Engineered Delivery Systems for Chinese Herbal Medicine in Peripheral Nerve Regeneration: Challenges, Advances, and Future Perspectives

Introduction

Peripheral neuropathy (PN) encompasses a range of disorders characterized by sensory, motor, and autonomic dysfunctions, representing a substantial and growing global health challenge.¹ It is estimated that PN affects millions of people worldwide, with a prevalence rising sharply with age to afflict up to 20% of the population over the age of 60.^2–4 Diabetic peripheral neuropathy (DPN), one of the most common forms, exemplifies this burden, impacting approximately 50% of all diabetic patients over their lifetime.⁵ These patients experience symptoms such as numbness, neuropathic pain, and impaired bodily functions.^6,7 With various etiologies including metabolic disorders, trauma, infections, and toxins, PN constitutes a significant therapeutic burden for which current therapies cannot cure.^8,9 The existing strategies mainly focus on symptomatic treatment. For DPN), glycemic control has shown inconsistent effectiveness, and first-line pain relievers such as duloxetine and pregabalin offer limited symptom relief, complicated by side effects and dosing issues.^10,11 The treatment of chemotherapy-induced peripheral neuropathy (CIPN) is similarly challenging, with a significant proportion of patients enduring symptoms even after discontinuing chemotherapy.¹² Current management strategies emphasize the importance of early detection and dose adjustment to minimize neurotoxicity, while maintaining anti-tumor efficacy.¹³ Novel approaches, such as high-concentration capsaicin and intravenous lidocaine infusions, have shown potential in reducing neuropathic pain from CIPN, but further studies are necessary to ensure their long-term safety.^14,15 This therapeutic deficiency is due to core biological challenges: the pathological heterogeneity demanding etiology-specific interventions;¹⁶ and the dynamic spatiotemporal evolution of nerve injury microenvironments through conflicting inflammatory, reparative, and fibrotic phases.^17–19 In addition, the systemic administration of drugs is a conventional approach for delivering therapeutic agents to injury sites. However, the effectiveness of this approach is substantially limited by the blood-nerve barrier (BNB), which is a highly selective protective shield that prevents most drugs from entering the injury site via the circulatory system.^20,21 In addition, severe structural damage (>3 cm nerve gaps) exceeds the endogenous regenerative capacity, whereas conventional reconstructive techniques have biological and mechanical limitations.^22,23

Chinese herbal medicine (CHM), traditionally administered as decoctions or oral powders, has long been recognized for its therapeutic potential for a wide range of medical conditions.²⁴ It offers a promising multifactorial alternative through neuroprotective, anti-inflammatory, anti-oxidant, and microenvironment-modulating bioactivities.²⁵ Yet, these conventional forms inherently suffer from substantial pharmacological drawbacks that severely limit their clinical translation, such as poor oral bioavailability of bioactive constituents,²⁶ non-targeted biodistribution resulting in insufficient nerve accumulation,²⁷ BNB impermeability to complex phytomolecules,²⁸ and physicochemical incompatibilities in multi-component formulations that undermine their ability to work synergistically.²⁹ Recent advancements in engineered delivery systems, particularly those utilizing nanotechnology, hydrogels, degradable membranes, and functionalized nerve conduits, offer promising solutions to these challenges.

Furthermore, this review provides a novel conceptual framework that systematically links the multifaceted therapeutic challenges of PN, such as pathological heterogeneity, spatiotemporal dynamics, BNB impermeability, microenvironment imbalance, and structural defects, with the design features of engineered delivery systems. We propose a structured decision classification to guide the selection of appropriate delivery platforms, such as nanocarriers, hydrogels, degradable membranes or functionalized nerve conduits, according to specific pathological phases and biological barriers. This framework is cross-referenced throughout the review to clarify when and why a particular system is advantageous, providing a strategic roadmap for the rational design and clinical translation of CHM-based therapies for PN.

This review comprehensively examined the burgeoning field of engineered CHM delivery systems for PN. As a narrative review, it aims to provide a broad, critical, and synthesized overview of the current state of the art, identify key challenges, and propose a novel conceptual framework for the field, rather than to systematically evaluate all available evidence. We first outline the complex therapeutic challenges that current strategies fail to adequately resolve, and critically analyze the inherent limitations of CHM that impede its clinical efficacy despite its therapeutic potential. The core focus is on innovative engineering solutions, where we dissect the design principles, mechanisms of action, preclinical efficacy, and persisting translational hurdles of key delivery platforms, including nanocarriers, hydrogels, degradable slow-release membranes, and functionalized nerve conduits. By integrating these advanced techniques and materials, our review not only presents a distinct viewpoint but also provides a clear strategy for addressing existing challenges in conventional CHM delivery. By synthesizing promising preclinical cases, this review aims to propose a conceptual framework that could establish a new standard for future research (Figure 1).

Figure 1 Engineered delivery systems overcoming multifaceted barriers in PN treatment. Current therapeutic challenges in nerve repair involve: pathological heterogeneity (requiring distinct strategies for multiple etiologies); spatiotemporal dynamics (demanding precise regulation due to stage conflicts); the blood-nerve barrier (BNB, hindering effective drug penetration into nerves); microenvironmental imbalance (delaying regenerative processes); and long-gap defects (impeding regeneration). Limitations of CHM therapy include: low bioavailability (compromising therapeutic efficacy); non-targeted distribution (causing off-target accumulation); BNB-mediated exclusion of large therapeutic molecules; and physicochemical incompatibilities of multi-herb formulas (reducing synergistic effects). Engineered delivery systems resolve these barriers through four synergistic strategies: (1) Nanocarriers enhance bioavailability, penetrate the BNB, enable targeted delivery, and allow stimuli-responsive spatiotemporal control; (2) Hydrogels degrade synchronously with nerve repair timelines, enable stimuli-responsive spatiotemporal regulation, enhance bioavailability, and modulate the local microenvironment; (3) Slow-release membranes also degrade in sync with nerve repair timelines, improve bioavailability, and optimize the microenvironment; (4) Nerve conduits bridge long-gap defects and ameliorate the microenvironment. Collectively, these engineering approaches overcome critical barriers to promote nerve repair. (By Figdraw).

