Nanofibrous Platforms for Advanced Wound Therapeutics: Mechanisms, Materials, and Clinical Horizons

ABSTRACT

Nanofibrous dressings have emerged as a transformative solution for acute and chronic wound management. This review evaluates recent advancements in wound therapeutics utilizing nanofiber platforms incorporating various bioactive agents. A comprehensive overview of the wound healing process across distinct physiological phases is provided, alongside a discussion of nanofiber fabrication methodologies and their underlying mechanisms. Particular emphasis is placed on the constituent materials and their inherent biofunctionality, specifically regarding how cellular interactions with these substrates enhance wound healing outcomes. Furthermore, we examine the clinical utility of these scaffolds in achieving hemostasis, modulating inflammation, preventing infection, and promoting angiogenesis. The review concludes by highlighting the cost‐effectiveness and tunable architecture of nanofibers, underscoring their potential as the next generation of efficacious wound care systems.

The fabrication of nanofiber dressings from diverse natural and synthetic polymers offers a robust platform for chronic wound management. By encapsulating therapeutic drugs and bioactive agents, these dressings can simultaneously attenuate chronic inflammation and microbial infection while actively stimulating angiogenesis, addressing the primary physiological disturbances that impede chronic wound repair.

Abbreviations

(A‐FRP)

Advanced fibrin‐rich platelet

(CaO2 NP)

Calcium peroxide nanoparticle

(CNF)

Carbon nanofiber

(CS)

Chitosan

(DHM)

Dihydromyricetin

(Epi‐1)

Epinecidin‐1

(EuCP)

Europium(III) coordination polymers

(GelMA)

Gelatin methacryloyl

(GSV)

Gelsevirine

(HHA)

High molecular weight hyaluronic acid

(HNT)

Halloysite nanotubes

(MnO₂ NS)

Manganese dioxide nanosheet

(OI)

4‐Octyl itaconate

(PBChi)

Chitosan stabilized Prussian blue nanoparticle

(PDA)

Polydopamine

(PEO)

Poly(ethylene oxide)

(PLATMC)

Poly‐(lactic acid‐co‐trimethylene carbonate)

(PLCL)

Poly(L‐lactide‐co‐ε‐caprolactone)

(PLLA)

Poly‐L‐lactic acid

(PNVCL)

Poly(N‐vinylcaprolactam)

(PR1P)

Prominin1‐derived peptide

(PTEF)

Polytetrafluoroethylene

(PU)

Polyurethane

(PVB)

Polyvinylbutyral

(PVP)

Polyvinylpyrrolidone

(SA)

Sodium alginate

(SPC)

Sodium per carbonate

(TA)

Tannic acid

1. Introduction

The skin functions as far more than a passive anatomical casing; it is the human body's primary interface with the external world, a metabolically active organ essential for survival [1]. As the outermost barrier, it orchestrates a complex defense against environmental stressors, ranging from mechanical trauma and ultraviolet radiation to the pervasive threat of microbial invasion [2]. Beyond simple protection, the skin is central to physiological homeostasis, regulating core body temperature through thermoregulation and preventing systemic dehydration via the control of transepidermal water loss. Consequently, any disruption to this structural integrity, whether through acute trauma, surgical intervention, or underlying pathology, represents a critical dysfunction of these vital biological functions, necessitating an immediate and coordinated repair response to restore homeostasis [3].

Wounds occur due to various causes, including surgery, injuries, mechanical trauma, diabetes, or vascular diseases [4]. These types of damage are classified as either acute or chronic wounds, depending on their underlying causes and consequences. Acute wounds, such as those caused by traumatic and surgical damage, recover to normal anatomical structures after an organized tissue repair process. Conversely, chronic wounds are characterized by impaired healing within the expected timeframe due to factors such as obesity, diabetes, poor circulation (e.g., arterial and venous insufficiency), radiation exposure, or immune system dysfunction [5]. This impaired healing may result in further complications, including infection, necrosis, or sepsis [6]. Chronic ulcers that occur on the leg and foot are known to persist for 12 to 13 months and frequently tend to recur [7]. For patients with diabetic foot ulcers, an estimated 40% experience recurrence within 1 year after ulcer healing, nearly 60% within 3 years, and 65% within 5 years [8]. This can result in loss of function and a decline in quality of life, representing a significant cause of morbidity [9].

The epidemiology of chronic wounds reveals a staggering public health challenge. The 1%–2% of the global population will suffer from a chronic wound during their lifetime. In the United States, this health burden is significant, affecting nearly 2.5% of the total population and an estimated 10.5 million U.S. Medicare beneficiaries [9]. The economic implications of this silent epidemic are profound. In 2022, the cost of wound care in the United States was estimated to exceed $148.65 billion, a figure that eclipses the budgets of many other major health conditions [10]. This financial strain is mirrored globally, with China and Japan reporting wound care expenditures of $42.78 billion and $22.91 billion, respectively. As the global population ages and the prevalence of metabolic diseases such as diabetes continues to rise, the incidence of chronic wounds is projected to increase, solidifying wound care as a persistent and escalating public health concern. Given the escalating clinical and economic crisis, the limitations of conventional wound management strategies have become increasingly apparent. Traditional dressings, while ubiquitous, often fail to address the complex microenvironmental requirements of chronic wounds, leading to suboptimal outcomes and prolonged suffering. This understanding has necessitated a paradigm shift toward advanced therapeutic materials.

