remyelinate regenerated axons to restore neuroelectric signaling [34,35]. During this process, cytokines derived from SCs, including the leukemia inhibitory factor (LIF), interleukin-6 (IL-6), and interleukin-17B (IL-17B), promote the recruitment of macrophages to enhance clearance of axonal and myelin fragments [36,37]. In addition to SCs, neural stem cells/progenitor cells (NSCs) and mesenchymal stem cells (MSCs) also contribute to neural repair [38]. NSCs have been reported to facilitate the repair of small-sized sciatic nerve defects in rats [39], and both transplantation and mobilization of endogenous NSCs have shown favorable outcomes in spinal cord injury (SCI) repair [40]. MSCs secrete cytokines that provide neuroprotection and differentiate into Schwann-like cells that can replace damaged glia, thereby enhancing axonal regeneration and neuronal survival [41].
After PNS dysfunction caused by tissue injury or lesion, full recovery to the original functional level remains limited, despite the intrinsic ability of axons to regenerate and re-innervate following SC activation. At this stage, NTFs expressed and secreted by cells surrounding the injury site play a multidimensional regulatory role, synergistically promoting repair and functional remodeling of damaged structures by targeting neurons, glial cells, and the microenvironment [42,43]. For example, nerve growth factor (NGF) binds to its high-affinity tropomyosin receptor kinase A (TrkA) receptor. It then activates specific downstream signaling cascades that suppress the expression of apoptosis-associated proteins, thereby reducing neuronal death in injured areas [44]. In addition, NGF enhances the autophagy clearance capacity of SCs for myelin debris through the p75NTR/AMPK/mTOR signaling pathway [45]. Furthermore, neurotrophin-3 (NT-3) significantly contributes to peripheral nerve regeneration after denervation and maintains the repair phenotype of SCs through the TrkC/ERK/c-Jun pathway [46]. For the CNS, NTFs also play a crucial role in repair. Brain-derived neurotrophic factor (BDNF) promotes the differentiation, proliferation, and maturation of oligodendrocyte progenitor cells through TrkB receptor activation, which triggers PI3K/Akt and MAPK/ERK and upregulates myelin-related protein expression. This process mediates signaling interactions between axons and glial cells, ultimately facilitating myelin regeneration and restoring electrophysiological function [47]. In a previous study, Huang et al. reported a successful treatment of SCI in rats and monkeys using small extracellular vesicles loaded with BDNF [48]. Similarly, glial cell line-derived neurotrophic factor (GDNF) can activate Rearranged during Transfection (RET) by binding to its receptor GFRα1, forming a GDNF–GFRα1–RET tyrosine kinase ternary complex. This complex triggers RET dimerization and autophosphorylation, thereby activating downstream signaling pathways to support neuronal survival and axonal growth [[49], [50], [51]]. Owing to this characteristic, it has been widely used in the repair of CNS injury. For instance, Aderson et al. utilized the delivery of GDNF to achieve axonal regeneration in the lesion site of SCI [52]. Ma et al. engineered GDNF-loaded nanoparticles for SCI repair, wherein GDNF effectively promoted the regeneration of corticospinal tract motor axons and recovery of motor function [53]. Furthermore, Liu et al. combined decellularized spinal cord extracellular matrix with cross-linked GDNF to design a hydrogel for SCI. This approach promoted neuronal regeneration at 4 weeks post-operation, followed by enhanced axonal growth and myelination at 8 weeks, ultimately leading to significant motor function recovery [54].
In addition, neurotransmitters and neuropeptides play key roles in neural repair after tissue injury or lesions by synergistically regulating neuronal signaling, cell survival, and microenvironmental stability. For example, Zhang et al. discovered that vasoactive intestinal peptide (VIP) upregulates NTF levels, enhances neuronal survival, and contributes to the reconstruction of related neural pathways [55]. Kim et al. reported that calcitonin gene-related peptide (CGRP) acts as signaling mediator in glial cells and neuromuscular interactions during facial nerve regeneration [56].
Neurogenic inflammation and immune regulation are the processes through which the nervous system actively modulates inflammatory and immune responses, involving both the regulatory effects of relevant NTFs on the local microenvironment and the interactions between neuropeptides released by immune cells and sensory neurons. NTFs promote tissue repair by establishing an anti-inflammatory, pro-regenerative milieu. For instance, mesencephalic astrocyte-derived neurotrophic factor (MANF) modulates inflammatory environment by inhibiting the JAK/STAT pathway or negatively regulating NF-κB signaling [[57], [58], [59]]. Regarding immunomodulation, MANF has been proven to accelerate tissue repair and regeneration in retina [60]. Furthermore, sensory nerve fibers in the immediate vicinity of immune cells exert local immunomodulatory effects by releasing neuropeptides. Among them, CGRP promotes tissue healing through neuro-immune interactions. By binding to receptor activity-modifying protein 1 (RAMP1) on immune cells, CGRP induces the release of thrombospondin-1 (TSP-1). Subsequently, through the autocrine and paracrine actions from TSP-1, macrophages at the injury site undergo polarization from pro-inflammatory M1 phenotype to pro-repair M2 phenotype, thereby supporting tissue repair [[61], [62], [63]]. In parallel, in the presence of inflammatory factors, CGRP limits excessive recruitment of neutrophils and monocytes, thereby preventing uncontrolled inflammation. Moreover, CGRP promotes neutrophil apoptosis, reduces their persistence, and minimizes chronic inflammatory responses [61]. Collectively, these mechanisms balance the inflammatory responses with the repair process, enabling efficient tissue regeneration under the regulation of the neuroimmune microenvironment. Substance P (SP) is another key neuropeptide involved in wound healing, with well-documented therapeutic effects in diabetic wound models [[64], [65], [66]]. Pro-inflammatory cytokines (such as IL-8 and IL-6) are both locally and systemically expressed in a manner that is enhanced by SP, which initiate acute inflammation in the early stages of injury and facilitate tissue regeneration. In addition, SP modulates macrophage activity at the wound site by promoting a shift from the M1 to the M2 phenotype, thereby accelerating tissue repair and wound closure [65,67].
Nerves and blood vessels cooperate to promote tissue regeneration through complex signaling interactions and synergistic structural mechanisms [28,68]. It has long been recognized that SC-conditioned medium (SC-CM) enhances the migration and proliferation capabilities of induced endothelial cells (IECs), while promoting the secretion of vascular endothelial growth factor (VEGF). This, in turn, facilitates angiogenesis. Zhang et al. co-localized SCs, IECs, and induced osteoblasts (IOBs) within a scaffold to create pre-vascularized tissue-engineered bone. When implanted into rat femoral defects, this construct significantly accelerated bone formation, thus revealing that the process of promoting bone regeneration is significantly influenced by neurovascular coupling [69]. Moreover, NTFs, such as BDNF and NGF, along with neuropeptides, such as SP secreted by neurons, stimulate vascular endothelial cell proliferation, migration, and neovascularization. For example, NGF becomes a pro-angiogenic factor after binding to the high-affinity receptor TrkA. This directly promotes neovascularization and indirectly enhancing it by upregulating VEGF expression. Similarly, elevated SP levels in peripheral blood after ischemic injury promote the movement of angiogenic progenitor cells (PCs) from their usual location, thereby contributing to revascularization [[70], [71], [72], [73]]. Conversely, factors secreted by vascular endothelial cells and pericytes, such as VEGF, support neuronal survival and axonal regeneration, creating reciprocal neurovascular trophic interactions [74,75]. Adhesion molecules further reinforce the physical connection between nerve endings and vascular endothelial cells, guiding blood vessel growth along neural pathways, facilitating the formation of functional microvascular networks, and accelerating tissue repair and functional recovery. Notably, blood vessels also serve as tracks for SCs and axon regeneration, as SC migration along vascular structures and the establishment of regeneration channels are essential for PNS repair and reconstruction [76].
Transforming growth factor-β3 (TGF-β3), expressed in dedifferentiated Schwann cells (dSCs), is vital for the healing of nervous system wounds. The process of tissue repair is promoted by the directed migration of multiple cell types (such as dermal and epidermal cells) when TGF-β binds to its receptor [77]. Similarly, GDNF binds to the GFRA1/RET receptor complex, activating RET signaling. It encourages the growth of stem cells in the bulge of hair follicles, their transformation into epidermal cells, and their migration to wound sites, where they directly contribute to epidermal regeneration [78]. Moreover, Zhu et al. reported positive outcomes from combining SP with epidermal stem cells to treat wounds in patients with diabetes [66]. In this context, one mechanism through which SP enhances wound healing involves stimulating the migration, differentiation, and proliferation of epidermal stem cells, while promoting the adhesion of epithelial cells to the extracellular matrix, thereby facilitating tissue repair [79]. In addition, NGF promotes the healing of skin wounds by accelerating the migration of dermal fibroblasts through the PI3K/AKT-Rac1-JNK and ERK pathways [80]. Collectively, neuromodulation regulates stem cell proliferation, differentiation, and migration to enhance tissue repair. Proliferation expands the pool of cells available for repair, differentiation generates specialized cells to restore function, and migration ensures precise delivery of these cells, thereby supporting the restoration of damaged tissues in terms of both structure and function.
The selection of biomaterials is crucial in tissue repair. as they must mimic the appropriate microenvironment of extracellular matrix to provide structural support for damaged or diseased tissues while promoting cell differentiation, proliferation, and adhesion, ultimately enabling tissue regeneration [81]. Several factors guide material selection. First, biomaterials must be biocompatible with adjustable degradability. Second, scaffolds should be highly porous to facilitate cell infiltration, nutrient exchange, and angiogenesis, with pore size tailored to the target tissue. Third, mechanical properties should match those of the host tissue, ideally achieved through novel fabrication techniques such as electrospinning. Finally, scaffolds should permit functionalization (such as incorporation of bioactive factors, seed cells, or conductive additives) to enhance tissue-specific regenerative outcomes. Notably, 3D biomaterial scaffolds for tissue repair are generally classified into three (1) natural polymers, which provide excellent biocompatibility and promote cell adhesion but exhibit limited mechanical strength [82]; (2) synthetic polymers, which have lower inherent biocompatibility but can be chemically modified to achieve controlled degradation and desirable mechanical properties [83]; and (3) composite biomaterials, which integrate the advantages of synthetic and natural materials to balance bioactivity, mechanical stability, and degradation kinetics [84].
Silk protein, Collagen, hyaluronic acid, and chitosan are commonly used natural polymers in tissue engineering [[85], [86], [87]]. Collagen, a natural structural protein, forms a 3D mesh. Cell migration, adhesion and proliferation are all supported by this. Notably, the natural structural protein can promote angiogenesis through bioactive signaling. As the significant component of the extracellular matrix in skin, tendon, and bone, type I collagen is pivotal in tissue repair [88,89]. Clinically, collagen is processed into dressings, hydrogels, sponges, or composite scaffolds and is widely applied in treating skin burns, bone defects, and cartilage repair. Its excellent biocompatibility, degradability, and low immunogenicity enable effective wound sealing and reduced infection risk as well as facilitate gradual replacement by host tissues, making it an essential carrier in regenerative medicine [[90], [91], [92]]. Silk, primarily composed of fibroin and sericin, also exhibits excellent biocompatibility. Its nanofiber structure closely resembles the natural extracellular matrix. This provides a biomimetic scaffold that facilitates cell adhesion, proliferation, as well as oriented differentiation [93,94]. Simultaneously, it is possible to load silk proteins with growth factors so that controlled release can be achieved. Thereby, this process instigates the activation of cell signaling pathways that regulate tissue repair-related gene expression, suppress inflammation, and induce neovascularization [95]. In addition, silk has tunable mechanical properties and can be processed into nanofiber membranes, hydrogels, and porous sponges according to the requirements of different tissues [96]. Chitosan, which is derived from exoskeletons and shells, is one of the most widely used polysaccharide biomaterials for tissue repair [97]. Its intrinsic hemostatic, antibacterial, and antifungal properties make it highly effective for wound healing and bone regeneration [98]. As a natural component of the extracellular matrix, hyaluronic acid is rich in hydrophilic groups, enabling it to absorb 1000-fold its own weight in water. This strong hydration capacity maintains a moist microenvironment conducive to cell proliferation and migration during tissue repair [99]. Hyaluronic acid also exhibits excellent biocompatibility, minimizing foreign body reactions, and is biodegradable through enzymatic degradation by hyaluronidase, which ensures its natural turnover in tissues [100]. When it binds to receptors such as CD44, it activates multiple signaling pathways that regulates cell growth, adhesion, and migration, for instance, by promoting angiogenesis in vascular endothelial cells and collagen synthesis in fibroblasts [101]. Fragmented hyaluronic acid can also bind Toll-like receptors (TLRs), thereby contributing to immune response regulation.