Therapeutic Challenges and Current Treatment Limitations in Peripheral Neuropathy

Pathological Heterogeneity Challenge

PN encompasses diverse etiologies such as mechanical trauma, metabolic dysfunction, inflammation, infection, malignancy, inherited disease, drugs, and toxins, each demanding a distinct therapeutic strategy.¹⁶ However, current treatments have failed to address this heterogeneity. Similarly, drug treatments such as methylcobalamin injections cannot concurrently resolve different pathologies. However, they do not sufficiently accelerate axonal regeneration to prevent irreversible muscle atrophy.³⁰

Spatiotemporal Dynamics Challenge

The injury microenvironment evolves through three phases with conflicting therapeutic requirements: an early inflammatory phase dominated by TNF-α/IL-1β rise sharply,¹⁷ a repair phase dependent on Schwann cells (SCs) dedifferentiation and neurotrophic factors secretion,¹⁸ and a late fibrotic Phase In which collagen deposition blocks regeneration.¹⁹ Current therapies lack spatiotemporal adaptability. For example, early anti-inflammatory drugs may alter the inflammatory nerve microenvironment, inadvertently inhibiting SCs dedifferentiation, and weakening repair mechanisms.³¹ Systemically administered neurotrophic factors have low bioavailability at lesion sites owing to rapid clearance and missing critical phase-specific windows.²⁸ However, single-type drug formulas cannot be adapted to these shifts.

Biological Barrier Challenge

BNB is a critical component of the peripheral nervous system and serves as a protective shield that maintains homeostasis of the neural environments.³² However, its highly selective nature poses significant challenges for the delivery of large therapeutic molecules, which has garnered considerable attention in recent research. BNB is composed of highly specialized endothelial cells, perineurial cells, SCs, pericytes, basement membranes, and invested axons.³³ The more extensively studied blood-brain barrier (BBB) is characterized by its unique structural and functional properties, including specialized tight junctions and transport mechanisms that restrict the passage of substances.³⁴ This selectivity is crucial for protecting peripheral nerves but simultaneously complicates therapeutic interventions aimed at treating neuropathies and other nerve-related conditions. For example, it prevents the transmission of large molecules such as IgG antibodies, nerve growth factors, and albumin through the body.³³ This limits the effectiveness of the traditional treatments. For example, neurotrophic agents such as methylcobalamin have no significant effect on the treatment of painful diabetic neuropathy.³⁵ Oral analgesics such as gabapentin accumulate predominantly in the liver and kidneys,³⁶ causing systemic toxicity without overcoming BNB exclusion.

Microenvironmental Imbalance Challenge

Chronic conditions, such as diabetes, induce severe microenvironmental dysregulation, elevating M1/M2 macrophage ratios to delay repair.³⁷ SCs under hyperglycemic conditions show impaired production and secretion of neurotrophic factors, thereby hindering regeneration.³⁸ Although stem cells improve neurological function and alleviate clinical symptoms through secretion of neurotrophic factors and immune regulation, different stages of the disease may require different types, doses, and injection routes of stem cells, and long-term risks require more comprehensive observation.³⁹ This dysregulation fundamentally undermines treatment efficacy. Even for specific symptoms of neuropathic pain in DPN, where pharmacological options exist, significant limitations persist, and Duloxetine and Pregabalin are the only drugs approved by the US Food and Drug Administration (FDA) for treating painful DPN. However, these drugs are only effective in approximately 40% of patients.⁴⁰ This exemplifies the broader challenge of achieving adequate symptom control, let alone modifying the underlying disease process, across neuropathies characterized by microenvironmental derangement.

Structural Repair Challenge

Long-gap defects (>3 cm) make it impossible to repair them with internal regeneration ability.⁴¹ Autologous remains the gold standard for gaps exceeding 3 cm, but is hindered by complications such as scarce donor tissue, neuroma formation, nerve distortion or dislocation, and nerve diameter mismatch.^41,42 Although nerve guiding catheters (NGCs) can bridge nerve defects, repair is almost impossible by autologous nerve transplantation. In addition, substances such as cells and neurotrophic factors must be added to provide nutrition and oxygen; however, to date, this integrated biomaterial remains elusive.⁴² This highlights the inadequacy of the current reconstructive approaches for treating severe nerve injuries.

Overall, these multifaceted obstacle barriers underlie the poor efficacy of current peripheral neuropathy treatments, making most treatments palliative rather than restorative, and highlighting the urgent need for innovative approaches.

CHM Treatment and Its Limitations

CHM and its active ingredients have multifactorial advantages in treating peripheral neuropathy, including neuroprotective and anti-inflammatory effects, enhanced anti-oxidant activity, improved microenvironment, promotion of angiogenesis, improved blood circulation, as well as their abundant sources, and cost-effectiveness. The detailed mechanism and application of CHM in the treatment of peripheral neuropathy can be found in Chen’s review (Figure 2).²⁵

Figure 2 Mechanisms of CHM in regulating peripheral nerve repair and regeneration following PN. CHM promotes peripheral nerve regeneration after PN through various mechanisms, including: exerting anti-oxidant effects to mitigate oxidative stress and protect neurons; inhibiting neuroinflammation to suppress inflammatory reactions and reduce cellular injury; mediating neurotrophic and neuroprotective activities to induce neurotrophic factors; reconstructing the local microenvironment to optimize conditions for nerve regeneration; and promoting angiogenesis and circulation to enhance blood flow and nutrient supply.25 Copyright 2025, Springer.

