Peripheral Nerve Injury (PNI) refers to structural or functional abnormalities in the peripheral nervous system caused by various factors (1, 2), leading to sensory, motor, and autonomic nervous dysfunction (3). These injuries often result from traumatic events such as car accidents, gunshot wounds, sports injuries, and surgical accident (4). Peripheral nerve injury often leads to sensory and motor dysfunction in patients (5). The treatment period following peripheral nerve injury typically exceeds three months (6), accompanied by chronic neuropathic pain, which reduces patients’ quality of life and significantly impacts social productivity (6). Currently, PNI impacts over 20 million people, and the annual medical expenditure for PNI treatment in the United States reaches $150 billion, imposing a huge economic burden on society (4). Due to high treatment costs and suboptimal efficacy, PNI treatment remains a critical issue requiring attention.

Clinical treatment strategies for PNI usually depend on the injury type, severity, etiology, and individual differences (7). Peripheral nerve injuries (PNI) were classified into five grades based on the severity of damage and functional loss. Grades I and II can recover slowly and spontaneously, while Grades III, IV, and V involve injuries to the endoneurial tubes, perineurium, and epineurium, respectively, typically requiring surgical intervention (8, 9). For peripheral nerve atrophy caused by underlying pathologies, pharmacotherapy and physical therapy are commonly employed, though these approaches lack reparative effects on structural defects such as nerve transection (10). To address nerve transection injuries, key techniques for PNI repair and regeneration include: direct neurorrhaphy, autologous nerve grafting, allogeneic nerve transplantation, and bioengineered neural scaffolds (11). For short-distance defects (<3 mm), clinical practice employs nerve anastomosis to suture the nerve (12). The “gold standard” for treating long-distance peripheral nerve defects is autologous nerve grafting (13), where grafts are harvested from other body parts. Optimal outcomes occur when the diameter and axon density of the donor and recipient nerves match (14). Autologous nerves exhibit excellent biocompatibility and support the proliferation of Schwann cells and other cells (15). However, their clinical application is limited by factors such as limited supply, the need for a second surgery, and mismatches between donors and recipients (16). Over the past few decades, researchers have explored biological tubular grafts, such as blood vessels, arteries, muscle fibers, bone conduits, and allogeneic nerve grafts. However, concerns about limited availability, poor mechanical properties, and high morbidity at the donor site have restricted their use (17). Therefore, by integrating multidisciplinary perspectives from materials science, immunology, and regenerative medicine, this review aims to provide innovative theoretical frameworks and technical pathways for efficiently repairing PNI, advancing clinical translation ( Figure 1 ).

Over the past few decades, researchers have focused on developing artificial nerve grafts to replace autologous nerve grafts (18). The U.S. Food and Drug Administration (FDA) has approved several nerve conduits made from materials like type I collagen, chitosan, and polyvinyl alcohol for nerve defect repair (19). However, these approved grafts are only suitable for defects shorter than 3.0 cm and ineffective for severe nerve defects (20). A nerve conduit is a tissue-engineered artificial pipeline used to bridge peripheral nerve defects, guiding axonal and neural regeneration and functional recovery by connecting the two ends of the transected nerve (21).

When axons are damaged in PNI, the distal part undergoes axonal and myelin decomposition, a process known as Wallerian degeneration (22). A key event in Wallerian degeneration is the intrinsic degeneration of the injured axon (23). However, PNI is not an isolated event; damaged axons trigger a multicellular response involving multiple components, often accompanied by Schwann cell dedifferentiation and immune response activation (24). Substantial evidence shows that immune cells are recruited to the injury site within hours to days after PNI, playing a critical role in nerve regeneration (25). Macrophages are the most well-studied cell type: they not only remove phospholipid debris and regulate Schwann cell activity but also release numerous axon regeneration factors to promote axonal growth under microenvironmental guidance (26). Macrophages exhibit high plasticity, differentiating into pro-inflammatory (M1) and anti-inflammatory (M2) phenotypes under different stimuli. M1 macrophages activate immune responses to eliminate cells and microbes, while M2 macrophages are immunosuppressive, promoting tissue remodeling and repair (27). After PNI, CD4⁺ Th2 cells also secrete interleukin-4 (IL-4) and interleukin-13 (IL-13), promoting the polarization of M2-type macrophages, releasing neurotrophic factors, and inhibiting inflammatory cytokines. Evidence indicates that M2 polarization promotes peripheral nerve regeneration, unlike M1 polarization (28).

