Coaxial nanofiber design for traumatic brain injury: a review and framework for phase-responsive therapeutics

Abstract

Traumatic brain injury (TBI) is characterized by an evolving pathophysiology spanning acute, subacute, and chronic stages, each demanding temporally tailored therapeutic interventions. However, conventional drug delivery systems lack the capacity to adapt to these shifting biological windows, limiting therapeutic precision and clinical efficacy. Coaxial electrospun nanofibers, with their spatially distinct core–shell architecture, offer a platform for phase-responsive drug delivery, enabling temporal modulation of therapeutic release. In this review and design framework, we explore how the structural logic of coaxial nanofibers can be leveraged to address the temporally distinct therapeutic needs of TBI. We align polymer composition, fiber geometry, and degradation kinetics with the molecular and cellular hallmarks of each injury phase, emphasizing design strategies that synchronize scaffold behavior with evolving oxidative, neuroinflammatory, and regenerative processes. The review also outlines a roadmap for programmable coaxial electrospun nanofiber design logic for TBI application. By structurally aligning therapeutic delivery with TBI’s temporal dynamics, phase-responsive nanofiber scaffolds may advance the field toward more precise, adaptive, and effective neurotherapeutics.

Introduction

Traumatic brain injury is a leading global cause of death and long-term disability worldwide, with an estimated incidence of 939 per 10,000 individuals [1–3]. Often described as a “silent epidemic”, its burden is frequently underdiagnosed and evolves cumulatively over time [4, 5]. Traumatic brain injury results from external mechanical forces such as blunt trauma, rapid acceleration-deceleration, or penetrating injury that disrupt normal brain structure and function [6–8]. Clinical outcomes are highly variable, owing to the complex and evolving pathophysiology, and the inherently limited regenerative capacity of central nervous system (CNS) tissues [9, 10].

The pathobiological response to TBI follows a biphasic trajectory, initiating with primary mechanical insult and progressing into a delayed secondary injury cascade (Fig. 1). This secondary phase unfolds over hours to weeks, encompassing excitotoxicity, oxidative stress, neuroinflammation, mitochondrial dysfunction, and blood–brain barrier (BBB) disruption, and progressive neuronal death [11–14]. These temporally distinct molecular events present shifting therapeutic targets, yet current interventions, however, current pharmacological interventions lack the temporal specificity to align with these biological windows, leading to suboptimal clinical outcomes.

Fig. 1.

Temporal progression of TBI and scaffold matching opportunities. The schematic illustrates biphasic injury progression: (i) primary injury with immediate structural damage; (ii) secondary cascades including excitotoxicity, oxidative stress, neuroinflammation, and apoptosis. Defined therapeutic windows provide targets for phase-responsive scaffold intervention. Image reused from Aqel et al. [3]. Biomaterials in traumatic brain injury: perspectives and challenges. Biology, under the terms and conditions of the Creative Commons Attribution (CC BY) license (https://creativecommons.org/licenses/by/4.0/)

Traumatic brain injury severity is typically classified as mild, moderate, or severe using tools such as the Glasgow Coma Scale and neuroimaging modalities [15–17]. However, initial severity grading does not reliably precdict long-term functional recovery outcomes [18]. More comprehensive classification frameworks, such as those from the U.S. Department of Defense, have attempted to improve prognostic accuracy by integrating neuroimaging, consciousness level, and clinical parameters [18, 19]. Nonetheless, a persistent mismatch remains between existing therapeutic protocols and the evolving neurobiology of TBI.

Despite decades of research, no pharmacological agent has received U.S. Food and Drug Administration (FDA) approval specifically for the treatment of TBI [3, 10, 14, 20–22]. Over 30 clinical practice guidelines (CPGs) have been proposed globally [23], yet none have led to universally effective treatments [22, 23]. Moreover, 23 out of 24 TBI guidelines originate from high-income countries, raising ocncerns about their relevance to low- and middle-income countries (LMICs) where the majority of TBI cases occur [23–28].

This review addresses that translational gap. We present a literature review and design framework for engineering coaxial nanofiber scaffolds to match TBI’s temporally distinct therapeutic demands. We map polymer composition, fiber architecture, and release kinetics to injury stage-specific pathophysiology. Additionally, we evaluate current work and propose a roadmap for programmable coaxial electrospun nanofiber design logic for advancing programmable, phase-aligned nanofiber systems toward next-generation neurotherapeutics.