Despite the promising application prospects of CHM, its clinical translation has several pharmacological obstacles. First, oral administration has low bioavailability. For example, Salvia miltiorrhiza undergoes extensive first-pass metabolism (<10% systemic availability),⁴³ whereas the water solubility of hydrophobic agents, such as tanshinone IIA, can be ignored.⁴⁴ Second, non-targeted distribution leads to off-target accumulation, which wastes the dosage and has the potential to cause hepatotoxicity.²⁷ Third, BNB excludes large therapeutic molecules and reduces the effective nerve concentrations below the regenerative thresholds.³³ Fourth, the synergistic interactions in multi-herb formulations are compromised by physicochemical incompatibilities.²⁹ For example, hydrophilic paeoniflorin⁴⁵ and hydrophobic tanshinone⁴⁶ cannot coexist optimally as standard carriers. Collectively, these limitations require engineered delivery systems to unlock CHM’s therapeutic potential.

Engineered Chinese Herb Delivery Systems

Nanocarriers

Nanocarriers represent a first-line strategy in our proposed framework to overcome the dual challenges of BNB impermeability and the demand for spatiotemporal precision in PN therapy. By encapsulating CHM active ingredients, these engineered systems directly address the pharmacological limitations of poor bioavailability and non-targeted distribution. Their subcellular size enables penetration of the BNB through passive diffusion or active ligand-mediated targeting, thereby ensuring sufficient drug accumulation at the nerve injury site. Furthermore, the design flexibility of nanocarriers allows for the incorporation of stimulus-responsive elements, making them uniquely capable of releasing their payload in response to specific phase-specific microenvironmental cues (eg, pH, enzymes). This capability is crucial for addressing the conflicting demands across the inflammatory, repair, and fibrotic phases of nerve recovery.

The utility of this approach is demonstrated by several encapsulation strategies designed to enhance bioavailability and stability. For instance, curcumin-loaded polycaprolactone nanofibers increased cellular uptake and extended neurite growth by 1.7-fold compared to pure polymer controls, as demonstrated in neuronal models.⁴⁷ Lycium barbarum polysaccharide (LBP) encapsulated in poly (lactic-co-glycolic acid) (PLGA) core-shell fibers achieved sustained release for over 60 days, enhancing SCs proliferation by 50%.⁴⁸ Electrospun PCL/gelatin nanofibrous scaffolds enable sustained Ginkgo biloba extract (GBE) release (73.37% over 10 days), which exerts its neuro-regenerative effects through its active compounds, primarily flavonoids and terpenoids. These compounds are responsible for the observed significant reduction in pro-inflammatory cytokines (IL-6, TNF-α) via modulation of macrophage activity. When this GBE-loaded scaffold is combined with adipose-derived stem cells (ADSCs), a powerful synergy emerges. The ADSCs further contribute to immunomodulation and secrete essential neurotrophic factors. Together, this combination treatment leads to a significant upregulation of NGF and BDNF in the sciatic nerve tissue, creating a balanced inflammatory microenvironment and enhancing axonal regeneration.⁴⁹

A key advantage of nanocarriers within our framework is their versatility in overcoming biological barriers and enabling precise intervention. They penetrate the BNB via passive diffusion,^47,48 whereas conductive components such as polyaniline-doped nanofibers reduce the charge transfer resistance to enhance ion exchange and nutrient delivery.⁴⁷ Ligand-mediated targeting strategies enhance site-specific delivery: gambogic acid-conjugated nanoparticles improve systemic bioavailability for DNP treatment,⁵⁰ while other functionalized systems achieve receptor-mediated transport to the dorsal root ganglia (DRG), reducing off-target effects in neuropathic pain models.^51,52

This targeted capability is complemented by spatiotemporal control mechanisms. Stimuli-responsive designs allow for precise drug release aligned with pathological phases; for example, chitosan/hyaluronic acid nanoparticles released 90% payload at pH 6.14 (inflammatory phase) versus 19% payload at physiological pH.⁵³ Aligned electrospun nanofibers, such as poly(3-hydroxybutyrate-co-3-hydroxyvalerate) blended with aloe vera, provide topological guidance for directional axonal regeneration.⁵⁴

These engineered systems collectively demonstrate superior efficacy in preclinical studies, with their functional benefits supported by distinct molecular mechanisms. Conductive nanofibers upregulate neuronal differentiation markers (βIII-Tubulin 2.1-fold; TREK-1 1.8-fold) in PC12 cells.⁴⁷ Crucially, nano-curcumin formulations produce significant functional recovery by targeting specific pathological pathways: they alleviate neuropathic pain by inhibiting P2X₃ and P2Y₁₂ receptors in the DRG and suppressing associated ERK1/2 and Akt signaling,^51,52 and mitigate neuroinflammation by reducing NLRP3 expression and IL-1β maturation, alongside a ~50% reduction in macrophage infiltration.⁵⁰ Furthermore, they address oxidative stress via Nrf2 pathway activation, leading to enhanced remyelination. These actions translate to key functional outcomes, including the restoration of 85% sciatic nerve conduction velocity and a 65% increase in motor nerve conduction velocity in Charcot-Marie-Tooth (CMT) models,⁵⁵ as well as the normalization of intraepidermal nerve fiber density in DNP at half the dose of free curcumin.⁵⁰ Electrical stimulation of conductive scaffolds extended the DRG neurite length by 31%.⁵⁶ In vivo, PLGA-LBP scaffolds promoted neurite extension for 3.4-fold longer than non-engineered groups⁴⁸ (Table 1 and Table 2).

Table 1 Comprehensive Summary of in vivo Animal Studies on Engineered CHM Delivery Systems for PN

Table 2 Comprehensive Summary of in vitro Studies on Engineered CHM Delivery Systems for PN

Despite these advances, its clinical translation has faced significant hurdles. Scalability issues arise from batch inconsistencies in coaxial electrospinning, where fiber diameter variations (295–896 nm) compromise drug release consistency.⁴⁸ Biocompatibility concerns include inflammatory responses triggered by acidic degradation byproducts of conductive polymers⁷³ and unassessed chronic toxicity of metallic nanoparticles.⁷⁴ Dosage optimization remains challenging, as high LBP concentrations reduce scaffold tensile strength by 22%,⁴⁸ whereas industrial processing diminishes bioactive components such as aloe vera polysaccharides.⁵⁴ Targeting precision requires refinement to prevent off-target effects,^51,52 and regulatory pathways lack standardization for nanocarrier characterization in large-animal models.^49,51

The compelling evidence for BNB penetration, targeted delivery, and spatiotemporal control, as summarized in this section, is exclusively in preclinical investigations. The hypothesis that nanocarriers can serve as a primary strategy to overcome pharmacological barriers in human PN patients awaits validation in clinical settings.