Synthetic polymers primarily include polyethylene glycol (PEG), polylactic acid (PLA), and polycaprolactone (PCL). With excellent mechanical properties and biocompatibility, PCL is a biodegradable polyester. It is fabricated into 3D scaffolds using melt molding, electrospinning, and other techniques, providing excellent structural support for cell adhesion and other behaviors [102]. Moreover, PCL serves as a delivery vehicle for growth factors or cells, enabling them to be released in a sustained and localized way. This reduces drug side effects, protects grafted cells, and enhances implantation efficiency [103,104]. Although PCL exhibits a long degradation cycle, Yeingst et al. showed that its degradation rate can be significantly accelerated by high-intensity focused ultrasound [105]. PEG, a hydrophilic polymer, can be photochemically or chemically cross-linked to form highly permeable hydrogels with 3D porous structures that create a favorable microenvironment. Its hydrophilic network also facilitates nutrient and metabolite transport [106]. PEG is further used as a carrier for growth factors; functionalization of its terminal groups with targeting ligands can prolong circulation time, enable cell- or tissue-specific delivery, and even achieve intracellular localization at the organelle level, thereby improving drug bioavailability, while minimizing side effects [107]. PLA exhibits diverse mechanical properties, structures, and geometries suitable for biomedical applications and is frequently used to construct 3D porous scaffolds, particularly for bone regeneration [[108], [109], [110]]. However, PLA is highly hydrophobic, which limits its utility in drug delivery, and its low impact toughness restricts its use as an implant material in mechanically demanding environments such as bone grafting. Therefore, improving PLA performance through the incorporation of modified additives or mixing with other polymers is essential for successful application in tissue repair and regeneration [111].
Composite biomaterials are important in tissue engineering because they can be designed as polymer/bioactive filler, polymer/polymer filler, or polymer/metal filler composites, each providing distinct advantages for tissue regeneration [84]. Mechanical reinforcement, for instance, can be achieved by incorporating mineral components into polymer matrices. Studies have shown that 3D-printed scaffolds composed of hydroxyapatite and PCL or poly(lactic-co-glycolic acid) (PLGA) yield highly resilient constructs that promote bone growth and are easy to manipulate during surgery [84]. Owing to its unique ability to form gels and its excellent biocompatibility, alginate is another widely used biomaterial in biomedical and pharmaceutical applications [112]. However, alginate tends to form soft gels with poor mechanical strength in physiological environments, limiting its use in soft tissue regeneration. To overcome this limitation, synthetic polymers (such as PLGA) or natural polymers (such as cellulose and chitosan) are often incorporated into alginate matrices to form composite hydrogels. These modifications significantly improve mechanical properties, while providing synergistic functions that enhance cell adhesion and accelerate tissue regeneration [[113], [114], [115]]. Similarly, the hydrophilic silk glycoprotein is widely used in composite preparation. Although it lacks the strong mechanical properties of silk fibroin, it possesses outstanding antimicrobial, antioxidant, and ultraviolet (UV)-resistant properties. Consequently, silk glycoproteins are often copolymerized or blended with other polymers to improve mechanical strength, while retaining biofunctional advantages [110]. Composite biomaterials are becoming increasingly significant because they mimic the natural composite structure of tissues. Their dual advantages (1) reproducing the complex properties of native tissues through structural design, for example, constructing fiber networks resembling the extracellular matrix and tailoring mechanical properties, and (2) introducing specific functionalities, such as bioactivity, antimicrobial activity, or electrical conductivity. These characteristics make composite biomaterials highly valuable for diverse applications, including tissue regeneration [[116], [117], [118]].
In neural tissue engineering, ideal biomaterials must integrate multiple functions, such as inducing nerve regeneration and promoting the release of NTFs, and composite biomaterials are particularly advantageous in this context. The fibrous topology of the extracellular matrix can be mimicked by natural biomaterials, which in turn guide axonal orientation and provide excellent biocompatibility. However, they are often limited by low mechanical strength and batch-to-batch variability, which restricts their ability to provide long-term structural support. Conversely, synthetic biomaterials allow precise modulation of degradation rates and electro-responsive properties but lack intrinsic bioactive signals [119]. Composite biomaterials overcome these limitations by integrating the bioconductivity of natural components (such as collagen to replicate extracellular matrix features) with the mechanical tunability of synthetic components (such as PCL to enhance scaffold strength). Furthermore, functional modules such as conductive additives can be incorporated to mimic nerve impulse conduction, thereby enhancing neuronal activity, promoting axonal guidance, and supporting electrical signal coupling [[120], [121], [122]]. In this way, composite scaffolds enable the construction of biomimetic fibrous topologies to direct axonal growth, while simultaneously allowing spatiotemporally controlled regulation of the electrophysiological microenvironment and release of bioactive molecules. This multifunctional approach effectively overcomes the limitations of single-component materials by balancing mechanical stability, biocompatibility, and complex signaling requirements.
Biomaterial-assisted neuralization strategies primarily involve the rational design of scaffold structures and mechanical properties to provide a pro-neuroregenerative framework. Such scaffolds guide neuronal orientation, axonal extension, and the migration of supporting cells, thereby mimicking the natural neural matrix microenvironment [123]. Building on this structural foundation, intelligent delivery systems have been engineered to achieve spatiotemporally controlled release of bioactive cues, such as NTFs, neurogenic cells, and therapeutic genes, to precisely regulate biochemical signaling in the local microenvironment and promote neuralized tissue repair [124]. In parallel, the construction of electrophysiologically active microenvironments is essential to support intrinsic neuroelectric signaling. This can be achieved by integrating conductive or piezoelectric additives to endow scaffolds with electroactivity, or by designing interfacially coupled exogenous electrical stimulation devices. These approaches enhance the transmission of physiologically relevant electrical signals, thereby promoting neuronal excitability, increasing axonal conduction velocity, and supporting functional integration of neural networks [120]. An emerging trend is the incorporation of smart-responsive materials into neuralization strategies. These materials can sense and respond to physical or chemical signals from the damaged microenvironment, enabling on-demand release of therapeutic payloads, dynamic remodeling of scaffold architecture, and in situ modulation of conductivity. Such adaptive properties allow precise regulation of the regeneration microenvironment, advancing dynamic and spatiotemporally controlled neural tissue repair [125].
When constructing a physical microenvironment for pro-neural regeneration, biomaterials play a central role in preventing disordered regrowth by providing spatial guidance and channels for axonal extension and SC migration through precisely engineered structural features. Scaffold design has long been inspired by the fibrous architecture of extracellular matrix components in nerve cells. Incorporating nanotechnology further refines these designs, as nanofibers can induce NSC differentiation and promote axonal extension through the contact guidance effect [[126], [127], [128]]. For example, Fan et al. reported that PCL nanofibers integrated with magnetic nanoparticles, combined with a polypyrrole (Ppy) coating, established a favorable microenvironment for neural regeneration under a static magnetic field [129]. Porous nerve guidance conduits (NGCs) are increasingly being recognized as viable alternatives to nerve autografts. Their biomimetic structures can substitute for the epineurium and perineurium of damaged nerves, providing a dynamic physical-biological microenvironment that supports nerve regeneration [[130], [131], [132]]. Engineered NGCs must promote axonal growth from proximal to distal nerve segments as well as minimize host inflammatory responses. Natural biomaterials are advantageous in this context, as they exhibit minimal immunogenicity and can achieve repair outcomes comparable to autologous grafts, thereby facilitating functional nerve regeneration [133]. Modifications in NGCs microarchitecture, such as the ratio of internal surface area to volume, strongly influence axonal growth. Structures with a high internal surface area-to-volume ratio enhance the adsorption density of proteins, promote integrin clustering on neurons, and thereby facilitate axonal extension. However, excessively high ratios may lead to localized inflammation because of protein overload adsorption, ultimately impairing regeneration processes [134]. NGCs design has progressed from simple hollow silicone tubes to sophisticated constructs incorporating electrospinning, topographical cues, cellular inoculation, and bioactive factor modification to promote neural regeneration [135]. Tong et al. designed a novel three-layer biomimetic NGC using a combination of electrospinning and melt-spinning techniques (Fig. 1A) [130]; it featured an inner layer of oriented melt-spun fibers that mimicked the nerve endothelium topology to guide axonal extension and an outer helical layer that provided mechanical stability, collectively promoting nerve regeneration (Fig. 1B and. C) [130]. In recent studies, advances in tissue engineering have been accelerated by the advent of 3D printing, with the field showing significant promise for nerve repair [131,136]. Zhang et al. used gelatin-methacryloyl (GelMA) to fabricate customized NGCs by 3D printing, achieving both peroneal nerve repair and partial preservation of donor tibial nerve function [137]. To further improve conduit performance, Namhongsa et al. combined electrospinning and 3D printing to create PPy-coated electrospun scaffolds using poly(l-lactide-co-ε-caprolactone) as the extruded material and PLGA as the surface layer [138]. This hybrid scaffold exhibited a well-optimized architecture with both nanopores and micropores, whereas the PPy coating improved electrical conductivity and hydrophilicity [138]. Notably, it showed high cytocompatibility, reduced membrane leakage, and minimized necrotic tissue formation, highlighting its potential for nerve repair [138].Fig. 1Application of a novel NGC in neural tissue engineering. A) Schematic diagram of the NGC structure. The inner layer produces fine electrospun fibers, while the outer surface prepares spiral melt-spun fibers. B) Schematic diagram of three-point bending test. NAH and NRH showed no kinking during implantation or in the bent state post-implantation, whereas NR exhibited significant kinking. The NAH consisted of aligned electrospun fibers in the lumen and helical melt-spun fibers on the outer surface. The NRH consisted of random electrospun fibers in the lumen and helical melt-spun fibers on the outer surface. The NR consisted of only random electrospun fibers. C) Cross-sectional staining images with S100 antibody and co-stained longitudinal section images using S100 and NF200 antibodies. S100, Alexa Fluor 488, green; DAPI, blue; NF200, Alexa Fluor 594, red. Adapted reprinted with permission from Ref. [130], based on CC BY License, Copyright © 2025 Wolters Kluwer Medknow Publications.Fig. 1
The mechanical properties of biomaterials are also crucial in regulating cellular activity, influencing differentiation, migration, proliferation, and apoptosis in both the CNS and PNS. For example, rat NSCs preferentially differentiated into neurons and astrocytes on soft methacrylamide chitosan hydrogels (<1000 Pa), whereas stiffer substrates (>7000 Pa) favored oligodendrocytic differentiation, with maximal proliferation occurring at intermediate stiffness (approximately 3500 Pa) [139]. Neuronal differentiation generally favors soft substrates (<1000 Pa), likely reflecting the native brain tissue stiffness (approximately 600 Pa). In contrast, the PNS spans a broader elastic modulus range of approximately 0.5 kPa to 0.5 MPa. SCs exhibit optimal growth on substrates with an elastic modulus of approximately 7–20 kPa but can also tolerate softer or stiffer materials. Thus, as long as the modulus approximates the natural range of peripheral nerves, it does not significantly impair neuralization [140]. These findings emphasize the significance of adapting the mechanical properties of biomaterials so that they match the properties of the target tissue. Such tuning is achieved by adjusting material parameters, such as degree of methacrylation or hydrogel cross-linking density. In addition, optimizing viscoelasticity, for instance, by leveraging the stress-relaxation behavior of dynamically covalently bonded hydrogels, can further reduce resistance to cellular migration and accelerate nerve regeneration.