The temporal pathophysiology of TBI: implications for scaffold phase-matching

Primary injury: immediate structural disruption

The primary phase of TBI is initiated by mechanical forces such as blunt impact, penetration or acceleration-deceleration, that cause irreversible disruption of neuronal, glial, and vascular structures [2, 29, 30]. Resulting pathologies such as diffuse axonal injury, cerebral edema, and vascular rupture manifest within seconds to hours post injury [11, 12, 29–35, 35]. This phase is largely non-modifiable once initiated. However, it dictates scaffold timing: implants must be biocompatible and safely insertable post-injury, without exacerbating mechanical trauma. While not suitable for direct repair of structural damage, scaffolds can be strategically deployed to target the evolving secondary pathology that ensues (Table 1).

Table 1.

Phase-matched therapeutic opportunities for scaffold-mediated intervention in TBI

TBI phase / timing	Dominant pathophysiology	Therapeutic targets	Scaffold design implications
Primary injury (seconds—hours)	Immediate mechanical disruption (neuronal/axonal shearing, vascular rupture, edema)	Intracranial pressure (ICP) control, neuroprotection (largely supportive care)	Direct repair unlikely; scaffolds must be biocompatible and implantable post-injury without exacerbating damage
Early secondary injury (hours—days)	Excitotoxicity (glutamate surge, Ca2⁺ overload)	NMDA/AMPA antagonists, antioxidants	Scaffold should provide rapid release from the shell to counter excitotoxicity. Use hydrophilic polymers (e.g., polyvinyl alcohol (PVA), polyethylene glycol (PEG)) or burst-release coatings
Intermediate secondary injury (days)	Oxidative stress and mitochondrial dysfunction (ROS, lipid peroxidation, BBB disruption)	ROS scavengers (N-acetylcysteine, edaravone), mitochondrial stabilizers	Scaffold core compartment loaded with antioxidant cargo for sustained release; polymers with intrinsic ROS-scavenging properties (e.g., poly(thioglycidyl glycerol) [54])
Secondary injury – inflammatory phase (days–weeks)	Neuroinflammation, glial scar formation (IL-1β, IL-6, TNFα, astrocyte reactivity)	Anti-inflammatory agents, glial scar modulators, growth factors	Scaffold requires delayed-release profile: hydrophobic shell polymers (e.g., polycaprolactone (PCL), Poly(lactide-co-glycolide) (PLGA) to prolong delivery of anti-inflammatory drugs or trophic factors
Secondary injury—late phase (weeks)	Apoptosis and chronic degeneration (caspase activation, oligodendrocyte loss)	Anti-apoptotic agents, neurotrophic factors, stem cell support	Long-term release from core polymer (slow-degrading PLGA, collagen blends) to promote cell survival and neuroregeneration
Chronic phase (weeks—months)	Persistent inflammation, limited repair, cognitive deficits	Neuroregeneration support, axonal growth promotion	Scaffold must support structural regeneration (biodegradable ECM-mimetic fibers, heparin—collagen blends[61]), combined with sustained trophic factor delivery

Mapping of TBI phases to dominant pathological events, therapeutic targets, and implications for coaxial scaffold design

Secondary injury: biochemical and cellular sequelae

Secondary injury mechanisms emerge within minutes and may persist for weeks, progressively exacerbating neural damage through excitotoxic, oxidative, inflammatory, and apoptotic pathways [11, 12, 22, 36–41]. These events offer therapeutically actionable windows, requiring scaffold designs that match their onset and persistence profiles (Table 1). Table 1 contrasts coaxial fiber architectures across polymer combinations, highlighting how release characteristics can be strategically aligned with acute, subacute, and chronic TBI phases. For example, hydrophilic PEG sheaths facilitate burst release appropriate for the excitotoxic window, while PLGA cores enable sustained trophic factor support.

Excitotoxicity

Excess glutamate release overstimulates N-methyl-d-aspartate (NMDA) and α-amino-3-hydroxy-5-methyl-4-isoxazolepropionic acid (AMPA)-type glutamate receptors, driving intracellular calcium (Ca²⁺) overload, calpain activation, and mitochondrial damage[12, 21, 42–50]. This cascade peaks within the first few hours post-injury (Fig. 1), necessitating rapid counteraction. Coaxial scaffolds can be engineered to respond via hydrophilic shell layers (e.g., polyvinyl alcohol (PVA), polyethylene glycol (PEG)) that enable immediate or burst-phase release of neuroprotectants such as NMDA antagonists, antioxidants, or calcium modulators. These polymers may facilitate aqueous solubility and fast diffusion, ideal for the ultra-acute therapeutic window. However, process limitations (e.g., premature core mixing or polymer dissolution during electrospinning) must be managed to avoid payload leakage or poor shell integrity. (Table 1).