Hydrogels

Within our therapeutic framework, hydrogels emerge as a uniquely suited platform for tackling the challenge of microenvironmental imbalance in PN. Their fundamental advantage lies in the use of biocompatible natural polymers that undergo synchronous degradation with the native process of nerve regeneration. For instance, alginate/chitosan hydrogels demonstrate principle by losing approximately 80% of their mass over 14 days, a timeline that aligns with key phases of peripheral nerve repair.⁵⁷ This timed degradation is designed to supply sustained support throughout the vulnerable regeneration phase.

The functional capacity of hydrogels is significantly advanced by engineering complex stimuli-responsive systems that enable spatiotemporally precise drug release. This design directly addresses the dynamic pathological shifts in PN. NIR-activated curcumin hydrogels utilize π-π stacking with synthetic melanin nanoparticles to achieve on-demand payload release,⁵⁸ whereas collagen-based hydrogels loaded with naringin deliver an initial 22.5% burst release within 12 h for acute inflammation control,⁵⁹ followed by sustained delivery of 72.5% over 14 days to support chronic regeneration.⁵⁹ Conductive hydrogels incorporating reduced graphene oxide exhibit electrical conductivity of 5.27×10⁻⁴ S/cm, enhancing neural differentiation under electrostimulation.⁷¹ Multifunctional designs combine drug delivery with auxiliary therapies: supramolecular hydrogels co-deliver ginsenoside Rg1, Stromal cell-derived factor-1α (SDF-1α), and stem cells to synergistically regulate neurogenesis and angiogenesis,⁶⁰ whereas boronic ester-based hydrogels enable long-term curcumin release (71.4% over 28 days) via dual cross-linking.⁶¹

A pivotal function of hydrogel systems is their ability to simultaneously overcome CHM bioavailability issues and actively correct the dysregulated microenvironment. For hydrophobic compounds like curcumin, encapsulation in lipid nanocapsules achieves 92.7% efficiency,⁶² whereas berberine and naringin nanoparticles in chitosan/alginate matrices improved solubility by 51.6% and 47.1%, respectively.⁶³ Porous architectures with pore dimensions of 22–32 μm in chitosan/alginate hydrogels,⁶³ 50–250 μm in chitosan/oxidized cellulose composites,⁷¹ and 90 μm in collagen scaffolds⁵⁹ facilitate nutrient diffusion and SCs infiltration. Microenvironment correction occurs through two mechanisms: hesperidin-loaded hydrogels sustain 75.8% anti-inflammatory release over 14 days,⁵⁷ whereas curcumin systems scavenge 82.53% of DPPH radicals and shift macrophages toward anti-inflammatory M2 phenotypes via CD206 upregulation.⁵⁸

Hydrogel-based delivery systems have demonstrated significant therapeutic advantages in both in vitro and in vivo studies, supported by their multi-targeted molecular actions. In vitro studies have shown that engineered hydrogels profoundly enhance nerve regeneration. Collagen hydrogels loaded with naringin double SCs proliferation compared to control surfaces, achieving an optical density value of 0.4 versus 0.2 after 72 hours of culture.⁵⁹ Conductive hydrogels incorporating reduced graphene oxide elevated neuronal differentiation to 77.96% in PC12 cells under electrical stimulation, leveraging electroactive properties to accelerate nerve maturation.⁷¹ The regenerative benefits of these hydrogel systems are driven by their capacity to coordinately regulate multiple cellular processes. Key mechanisms include the promotion of Schwann cell proliferation and migration through ERK1/2 pathway activation and MMP-9 upregulation,⁶⁵ the suppression of apoptosis via caspase-3/9 downregulation,⁶⁵ and the enhancement of endogenous antioxidant defenses through Nrf2 pathway activation, elevating levels of SOD and GSH by 30% and 25%, respectively.^57,59,61 A central function is the effective polarization of macrophages from a pro-inflammatory M1 state to a regenerative M2 phenotype, evidenced by reductions in M1 markers of up to 50% and increases in M2 markers by as much as 2.5-fold.^58,60,63,64 Concurrently, these systems dampen neuroinflammation by inhibiting NF-κB signaling^60,64 and directly promote nerve repair by upregulating critical structural proteins such as MBP, βIII-tubulin, and GAP-43, thereby enhancing myelination and axonal extension.^{57,59,60,63,65}

In animal models, functional and structural recovery is markedly superior to non-engineered interventions. Photoresponsive curcumin hydrogels triple myelinated axon density in sciatic nerve injuries, generating 360 fibers/μm² versus 104.67 fibers/μm² in crush-injury controls.⁵⁸ Functional restoration is equally striking: berberine-loaded alginate/chitosan hydrogels improve sciatic functional index (SFI) to −20.23 ± 1.35 after 8 weeks, dramatically surpassing untreated groups (−76.90 ± 4.12) and pure hydrogel controls (−46.00 ± 1.60).⁶⁴ Pain management capabilities are evidenced by curcumin nanocapsule hydrogels reducing thermal hyperalgesia and shortening paw withdrawal latency to 6 s compared to 12s in baseline models.⁶² Chronic neuropathy interventions have shown sustained efficacy, with boronic ester-based hydrogels enhancing nerve conduction velocity to 25 m/s versus 10 m/s in constriction injury models.⁶¹ Microenvironmental corrections underlie these outcomes. Curcumin systems shift macrophage polarization toward regenerative M2 phenotypes, elevating CD206 expression while suppressing pro-inflammatory cytokines.⁵⁸ Keratin-chitosan/curcumin hydrogels reduced muscle atrophy to 5.7% weight loss versus 40% weight loss in denervated controls, confirming the integrated tissue protection.⁶⁵ These multi-mechanistic effects collectively validate hydrogel carriers as potent platforms for nerve repair (Table 1 and Table 2).