Additionally, T cells play a dual role in the repair of PNI, where their dynamic balance directly influences the regulation of neural regeneration and the immune microenvironment. Following peripheral nerve injury, Th1 cells secrete pro-inflammatory cytokines such as IFN-γ and TNF-α, which activate microglia and disrupt the blood-nerve barrier (BNB), exacerbating neural damage. Th17 cells release interleukin-17A (IL-17A), inducing neuroglial cells to produce pro-inflammatory factors that lead to demyelination and axonal injury. Post-injury, dorsal root ganglion (DRG) neurons secrete chemokines like CXCL13 to recruit CXCR5⁺ CD8⁺ T cells into the neural tissue, establishing a chronic inflammatory microenvironment (29). In contrast, regulatory T cells (Tregs) secrete interleukin-10 (IL-10) and TGF-β, suppressing the activity of Th1/Th17 cells, alleviating neuroinflammation, and promoting the expression of neurotrophic factors to support neuronal survival (30, 31). Therefore, nerve conduit technology provides a novel strategy for nerve repair by utilizing biomaterials to deliver T cells or their regulatory factors to modulate the immune microenvironment.

Nerve conduits physically guide nerve regeneration by providing directional channels for axonal growth, preventing disordered axonal growth and scar tissue invasion, and isolating the regenerating nerve from surrounding fibrosis. They can also create a microenvironment conducive to M2 polarization of macrophages through microstructural design, cytokine loading, and drug release (32, 33). Various nerve conduits have been developed, optimizing mechanical properties, macrostructure, and luminal microstructures to better mimic the natural extracellular matrix (ECM) of neural tissue (34, 35). Additionally, Furthermore, nerve conduits can regulate the immune microenvironment through the incorporation of immunomodulatory agents. For example, conduits loaded with neurotrophic factors (e.g., nerve growth factor, NGF) promote CD4⁺ Th2 polarization, inhibit macrophage M1 phenotype transition, and enhance Treg infiltration. Additionally, nerve conduits loaded with immune-modulating factors (e.g., IL-10) suppress NF-κB activation and reduce the release of pro-inflammatory cytokines (36, 37).

Advances in the isolation, purification, and manufacturing of natural biomaterials have enabled the production of robust acellular conduits (38, 39). Natural biomaterials are derived from non-mammalian macromolecules (e.g., chitosan, silk fibroin) or classical components of the mammalian ECM (e.g., collagen, elastin) (40). Due to their biological origin, these materials exhibit high compatibility with human tissues, significantly reducing inflammation, foreign body rejection, and enabling active regulation of the nerve regeneration microenvironment (41). Natural biomaterials have a reduced risk of immune rejection due to their compositional similarity to human tissues (42). Additionally, natural biomaterials possess inherent bioactivity, containing native cell recognition sites (e.g., Arg-Gly-Asp sequences) that promote cell adhesion, proliferation, and differentiation (43). Most importantly, they exhibit controllable degradability, with non-toxic degradation products metabolized via enzymatic or hydrolytic pathways (44). However, conduits fabricated from natural biomaterials often suffer from insufficient mechanical properties to adequately support tissue regeneration. Furthermore, certain materials (e.g., chitosan) may induce chronic inflammation, and since most natural biomaterials are derived from animal sources, they carry a risk of pathogen residue contamination (45). Nonetheless, as “biologically friendly” materials, they provide a basis for promoting M2 polarization of macrophages and offer more physiologically relevant solutions for nerve conduit design.