Oxidative stress

Mitochondrial Ca²⁺ influx accelerates reactive oxygen species (ROS) production, triggering lipid peroxidation, DNA damage, and BBB tight-junction disruption (Fig. 1). This phase spans hours to days and requires sustained antioxidant presence in the injury site. Scaffold cores loaded with N-acetylcysteine, edaravone [51–53](encapsulated within hydrophobic, diffusion-resistant polymers (e.g., PLGA) or constructed from ROS-scavenging polymers such as poly(thioglycidyl glycerol) [54] Such a design approach may potentially provide a delayed and prolonged release profile. Some platforms may also incorporate ROS-responsive degradation (e.g., thiol-rich or boronic ester polymers) to trigger on-demand drug release. However, few studies report real-time oxidative biomarker clearance, and scaffold degradation kinetics remain under-characterized in vivo (Table 1, 2, 3).

Table 2.

Illustrative failures in TBI drug development: temporal mismatch, translational limitations, and scaffold opportunities

Therapies	Clinical outcome/limitation	Underlying issue	Design gaps in clinical application	Opportunity for coaxial scaffolds
Progesterone (anti-inflammatory, neuroprotection)	Phase III RCTs failed to improve outcomes [72]	Short plasma half-life; systemic dosing misaligned with 24–72 h therapeutic window	No delivery strategy enables localized, temporally controlled anti-inflammatory release. No biomaterial system evaluated in TBI settings to bypass systemic clearance	Core-sheath fiber enabling rapid shell-based burst release and core-sustained delivery during subacute phase
Magnesium sulfate ((NMDA antagonist, Ca2⁺ modulator))	No neurological benefit; systemic toxicity [64, 65]	Required rapid systemic dosing within minutes; toxicity at higher doses	Incompatible with emergency care workflows. CNS penetration insufficient. No localized formulation validated in vivo	Immediate-release hydrophilic shell for excitotoxicity mitigation without systemic burden
Corticosteroids (broad anti-inflammatory)	CRASH trial (10,008 patients, 49 countries) showed no mortality benefit; increased adverse effects [73]	Non-specific immunosuppression; systemic side effects	No scaffold-enabled local delivery for inflammation-specific modulation. Broad systemic suppression leads to high risk-to-benefit ratio	Scaffold-mediated localized, slow-release to modulate local immune microenvironment while avoiding systemic side effects
Erythropoietin	Mixed results; safety concerns at high systemic doses [69]	BBB impermeability, poor PK	Scaffold-biologic combo not validated in TBI. Lacks clinical data on CNS delivery, dose retention, or repair outcomes. Regulatory hurdles for combination products	Core compartment for localized, sustained delivery post-injury to support repair while minimizing systemic dosing
N-acetylcysteine (antioxidant, ROS scavenger)	Some positive preclinical/early clinical findings [70, 88]	Rapid clearance, poor brain penetration	No in vivo scaffold study addresses oxidative stress kinetics or tracks antioxidant bioavailability in CNS tissue	Scaffold core as a controlled antioxidant reservoir tailored to oxidative phase duration
Amantadine (dopaminergic agonist)	Modest improvements in consciousness; variable efficacy [71]	Limited brain concentrations; systemic side effects	No long-term local delivery solution assessed. Chronic phase support not targeted via implants. Workflow, device retention and safety unaddressed	Chronic-phase compartment for slow dopaminergic support aligned with recovery trajectory
Neurotrophic factors (BDNF, NGF, VEGF, bFGF) (promote axonal growth, regeneration)	Preclinical promise; clinical translation poor [68, 70]	Instability, immunogenicity, rapid degradation	Poor data on release kinetics, bioactivity retention in vivo, and receptor engagement. Complex regulation for protein-device combo…	ECM-mimetic coaxial scaffolds for gradual, protected release over weeks with mechanical cues promoting axonal repair
MSC-derived exosomes (repair signaling, immunomodulation)	Preclinical benefit; translational barriers remain [75, 77, 78]	Instability, heterogeneity, delivery inefficiency	Scaffold-based delivery not shown to preserve vesicle integrity, biodistribution, or function. No validated, standardized, and FDA-approved Good Manufacturing Practice (GMP)-compatible delivery route	Core-loaded scaffolds acting as protective depot for staged exosome release, improving delivery precision and safety

Table 3.