As a key link between biomaterials, tissue repair, and regeneration, delivery systems are designed to controllably transport bioactive molecules, neurogenic cells, and other neuromodulatory agents to the site of injury or lesion. By precisely engineering carrier structures and release mechanisms, these systems regulate the local microenvironment to promote regeneration. Currently, multifunctional delivery systems based on hydrogels, nanoparticles, microparticles, and gene carriers have shown significant potential. Their core advantages include mimicking extracellular matrix structures, enabling cross-barrier targeted delivery, and providing temporal control over therapeutic release.
NTFs (such as GDNF, NGF, and BDNF) and neuropeptides (such as CGRP and SP) are indispensable components in delivery systems for neural tissue engineering. Thus, identifying suitable platforms for their controlled release is crucial for effective neural repair [141]. Among these, electrospun fiber scaffolds have attracted increasing attention owing to their micron-to-nanoscale structures, high porosity resembling the extracellular matrix, and large surface-area-to-volume ratios [142,143]. For instance, Pi et al. designed a BDNF-loaded aligned conductive scaffold composed of PCL and carbon nanotubes (CNTs). The sustained release of BDNF was achieved for a period of 28 d [144]. Similarly, Gregory et al. incorporated GDNF into aligned PCL nanofibers, significantly enhancing neurite outgrowth in rats and highlighting its potential as a pro-regenerative biomaterial [145]. Shafiq et al. fabricated PCL/collagen type I patches incorporating SP through co-electrospinning, revealing promising outcomes in cardiac repair [146]. In addition to single-factor delivery, electrospinning techniques enable the fabrication of scaffolds with diverse morphologies and architectures tailored for multifunctional delivery. For example, multilayered scaffolds constructed through combinations of emulsion electrospinning, coaxial electrospinning, high-speed electrospinning, and dual-source dual-power electrospinning enable the simultaneous delivery of several bioactive factors, such as NGF and GDNF [147,148].
Owing to their hydrophilic nature and 3D porous structure, hydrogels represent another versatile platform for controlled delivery of NTFs and neuropeptides. By encapsulating bioactive substances, hydrogels can modulate cellular behavior, thereby promoting tissue repair. This can be applied to diverse areas, including bone, neural and skin tissue engineering [[149], [150], [151], [152]]. For example, Torres-Ortega et al. developed a multifunctional hydrogel mechanically compatible with brain tissue, which enabled sustained release of GDNF, while protecting it from degradation [153]. Liu et al. reported that a DNA hydrogel incorporating VEGF/NGF promoted SC proliferation, migration, and myelin formation, highlighting its therapeutic value in neural tissue repair [154]. An et al. designed NGF-loaded hyaluronic acid hydrogel coatings on polyetheretherketone, which allowed gradual degradation and sustained NGF release, significantly enhancing osteogenesis and angiogenesis in bone tissue engineering [151]. In cartilage repair, Kim et al. created self-assembled peptide–SP hydrogels. It reduced inflammation and promoted cartilage regeneration in knee osteoarthritis, with longer intra-articular retention than that of SP alone [155]. Similarly, carboxymethyl chitosan–gelatin hydrogels prepared by chemical cross-linking for SP co-delivery showed efficacy in enhancing wound healing in skin tissue engineering [156]. Notably, release kinetics can be finely tuned. For instance, NGF release from cryogels was directly correlated with the heparan sulfate ratio, allowing customization of cryogels for higher NGF loading and controlled delivery [157].
Microparticles represent another promising platform for neuralized tissue engineering, capable of efficiently encapsulating bioactive substances through structural designs such as porous microspheres (PMs) and functionalized nanoparticles. These systems provide protection, controlled release, as well as targeted delivery of therapeutic agents. Hellenbrand et al. showed that mineral-coated microparticles enabled sustained release of biologically active GDNF and NGF, significantly improving sciatic nerve axon growth and hindlimb function in rats [158]. Similarly, Wang et al. developed CGRP-loaded PLGA nanoparticles that achieved sustained in vivo release, thereby modulating vascular and inflammatory functions [159]. Among NTFs, GDNF stimulates SC proliferation and promotes neuronal survival, whereas NGF supports neuronal growth and survival [159]. However, supraphysiological doses of either factor may induce excessive encapsulation of axons and SCs, thereby impairing regeneration. Conversely, their short in vivo half-life of these molecules limits their potential for single-dose administration. To overcome these challenges, Lackington et al. designed a novel dose-controlled release system by encapsulating GDNF and NGF in PLGA microparticles, which significantly enhanced nerve regeneration [160].
However, in neuralized tissue engineering, these neuropeptides and NTFs loaded onto biomaterials may also excessively activate neurogenic inflammation and immune regulation. This subsequently recruits additional inflammatory cells leading to chronic inflammation, promotes fibroblast activation resulting in implant encapsulation by scar tissue, and disrupts newly formed blood vessels and nerves [161]. Ultimately, this compromises the integration and function of the implanted scaffold. Considering neuropeptides (such as CGRP and SP) as an example, while they possess unique advantages in promoting tissue repair, we cannot deny that their excessive activation of neurogenic inflammation and immune responses may also exert certain negative effects. CGRP has long been established as a central mediator in migraine pathogenesis, capable of inducing dural vasodilation, neurogenic inflammation, and pain signaling [162]. Furthermore, CGRP released from sensory nerves contributes to abnormal cutaneous vasodilation, excessive keratinocyte proliferation, and immune cell infiltration, driving the progression of inflammatory skin diseases [163]. Similarly, synovial fibrosis and hyperalgesia observed in osteoarthritis are inextricably linked to CGRP [164]. The adverse effects of SP primarily arise through the neurokinin-1 receptor (NK-1R). Existing evidence indicates that excessive SP-NK1R signaling contributes to the pathogenesis of various inflammatory diseases and their associated organ damage, such as sepsis-associated lung injury, acute pancreatitis-associated lung injury, burn-associated lung injury, and chronic urticaria-related skin lesions [165]. Beyond its well-known ability to directly upregulate collagen synthesis in fibroblasts via the NK-1R signaling pathway, it has also been demonstrated to induce fibrosis by activating the RhoA/ROCK pathway [166]. Regarding NFTs, it is inevitable that some biomaterials initially used in tissue engineering may undergo rapid and massive release of NGF during the early stages of implantation. This can cause local concentrations to momentarily far exceed the therapeutic window, leading to severe inflammatory reactions and hyperalgesia in the damaged tissue [167,168]. When NGF levels are abnormally elevated, it directly acts on immune cells expressing its high-affinity receptor TrkA, promoting their activation, proliferation, and release of inflammatory factors [169]. Additionally, it can act on sensory nerve endings, enhancing their sensitivity to cause hyperalgesia and inducing neurons to release SP and CGRP, which indirectly activate nearby immune cells [170]. These limitations imposed by neurogenic inflammation and immune responses remain unresolved. However, their innovative significance in achieving neurogenic tissue repair has spurred the development of relevant biomaterials.
Delivering SCs to regulate neural regeneration represents a promising strategy in neurogenic tissue engineering. However, two major challenges must be addressed. First, shear-induced damage often causes extensive cell death when SCs are directly injected. Second, the injected cells typically rapidly disappear from the delivery site, leading to low in situ survival rates [171]. Therefore, identifying an appropriate delivery platform is crucial for enhancing therapeutic outcomes. Hydrogels have emerged as an effective means of delivering SC. Their polymer chains provide shear resistance and protect cells from mechanical rupture, whereas high porosity and permeability support nutrients, oxygen, and cytokine exchange with surrounding tissues. These can promote cell survival and growth [172]. Consequently, hydrogels are considered highly suitable delivery platforms for SCs [173]. For example, Liu et al. encapsulated SCs within 3D GelMA hydrogels for SCI repair, revealing favorable SC proliferation and survival in vivo [174]. In a recent study, microgels derived from natural or synthetic hydrogel materials also showed notable promise as multifunctional cell carriers. Compared to bulk hydrogels, microgels exhibit enhanced porosity, smaller size, and high flexibility, enabling their assembly into aggregates. In addition, multi-cell combination therapies can be supported by microgel-based scaffolds, whereas each cell type is maintained in its own microenvironment [175]. The delivery of SC has also been explored using other biomaterials. For example, Meng et al. showed that chitosan films can guide skin precursor-derived SCs (SKP-SCs) toward a repair phenotype, thereby promoting nerve regeneration (Fig. 2A) [176]. SKP-SC exhibits ordered arrangement and elongated morphology through topographic cues induction, with the 30 μm surface demonstrating superiority over the 10 μm and 50 μm surfaces (Fig. 2B) [176]. Simultaneously, through appropriate topographical cues, SKP-SCs exhibit smaller cell areas, longer extensions, and smaller cell angles, which are conducive to neural regeneration (Fig. 2C) [176]. Moreover, chitin oligosaccharides, intermediate products of chitin degradation, have been shown to induce autophagy in SCs, promote myelin clearance, and enhance peripheral nerve regeneration, highlighting their potential in developing innovative neurogenic cell delivery platforms. Nevertheless, with regard to existing neuralization strategies, achieving neuralized tissue repair through SC delivery still faces a significant cells implanted in neurotrophic scaffolds often exhibit reduced biological activity during the early stages of treatment [177]. A key reason for this is the limited oxygen diffusion and slow vascularization at the injury site, which leads to insufficient oxygen supply for SCs and significantly restricts their ability to support neuroregeneration. Concurrently, researchers have attempted to improve SC survival rates compromised by hypoxia through pre-hypoxic conditioning. While this approach enhanced post-transplantation vascular formation, it failed to significantly improve cell survival rates [178]. Therefore, establishing a pre-built, sustained oxygen delivery system within neuro-engineered scaffolds represents a critical future breakthrough for overcoming limitations in neuralized tissue engineering [179].Fig. 2Application of a chitosan scaffold in cell delivery. A) Schematic diagram of chitosan scaffold delivering SKP-SCs to promote nerve regeneration. B) Optical images and SEM images of micropatterned chitosan films and adhered SKP-SCs. C) Fluorescent microscopy images show the morphology of SKP-SCs after 36 h of culture on different surface substrates. Adapted reprinted with permission from Ref. [176], based on CC BY License, Copyright © 2024 Oxford University Press.Fig. 2
Stem cells, including NSCs and MSCs, are crucial in regulating neural repair and improving the neural regulatory microenvironment [180,181]. In designing delivery systems for NSCs and MSCs, their derived exosomes have attracted increasing attention. Exosomes released from NSCs and MSCs are known to accelerate nerve regeneration and tissue repair [182,183]. Exosomes are nanoscale extracellular vesicles that are secreted by mammalian cells and appear as flattened, spherical structures. They can be readily delivered through the bloodstream or other biological fluids owing to their nanoscale size. Neurogenic cell-derived exosomes can enhance survival and proliferation of NSCs, facilitate their differentiation into mature neurons, as well as promote integration into neural circuits [184]. Moreover, exosomes encapsulate various bioactive molecules, including NTFs, which actively participate in neural repair and regeneration [185]. In biomaterials, exosome-based delivery can be achieved through diverse functional integration strategies. One promising approach is surface grafting, in which exosomes are anchored to biomaterial surfaces through chemical bonding or physical adsorption to impart targeted recognition functions. For example, Lian et al. grafted NGF-stimulated MSC-derived exosomes onto 3D-printed layered porous scaffolds. When applied in a rat bone defect model, these constructs significantly promoted neurovascular structure formation and nerve-innervated bone regeneration [186]. Matrix encapsulation technology directly embeds exosomes within the bulk matrix or porous network structures of biomaterials. This approach preserves exosome bioactivity, while exploiting the intrinsic degradation properties of the carrier material (such as hydrolysis) to achieve sustained and controlled release of therapeutic components. For instance, in cross-linked hydrogels, exosomes can be uniformly dispersed within the hydrogel precursor solution. Upon gelation, a 3D carrier is formed, enabling gradual release of exosomes through diffusion or matrix degradation, thereby supporting long-term delivery at targeted sites [187]. Yang et al. exemplified this strategy by developing a thermosensitive PF-127 hydrogel encapsulating MSC-derived exosomes. Notably, local application in a diabetic rat model facilitated wound healing [188]. Finally, a new strategy is emerging that involves integrating exosomes with nanoparticles. This approach combines the targeted delivery, tunable release profiles, and protective functions of nanocarriers with the inherent biocompatibility and bioactivity of exosomes, resulting in multifunctional composite delivery systems [189].