Synthetic material-based nerve conduits are a research hotspot for repairing peripheral nerve defects, offering core advantages in customizable structure and function (72). Synthetic biomaterials can regulate mechanical properties (e.g., stiffness, porosity) and degradation rates through chemical modifications, and synthetic polymers exhibit minimal batch-to-batch variation, making them suitable for standardized production. Widely used 3D printing substrates like polylactic acid (PLA) and polycaprolactone (PCL) can form biomimetic microchannels or oriented nanofibers via 3D printing or electrospinning, mimicking ECM topological structures to guide Schwann cell migration and axonal orientation (73). Adjusting the molecular weight of polymers like poly (lactic-co-glycolic acid) (PLGA) or the ratio of lactic to glycolic acid in PLGA can synchronize degradation rate with nerve regeneration, avoiding secondary surgery (74). Integrating conductive materials or loading neurotrophic factors endows synthetic conduits with biochemical signals and electrical stimulation (75). However, pure synthetic materials lack bioactivity, often requiring additional modification for cell adhesion (76), and suffer from mechanical mismatch issues: PLA has high elastic modulus causing local stress concentration. The degradation byproducts of synthetic biomaterials may trigger inflammatory responses. For instance, materials like PLGA (poly(lactic-co-glycolic acid)) generate acidic degradation products during breakdown, leading to a local decrease in pH that exacerbates macrophage infiltration and inflammatory reactions. The implantation of polyethylene glycol (PEG) hydrogels may activate macrophages, leading to fibrosis and fibrous capsule formation. Studies have shown that the material’s surface charge, porosity, and degradation products collectively influence the behavior of immune cells (77). Thus, the long-term safety of such materials in vivo still necessitates comprehensive evaluation.

While natural materials offer excellent cell affinity and regenerative microenvironments, they often have low mechanical strength, uncontrollable degradation, and batch variability limiting clinical use. Synthetic materials, though customizable in mechanics and degradation, lack bioactive signals and their degradation products may acidify the microenvironment. Composite materials combine natural and synthetic materials to leverage synergistic effects, overcoming single-material limitations by integrating mechanical properties, bioactivity, conductivity, and controllable degradation to mimic the complex microenvironment of nerve regeneration (92). Composite strategies can create smart systems with biomimetic topology, dynamic degradation matching, and functional expansion, avoiding inherent defects like immunogenicity and poor mechanics (93). However, composites face challenges such as complex processing, poor interfacial compatibility, high production costs, and structural instability due to differential degradation rates. Nano-additives (e.g., graphene) and crosslinkers (e.g., glutaraldehyde) may also pose biocompatibility risks, potentially causing cytotoxicity or immune reactions (94).

Liu et al. prepared PLCL/SF conduits combined with NGF-loaded conductive TA-PPy-RGD hydrogel via electrospinning. The RGD-modified tannic acid (TA)-polypyrrole (PPy) hydrogel provided cell adhesion sites, a conductive microenvironment, and sustained NGF release, promoting nerve cell proliferation. Compared to traditional and composite conduits, PLCL/SF/NGF@TA-PPy-RGD showed superior nerve regeneration due to its innovative combination of biocompatibility, bioactivity, and conductivity, highlighting the need for continued research for clinical translation (34). Cheng et al. constructed a PPy-coated PCL/silk fibroin scaffold, finding that ES promoted nerve regeneration and M2 polarization of macrophages, with in vitro ES significantly enhancing M2 pro-regenerative polarization. Bioinformatics analysis revealed differences in signal transducer and activator of transcription expression related to M2 gene promotion, indicating ES’s impact on macrophage phenotype (95).

Nerve conduits, as critical supportive structures, guide and promote nerve regeneration. Functionalizing internal fillers, especially with bioactive molecules, is a key strategy to enhance repair efficacy (96, 97). Bioactive molecules include neurotrophic factors, cytokines, and growth factors, all playing vital roles in nerve repair (98). Loading these factors into hydrogels within conduits can significantly improve the microenvironment, regulate immune responses and inflammation, and promote axonal growth (99).

Neurotrophic factors like nerve growth factor (NGF), brain-derived neurotrophic factor (BDNF), and glial cell-derived neurotrophic factor (GDNF) are widely used in filler functionalization (100–102). They bind to neuronal surface receptors, activate downstream pathways, and promote axonal extension. Especially in the early injury stage, they enhance neuron growth, inhibit apoptosis, and improve functional recovery (103). Hydrogels, as carriers, provide a suitable microenvironment for regeneration and enable sustained factor release via controlled-release systems, maintaining long-term bioactivity.

Growth factors like vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF) are crucial for improving blood supply and angiogenesis. After nerve injury, blood supply significantly affects regeneration; VEGF promotes vascular regeneration in the injury area, providing oxygen and nutrients to support nerve growth (104). Combined with hydrogel carriers, growth factors achieve long-term stability.

Cytokines like transforming growth factor β (TGF-β) and interleukin-10 (IL-10) regulate the immune microenvironment during nerve repair (105, 106). They inhibit excessive immune inflammation, reduce fibrosis, and promote nerve regeneration and angiogenesis. Loading cytokines into hydrogels creates a regenerative-friendly immune microenvironment, avoiding adverse immune factors (107).