Representative coaxial nanofiber systems for neural applications and their translational relevance to TBI repair

Sheath polymer	Core precursor	Drug	Release profile	Cell lines / Model	Translational outcome	Reference	TBI phase relevance
PCL	Gelatin, PDMS, PES	None (drug-free)	Tunable degradation	OSU-2 patient-derived GBM cells	Aligned, low-modulus fibers guided migration; mimicked white matter topography	[92]	Chronic (structural guidance, axonal regeneration)
PCL	PVP	Mycophenolic acid (MPA)	Sustained, suppressed burst	U-87 MG glioblastoma cells	3–5 × antiproliferative efficacy; localized delivery	[91]	Subacute—chronic (sustained modulation of glial/tumor activity)
Cellulose acetate	PLA (and citalopram-loaded gelatin nanocarriers	Citalopram	Release not explicitly reported*	Primary rat Schwann cells (SCs)	Promoted neural cell viability and differentiation; demonstrated potential for sciatic nerve regeneration	[94]	Subacute (neuroprotection, inflammation modulation)
PCL-diol polyurethane	Poly(ε-caprolactone-diol)-based polyurethane	Temozolomide	Sustained, diffusion-limited	U-87 MG human glioblastoma cell line p	Targeted therapy potential for glioblastoma multiforme (GBM). higher antitumor activity of synthesized nanofibers against glioblastoma cells compared with pristine TMZ	[93]	Subacute (tumor/TBI overlap)
PCL	PVA	RAD001 (mTOR inhibitor)	pH-responsive, extended	Pediatric brain tumor models	Synergy of hydrophobic core and mucoadhesive sheath enabled nose-to-brain delivery	[96]	Acute—subacute (anti-proliferative, bypassing BBB)
	#Same polymers with TMZ or nimorazole	TMZ/nimorazole	Tri-phasic (lag → accelerated → plateau; > 95% release in 15 weeks)	Not specified	High encapsulation efficiency (> 85%); staged, programmabke delivery	[95]	Acute—subacute—chronic (temporal alignment with TBI pathology)

Polyvinylpyrrolidone (Mw = 360 kDa, PVP40); poly(ε-caprolactone) (PCL, Mn = 80 kDa); 2,2,2,-trifluoroethanol (TFE, 99.8% purity), trifluoroacetic acid (TFA), and dichloromethane (DCM); glioblastoma multiforme (GBM) tumor cell, Cellulose acetate (Mw = 30 kDa, acetyl content = 39.70% (w/w)); Poly (L-lactic acid) (PLA; Mw = 60 kDa); ^#Multilayer coaxial-electrospun nonwovens (Series B) with core–shell architecture and electrosprayed outer layer

Neuroinflammation

Activation of microglia and astrocytes, alongside BBB breakdown, drive a robust inflammatory cascade involving cytokine release (interleukin-1β (IL-1β), interleukin-6 (IL-6), tumour necrosis factor alpha (TNFα), and nitric oxide), glial scar formation, and axonal growth inhibition reactivity [3, 10, 42, 55–59]. Chronic glial reactivity impedes neural regeneration via scar formation. This neuroinflammatory phase dominates the subacute to chronic injury window. Scaffold should provide delayed, prolonged release of anti-inflammatory or glial-modulating agents from hydrophobic shells (e.g., PCL, PLGA) (Table 1). Coaxial fibers with hydrophobic polymer shells (e.g., PCL, high-molecular-weight PLGA) could be tailored for delayed, extended release of anti-inflammatory agents or glial-modulating compounds. Layer-by-layer or triple-layer scaffolds may separate anti-cytokine and anti-scarring agents. However, the reported studies (Table 3) quantify cytokine reduction acutely, without longitudinal microglial state profiling or astrocytic scarring markers. The lack of standardized immune engagement metrics remains a translational barrier (Table 3).

Apoptosis

Mitochondrial dysfunction, cytokine exposure, and caspase activation lead to widespread apoptosis of neurons and oligodendrocytes, limiting intrinsic repair [60]. This phase emerges from days to weeks post-injury. Long-acting core compartments loaded with anti-apoptotic agents, neurotrophic factors (e.g., BDNF, NGF), or stem cell-derived exosomes may be an suitable approach for this stage. (Table 1). Scaffold strategies that mimic the extracellular matrix (e.g., HA-modified or gelatin-blended cores) can enhance cellular uptake and protect against proteolytic degradation. However, bioactive half-life, diffusion limitations, and immune visibility of large molecules remain critical design constraints. Experimental studies rarely align scaffold degradation with apoptotic marker resolution, and functional recovery metrics (e.g., remyelination, axonal sprouting) are inconsistently reported (Table 3).