Gene therapy represents an innovative strategy for manipulating cells to produce therapeutic proteins and precisely regulate the expression and release of bioactive molecules. Adeno-associated viral vectors are widely used for peripheral nerve gene delivery owing to their low immunogenicity, minimal mutagenic risk, and high transduction efficiency [190]. For example, adenoviral vectors have been explored to specifically target SCs, enhancing their neurotrophic activity in injured or diseased tissues and thereby promoting nerve regeneration. Moreover, adenoviruses carrying NTF-encoding genes can stimulate axonal regrowth upon local administration, further contributing to functional recovery. However, this vector shares high homology with wild-type viruses commonly infecting humans, imposing significant limitations on its platform application in terms of immune responses [191]. Studies have confirmed that healthy donors exhibit both humoral and cell-mediated immune responses to wild-type adenoviruses, with anti-adenovirus antibodies demonstrated to, at least, potentially exert a major impact on gene transfer outcomes [191]. Furthermore, in a small number of animal models, integration of recombinant adenovirus vector transgenes into the host genome led to adverse events [192,193]. Although this phenomenon has not been observed in humans, further long-term follow-up data are still required. Our understanding of the safety profile of adenovirus-based gene transfer is now increasingly comprehensive. We have found that while multiple factors influence the overall immunogenicity of adenovirus vectors, the vector dose appears to be a key determinant of immune-mediated toxicity. Gene therapy using adenovirus vectors is relatively safe when lower vector dose thresholds (≤10^12^ vgs/kg/patient) are required [194]. In recent years, non-viral vector-based gene delivery has attracted increasing attention, particularly approaches leveraging nanomedicine [190]. Lackington et al. developed a gene-activated NGC by incorporating non-viral polyethyleneimine–plasmid DNA nanoparticles encoding NGF and GDNF, achieving efficient peripheral nerve repair and regeneration [195]. Electrospun fibers have also emerged as promising non-viral carriers. Zhang et al. fabricated a 3D hydrogel scaffold incorporating electrospun fibers (Fig. 3A) [196]. It provided both topographical guidance for axonal extension and sustained delivery of microRNAs. Collectively, these miRNAs are crucial in axonal regeneration (Fig. 3B and. C) [196].Fig. 3Applications of the 3D fiber‐hydrogel scaffold as a non-viral vector in gene delivery. A) Schematic diagram of scaffold preparation using electrospinning technology. B) Representative fluorescence images of NF200 staining in rats treated with Axon miRs and Neg miR 12 weeks post SCI. C) Representative fluorescence images of Tuj1 staining at 12 weeks SCI. Adapted reprinted with permission from Ref. [196], based on CC BY License, Copyright © 2021 John Wiley and Sons.Fig. 3
In the human body, bioelectricity is a core element for maintaining physiological functions, operating across regulatory networks from cells to tissues. Among its manifestations, cell membrane potential plays a fundamental role in dynamically regulating cell behaviors. For instance, neurons generate action potentials through membrane potential changes, enabling the rapid conduction of nerve signals. The precision of such electrophysiological activity depends on charge transfer and signaling interactions at both the cell–substrate and the cell–cell interfaces. By incorporating electroactive additives into biomaterials to create electroactive scaffolds, diseased or damaged tissues can be induced to generate or respond to local electrical signals. This, in turn, facilitates the construction of an electrophysiologically regulated microenvironment that promotes tissue repair and regeneration [120].
Incorporating conductive additives into scaffold materials enables closer mimicry of the natural bioelectrical microenvironment, efficient charge transfer at the cell–matrix interface, and regulation of intercellular interactions. Even without exogenous electrical stimulation, the intrinsic conductivity of such materials can drive stem cell self-renewal or directed differentiation, thereby facilitating extracellular matrix regeneration [197]. This conduction-based regulatory mechanism adds a novel dimension to tissue engineering in addition to conventional chemical signaling. By reproducing the electrophysiological characteristics of target tissues, electroactive biomaterials with conductive additives can guide neuronal axonal growth, form conduction networks in cardiomyocytes to support synchronized contraction, and enhance recovery after skeletal muscle injury, particularly in neural, cardiac, and musculoskeletal tissue repair [[197], [198], [199]]. Conductive additives generally fall into three polymers, carbon-based materials, and metal/metal oxides, typically incorporated into biocompatible matrices (such as hydrogels) to produce functional composites combining conductivity, biocompatibility, and tunable mechanical properties [120].
Conductive polymer additives can provide electrical conductivity and retain the favorable properties of conventional polymers, such as flexibility and low stiffness. Among them, polyacetylene (PA), polythiophene (PT), polyaniline (PANI), PPy, and poly(3,4-ethylenedioxythiophene) (PEDOT) are the most studied, with promising applications in neural and cardiovascular tissue engineering [[200], [201], [202]]. These polymers have been widely investigated as biomaterials for interfacing with NSCs in brain, spinal cord, and peripheral nerve regeneration. For example, PPy-based substrates enhance VEGF-A production in human NSCs under electrical stimulation. Furthermore, electrical stimulation on PPy substrates has been shown to encourage the differentiation of NSCs. In addition, it is preferable for NSCs to differentiate into neurons instead of glial cells, which can aid in reducing scar formation [203]. Striking a balance between electrical and mechanical properties remains a major challenge when designing conductive biomaterials [200]. Biomaterials with superior electrical properties can promote neuronal development and neural network reconstruction by optimizing the efficiency of electrical signal transmission [204]. Research by Kreysing et al. indicated that transthyretin (TTR) exhibits elevated expression levels on softer substrates. As a key downstream regulator of neuronal maturation following Piezo1 activation, TTR promotes synaptic formation [205]. However, enhancing electrical properties typically relies on introducing highly conductive fillers, such as metals, which tend to increase material hardness. Conversely, overemphasizing the flexibility of biomaterials often compromises electrical conductivity. Notably, processing conductive polymer additives into films and coatings applied over highly biocompatible biomaterials holds promise for achieving a balance between electrical and mechanical properties [206]. PPy exhibits high conductivity (up to 10^3^ S cm^−1^), can be synthesized in large quantities at room temperature in both aqueous and organic solvents, and exhibits good biocompatibility, supporting proliferation, differentiation, and adhesion of diverse cell types in vitro. However, its inherent drawbacks, crystallinity, rigidity, brittleness, and insolubility, remain significant [207]. To address these limitations, PPy is often blended with biodegradable synthetic or natural polymers (such as PLA, PLGA, chitosan, or filipin proteins) to compensate for its mechanical shortcomings [121]. For example, Namhongsa et al. demonstrated that compared to a scaffold without PPy coating, a PLGA scaffold coated with PPy exhibited significantly enhanced conductivity (electrical conductivity of 10.50 ± 0.08 S cm^−1^) while maintaining its original biocompatibility. Additionally, PEDOT is frequently used as a conductive additive to enhance the electrical conductivity of biomaterials [[208], [209], [210]]. For instance, a PEDOT nanocomposite hydrogel designed by Sun et al. simultaneously achieves exceptional stretchability (>800 %) and high conductivity (≈1.25 S cm^−1^) [211].
Carbon-based additives, composed of elemental carbon, primarily include graphene, CNTs, carbon fibers, and fullerenes. Among these, CNTs and graphene are the most applied because of their high electrical conductivity, exceptional mechanical strength, large surface area, and chemical stability [212]. These properties make them effective reinforcing agents in polymer matrices, enhancing both mechanical performance and electrical conductivity in tissue engineering scaffolds. Such improvements promote axonal regeneration and facilitate stem cell differentiation into neurons through neuroelectric signaling [213,214]. Their abundant functional groups and large surface area also support the loading and release of bioactive substances. However, the intrinsic hydrophobicity of numerous carbon-based materials hinders interactions with hydrophilic polymers, highlighting the need for hydrogel systems incorporating these additives. For example, Shin developed a conductive CNT-doped hydrogel that overcame dispersion challenges in aqueous environments, leading to significantly improved human induced pluripotent stem cell-derived neural progenitor cells and neuronal differentiation of human fetal neural stem cells [215]. Meanwhile, a conductive hydrogel developed by Chu et al. achieved exceptional conductivity (9.52 S m^−1^), low tensile modulus (∼100 kPa), and high extensibility (∼1000 %) even with only 0.33 wt% CNTs [216]. When transplanted into a sciatic nerve loss model, this hydrogel significantly enhanced nerve regeneration and accelerated functional recovery. In addition, CNT-based hydrogels strengthen electrical coupling in cardiomyocytes. Owing to their conductivity and ability to mimic the electrical signaling microenvironment, CNTs hold significant promise for cardiac tissue engineering, particularly in myocardial repair [217].
Metal and metal oxide additives are also extensively used owing to the exceptional mechanical strength, electrical conductivity, and fatigue resistance. Bulk metals and their alloys have shown outstanding performance as bone implants bone restoration [218]. Zhang et al. revealed that implant-derived Mg could enhance fracture healing in rats through the induction of local neuronal production of CGRP [219]. In recent studies, metallic biomaterials with micro-/nanostructures have been developed for biomedical applications. Among these developments, gold nanoparticles (AuNPs) are particularly promising for cardiac tissue engineering, as they can be readily synthesized with tunable shapes, sizes, and surface properties, thereby providing adjustable mechanical performance and electrical conductivity [220]. Similarly, Jaswal et al. designed a polycaprolactone (PCL) electrospun composite scaffold incorporating reduced graphene oxide (RGO)-coated AuNPs. This scaffold exhibited unique biological and topological features that effectively mimicked the natural extracellular matrix and promoted neuronal cell functional recovery [221]. In addition, Zhao et al. developed a novel nano-neuroimmune blocker by coating gold nanocages (AuNCs) with glioma cell membranes enriched in transient receptor potential vanilloid 1 (TRPV1) receptors [222]. This biomimetic strategy allowed the nanostructure to neutralize streptolysin S (SLS) toxin, a major virulence factor in streptococcal necrotizing fasciitis, which normally suppresses neuronal signaling and immune function by binding to TRPV1. By competitively blocking SLS, the engineered AuNCs restored CGRP release and alleviated streptococcus pyogenes infection, providing a new therapeutic strategy targeting neuroimmune communication.