The bioactive molecule functionalization strategy for hydrogel fillers in nerve conduits provides multi-level support for nerve repair, promoting neuron growth, survival, and differentiation while regulating immunity, inflammation, and the local microenvironment for long-term effective recovery (108) Yang et al. first established a pro-inflammatory model to validate the guiding and repairing role of M2 macrophages in long-distance PNI, then developed a self-assembling biomimetic peptide hydrogel scaffold encapsulating M2-derived regenerative cytokines and extracellular vesicles. This scaffold mimicked the immune microenvironment mildly, remodeling the local environment for M2 polarization and recruitment, facilitating long-distance peripheral nerve regeneration (109).

Microstructural functionalization of nerve conduit fillers is a key strategy for promoting nerve repair (37). Fillers must provide physical support and mimic the ECM microenvironment to guide neuron growth, differentiation, and axonal extension. Microstructures can be designed by adjusting porosity, fiber alignment, and surface properties to mimic the natural ECM, promoting nerve regeneration (110).

Porosity and pore size distribution are critical for repair efficacy. A reasonable pore structure provides space for cell adhesion, axonal extension, and angiogenesis (111). Techniques like gas foaming, solvent casting, and 3D printing enable precise control of porosity and pore size to optimize filler performance (112).

Incorporating nanostructures significantly enhances physiological functions of fillers (113). Nanofibers, particles, and meshes mimic the 3D network of natural ECM, providing an ideal microenvironment for nerve regeneration (114). Electrospun nanofibers, in particular, form collagen-like networks within conduits, offering adhesion sites for neurons and guiding axonal orientation via physical cues (115). Studies show that surface topography can influence macrophage phenotype by adjusting cytoskeleton: M2 macrophages are typically more elongated than M1 (116). Oriented electrospun nanofibers can regulate macrophage morphology to an elongated state, promoting M2 polarization and reducing pro-inflammatory factor secretion (117). Small surface pores also reduce inflammatory stimulation of macrophages (118).

Multilevel structural design combines micro- and macro-scale features for better repair effects (119). At the macroscale, ordered pores facilitate cell migration; at the microscale, grooved or oriented nanofibers guide neuron directionality (120). Multilevel structures act synergistically at different scales to enhance regeneration.

Sun et al. inspired by the ordered distribution of nerve fibers and endogenous electric fields, studied the synergistic effect of ES and topological cues on the immune microenvironment during peripheral nerve regeneration. They fabricated a nerve conduit with oriented electrospun nanofibers using a polyurethane copolymer containing conductive aniline trimer and degradable L-lysine. In vitro results showed the conduit promoted macrophage recruitment, induced M2 polarization, and enhanced Schwann cell migration and myelination. In a rat sciatic nerve injury model, it increased M2 macrophage proportion and improved regeneration, demonstrating potential applications from an immunomodulatory perspective (121).

Furthermore, accumulating evidence has elucidated the regulatory role of material stiffness in modulating innate immune cell responses. Mechanistic studies demonstrate that elevated substrate stiffness tends to promote macrophage polarization toward pro-inflammatory (M1) phenotypes, while uncoated high-stiffness substrates paradoxically induce anti-inflammatory (M2) polarization. Neutrophils exhibit reduced migration velocity yet enhanced spreading on rigid matrices, potentially exacerbating inflammatory responses through enhanced extracellular trap formation. Notably, natural killer (NK) cells leverage mechanosensing capabilities via surface stiffness receptors, with rigid substrates shown to upregulate cytotoxic markers and potentiate target cell elimination (122, 123). Fang et al. developed an electroconductive multiscale nerve conduit featuring multilayered stiffness gradients through architectural engineering (32). This design achieved coordinated macrophage repolarization toward M2 phenotypes via stiffness-dependent mechanotransduction. Additionally, aligned microfibers with stiffness gradients served as topographical guidance cues, facilitating directed migration of vascular endothelial cells and macrophages to accelerate neovascularization, thereby establishing metabolic support systems for functional regeneration.

Effective and sustained drug delivery is crucial for nerve regeneration. Nerve conduits, as supportive structures, can incorporate controlled-release systems in fillers to maintain local drug concentration and enhance repair (124). Controlled drug release has become a key functionalization strategy, enabling sustained, timed, and quantitative drug release to maximize efficacy (125).