Limitations and challenges of current therapies for TBI repair

Restricted therapeutic window

Therapeutic targets in TBI evolve dynamically, with transient surges in cytokines (e.g., IL-1β, TNF-α), excitatory neurotransmitters (e.g., glutamate), and reactive oxygen species (ROS). These biomolecular fluctuations create narrow and phase-specific therapeutic windows that require delivery systems capable of synchronizing release with molecular activity.

The secondary injury cascade progresses rapidly: glutamate excitotoxicity peaks within hours, while IL-1β expression rises within 1 h and peaks at 6–12 h before subsiding [62]. Agents administered outside these windows may lose efficacy or exacerbate pathology [21, 45]. Analogous lessons from stroke reinforce this principle: tissue plasminogen activator (t-PA) is beneficial only within hours of ischemia, beyond which delayed dosing induces henorrhagic risk [63]. Failed clinical trials of progesterone and magnesium sulfate exemplify the same challenge—temporal misalignment of drug exposure with evolving biology led to therapeutic futility [64, 65].

Despite the recognition of temporal dynamics, few therapies have integrated pharmacokinetic (PK) profiling with the molecular pathophysiology of TBI phases. Most clinical trials did not include time-stratified dosing arms or biomarker-guided intervention points, limiting interpretability and translational relevance (Table 3).

Beyond pharmacodynamics, systemic constraint further narrow therapeutic windows: diagnostic delays, absence of robust biomarkers, and logistical barriers in emergency care reduce the likelihood of achieving timely drug exposure [21, 66]. In this context, localized biomaterial systems, particularly coaxial scaffolds with programmable release, offer the opportunity to bypass systemic delays and sustain therapeutic presence directly at the injury site,, in synchrony with the spatiotemporal biology of TBI.

Limitations of current treatments

Despite decades of intensive investigation, no pharmacological agent has gained FDA approval for TBI-specific therapy [21, 67]. Existing interventions remain largely palliative and mom-regenerative, focused on intracranial pressure control or seizure prophylaxis rather than modulation of the multiphasic secondary injury cascade [3, 10, 42].

Candidate small molecules including corticosteroids, progesterone, erythropoietin, N-acetylcysteine, and calcium channel blockers have failed in translation due to short plasma half-lives, poor blood–brain barrier (BBB) penetration, and lack of phase specificity [22, 68–72],.The MRC CRASH trial, a multicentre RCT involving 10,008 patients across 49 countries, demonstrated no survival benefit from corticosteroids administered within 14 days, while increasing adverse effects, reinforcing the limitations of systemic, broad-acting immunosuppressants [73].

Critically, these clinical failures often resulted from generalized dosing schemes that did not account for the spatial and temporal heterogeneity of TBI. For instance, systemic delivery of progesterone in trials did not differentiate between diffuse versus focal injury types, or between early- and late-stage inflammation, which may have obscured potential phase-specific benefits.

Biologics, such as neurotrophic factors (e.g., vascular endothelial growth factor (VEGF), brain-derived neurotrophic factor (BDNF), nerve growth factor (NGF), chondroitin sulfate proteoglycans (CSPG)-cleaving enzyme chondroitinase ABC, and cell derived therapies (e.g., exosomes derived from mesenchymal stem cells (MSC-Exo)show preclinical promise, but are limited by instability, immunogenicity, scalability, and poor delivery efficiency [73–78]. The BBB remains a formidable barrier, excluding most hydrophilic or high molecular-weight drugs [79–81]. Attempts to bypass it, via osmotic disruption or ultrasound-mediated BBB opening carry the risk of secondary inflammation and cytotoxicity [82]. While intranasal delivery and exosome carriers represent progress [83–85], dose reproducibility and clinical scalability remain unresolved. Injectable hydrogels, microspheres, and liposomes offer local delivery but currently lack the capacity to precisely align release kinetics with TBI’s temporal pathophysiology [75, 86, 87].

To illustrate these translational barriers and highlight scaffold-based solutions, Table 2 summarizes representative drug candidates whose clinical failures can be traced to temporal mismatch and systemic limitations. Each case demonstrates how coaxial nanofiber architectures could theoretically overcome these barriers by enabling spatiotemporal control over release.