However, a recent in vitro study by Carrillo-Romero et al. evaluating the toxicity of carbon-based and metal-based nanomaterials challenged our understanding of the significant advantages of conductive biomaterials in neuralized tissue engineering [223]. In their experiment, ZnO nanomaterials exhibited high levels of cytotoxicity in skin and lung cell lines; Ag nanomaterials showed cytotoxic effects in gastrointestinal cell lines; and CNTs demonstrated similar effects in skin cell lines, albeit with weaker impact. Most studies indicate that despite their many advantages, the unique properties of carbon-based and metallic nanomaterials may also pose significant health risks. These materials can interact with tissue cells, leading to severe cytotoxic damage including cell membrane disruption, oxidative stress, and DNA damage [[224], [225], [226]]. Therefore, rigorous evaluation must be conducted prior to their use. Additionally, in neuralized tissue engineering, the immunocompatibility of carbon-based and metal-based materials also poses a significant challenge. Carbon-based materials primarily overactivate immune cells such as macrophages through their sharp edges, surface chemical properties, and potential residual impurities, triggering persistent inflammatory responses, foreign body giant cell formation, and chronic granuloma [227]. Metal-based materials primarily induce delayed-type hypersensitivity reactions and aseptic inflammation against the prosthesis owing to metal ions released through corrosion or wear particles from the material itself [228]. Notably, Yamaguchi et al. recently discovered that human sialic acid immunoglobulin-like binding lectin-14 (Siglec-14) recognizes CNTs and triggers inflammation [229]. Therefore, we may potentially overcome inflammation caused by carbon nanotube implantation in the future by blocking Siglec-14.
Conventionally, exogenous electrical stimulation has been applied directly to localized tissues, where it promotes neural signaling, activates neurons, regulates stem cell behavior, and facilitates tissue repair and regeneration [230]. However, direct electrical intervention has notable limitations, including poor selectivity and attenuation of voltage at the target site. As a result, increasing attention has been directed toward stimulation strategies based on conductive biomaterials. For instance, Yao et al. designed an implantable electrical stimulation device that can promote fracture healing [231]. Implanted at the fracture site, this device uses a triboelectric nanogenerator (TENG) that converts biomechanical energy from body movements (such as knee bending) into therapeutic biphasic electrical impulses. Its island–bridge electrode configuration and pyramidal microstructured arrays enable stable contact with irregular tissue surfaces, ensuring efficient power conversion. In a rat tibia fracture model, the device significantly accelerated healing, while enhancing bone mineral density (+27 %) and bending strength (+83 %). Similarly, Chen et al. designed an electroactive wound dressing fabricated using electrospinning technology [232]. When combined with exogenous electrical stimulation, the dressing remodeled the immune microenvironment and promoted endothelial and fibroblast cell proliferation and migration. This dual action suppressed early inflammation, while enhancing vascular regeneration and collagen deposition, highlighting its potential in skin tissue engineering.
Electrical stimulation using conductive biomaterials can be delivered in two ways, namely alternating current (AC) or direct current (DC). Among these, DC stimulation is the simplest to implement, as it can be generated using basic battery systems, and it has therefore been more widely applied in biomedical research. Studies have shown that 15 d of DC stimulation accelerated fracture healing in rats compared with controls [233], and subsequent work revealed that this effect was mediated by the modulation of local biological processes [234]. Nevertheless, the most significant drawback of DC stimulation lies in the occurrence of harmful electrochemical reactions at the electrode interface. Under continuous current application, this typically leads to depletion or transformation of electrode materials and generation of new chemical substances [235]. For instance, a strongly acidic environment forms near the anode, where such extreme pH shifts can directly poison cells and cause protein denaturation. Concurrently, the metal electrodes themselves are prone to corrosion and release toxic ions, further causing tissue damage. These interface issues triggered by irreversible Faradaic reactions severely compromise the long-term safety and reliability of DC stimulation technology [236]. To sustain the application of DC stimulation in neuralized tissue engineering, specialized electrode materials are required to reconcile two seemingly contradictory aspects during continuous charge releasing ions through the electrode while employing mechanisms that neither corrode or damage the electrode nor generate toxic concentrations of stimulation byproducts within the tissue [237,238]. AC stimulation enables increased flexibility, as parameters such as waveform, frequency, and pulse width can be tuned for specific applications [120]. Unlike DC stimulation, AC stimulation avoids extreme pH changes and electrode corrosion, but its core adverse reaction lies in local tissue burns caused by the Joule heating effect [239]. Notably, optimized stimulation parameters can modulate signal transduction and neuronal growth in neuralized tissue engineering. For example, PC-12 cells cultured on scaffolds under optimized stimulation (100–150 mV cm^−1^) exhibited enhanced neurite extension and synaptic growth [240]. In contrast, bone tissue engineering often uses electrical stimulation to modulate neuropeptide synthesis and activate nerve cells, typically using different parameters (such as 200 mV·mm^-^1 for 2–8 h/d) [241]. These findings emphasize the significance of tailoring electrical stimulation to the specific regenerative context, paving the way for electrically active biomaterial implants capable of promoting tissue repair with precision.
Piezoelectricity typically arises from the non-centrosymmetric properties of materials, which can be attributed either to crystal structure of inorganic piezoelectric additives or orientation and molecular structure of organic piezoelectric additives. When materials containing piezoelectric additives are exposed to mechanical stress, they containing piezoelectric additives generate electrical signals owing to the unique electromechanical coupling characteristics [242]. For instance, collagen proteins, as natural piezoelectric additives, exhibit distinct molecular structure that confers excellent piezoelectric properties, enabling them to produce electrical signals in response to mechanical forces. These signals are often transmitted through the extracellular matrix to the cell membrane and subsequently to the nucleus, where the translation of bone morphogenetic protein (BMP) and transforming growth factor β (TGF-β) is induced. Both TGF-β and BMP are crucial for cartilage and bone formation, thereby facilitating tissue repair [243]. In general, piezoelectric additives are classified into two inorganic or organic.
Organic piezoelectric additives typically consist of carbon chains or carbon rings and include both synthetic and natural materials. Although synthetic piezoelectric polymers generally exhibit relatively low piezoelectric coefficients, they are well-suited for tissue-engineered biomaterials, particularly neural applications, because of their high flexibility and low stiffness [244]. These polymers can be processed into various structures, including microspheres, nanofibers, hydrogels, and thin films, using techniques such as spin coating, electrospinning, and template methods, which has broadened their application in tissue-engineering biomaterials [245]. Among synthetic polymers, poly(vinylidene fluoride-trifluoroethylene) (P[VDF-TrFE]) and poly(vinylidene fluoride) (PVDF) have attracted significant interest because of their strong voltage electrical responses and excellent mechanical properties [246]. PVDF exists in five primary crystalline forms, with the β phase contributing most significantly to its piezoelectric performance. Optimizing processing parameters can increase the β-phase content, thereby enhancing piezoelectric activity. PVDF combines biocompatibility with piezoelectric activity; in bone regeneration, electrical signals generated by PVDF scaffolds under mechanical stress stimulate osteoblast activity, thereby promoting calcium deposition and new bone formation [247]. PVDF also promotes neuronal elongation, stimulates SCs to secrete NTFs, and drives NSC differentiation into neurons [248]. P(VDF-TrFE) is another commonly used flexible piezoelectric polymer whose properties can be adjusted by altering the parameters of the electrospinning process. For example, P(VDF-TrFE) scaffolds fabricated using nano-electrospinning exhibited a d31 piezoelectric coefficient of up to 16.17 pC N^−1^, revealing strong piezoelectric conversion. Fibroblasts seeded on these nanofiber scaffolds proliferated 1.6-fold faster because of the electrical charge generated by the piezoelectric effect, highlighting its potential in bone tissue engineering [249]. Biomolecules also often exhibit piezoelectric properties, as previously discussed for collagen. Natural materials are already been extensively studied as natural piezoelectric additives [250]. For instance, chitosan-based nerve conduits can bridge sciatic nerve gaps, promote axon and myelin regeneration, stimulate angiogenesis, and regulate collagen fiber deposition [246].
Inorganic piezoelectric additives generally comprise single-crystal materials, piezoelectric ceramics, and piezoelectric semiconductor nanomaterials. Here, we focus on the applications of piezoelectric ceramics. Piezoelectric ceramics exhibit higher piezoelectric coefficients than their organic counterparts. Their effect arises from crystal deformation under mechanical stress, which aligns electric dipoles and results in the accumulation of surface charges [251]. Common examples include potassium sodium niobate (KNN), BaTiO3, and lead zirconate titanate (PZT). Although PZT exhibits markedly high piezoelectric constants, its lead content poses strong cytotoxic risks [248]. BaTiO3, the first lead-free piezoelectric ceramic to be used in bone regeneration research, exhibits a high dielectric constant [252,253]. Its use in biomedical applications is well-documented, particularly in the development of novel repair materials. However, its high brittleness, low bioactivity, and limited conductivity restrict its utility in tissue engineering. Notably, BaTiO3 properties can be modulated by doping with single or multiple metal elements, such as Zr, Sr, and Ca. In addition, composites formed by combining BaTiO3 with other materials using approaches such as mixing or coating have revealed broad application prospects [253,254].
Piezoelectric composites integrate the advantages of both organic and inorganic additives, achieving good flexibility, while maintaining high piezoelectric coefficients. A common strategy involves dispersing piezoelectric ceramic nanostructures into a polymer matrix to create polymer-based composites [255]. PVDF can significantly enhance its piezoelectric properties by incorporating nanofillers such as ZnO and BaTiO3, which increase the β-phase content. For example, Shuai et al. incorporated BaTiO3 nanoparticles into PVDF scaffolds. These functionalized nanoparticles were distributed within the PVDF matrix, thereby enhancing the scaffold's piezoelectric performance. These composites exhibited excellent potential in tissue engineering applications [256,257]. In addition, PCL has been used as the dispersed phase in polymer composites owing to its ease of processing and low melting point. By uniformly co-mixing PCL with BaTiO3 to fabricate composite scaffolds, biomaterials can be obtained that have improved mechanical properties, higher dielectric constants, and enhanced support for cell growth [258].
Smart-responsive materials are defined as materials that dynamically respond to physical or chemical stimuli, thus exhibiting significant potential for biomedical applications. They are generally categorized as physically or biochemically responsive, depending on the type of stimulus. Physically responsive materials react to cues such as electricity, light, or ultrasound, whereas biochemically responsive materials respond to factors including pH, enzymes, or reactive oxygen species (ROS) [255]. These systems are often constructed using hydrogels that release therapeutic cargo, such as NTFs, in response to external triggers, thereby enabling spatiotemporally controlled neuromodulation to promote tissue repair and regeneration [259,260].
Among physical stimuli, electrical stimulation provides notable advantages, noninvasiveness, controllability, and rapid response, making it widely applied in neural tissue engineering. Most electrically responsive biomaterials rely on redox reactions in conducting polymers. For instance, PPy contains an oxidized carbon skeleton with positively charged sites that serve as active sites for electrostatic and hydrophobic interactions with negatively charged substances. Upon electrical reduction to a neutral state, the PPy matrix releases the loaded cargo [261]. However, the polymer's chargeability limits the loading and release capacity. To overcome this limitation, conductive polymer films can act as switches, whereas non-electrically active materials serve as cargo containers. In addition, nanoscale dispersed or colloidal systems, such as nanoparticles, micelles, and vesicles, have been used to increase the effectiveness of cargo loading and release [262]. Light-responsive materials have promising applications in neuralized tissue engineering. For example, a light-responsive H2S composite nerve conduit containing Zn^2+^, designed by Huo et al., released H2S under near-infrared irradiation, modulating the neural regenerative microenvironment through multiple pathways in the presence of H2S and Zn^2+^ (Fig. 4A) [263]. Ultrasound-responsive biomaterials also show significant potential for neuralized skin tissue engineering. For instance, an ultrasound-responsive hydrogel was developed by Wang et al. Its structure consists of an interpenetrating polymer network of gelatin/polyvinyl alcohol, which is doped with reduced graphene oxide and piezoelectric nanocrystals [264]. This material converts acoustic energy into electrical signals through the piezoelectric enhancement effect [264]. It upregulates the neurotrophic effects of SCs, supporting axonal growth and promoting nerve regeneration in diabetic wounds [264]. Notably, dual-responsive biomaterials have also made significant progress. Qin et al. designed a new dual-responsive OSPPB hydrogel in which polyethyleneimine-grafted protocatechuic acid efficiently adsorbs negatively charged exRNA, blocking recruitment of pro-neurovascular factors such as VEGF (Fig. 4B) [265]. Simultaneously, bevacizumab-loaded nanoparticles continuously release antibodies to directly inhibit VEGF activity [265]. The hydrogel can further sense high-ROS and acidic microenvironments through dynamic boronic ester (ROS-sensitive) and Schiff base (pH-sensitive) cross-links, achieving on-demand drug release and precisely inhibiting aberrant neurovascular formation at the bone-cartilage interface, thereby relieving pain and slowing osteoarthritis progression [265].Fig. 4Application of smart response biomaterials in neuralized tissue engineering. A) Schematic illustration of the construction of a light-responsive H2S compound system and its enhancement of the peripheral nerve regeneration microenvironment. Adapted reprinted with permission from Ref. [263], based on CC BY License, Copyright © 2025 John Wiley and Sons. B) Schematic illustration of OSPPB hydrogel inhibiting exRNA adsorption to VEGF. OSPPB hydrogel successfully adsorbs and immobilizes exRNA, thereby preventing them from recruiting factors such as VEGF in subsequent processes, ultimately achieving the therapeutic goal of tissue repair. Adapted reprinted with permission from Ref. [265], based on CC BY License, Copyright © 2025 Springer Nature Portfolio.Fig. 4
In this section, we detail four key structural and mechanical design, delivery of bioactive signals, electrophysiological modulation, and smart responsive materials. Subsequent sections will focus on mapping these strategies to neural, skin, skeletal, and cardiac tissues. Notably, these strategies can be combined and applied according to the physiological characteristics and repair objectives of different tissues. For instance, neural tissue repair heavily relies on electrophysiological regulation and the targeted delivery of bioactive signals [266], whereas skin tissue regeneration focuses on intelligent responsiveness and the time-controlled release of active signals [267]. Furthermore, bone tissue construction emphasizes the synergy between mechanical support and osteogenic signaling [268], and cardiac tissue engineering is highly dependent on electrophysiological matching and dynamic mechanical support [24].