Design of controlled-release systems relies on various materials and mechanisms, including hydrogel-based sustained-release systems (126), nanocarriers (127), and self-responsive release systems (128). Hydrogels, with good biocompatibility and adjustable degradation, are ideal carriers: crosslink density, hydrophilicity, and pore structure can be tuned for precise release control (129). They maintain a moist microenvironment and stable long-term drug release, ensuring sustained action on nerves. Anti-inflammatory drugs, for example, are released gradually via diffusion or hydrolysis, promoting repair (130).

Nanotechnology offers new approaches for controlled release: nanoparticles (e.g., lipid, polymer nanoparticles) efficiently load and target drugs, preventing degradation or rapid loss. Particle size, surface properties, and composition can be adjusted for precise release (92, 131). As reported by Wan et al., a nerve conduit featuring a dynamic 3D interconnected porous network was developed by integrating chitosan with multifunctional gelatin microcapsules. These gelatin microcapsules served as drug carriers and demonstrated efficient adsorption of positively charged insulin-like growth factor-1 (IGF-1) through electrostatic interactions. The degradation-controlled release mechanism of gelatin effectively regulated drug delivery kinetics, preventing both localized drug overdosage and premature depletion. The study revealed that IGF-1 synergized with the 3D porous architecture to promote macrophage polarization toward the M2 phenotype via activation of the PI3K/Akt signaling pathway. This immunomodulatory effect established an anti-inflammatory microenvironment, which significantly enhanced the regeneration of sciatic nerve defects in rat models (132).

Self-responsive systems, another emerging direction, adjust release rates based on microenvironmental changes (pH, temperature, electric field) during nerve repair. As pH and temperature change with regeneration, self-responsive hydrogels release drugs adaptively, optimizing efficacy (133, 134).

Multi-drug co-delivery is also promising: combining neurotrophic factors (135), antioxidants (136), etc., addresses the complex biology of nerve repair. Controlled-release systems enable synergistic release of multiple drugs at different rates and time windows, ensuring stage-specific effects. Dexamethasone, a common immunosuppressive small molecule, is widely used in implantable materials for its low cost and availability. Kumar et al. found silk hydrogels sustained release of dexamethasone and IL-4, promoting M2 polarization and maintaining tissue function (137).

Yu et al. regulated zein-induced immune responses by adjusting the size, pore structure, and dexamethasone loading of zein microspheres, inhibiting neutrophil recruitment and promoting M2 polarization via pore structure design. They then prepared high-porosity zein conduits loaded with zein microspheres for bridging 15 mm sciatic nerve defects in rats, showing better repair efficacy at the degradation stage, with porous microspheres enabling sustained drug release to enhance tissue repair (138).

The core goal of cell loading in nerve conduit fillers is to introduce cells or cell-derived materials that directly support nerve repair. Unlike traditional drug loading, cell loading participates in repair via multiple mechanisms: promoting neuron growth/differentiation, regulating the local microenvironment, and enhancing neural network reconstruction (139). Stem cell loading is a key direction, with mesenchymal stem cells (MSCs) (140) and induced pluripotent stem cells (iPSCs) (141) being promising due to their self-renewal and differentiation potential into neurons or glial cells. These cells differentiate into neural tissue at the injury site, secrete neurotrophic/growth factors, and promote repair (142). Glial cell loading, especially astrocytes, is another strategy. They secrete growth/nutritional factors and ECM, support neuron growth, participate in local immunity, clear cellular debris, and maintain microenvironmental stability, accelerating healing and axonal regeneration (143). Fibroblast loading provides structural support via collagen/ECM secretion, forming a regenerative matrix and releasing local growth factors to facilitate repair, improving conduit stability and microenvironment (144, 145). Yuan et al. inspired by spinal cord anatomy, developed a multi-channel fibrous conduit with multicellular distribution for spinal cord injury repair. Using directional freeze-casting, they created a conduit with hierarchical parallel channels and oriented layered structures, inoculating MSCs in central channels and Schwann cells in peripheral channels. In vitro, cell interactions promoted Schwann cell migration, MSC differentiation, endothelial cell angiogenesis/migration, and M2 polarization of macrophages. In vivo, it reduced glial scarring, promoted neuron regeneration, myelination, and functional recovery in rats (146).