Current coaxial electrospun nanofibers for brain tissue application

Material matters: engineering nanofiber biofunctionality and mechanics for neural drug delivery

Polymer selection is the primary determinant of coaxial electrospun nanofiber performance, governing scaffold mechanics, degradation, and drug release kinetics. Natural polymers such as hyaluronic acid, gelatin, silk fibroin, chitosan, and alginate provide biomimetic extracellular matrix (ECM) cues that promote neural adhesion and differentiation [89, 90]. However, their limited mechanical robustness and batch variability necessitate blending with synthetic polymers such as polycaprolactone (PCL), poly (lactic acid) (PLA), and polyurethanes to ensure durability and tunability. Polycaprolactone in particular, has been widely adopted due to its semi-crystalline structure, hydrophobicity, and slow degradation, supporting sustained release in aqueous brain microenvironment [91, 92]

Coaxial architectures enable precise exploitation of these polymer traits (Table 3). Han et al. [90] demonstrated that tuning sheath thickness in PCL-based coaxial fibers modulated diffusion-controlled release of mycophenolic acid (MPA), achieving sustained bioactivity against glioblastoma cells while suppressing burst release common in monolithic fibers [91]. However, the study by Han et al. [90], while demonstrating sustained MPA release, lacked in vivo validation or PK/PD correlation, limiting translational relevance. Similarly, Irani et al. [99] reported coaxial PCL-diol-based polyurethane fibers loaded with temozolomide (TMZ), where modulation of shell flow rate provided predictable 30-day release while retaining cytotoxicity in U-87 MG cells [93]. Collectively, these studies confirm that polymer crystallinity, hydrophobicity, and interfacial viscosity at the core-sheath boundary directly dictate localized drug availability, an essential requirement for temporally staged neural repair.

Beyond drug delivery, scaffold mechanics and architecture also modulate cellular behavior. Rao et al. [95] fabricated aligned PCL-sheathed coaxial fibers with gelatin, PDMS, or PES cores of variable stiffness and scaffold modulus was shown to regulate glioblastoma migration, while fiber alignment recapitulated anisotropic white matter topography [92]. Importantly, fibre alignment recapitulated the anisotropic topography of white matter tracts, facilitating directed migration and modelling the infiltrative behaviour of brain tumours. Such systems highlight how coaxial fibers can act as “bio-instructive matrices,” influencing mechanotransduction pathways such as focal adhesion kinase (FAK) and myosin light chain 2 (MLC2), in addition to serving as carriers.

Surface chemistry further adds another layer of tunability. Naseri-Nosar et al. (2017) coated coaxial cellulose acetate/PLA fibres with citalopram-loaded gelatin nanocarriers [94]. This modification transformed hydrophobic fibres into highly hydrophilic scaffolds (contact angle shift from approximately 124.5° to 0.0°), enabling improved Schwann cell adhesion and BDNF upregulation in vitro, and functional nerve repair in vivo (Table 3). Such hybrid approaches illustrate how surface coatings can potentially reconcile mechanical robustness with biofunctionality, a balance critical in neural implants.

Finally, coaxial designs enable multi-drug loading with programmable release. Musiał-Kulik et al. [101] fabricated multilayer coaxial nonwovens encapsulating (temozolomide) TMZ and nimorazole in distinct compartments, sealed with electrosprayed barriers. [95]. This architecture generated a tri-phasic release over 15 weeks, aligning therapeutic availability with acute, subacute, and chronic needs (Fig. 2). Compared with Gliadel™, which releases carmustine over ~ 7 days, these constructs demonstrate the feasibility of long-term, programmable drug delivery for neural repair.

Fig. 2.

Release of temozolomide and nimorazole from PLACap and PLAGATMC scaffolds (mean ± S.D., n = 3). Series A: monoaxial fibers; Series B: coaxial fibers with electrosprayed sealing. Reused from Bioresorbable, electrospun nonwoven for delayed and prolonged release of temozolomide and nimorazole. Eur J Pharm Biopharm 2021;161:29–36. [95], with permission from Elsevier

Bioactive release kinetics: strategies for temporal modulation

A recurring theme in coaxial nanofiber systems is the suppression of burst release and alignment of payload kinetics with biological demand. Han et al. [90] showed that thickened PCL sheaths slowed MPA efflux, maintaining antiproliferative concentrations over multiple culture cycles [91]. Irani et al. [99] confirmed that adjusting shell flow rates during fabrication allowed fine-tuning of TMZ release profiles[93]. In contrast, Naseri-Nosar et al. (2017) demonstrated a biphasic citalopram release, an initial therapeutic pulse followed by sustained delivery using gelatin nanocarriers [94].