Despite the compelling efficacy, four major challenges impede clinical adoption: dose optimization requires balancing efficacy and safety, as concentrations exceeding 1% naringin clog hydrogel pores⁵⁹ whereas 10% berberine induces cytotoxicity.⁶⁴ Degradation synchronization remains problematic, with alginate/chitosan hydrogels degrading too rapidly (80% mass loss in 14 days) for slow nerve regeneration,⁵⁷ whereas boronic ester hydrogels degrade only 20% over 28 days.⁶¹ Manufacturing complexity hinders the scale-up of multifunctional systems such as liposome-hydrogel hybrids.⁷¹ Technical dependencies on clinical-grade near-infrared/electrostimulation devices limit accessibility,^58,71 whereas long-term safety concerns persist regarding keratin immunogenicity⁶⁵ and metallic nanoparticle toxicity.⁷¹

The ability of hydrogels to correct microenvironmental imbalances and provide phase-adapted release, as demonstrated in these animal studies, positions them as a uniquely promising platform for PN treatment. Translating this microenvironment-modulating capacity into clinical benefit represents a central translational hypothesis moving forward.

Degradable Slow-Release Membranes

Within our therapeutic framework, degradable slow-release membranes emerge as a specialized platform for resolving the critical challenge of localized and sustained pharmacotherapy in PN. Engineered from biodegradable composites like chitosan-polyethylene oxide (PEO) and polylactic acid-glycolic acid-polyethylene glycol (PLGA-PEG), these membranes exhibit excellent biocompatibility and controlled degradation profiles. This design avoids severe immune responses while aligning with the extended timeline of nerve regeneration. For instance, chitosan-PEO membranes fabricated as thin films gradually degrade without inducing scar formation, providing a favorable microenvironment for axonal growth.⁶⁶ PLGA-PEG membranes, combining biodegradable PLGA with hydrophilic PEG, degrade over 85% in vitro within 10 weeks, matching the prolonged recovery process of nerve injuries, and their curved structure conforms to anatomical sites, such as the cavernous nerves, ensuring efficient localized drug delivery.⁶⁷

The core function of these membranes lies in their ability to maintain stable, long-term drug release directly at the injury site. They enable sustained and dose-dependent release of curcumin: chitosan-PEO membranes maintain local curcumin concentrations to avoid systemic toxicity, while PLGA-PEG membranes achieve stable release (72.5–89.8% over 45 days) without burst effects, with higher curcumin ratios (1.4%) accelerating release to meet dynamic healing needs.^66,67 Additionally, their structural adaptability, such as trimable chitosan-PEO films for sciatic nerves and conformable PLGA-PEG membranes for cavernous nerves, facilitates their precise application. Multifunctional integration, such as chitosan’s anti-fibrotic properties combined with curcumin’s anti-inflammatory effects or PLGA-PEG’s enhancement of curcumin solubility alongside nNOS upregulation, further improves therapeutic efficacy.^66,67

Beyond their pharmacokinetic benefits, these membrane systems actively contribute to healing by mitigating bioavailability issues and correcting a dysregulated microenvironment. PLGA-PEG membranes maintain local curcumin concentrations for 45 days, whereas chitosan-PEO membranes reduce systemic clearance and enhance accumulation at injury sites.^66,67 These systems correct dysregulated microenvironments by scavenging reactive oxygen species (ROS), inhibiting pro-inflammatory cytokines, and reducing collagen I expression.^66,67

In preclinical studies, engineered membranes have demonstrated superior efficacy compared to non-engineered membranes. For sciatic nerve transection in rats, chitosan-PEO + curcumin improves SFI (−64.9 ± 5.1 vs −75.3 ± 5.2 in the curcumin-only group) and increases myelinated fiber count (46.7 ± 7.4) and myelin thickness (2.19 ± 0.7 μm).⁶⁶ In cavernous nerve injury models, high-dose (1.4%) curcumin-loaded PLGA-PEG membranes improve erectile function (ICP/MAP: 0.68 ± 0.07) versus 0.51 ± 0.05 in the low-dose group and 0.34 ± 0.07 in the blank membrane group, while upregulating nNOS expression and reducing penile fibrosis⁶⁷ (Table 1).

Despite promising results, its clinical translation is challenging. Dose optimization is critical, as higher curcumin concentrations (1.4%) in PLGA-PEG membranes enhance efficacy but may increase local toxicity.⁶⁷ Membrane degradation rates must be synchronized with nerve regeneration; PLGA-PEG membranes degrade faster in vivo (~70% over 4 weeks) than in vitro, with the risk of premature loss of drug release.⁶⁷ Scalability and standardization of membrane fabrication, which ensures uniform drug loading and structural consistency, remain hurdles.⁶⁶ In addition, adapting the membrane to different types of nerve injuries (such as sciatic and cavernous nerves) needs to be validated in the human body to ensure clinical applicability.⁶⁷

The localized and sustained pharmacotherapy achieved by degradable membranes in preclinical nerve injury models highlights their potential for focal neuropathies. The critical question of whether these release profiles and functional outcomes can be replicated in human nerves remains to be clinically investigated.