Compared to the CNS, the PNS possesses greater regenerative potential, with peripheral nerves being capable of spontaneous regeneration even in cases of relatively minor injury. This is because, despite the presence of a transient inhibitory environment in the early stages of peripheral nerve injury, SCs rapidly dedifferentiate into a reparative phenotype. They phagocytose myelin debris to clear inhibitory substances, align themselves in an orderly fashion to form the Büngner band, and synthesize and secrete NTFs in large quantities [269]. Distinct microenvironments and inhibitory cues emerge after CNS injury. First, disrupted oligodendrocytes release myelin-associated inhibitory factors such as Nogo, myelin-associated glycoprotein, and oligodendrocyte myelin glycoprotein [270]. The increased levels of these molecules directly inhibit axonal regeneration. Second, the glial scar formed by astrocytic overproliferation impedes neural regeneration and secretes large quantities of molecules like chondroitin sulfate proteoglycans that obstruct axonal extension [271]. Therefore, given the differences in their microenvironments and inhibitory cues, the biomaterials applied to each also exhibit certain variations [272]. The core of peripheral nerve repair lies in bridging, and NGCs and decellularized extracellular matrix (dECM) precisely meet this requirement. As a physical bridge, NGCs connect the proximal and distal ends of the severed nerve, providing a protected pathway for regenerating axons [273]. dECM preserves the natural 3D structure of the extracellular matrix, basement membrane tubules, and associated bioactive molecules, offering an ideal natural scaffold for SC migration and axonal regeneration [274]. Recent studies have demonstrated that dECM also exhibits promising efficacy in the field of SCI treatment [275,276]. Additionally, it is noteworthy that electroactive materials, leveraging their ability to reconstruct the electrophysiological microenvironment of injured regions, have found extensive applications in both the PNS and CNS [23].
The primary goal of tissue engineering in the treatment of peripheral nerve injury is to provide guidance and promote nerve regeneration. NGCs, as core repair tools, help advance a biomimetic tubular structure that physically directs axonal growth. This is particularly critical for long-segment peripheral nerve injuries and has broad applications in peripheral nerve tissue engineering [273]. Its tubular walls were constructed from natural or synthetic polymers and optimized to create a regenerative microenvironment by loading NTFs, seeding with neurogenic stem cells, or embedding dECM. These strategies prevent scar invasion, promote axonal growth, and facilitate myelination [39,277]. For instance, Li et al. developed a flexible NGC using electrospinning technology that significantly enhanced sciatic nerve regeneration and mechanical performance [130]. The inner layer of this NGC is prepared using electrospun fibers, while the outer surface is formed by helically melt-spun fibers, demonstrating outstanding anti-tangling characteristics and excellent nerve regeneration induction properties [130]. Notably, the potential of smart NGCs incorporating electroactive biomaterials to promote axonal and neuronal regeneration has been shown to be significant. For example, a conductive NGC developed by Song et al. supports neural cell proliferation and M2 macrophage polarization, thereby promoting sciatic nerve regeneration and restoring motor function in rats [278]. Furthermore, composite NGCs can integrate responsive delivery systems (such as smart hydrogels) or biomimetic topologies (such as nanofibers) to better simulate neural microenvironment signaling, thereby enhancing axon density and functional recovery of regenerated nerves [[279], [280], [281]]. Xu et al. recently developed an NGC with an inner layer of barium titanate piezoelectric nanoparticle (BTNP)-doped polyvinylidene fluoride-trifluoroethylene electrospun nanofibers, which featured enhanced piezoelectric properties and aligned orientation, and an outer layer of thermoresponsive poly(N-isopropylacrylamide) hybrid hydrogel encapsulating bioactive drugs (Fig. 5A) [123]. This design enables directed neuronal growth and axonal guidance through ultrasound-triggered electrical stimulation. Notably, the contraction of the hydrogel under ultrasound enables controlled NGF release. Overall, the conduit was effective in promoting axonal repair and nerve regeneration (Fig. 5B and. C) [123].Fig. 5Application of a novel nanofibers derived hydrogel conduits. A) Schematic of an ultrasonic-responsive directional piezoelectric nanofiber hydrogel conduit for peripheral nerve regeneration. B) Schematic diagram of the gastrocnemius muscle. At 8 weeks postoperatively, the operated side (right side in the image) exhibited muscle atrophy compared to the normal side (left side in the image), which is considered a key indicator for the functional recovery of the damaged sciatic nerve. C) H&E-stained cross-sectional image of the gastrocnemius muscle. Muscle fibers were stained using H&E, and analysis of gastrocnemius morphology and fiber diameter revealed that fiber diameter was greater in the BPN US (+) group compared to the BPN US (−) group, indicating its efficacy in sciatic nerve regeneration. Adapted reprinted with permission from Ref. [123]. Copyright © 2024, John Wiley and Sons.Fig. 5
dECM serves as a vital bridge between natural tissues and synthetic materials by mimicking the natural neural microenvironment; it is also widely applied in peripheral nerve tissue engineering [282,283]. Obtained by removing cellular components using physical, chemical, or enzymatic methods, dECM preserves the 3D structure, bioactive components (such as collagen, elastin, fibronectin, and laminin), and the mechanical properties [274,284]. By retaining these biomimetic signals, dECM regulates the stem cells migration and proliferation, including that of SCs, and promotes myelination, providing a key biological basis for nerve repair [282]. For example, fibronectin supports SC growth and migration, whereas laminin enhances SC proliferation [285,286]. Notably, dECM scaffolds can also shift macrophage phenotypes from pro-inflammatory M1 to anti-inflammatory M2. This stabilizes the microenvironment and promotes tissue repair [274]. In addition, dECM can repair damage to the CNS by delivering bioactive molecules such as loaded NTFs. Liu et al. developed a multifunctional decellularized spinal cord-derived extracellular matrix platform cross-linked with GDNF to support myelin regeneration [54]. In this platform, extracellular matrix retains key protein components required for supporting stem cell proliferation and promoting differentiation, while enabling sustained release of GDNF. Activation of the PI3K/Akt and MEK/ERK pathways promotes axon regeneration, enhances neural signal transmission, and holds potential for improving motor function following SCI. Furthermore, through additional processing, such as hydrogelation for use as a lumen filler in NGC or a transplant carrier for stem cells, including SCs, dECM can be fabricated into various biomaterial forms, optimizing its effects on nerve repair [287].
Owing to the electrophysiological microenvironment of the nervous system, electroactive biomaterials provide unique advantages for treating nerve injuries. For example, a PPy self-healing conductive hydrogel designed by Xuan et al. has been successfully applied to peripheral nerve repair [288]. It targets the interleukin-17 receptor A to enhance the expression of genes and proteins involved in myelin formation in SCs, thereby promoting myelin regeneration. Vijayavenkataraman et al. developed a PCL/PPy conductive scaffold as a nerve graft carrier using 3D printing technology, which was successfully applied to repair peripheral nerve injury [289]. Additionally, PVDF scaffolds with piezoelectric activity are also considered promising for peripheral nerve repair [290]. Unlike peripheral nerve repair, materials used for CNS repair are typically constrained by the microenvironment and inhibitory cues. Furthermore, the hydrophobicity and inadequate mechanical properties of electroactive biomaterials pose significant obstacles to CNS repair. Therefore, integrating electroactive biomaterials with hydrogels has emerged as a current research focus in CNS repair. For instance, a conductive hydrogel based on PT designed by Rinoldi et al. supported the differentiation and proliferation of NSCs and MSCs. The hydrogel also exhibited good 3D printing potential, making it a perfect candidate for scaffolds in CNS repair [291]. In addition to PT, other materials commonly used to construct conductive hydrogels for CNS repair include the conductive polymer PPy and carbon-based materials [121]. PPy offers numerous advantages such as high conductivity, environmental stability, and ease of modification. However, its physical characteristics make it unsuitable for soft natural spinal cord tissue, a significant drawback [292]. As such, researchers typically incorporate it into hydrogels for application. For instance, a conductive hydrogel was created by Wu et al. using a bioactive collagen/hyaluronan hydrogel to embed dual-functional PPy nanoparticles [293]. This design protects encapsulated bone marrow mesenchymal stem cells (BMSCs) from oxidative damage and promotes neuronal differentiation through the combined effects of electrical stimulation and conductivity, thereby promoting the repair of SCI. Regarding carbon-based materials, both graphene and CNTs have been extensively applied in SCI repair [294]. Agarwal et al. successfully prepared a graphene-crosslinked neural conduit using cryogel technology, which effectively supported cell growth and migration within the adult rat spinal cord [295]. Additionally, Yao et al. transplanted CNT/GelMA hydrogel fibers into a T9 transection rat SCI model. The results demonstrated that these fibers functioned as scaffolds to induce nerve regeneration, thereby promoting recovery from SCI [296].
Based on the above discussion, the core of material design in neural tissue engineering adheres to the dual principles of functional adaptation and microenvironment reconstruction. NGCs with directional topologies and active factor delivery capabilities are designed for bridging and guiding PNS repair. For the inhibitory microenvironment of the CNS, the focus shifts to integrating spatiotemporal release of electrophysiological and bioactive signals to overcome regenerative barriers. However, we have yet to provide a perfect solution for achieving high compatibility between material properties (such as conductivity and mechanical performance) and biological soft tissues. In the short term, breakthrough opportunities lie in combining advanced technologies like 3D printing and electrospinning to construct multifunctional bionic scaffolds that mimic natural neural repair.
The skin is densely innervated by the PNS. Functional skin regeneration requires reconstruction of the epidermal and dermal structures as well as restoration of the neural network [297]. Nerves are essential for both skin healing and neurons interact with fibroblasts to induce their differentiation into myofibroblasts, thereby influencing collagen formation and organization. SCs further support myofibroblast differentiation through TGF-β signaling, a pathway critical for myelin formation and nerve repair during wound healing [77,180]. Neuroactive substances, including neurotransmitters, neuropeptides, and NTFs, also modulate wound healing. For example, CGRP acts on macrophages and neutrophils through RAMP1, promoting neutrophil apoptosis and polarizing macrophages toward a pro-repair phenotype, which accelerates wound healing and enhances muscle regeneration [61]. Similarly, NGF can upregulate VEGF and fibroblast growth factor expression in endothelial cells, thereby promoting angiogenesis and fibroblast-to-myofibroblast differentiation. In addition, NGF stimulates the proliferation of keratinocytes and the outgrowth of neurites through the TrkA-PI3K/Akt pathway [298]. In patients with severe burns, traumatic defects, or diabetic ulcers, chronic wounds are always unable to heal because of persistent inflammation, in some instances leading to irreversible nerve damage. Although nerve repair has historically been overlooked in wound healing, growing recognition of its importance has driven the development of biomaterials designed to support neurogenesis.