Innovations have extended beyond linear control. Louis et al. (2024) designed a pH-responsive coaxial PCL/PVA system carrying RAD001, which released up to 90% of its payload at pH 5.5 (tumor-like microenvironment) but remained largely quiescent at physiological pH [96]. Musiał-Kulik et al. [101] further advanced the field by demonstrating tri-phasic release kinetics in multilayer poly(L-lactide-co-glycolide-co-trimethylene carbonate) (PLAGATMC) and poly(L-lactide-co-ε-caprolactone) (PLACap) (PLACap/PLAGATMC systems), initial lag, accelerated hydrolytic release, and terminal plateau, closely mirroring TBI’s temporal progression. [95]. These strategies provide a rational blueprint for scaffolds capable of synchronizing neuroprotective, anti-inflammatory, and regenerative release phases.

Translational potential of localized coaxial implants

Evidence from both in vitro and in vivo models supports the translational relevance of coaxial scaffolds for neural applications. Rao et al. [95] demonstrated that drug-free aligned fibers alone could direct glioblastoma migration via mechanochemical cues, highlighting their role as instructive biomaterials [92]. Irani et al. [99] confirmed that coaxial TMZ fibers maintained long-term cytotoxicity while leveraging fiber alignment to increase scaffold-cell interaction [93]. Naseri-Nosar et al. (2017) validated functional recovery in vivo in a rat sciatic nerve defect model, where citalopram-loaded coatings outperformed controls [94]. Louis et al. (2024) illustrated that intranasal RAD001-loaded scaffolds suppressed paediatric brain tumour growth in a time- and dose-dependent manner, although toxicity of PVA at higher doses highlighted safety considerations [96].

Most notably, Musiał-Kulik et al. [101] established a proof-of-concept for bioresorbable multilayer coaxial implants, showing staged TMZ and nimorazole release for 15 weeks, significantly extending the therapeutic coverage compared to Gliadel™ (7 days) [95]. However, this study did not evaluate scaffold performance in TBI-specific models, nor did it assess cognitive recovery, immune compatibility, or degradation kinetics in neural tissue, factors essential for clinical translation. While these programmable implants demonstrate the feasibility of spatial and temporal drug modulation, their translation to TBI settings requires additional validation in models replicating blood–brain barrier disruption, neuroinflammation, and tissue remodeling. These gaps underscore the need for future work to integrate mechanistic biomarkers (e.g., cytokine resolution, axonal regeneration) and patient-relevant outcomes (e.g., behavioral recovery, implant safety) into scaffold testing protocols.

Mapping fiber design to TBI phases and future outlook: synthesis after the evidence

The evidence surveyed herein across coaxial nanofiber systems reveals patterns that can be explored further into design principles for phase-matched repair of traumatic brain injury (TBI). Rather than isolated case studies, these trends highlight the structural logic that determines whether scaffolds succeed in synchronizing therapeutic delivery with the evolving pathophysiology of TBI.

We propose three interdependent design axes as critical determinants of therapeutic performance in coaxial scaffolds for TBI: (1) core-sheath independence and tunable release kinetics, (2) mechanical anisotropy and compliance with brain tissue viscoelasticity, and (3) multi-layer or modular programmability for sequential phase-specific targeting. These principles map scaffold design to biomarker-defined temporal injury windows.

Success modes frequently pair hydrophilic outer sheaths (e.g., PVA, PEGDA, or blends thereof) with core compartments entrapping anti-inflammatory or trophic agents. These systems achieve early-phase burst release followed by sustained delivery, aligning with the biphasic nature of TBI pathology. Core–sheath uncoupling enables spatiotemporal control of drug exposure.

Failure modes arise when monolithic fibers or ultrathin sheaths permit unbuffered burst release (< 4 h), leading to subtherapeutic profiles or premature depletion of neuroprotective agents. Conversely, highly glassy or crystalline sheaths (e.g., high-PCL, PLGA, or polyurethane without plasticizers) may entrap bioactive proteins, impeding release. Furthermore, process-induced degradation (e.g., organic solvent exposure, high electric fields, thermal denaturation) during electrospinning frequently compromises protein stability, a factor underreported in the literature. Scaffold design for TBI must balance release kinetics, payload sensitivity, and injury-phase matching to avoid these recurrent pitfalls.

First, release kinetics are governed by the core-sheath interface. Thickened or hydrophobic sheaths (e.g., PCL, PLGA variants) predictably suppress burst release and shift kinetics into diffusion-limited modes appropriate for acute-to-subacute windows. In contrast, hydrophilic or miscible cores (e.g., PVP-rich formulations) erode temporal fidelity through early leakage, demonstrating that interfacial stability is a critical attribute.