Functionalized Nerve Conduits

Within our therapeutic framework, functionalized nerve conduits constitute a definitive strategy for overcoming the challenge of structural repair in long-gap nerve defects. These conduits, crafted from polymers, such as chitosan, polyurethane (PU), polyacrylonitrile/chitosan (PAN/CS), and conductive composites, are engineered to align with the complexities of nerve repair. Chitosan-based conduits, for instance, feature tapered ends with differing inner diameters (0.8 mm at the thinner end and 1.8 mm at the thicker end), a design that addresses the size mismatch between proximal and distal nerves, which is an issue that plagues traditional uniform-diameter conduits. This allows tension-free suturing and reduces the risk of neuroma formation when bridging the proximal common peroneal nerve and distal mixed nerves (common peroneal and tibial).⁶⁸

The functionality of these conduits is greatly enhanced by integrating drug delivery and active signaling capabilities. Gastrodin-modified PU conduits integrate biodegradability with sustained drug release: gastrodin is stabilized via covalent binding, avoiding rapid metabolism and maintaining local concentrations (with approximately 83.3% released over 18 days), while their conductivity (around 1.23×10⁻³ S/cm) supports electrical stimulation to boost nerve regeneration.⁶⁹ Other systems, like conductive L-ZBZPGM conduits incorporate graphene oxide for conductivity and feature interconnected pores (100–200 μm) that facilitate nutrient exchange, creating a favorable environment for SCs infiltration and axonal growth.⁷² PAN/CS conduits, which combine the mechanical strength of synthetic PAN with the biocompatibility of chitosan, serve as dual carriers for berberine and stem cells, thereby enhancing cell adhesion and targeted drug delivery.⁷⁰

These integrated systems produce their restorative effects through multi-mechanistic actions on the nerve regeneration process. Berberine inhibits TNF-α and iNOS,⁷² gastrodin upregulates BDNF/GDNF to enhance myelination,⁶⁹ and chitosan conduits loaded with the modified formula Radix Hedysari (MFRH) stimulate the formation of myelinated fibers (approximately 46.7 fibers/μm²).⁶⁸ Functionally, chitosan/MFRH conduits improve the SFI (−64.9 versus −78.5 in controls) and nerve conduction velocity (approximately 35 m/s vs 20 m/s in controls).⁶⁸ PAN/CS conduits loaded with berberine and endometrial mesenchymal stem cells (EnMSCs) enhanced new nerve formation (18.5% vs 0.2% in controls) and reduced muscle atrophy, with denser and more organized myelinated fibers⁷⁰ (Table 1 and Table 2).

Despite promising preclinical results, its clinical translation has been hindered by several challenges. Optimization of the dose is crucial. Berberine at 200 μM slightly reduced EnMSC viability; therefore, precise adjustment is necessary, with 100 μM recognized as the optimal level.⁷⁰ Synchronizing degradation with nerve regeneration remains difficult; chitosan conduits degrade by approximately 80% within 16 weeks, whereas PAN/CS conduits degrade too slowly.^68,70 Additionally, conductive conduits rely on external electrical stimulation devices, which limits their use in resource-constrained settings.^69,72

While functionalized conduits have demonstrated superior efficacy to autografts in specific rodent models of long-gap defects, these findings constitute a preclinical proof-of-concept. Whether they can become a definitive solution for human nerve injuries is a remarkable but unconfirmed hypothesis.

Challenges in Clinical Translation

Although engineered CHM delivery systems present a promising avenue for PN, several challenges impede their translation from laboratory research to clinical application. Understanding these challenges is crucial to develop feasible and effective engineering therapies for PN.

Material Biocompatibility and Long-Term Safety

The long-term biocompatibility of engineered materials is a foundational challenge that directly impacts the practical utility of all platforms in our therapeutic framework. Ongoing concerns exist regarding the chronic inflammation caused by nanocarriers and conduits.^75–77 Nanoparticles interacting with the immune system may provoke adverse immune responses, as they can stimulate immune cells and lead to inflammation or severe immune reactions.⁷⁸ The potential immunogenicity of hydrogels requires comprehensive long-term human studies.⁷⁹ Additionally, complex multi-component systems, such as liposome-hydrogel hybrids or drug- and cell-integrated functionalized conduits, require thorough research beyond preliminary biocompatibility assessments. Furthermore, nanoparticles can accumulate in different organs, including the liver, spleen, and kidneys, causing organ-specific toxicity.⁸⁰ Consequently, the long-term effects of these materials remain unclear and require thorough preclinical toxicity assessment.

Precise Dosage Optimization and Controlled Release

Determining the optimal therapeutic dose within an engineered system is a complex challenge that sits at the heart of the spatiotemporal control principle.⁸¹ A dosage that is effective during the inflammatory phase may be insufficient or even counterproductive in the regenerative phase, highlighting the need for phase-specific dosing strategies. Although high concentrations of active compounds may enhance efficacy, they can also induce cytotoxicity, disrupt the delivery system framework, and cause local toxicity.^82,83 Conversely, insufficient dosing failed to achieve the therapeutic threshold.⁸⁴ Aligning the degradation rate of the carrier material with the nerve repair timeline is equally challenging, but vital for sustained, phase-appropriate drug release. Predictable and linear release profiles are still hard to achieve.

Targeting Specificity and Overcoming Biological Barriers

The challenge of targeting specificity directly determines whether the strategic positioning of nanocarriers in our framework is feasible. Although ligand-mediated strategies are promising for improving targeted delivery, further refinement is required to ensure precision and minimize off-target accumulation and systemic toxicity. This imperfect targeting undermines the core advantage of nanocarriers—their ability to provide high local drug concentrations at the nerve injury site while minimizing side effects. The effectiveness of these systems in crossing the BNB complex in human PN needs rigorous validation in more clinically relevant models to confirm their role as a primary solution for BNB impermeability. The effectiveness of these systems in crossing the BNB complex in human PN needs rigorous validation in more clinically relevant models to confirm whether they can serve as the primary solution for BNB impermeability.

Integration with Clinical Procedures and Accessibility

The dependency of many advanced systems on external stimuli creates a significant gap between their theoretical potential within our framework and their practical clinical utility.^85,86 Conductive scaffolds and conduits require integration with clinical-grade electrical stimulation devices, which may not be readily available or practical in a clinical setting.⁸⁷ Similarly, photoresponsive systems rely on specific NIR light sources for controlled drug release.^88,89 The technical complexity and costs associated with these dependencies limit their clinical accessibility, particularly in areas with limited resources.