The use of biomaterials to promote skin healing through NTFs primarily involves providing physical guidance for neural innervation [21]. Oriented nanostructured scaffolds can mimic the topological structure of basement membrane tubules within the neurofibrillary network, guiding neurons and SCs to migrate along fiber directions and form Büngner bands, thereby creating tracks for axonal extension. For example, Zhang et al. implanted dimethyloxalylglycine-loaded poly(ε-caprolactone) fiber (PCLF/DMOG) meshes prepared by electrospinning into a diabetic rat model, which increased the number of marker-positive neurons [299]. In addition to structural guidance, biomaterials can serve as spatiotemporal delivery carriers for bioactive molecules involved in biochemical signaling. In particular, hydrogels and PLGA microspheres are advantageous for NTF delivery because of their tunable degradation rates and controlled release profiles. Keykhaee et al. designed a multifunctional gum arabic–alginate hydrogel incorporating mesoporous silica nanoparticles. This hydrogel maintained NGF release for over 21 d, and histological analysis indicated improved re-epithelialization and angiogenesis, thereby effectively enhancing healing in diabetic foot ulcers [300]. Similarly, Li et al. developed a thermosensitive heparin-poloxamer hydrogel that co-delivers basic fibroblast growth factor and NGF. This was used to treat diabetic rats with sciatic nerve compression and effectively promoted SC proliferation and myelin regeneration [301]. In addition, a PLGA/collagen–chitosan scaffold has been developed for the delivery of NGF using PEI-conjugated plasmid nanoparticles. This scaffold increases nerve density and achieves complete wound healing [302]. Notably, electroactive regulation represents another promising approach in neuro-engineered skin tissue. Bi et al. designed a multifunctional conductive hydrogel that can release curcumin, enhance SC migration and proliferation, and promote myelin regeneration (Fig. 6A) [303]. This supports nerve repair in diabetic wounds and facilitates overall recovery [303]. Similarly, Liu et al. developed an adjustable biomimetic composite neural scaffold with conductive properties and directional architecture (Fig. 6B) [304]. Its structural layer was made up of an oriented chitosan/PCL nanofiber membrane, whereas the functional layer incorporated an ion-conductive chitosan/sodium alginate hydrogel. This scaffold effectively improved the neural microenvironment and promoted peripheral nerve damage [304].Fig. 6Applications of biomaterials in skin tissue engineering. A)Schematic diagram of hydrogel preparation and its reparative effects on peripheral nerve injury and diabetic wound healing. Adapted reprinted with permission from Ref. [303]. Copyright © 2024, Elsevier. B)Design and application of bionic composite neural scaffolds. Adapted reprinted with permission from Ref. [304]. Copyright © 2025, Elsevier. C)Schematic of the preparation and application of PP-MgSi composite patches in diabetic wound repair. Adapted reprinted with permission from Ref. [305], based on CC BY License, Copyright © 2025 John Wiley and Sons.Fig. 6
As research on neurovascular coupling progresses, poor integration between grafts and the host neurovascular system has emerged as a critical challenge in the field of neurogenic skin tissue engineering. To address this, Xu et al. developed a PCL-PLLA-MgSiO3 (PP-MgSi) patch using electrospinning. This material locally released Mg^2+^ and SiO3^2−^ ions, modulating inflammation by regulating TNF-α, iNOS, and CD206 expression, while simultaneously increasing CD31 and β3-tubulin levels to enhance neurovascular formation. In type 2 diabetic mice, the PP-MgSi patch achieved neovascularization comparable to three control products but showed superior peripheral nerve regeneration (Fig. 6C) [305]. In recent studies, 3D printing technologies have further advanced this field by enabling precise control of scaffold structure, thereby enhancing integration with host neurovascular systems. Consequently, to address the critical challenge of neurovascularization in skin tissue engineering, 3D printing holds significant potential [306].
In summary, the primary challenges in designing materials for neurotrophic support in skin tissue engineering currently lie in overcoming the complex inhibitory microenvironment of chronic wounds and achieving rapid, functional integration between the newly formed neural network and host sensory nerve endings. Notably, developing smart materials capable of dynamically responding to changes in the wound microenvironment holds promise for advancing skin repair and sensory function restoration.
In bone tissue engineering, the role of peripheral nerve regeneration cannot be overlooked, as it directly regulates the osteogenic differentiation [307]. For example, sensory nerves promote osteogenesis through neurotransmitters and NTFs, such as CGRP and semaphorin 3A (Sema3A). CGRP activates calcitonin gene-related peptide-like receptors and the RAMP1 complex on cell membranes, thereby enhancing the capacity of stem cells/mesenchymal cells to differentiate into osteogenic cells. In contrast, Sema3A exerts an anti-osteoclastogenic effect by binding to neurofilament protein-1 (Nrp1) and inhibiting tyrosine-based immune receptor activation motifs and Ras homolog family member A (RhoA) signaling [219,308,309]. Nerve–vascular coupling also facilitates the reconstruction of the skeletal vascular network. For example, NGF expression significantly increases after bone injury, stimulating angiogenesis. Conversely, neurons can express VEGF receptors or enhance axonal regeneration and functional innervation in peripheral tissues to regulate neurogenesis [310]. Previous studies in the skin, intestines, and adipose tissue have revealed that neuroimmune interactions modulate inflammation. Although evidence in bone regeneration remains limited, the possibility that nerve–immune cell crosstalk indirectly promotes bone healing by altering the local microenvironment cannot be ignored [311]. Historically, bone tissue engineering has focused more on developing materials that outperform autologous or allogeneic bone grafts [312]. Currently, a wide range of materials is available. However, few strategies can replicate the neural functional units within the bone. As the nervous system serves as a primary regulator and coordinator of immune, inflammatory, and vascular activities in the bone microenvironment, it plays an indispensable role in bone repair. Therefore, developing neuralized biomaterials could significantly advance bone tissue engineering. There are three broad approaches to designing neurogenic bone tissue engineering (1) selection of scaffold material and structural design; (2) construction of co-culture systems to facilitate interactions between neural and bone cells; and (3) incorporation of NTFs, neuropeptides, and bioactive ions to promote bone healing.
The structural characteristics of bio-scaffolds, particularly surface topology and layered architecture, are crucial in regulating cell growth [313]. Taking inspiration from the way tree leaves are arranged, 3D printing has been used to create tree-like scaffolds with distinct leaf patterns consisting of a central trunk and layered leaves, which show notable advantages in neurogenic bone tissue engineering. Scaffolds with optimized features (leaf spacing of 0.8 mm, divergence angle of 20°, and stem diameter of 2.4 mm) can be produced by precisely controlling parameters such as leaf spacing, divergence angle, stem diameter, and leaf-surface microstructure. These scaffolds provide an ideal surface area-to-volume ratio and significantly promote cell growth. Furthermore, BMSCs and SCs have shown excellent adhesion, proliferation, and differentiation on the leaf surfaces of these scaffolds [314]. Owing to the uneven distribution of blood vessels and nerves within natural bones, they exhibit an inherent structural complexity that is difficult to replicate. However, mimicking key natural processes, such as periosteum-mediated healing, can facilitate multi-tissue regeneration. For example, the incorporation of matrix metalloproteinase-degradable hydrogel-modified constructs into the body accelerates the healing of allogeneic grafts by supporting the endogenous cells migration and the neurovascular networks formation, thereby effectively simulating periosteal healing [315]. In addition, of critical importance to both osteogenic and neural differentiation are the physical properties of the extracellular matrix. For example, the stiffness of matrix is one decisive neural stem cells are prone to differentiating into neurons on softer matrices, whereas stiffer matrices favor differentiation into astrocytes or oligodendrocytes. Topography also plays a critical role; for instance, increasing surface roughness significantly reduces neuronal cell viability, possibly because of enhanced hydrophilicity, which reduces adhesion protein adsorption on rough surfaces [316]. Therefore, optimizing the physical properties of bone biomaterials is essential for promoting both bone and nerve tissue growth. Notably, studies have shown that precisely controlling the roughness and spatial arrangement of nanoscale features within bone biomaterials is crucial for regulating neuronal cell behavior within the bone matrix. For example, oriented nanofibers can mimic fibronectin signals to promote cell alignment and phenotype, uniform microchannels and circular nanocolumn arrays can guide axonal extension, and sericin nanofiber scaffolds can support directed neuronal growth and endothelial cell activity [317].