Second, biomimicry requires mechanical and topographical fidelity, not just biochemical resemblance. Aligned, low-modulus fibers replicate white-matter anisotropy and support axonal guidance, while isotropic or stiff matrices bias toward astrocytic scarring. Thus, scaffold critical properties include modulus (0.1–1 kPa) and anisotropy, alongside biochemical signals.

Third, multilayered or stacked systems (coaxial with electrosprayed barriers) demonstrate programmable, triphasic delivery (lag—accelerated—plateau), aligning with the 10–15 week horizon for chronic support. These architectures demonstrate that timing can be structurally encoded without exotic chemistries, a feasible pathway for scaling under GMP conditions offering reproducibility and scalability favourability. A recurring insight is that hydrophobic sheaths act as predictable diffusional resistors, where flow-rate tuning during electrospinning provides a reproducible and regulator-friendly process dial. Surface hydrophilization through coatings (gelatin, HA) enhances biological interfacing without destabilizing deeper cores, offering a rational compromise between biofunctionality and controlled kinetics.

Translational gaps remain, most reports provide cumulative release curves without anchoring them to TBI biomarker timelines (e.g., glutamate/Ca²⁺ minutes-hours; ROS/cytokines hours-days; glial remodeling days-weeks; trophic insufficiency months). Few establish PK/PD correlations with functional outcomes such as IL-1β suppression, BBB integrity recovery, or glial scar attenuation. Likewise, neuroimmune interactions are rarely quantified beyond acute toxicity, leaving chronic microglial and astrocytic modulation as an unmet frontier.

Beyond material process logic, key translational challenges must also be addressed. Neuroimmune interactions remain insufficiently characterized, with most studies reporting acute toxicity but not chronic modulation of microglial or astrocytic states. Manufacturing reproducibility poses another barrier: flow-rate tuning and sheath-thickness control are technically feasible, but rarely benchmarked against good manufacturing practice (GMP) standards. Finally, regulatory classification looms as a decisive hurdle. Lessons from Gliadel® wafers illustrate how drug-device combinations are subjected to dual scrutiny in sterility, stability, and dosing reproducibility, underscoring the need for early regulatory foresight in scaffold development.

Overall, these translational hurdles point to a preliminary roadmap for next-generation programmable nanofibers: (i) embed pharmacokinetic/pharmacodynamic (PK/PD) modeling into scaffold design to anchor release profiles against biomarker-defined targets; (ii) validate in vivo not only for safety but also for measurable engagement with TBI pathophysiology (e.g., BBB repair, cytokine suppression, glial scar reduction); and (iii) benchmark scaffold performance against clinical standards such as Gliadel® wafers to demonstrate added value in temporal precision and therapeutic durability. These dimensions are synthesized in Fig. 3, which contextualizes how design parameters, engineering strategies, and biofunctional exemplars converge into a translational framework.

Fig. 3.

Scaffold-centric framework summarizing design parameters (blue), engineering strategies (yellow), and biofunctional exemplars (green) for neural tissue engineering. The schematic illustrates how features such as nanoporosity, anisotropy, and degradability, along with surface chemistries (e.g., arginine-glycine-aspartic acid (RGD), hyaluronic acid, growth factors), may be integrated to support temporal repair following TBI. Reused from Villanueva-Flores et al. [106]. Toward a New Generation of Bio-Scaffolds for Neural Tissue Engineering: Challenges and Perspectives, Pharmaceutics. 2023 Jun 16;15(6):1750. under the terms and conditions of the Creative Commons Attribution (CC BY) license (https://creativecommons.org/licenses/by/4.0/) [97]

Author contributions

GDM: Conceptualization, methodology, data organization and synthesis, data analysis, figure preparation, writing-original draft preparation, writing-review and editing, visualization, YEC: Conceptualization, writing-review and editing, visualization, supervision, project administration, funding acquisition. All authors have read and agreed to the published version of the manuscript.

Funding

This work was funded by the National Research Foundation (NRF) South African Research Chairs Initiative (SARChI), NRF Competitive Programme for Rated Researchers, and the South African Medical Research Council (SIR Grant). The support of Carnegie Diversifying the Academy (CTDA) program at the University of the Witwatersrand, Johannesburg in, the Republic of South Africa, towards this research is hereby acknowledged. Opinions expressed, and conclusions arrived at, are those of the author and are not to be attributed to the CDTA.

Data availability

No datasets were generated or analysed during the current study.

Declarations

Ethics approval and consent to participate.

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Clinical trial number

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Competing interest

The authors declare no competing interests.