Scalability, Manufacturing Consistency, and Quality Control

Reproducible and scalable manufacturing is a translational precondition that determines whether the precise design features central to our framework can be reliably replicated. This process requires reliable acquisition of precise design parameters, which are crucial for therapeutic efficacy. But when transitioning from bench-scale synthesis to industrial production under Good Manufacturing Practice (GMP) standards, achieving this becomes even more challenging. A primary obstacle lies in the inherent variability of precision manufacturing techniques. Critical processes such as coaxial electrospinning for core-shell fibers and microfluidics for nanocarrier synthesis are easily affected by batch-to-batch inconsistencies. These variations are reflected in key physical attributes, including fiber diameter and nanoparticle size distribution, ultimately compromising the uniformity of drug loading and release kinetics and leading to unpredictable biological effects.^90–92 Scaling these processes while maintaining strict control over these critical quality attributes represents a significant engineering hurdle. Beyond process control, the consistent sourcing of biomedical-grade natural and synthetic polymers is fundamental. Moreover, industrial-scale processing may degrade the sensitive bioactive compounds in CHM, as their complex chemical profiles can be altered by factors such as shear stress, exposure to solvents, or elevated temperatures encountered during manufacturing.^93–95 Finally, establishing robust and validated analytical methods is crucial. The successful transition to industrial production requires the adoption of a “Quality by Design” framework. This approach not only requires in-process monitoring and control of the carrier’s physical properties, but also the chemical consistency and quality of the encapsulated CHM actives, which can be compromised by physicochemical interactions within the formulation^96,97 and processing variables.⁹⁸ The comprehensive standardization of these complex manufacturing processes to meet GMP standards constitutes a major obstacle, directly affecting product safety, efficacy, and ultimate clinical success.

Regulatory Pathways and Standardization

The clinical translation of engineered CHM delivery systems is further complicated by their status as combination products, integrating device, drug, and sometimes biological components, which places them within a complex and evolving regulatory landscape. A foundational challenge is their classification by major regulatory agencies such as the FDA and European Medicines Agency (EMA). A pivotal first step involves determining the product’s Primary Mode of Action (PMOA)—that is, whether the principal therapeutic action is attributable to the physical scaffold or the encapsulated herbal active ingredient. This classification dictates the lead regulatory center and the specific set of guidelines for approval, a decision that is often ambiguous for novel platforms where the scaffold itself exerts bioactive effects.^99,100 This classification challenge is compounded by a lack of standardized preclinical testing protocols. There is a pressing need for harmonized methods to evaluate long-term biodegradation, chronic toxicity of degradation byproducts, and biodistribution.^99,101,102 The multicomponent nature of CHM formulations introduces additional complexity, requiring novel regulatory science approaches to ensure consistent quality, safety, and efficacy, particularly concerning potential physicochemical interactions and synergistic effects.^96,97 Consequently, the Chemistry, Manufacturing, and Controls (CMC) documentation required for a regulatory submission is exceptionally demanding. It must provide a comprehensive and validated account of the entire manufacturing process, define all critical quality attributes, and demonstrate product stability.^99,100 Therefore, generating necessary long-term safety and batch to bioequivalence data for these complex products is a resource intensive and time-consuming task, posing significant barriers to their market access.

Bridging the Preclinical-Clinical Gap

Translating the success from rodent models to humans remains the ultimate validation hurdle for the entire proposed framework. It is important to verify the effectiveness and safety of these delivery systems in large animal models that can better replicate human nerve anatomy, regeneration rate, and disease pathophysiology; however, resources are often insufficient.^100,102 Moreover, proving that there is a significant improvement in functional recovery compared to current clinical standards, considerable time, and investment are required. A delivery system designed for specific types of nerves, such as the sciatic nerve conduit and sponge-like nerve membrane, must be carefully clinically validated to adapt to the different anatomical and pathological environments encountered in human PN.

Clinical Evidence and Gaps

Current Clinical Status: A Preclinical Landscape

A systematic search of international and domestic clinical trial registries (including ClinicalTrials.gov and the Chinese Clinical Trial Registry) confirms that, to date, no clinical trials have been initiated to evaluate the engineered CHM delivery systems discussed in this review for the treatment of peripheral neuropathy in humans. While clinical trials exist for oral CHM extracts and for synthetic drug-loaded advanced delivery systems, the specific combination of CHM bioactives with engineered carriers (nanocarriers, hydrogels, membranes, conduits) remains an unknown field in clinical research. The convincing therapeutic outcomes detailed throughout this review are entirely derived from preclinical models. Consequently, all assertions regarding the efficacy, optimal design parameters, and mechanisms of action of these systems must be anchored in their preclinical nature. The conclusions drawn serve to generate hypotheses for their future clinical application rather than to confirm their therapeutic value in humans. This clearly highlights a significant and critical translational gap.

Mapping the Minimal Clinical Package

Translating promising preclinical findings into clinical application requires a clear, step-by-step development pathway. Making this transition depends on generating a core set of evidence that addresses the special challenges of these combination products. The first and most crucial step is to conduct in-depth safety evaluations according to GLP standards. These studies, ideally performed in two different species, must go beyond standard toxicity screening to specifically investigate the long-term fate of the delivery systems. Critical questions that need answers include: whether the implantation site might cause long-term local inflammation, the specific patterns of nanocarrier accumulation and clearance rates in organs, and the safety of the materials themselves and their breakdown products for the nervous system.

At the same time, it is essential to fully understand how these systems behave in the body, namely their pharmacokinetics and distribution, in large animal models relevant to human medicine. Data from rodent studies, although valuable, are not sufficient to predict how these systems will perform in humans. Research must precisely measure how efficiently they cross the BNB, confirm whether the targeted delivery claimed for functionalized nanocarriers truly reduces accumulation in non-target organs, and clarify the full metabolic pathways of these complex systems. This stage is vital for reducing the risks of human trials, as it checks whether the basic pharmacological principles observed in small animals still hold true in a physiological environment more similar to ours.