Seed cells are vital components in neurogenic bone tissue engineering. To date, various seed cells, including SCs and MSCs, have been shown to support neurogenesis-assisted bone repair and regeneration. A model developed by Jones et al. validated the neurodependence of mouse skeletal stem cells (mSSCs): following denervation of the mandibular alveolar process, mandibular repair was significantly impaired because of mSSC dysfunction, as these cells could no longer participate in skeletal development and repair [318]. Based on these findings, numerous authors proposed using transplanted SCs to overcome bone repair deficits caused by denervation. In neuralized bone tissue engineering, the extraction of different seed cells has been largely resolved; thus, we should pay more attention to designing cell delivery systems in advancing the field. A widely applied strategy for bone regeneration is cell implantation using 3D bioprinted scaffolds [319]. Zhang et al. fabricated Haversian canal-mimicking bioceramic scaffolds containing Volkmann canals, Haversian canals, and cancellous bone structures using digital light processing-based 3D printing (Fig. 7A) [320]. The scaffold encompasses a multi-cell delivery MSCs were delivered through cancellous bone-like structures, whereas endothelial cells and SCs were delivered through the Haversian canal network. This approach enabled the precise localization of neurogenic, angiogenic, and osteogenic cells within specific scaffold regions, thereby providing an optimized strategy for bone regeneration. Neurogenic cell-derived exosomes have also emerged as a promising tool for enhancing bone repair. For example, Su et al. integrated SC-derived exosomes (SC-EXOs) into electrospun PCL membranes with axially aligned microchannels through electrostatic interactions, generating a polycaprolactone–polyethyleneimine–exosome (PPEA) membrane with enhanced regenerative potential (Fig. 7B) [321]. This system specifically targeted SCs at axonal injury sites, promoting neuropeptide (SP and CGRP) secretion, neuronal differentiation, axonal elongation, endothelial cell osteogenic differentiation, and proliferation, thereby enhancing innervation and vascularized bone regeneration in vivo. Hao et al. revealed that encapsulating SC-EXOs in a hydrogel further improved cranial regeneration by promoting neural innervation in vivo (Fig. 7C) [322]. These findings emphasize the promise of SC-EXOs for designing biomaterials in the field of neurogenic bone tissue engineering.Fig. 7Applications of biomaterials in bone engineering. A)Schematic diagram of a 3D-printed Haversian bone-mimetic scaffold. It integrates Haversian canals, Volkmann canals, and cancellous bone structures to deliver osteoblasts and angiogenic cells. Adapted reprinted with permission from Ref. [320], based on CC BY License, Copyright © 2020 The American Association for the Advancement of Science. B)Schematic diagram of the construction of a biomimetic periosteum containing exosomes. Adapted reprinted with permission from Ref. [321], based on CC BY License, Copyright © 2022 Elsevier. C)Schematic diagram of the novel hydrogel structure. Adapted reprinted with permission from Ref. [322], based on CC BY License, Copyright © 2022 Elsevier.Fig. 7
Furthermore, incorporating various NTFs and neuropeptides can enhance bone repair and regeneration by providing neurotrophic support. For example, NGF promotes the differentiation and proliferation of osteoblasts, inhibits apoptosis, and accelerates nerve regeneration, thereby indirectly enhancing bone regeneration. CGRP binds to the TRPV1 receptor, inhibiting osteoclast formation through the RANKL/OPG pathway, while simultaneously promoting osteoblast differentiation and MSC migration through the p38-MAPK signaling pathway. SP acts on the NK1R receptor in a concentration-dependent manner, stimulating MSC proliferation, mineralization, and osteoblast-related protein expression (such as Runx2), while also enhancing osteoclast activity by activating the NF-κB and RANKL signaling pathways [323]. Therefore, one of the primary objectives in the field of neurogenic bone tissue engineering is identifying suitable materials for loading bioactive molecules. To this end, Fitzpatrick et al. used 3D-printed silk–hydroxyapatite scaffolds loaded with NGF, which promoted bone regeneration by enhancing both neurotrophic activity and vascularization [324]. Similarly, Luo et al. developed CGRP-loaded PMs for periodontitis treatment, revealing their ability to protect BMSCs from inflammatory mediators and significantly enhance osteogenic activity, thereby promoting periodontal bone regeneration [325]. To further promote cell differentiation, bioactive factors are often co-delivered with cell-loaded scaffolds to facilitate neurogenesis. PMs co-loaded with CGRP and BMSCs achieved significantly greater alveolar bone regeneration than that of PMs loaded with either CGRP or BMSCs alone. In addition to growth factors and peptides, certain ions (such as Mg, Ca, Si, Li, Cu, and P) also exhibit biological activity, regulating key cellular behaviors, thereby contributing to neurogenic bone healing. Jing et al. proposed a biomaterial that is designed to stimulate the growth of skeletal-related neural networks and encourage bone regeneration in the treatment of infectious bone defects. Therefore, they developed a photosensitive conductive hydrogel by incorporating magnesium-modified black phosphorus (BP@Mg) into GelMA (Fig. 8A) [199]. This system released conductive nanosheets and bioactive ions that enhanced SC secretion and migration, promoted neurite outgrowth, and facilitated neurogenic bone regeneration (Fig. 8B) [199]. Additionally, it demonstrates excellent antibacterial effects, thereby more effectively promoting bone regeneration (Fig. 8C) [199]. Similarly, Xu et al. fabricated a biohybrid biodegradable hydrogel. This contained copper ion-modified germanium phosphate (GeP) nanosheets (Fig. 9A) [326]. It exhibited excellent antibacterial efficacy (Fig. 9B) [326]. Furthermore, the integrated hydrogel promoted BMSC differentiation and upregulated neurogenesis-related proteins in neural stem cells, ultimately enhancing bone regeneration (Fig. 9C) [326].Fig. 8Application of a photosensitive and conductive hydrogel in bone tissue engineering. A) Schematic illustration of GelMA-BP@Mg (GBM) hydrogel-assisted nerve regeneration in bone defects. B) Immunofluorescence images of PC12 cells on day 5, percentage of cells exhibiting neurites on day 5, and quantitative analysis of neurite length. PC12 cells in all groups exhibited neurite outgrowth, with the best results observed in the GBM group. This may be because the SC-conditioned medium in the GBM group contained higher levels of NGF, which effectively promoted neural repair. Magnesium ions released during BP@Mg degradation may also offer potential benefits for neural repair. C) Plate counts of Escherichia coli and Staphylococcus aureus on three types of hydrogels, showing untreated (left) and near-infrared irradiated (right) samples. Adapted reprinted with permission from Ref. [199], based on CC BY License, Copyright © 2023 John Wiley and Sons.Fig. 8Fig. 9Application of a novel electroactive composite hydrogel in bone tissue engineering. A) Schematic diagram of the preparation of GelMA/GeP@Cu electroactive hydrogel and its support for bone regeneration. B) Plate count results from co-culturing with Escherichia coli and Staphylococcus aureus cultures in three types of hydrogels. C) Typical micro-CT images obtained at 6 and 12 weeks after placement of four hydrogel scaffolds in the cranial defect. Adapted reprinted with permission from Ref. [326], based on CC BY License, Copyright © 2023 John Wiley and Sons.Fig. 9
Overall, the core of neuralization strategies in bone tissue engineering lies in bionically constructing structures that actively guide nerve ingrowth and coordinate the interactions among nerves, blood vessels, and bone. For instance, guiding directed cell growth and differentiation through structural biomimicry and physical modulation, achieving neural-bone cell coupling via co-culture systems or functional carriers such as exosomes, and loading NTFs with bioactive ions to synergistically promote innervation and osteogenesis. The primary challenge we currently face is maintaining neural signal homeostasis during long-term repair cycles. Therefore, developing scaffolds with spatiotemporally controlled release holds promise for advancing clinical translation in this field.
For cardiac tissue engineering, researchers have early proposed cell-based strategies to address the inherent limitations of cardiac muscle regeneration. However, the therapeutic efficacy remains limited by factors such as poor tissue integration and insufficient cell maturity. In most cases, implanted stem cells fail to differentiate into the desired cell lineages [327,328]. Therefore, stem cell therapy for myocardial infarction requires further research and refinement, indicating that a single-cell strategy may not be sufficient for effective cardiac tissue engineering. Emerging evidence highlights the unique role of neural regulation in myocardial regeneration. Overactivation of the sympathetic nervous system may significantly exacerbate inflammation and fibrosis, whereas the parasympathetic nervous system suppresses immune responses and promotes myocardial cell survival and regeneration [329]. Notably, activation of the IL-10/STAT3 pathway has been shown to be stimulated by stimulation of the cardiac vagus nerve. This promotes the polarization of cardiac macrophages toward the M2 phenotype, thereby optimizing the regenerative microenvironment after myocardial infarction and ultimately improving cardiac function [330]. In addition, NTFs and neuropeptides are crucial in guiding myocardial cell proliferation and survival. For example, overexpression of mesencephalic astrocyte-derived NTF (MANF) in cardiomyocytes reduces endoplasmic reticulum stress-induced apoptosis. Notably, pharmacological disruption of cholinergic signaling reduces the proliferation of cardiomyocytes in the injured hearts of zebrafish and newborn mice. Conversely, administering NGF stimulates cardiac regeneration by reactivating cardiomyocyte proliferation [331,332]. Meloni et al. further confirmed that NGF promotes cardiac regeneration primarily by activating the pro-survival and pro-angiogenic Akt/Foxo pathway and upregulating stem cell factors (c-kit receptor ligands) to recruit c-kitpos PCs to the myocardium [333]. These findings indicate that investigating cardiac regeneration from the perspective of neural modulation is a promising strategy for overcoming the current bottlenecks in cardiac tissue engineering. Building on this concept, Wang et al. used 3D bioprinting technology to construct a novel pre-neuralized scaffold by combining strontium silicate (SrSiO3, SS) microparticles with NSCs (Fig. 10A) [24]. SEM images reveal the porous structure of the scaffold, with SS particles well distributed throughout the GelMA matrix (Fig. 10B) [24]. The pre-neuralized scaffold may activate four typical circadian rhythm-related genes associated with cardioprotection, thereby enhancing myocardial repair after infarction. Within the scaffold, NSCs predominantly differentiate into neuron-like cells and secrete NTFs, which contribute to functional cardiac repair. Simultaneously, the scaffold released Sr and Si ions, which further promoted neuronal differentiation and maturation of NSCs, ultimately restoring cardiac muscle function (Fig. 10C) [24].Fig. 10Application of a pre-neuralized scaffold in cardiac tissue engineering. A) Schematic diagram of the design and preparation of pre-neuralized scaffolds and their application in myocardial regeneration. B) Scanning electron microscope images of cross-sections of grid-like scaffolds after immersion in cell culture medium with different concentrations of SS particles, and energy-dispersive X-ray spectroscopy elemental distribution maps of 5SS-GelMAs scaffolds. C) Schematic diagram of Ca^2+^ transient changes. The waveforms revealed that the CMS-co-5SS-NSCS group exhibits reliable calcium signal synchronization. Adapted reprinted with permission from Ref. [24], based on CC BY License, Copyright © 2025 John Wiley and Sons.Fig. 10
Presently, the development of materials for supporting neurogenesis in cardiac tissue engineering remains in the early stages of research, facing highly complex challenges. The question of how to achieve stable, controllable, and physiologically functional innervation within the intricate microenvironment of myocardial infarction remains to be resolved. Concurrently, mitigating the risk of arrhythmias caused by excessive sympathetic nerve activation and ensuring precise electromechanical coupling between the newly formed neural network and the host myocardium are critical considerations that require our focused attention.
Although biomaterial-assisted neuralization in tissue engineering holds significant promise, several challenges remain. First, the involvement and regulatory role of neural mechanisms in tissue regeneration have only recently gained attention, and the underlying biological pathways remain incompletely defined. Although knowledge of neural responses to tissue injury has advanced, the roles of neurogenic inflammation, immunomodulation, and neuromodulation in regulating stem and progenitor cell behavior remain insufficiently explored. Similarly, the contribution of neurovascular coupling to tissue repair, an emerging area of study, requires validation through large-scale investigations. Second, following in vivo implantation, the mechanisms governing interactions among biomaterials, innervation, and target cells or tissues remain unclear, making it difficult to establish a direct causal relationship between innervation and functional tissue regeneration. Developing 3D in vitro pathophysiological models through biomanufacturing technologies may help address this gap. Third, compared with conventional tissue engineering strategies, biomaterial-assisted neuralization is relatively young and has not yet progressed to clinical application. A key challenge is the establishment of functional neural connections, which requires effective extension and precise targeting of long-distance axons, functional synapse formation with host nerves, and adequate myelination across injured regions. Fourth, accurate modeling of the neural microenvironment remains technically demanding. An ideal neuralized scaffold should integrate a well-defined structure, appropriate mechanical strength, and a controlled degradation rate, while supporting neurogenic cells and enabling the spatiotemporally regulated release of bioactive molecules such as NTFs. In addition, host immune responses and glial scarring can inhibit regeneration, whereas chronic inflammation or foreign body reactions further impair neural integration. Insufficient vascularization also leads to nutrient deprivation and metabolic waste accumulation in the central region, thereby limiting the survival and function of large-sized neural grafts. Finally, the construction of electrophysiological microenvironments, arguably the most critical aspect of neuralization, still raises long-term biosafety concerns. The cytotoxicity and metabolic fate of electroactive biomaterials remain unresolved. For instance, carbon-based additives may induce protein aggregation and compromise cell membrane integrity, whereas metal-based additives can generate hydroperoxide radicals, leading to cytotoxic effects and adverse metabolic outcomes.
In the future, continued advances in manufacturing technologies such as 3D printing, nanotechnology, and electrospinning are expected to facilitate the engineering of more sophisticated neuralized scaffolds with improved biocompatibility, patient specificity, and safety for clinical application. In addition, the integration of artificial intelligence (AI) is expected to revolutionize biomaterials and tissue engineering. AI-assisted scaffold design and fabrication can enhance precision, reduce production time and cost, and significantly accelerate progress in 3D printing. Furthermore, the development of smart-responsive biomaterials represents a promising direction for assisted neuralization. By dynamically modulating their physicochemical properties or releasing bioactive factors in response to the local microenvironment, these materials may support tissue repair and neuromodulation in a highly controlled manner.
Jiahao Ye: Writing – review & editing, Writing – original draft, Investigation. Lei Ji: Visualization, Validation, Investigation. Liangle Liu: Supervision, Resources, Funding acquisition. Keyu Zhou: Supervision, Investigation. Rui Zhu: Visualization, Validation. Chuchu Sun: Validation, Investigation. Lanjie Lei: Writing – review & editing, Project administration, Conceptualization. Minghai Dai: Writing – review & editing, Visualization, Funding acquisition.
The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.