The design of the first human trials should be guided by a careful, step-by-step approach to managing risks. A prudent strategy for a First-in-Human study would focus on a clearly defined group of patients where the delivery system can be applied directly to a specific area. An example would be during surgery to repair a localized nerve injury. This method confines the treatment to a precise location, greatly reduces exposure to the rest of the body, and makes it easier to evaluate how well the local area tolerates the implant. The main goal of such a Phase I study would be to build a complete safety profile, while measurements of functional recovery could serve as initial indicators to look for signs of effectiveness. By outlining this route, which encompasses rigorous safety and distribution studies in large animals and a carefully planned initial clinical investigation, the field can build a solid bridge connecting the current state of preclinical research to future therapeutic application.

Future Perspectives

The development of engineered traditional CHM delivery systems has brought great hope for PN treatment. To bridge the gap between preclinical promise and clinical reality, future efforts must be strategically guided by the framework proposed in this review, focusing on enhancing the intelligence, scalability, and personalization of these platforms.

We hypothesize that the next generation of CHM delivery systems should evolve towards intelligent multifunctionality to fully realize the spatiotemporal control principle of our framework. By combining stimulus response mechanisms such as enzyme-, redox-, or biomarker-triggered release with real-time biosensing, it is possible to autonomously adapt to dynamic nerve injury microenvironments.^103–106 For instance, systems that detect elevated TNF-α or ROS levels to precisely release anti-inflammatory agents or neurotrophic factors during the inflammatory or regenerative phases would overcome current limitations in spatiotemporal control. By combining CHM with advanced models, such as gene therapy or immune modulators, multiple pathological pathways can be synergistically targeted to enhance regeneration outcomes and address heterogeneity issues.^42,107–110

Scalable production and rigorous validation of long-term safety are crucial for clinical implementation. Scalable and repeatable production requires advanced technologies, such as continuous flow microfluidics for nanocarriers and automated 3D bioprinting for hydrogels/conduits.^111–113 This ensured consistency in drug loading, structural characteristics (such as pore size and fiber diameter), and degradation rates across different batches.¹¹⁴ Long-term biocompatibility must be rigorously validated through large-scale animal studies, such as primate models, with a focus on the chronic toxicity analysis of degradation byproducts and immunogenicity.^115–117 There is an urgent need for standardized regulatory frameworks for combination products, including devices, drugs, and biologics, along with harmonized protocols for assessing biodistribution, bioaccumulation, and immunotoxicity.

Personalization is proposed as a critical frontier for optimizing therapeutic efficacy and safety in future clinical translation.^118,119 Using clinical imaging data, 3D-printed conduits and hydrogels can be customized to match the anatomical structure of nerve defects and support individualized CHM dosing.^120–122 Combining machine learning with multi-omics data shows potential for predicting optimal CHM combinations and doses, mitigating physicochemical incompatibilities and cytotoxicity risks.^123,124

Improving accessibility and clinical practicality are crucial for its widespread adoption. By creating self-powered systems and reducing the dependence on external stimuli, such as clinical-grade near-infrared devices and electrical stimulators, their usefulness will greatly increase, particularly in environments with limited resources.^125,126 Cost-effective solutions, such as optimizing CHM extraction efficiency or utilizing biosynthetic analogs of rare herbal compounds, could enhance affordability.^127,128 Ultimately, successful translation depends on sustained interdisciplinary collaboration among material scientists, neural engineers, herbal pharmacologists, clinicians, and regulatory experts.

Limitations

As a narrative review, this work has certain inherent limitations. The selection of literature, while aimed at representing key advances and illustrating the proposed framework, was not based on a systematic, reproducible search of all databases. Consequently, it may be subject to selection bias, and some relevant studies might have been overlooked. Furthermore, narrative reviews do not typically include a formal risk-of-bias assessment of the included preclinical studies, which is a standard component of systematic reviews. Therefore, the findings and conclusions presented here should be interpreted as a critical narrative synthesis and a proposal for a new perspective, intended to guide and inspire future research, rather than as an exhaustive or quantitatively definitive summary of the evidence.

Conclusions

Preclinical evidence positions engineered CHM delivery systems as a potentially transformative approach to overcoming the multifaceted therapeutic challenges of PN. This review has not only organized these advances but also established a conceptual framework. It systematically links the core pathological challenges of PN—including pathological heterogeneity, spatiotemporal dynamics, BNB impermeability, microenvironmental imbalance, and structural defects—to the functional capabilities of specific delivery platforms. Through this perspective, we have organized a clear taxonomy: nanocarriers are the primary solution for BNB penetration and targeted, stimuli-responsive delivery; hydrogels excel at modulating the microenvironment and providing sustained, phase-adapted release; degradable membranes offer localized and prolonged pharmacotherapy for focal injuries; and functionalized nerve conduits are indispensable for bridging long-gap structural defects.

By addressing the inherent limitations of CHM, these engineered platforms have unlocked its multifactorial therapeutic potential, demonstrating significant preclinical efficacy in enhancing nerve regeneration and functional recovery. However, significant translational hurdles remain to be overcome. Key challenges include ensuring the long-term biocompatibility of materials, maintaining consistent scalable manufacturing, and optimizing the dosage to balance efficacy with cytotoxicity. Dependence on external triggers restricts accessibility and the absence of standardized regulatory frameworks makes clinical adoption more complex. Crucially, successful translation requires validation in large animal models that replicate human neuroanatomy and disease progression, which demands interdisciplinary collaboration and substantial investment.

In the future, combining smart multifunctionality, personalized design, and affordable production will be important. Collaboration among material scientists, clinicians, herbal pharmacologists, and regulatory authorities is crucial to move these systems from preclinical to clinical applications. If these challenges are resolved, engineered CHM delivery can redefine PN management, shift from palliative care to restorative therapy, and establish a new standard for precision neuromedicine.

In summary, this review synthesizes a compelling preclinical data for engineered CHM delivery systems. The framework and conclusions presented herein are primarily hypothesis-generating, outlining a roadmap for future research that must now be validated through rigorous clinical investigation.