Rapid technological advancements have significantly expanded the repertoire of available wound dressing materials. Among modern solutions, nanofibers have gained significant attention due to their intrinsic properties, such as a high surface‐to‐volume ratio, tunable porosity, and nano‐sized diameter, which mimic the native extracellular matrix (ECM) [11, 12]. These scaffolds demonstrate remarkable therapeutic potential by promoting cell migration and proliferation while supporting three‐dimensional tissue formation. Additionally, intrinsic properties such as nano‐sized diameter, high surface‐to‐volume ratio, and tunable porosity make them excellent candidates for wound management. In this context, diverse nanofiber dressings have been engineered to incorporate bioactive agents, such as growth factors, antioxidants, and antimicrobial compounds, designed to mitigate the complications associated with impaired wound healing and to actively facilitate the regenerative process. This review elucidates nanofiber fabrication methodologies and constituent polymer matrices, providing a comprehensive analysis of their biofunctional properties for advanced wound healing applications.

2. Processes of Wound Healing

Wound healing is a complex and highly coordinated biological process involving interactions among numerous types of cells, cytokines, mediators, and extracellular matrices (ECMs) to repair damaged tissue [13]. After tissue damage, wound healing proceeds through four distinct but overlapping phases: hemostasis, inflammation, proliferation, and remodeling (Figure 1). The events occurring in each phase are not mutually exclusive and occur in a precise and sequential manner.

FIGURE 1.

Schematic representation of the wound healing process, highlighting the spatiotemporal roles and phenotypic changes of key cell populations across different stages of repair.

2.1. Hemostasis

The first phase of hemostasis begins with the formation of a platelet plug immediately following the disruption of vessels. When a blood vessel is damaged, vasospasm occurs and subsequently triggers vasoconstriction, which slows blood flow. Vasoconstriction causes platelets to leak into the intravascular space and adhere to the exposed subendothelial matrix with the help of platelet membrane receptors and plasma proteins called von Willebrand factor [14]. The adhered platelets undergo activation, resulting in morphological changes, accumulation, degranulation, and the release of chemokines and growth factors that drive clot formation. Platelet alpha granules secrete a variety of growth factors, including platelet‐derived growth factor (PDGF), transforming growth factor beta (TGF‐β), transforming growth factor alpha (TGF‐α), basic fibroblast growth factor (bFGF), insulin‐like growth factor‐1 (IGF‐1), and vascular endothelial growth factor (VEGF), all of which support the wound healing process that follows hemostasis [4]. Wound healing begins with hemostasis and platelet deposition, followed by an influx of neutrophils and monocytes via signals such as PDGF, TGF‐β, C5a, and bacterial peptides, while simultaneously vascular endothelial cells are activated by VEGF, TGF‐α, and bFGF to initiate angiogenesis [15]. After the hemostasis phase is complete, fibrinogen undergoes a series of enzymatic activation processes to form a fibrin network, which forms a fibrin clot. This clot is a temporary matrix composed of fibrin, fibronectin, vitronectin, and thrombospondin that not only stops bleeding but also promotes cell migration to aid in early wound healing [16].

2.2. Inflammation

The inflammatory phase initiates in the late stages of hemostasis, promoting the infiltration of inflammatory cells to mitigate microbial infection. Neutrophils are the first white blood cells to infiltrate the wound site and are primarily attracted by various chemotactic factors. These factors include complement components such as C3a and C5a, bacterial formylmethionyl peptides, and platelet products [17]. Following the clearance of contaminants, neutrophil activity diminishes, and the cells are subsequently removed by the intrinsic and extrinsic apoptosis pathway [18]. The complete clearance of these neutrophils occurs without harming adjacent tissue or prolonging inflammation, thereby allowing healing to progress to the subsequent stage. Subsequently, exposure to pro‐inflammatory cytokines, interferons, lipopolysaccharides, or damage‐associated molecular patterns (DAMPs) induces the activation and differentiation of macrophages to a pro‐inflammatory phenotype (M1) [19]. This M1 phenotype performs vital phagocytosis to remove apoptotic debris and simultaneously secretes a wide array of mediators and cytokines, including IL‐1, IL‐6, IL‐12, TNF‐α, and inducible nitric oxide synthase (iNOS) [20]. These factors collectively amplify and regulate the inflammatory cascade while simultaneously promoting further leukocyte recruitment to the site of injury. Macrophages in the alternative (M2) phenotype are polarized by the exposure to IL‐4 or IL‐13 cytokines [21]. The M2 phenotype produces anti‐inflammatory mediators such as the IL‐1 receptor antagonist (IL‐1ra), the decoy IL‐1 receptor II (IL‐1RII), and IL‐10. Additionally, the various growth factors derived from these cells, including TGF‐β, PDGF, VEGF, and insulin‐like growth factor 1 (IGF‐1), support the migration and proliferation of fibroblasts, keratinocytes, and endothelial cells, thereby promoting granulation tissue formation and angiogenesis [19, 22].

2.3. Proliferation

During the proliferative phase of wound healing, several sequential steps occur: granulation tissue formation, angiogenesis, re‐epithelialization, and collagen deposition. These are the main processes that facilitate ongoing tissue repair and overlap with other wound healing phases [23]. As the proliferation phase continues, the fibrin clot is gradually replaced by highly vascularized granulation tissue, and fibroblasts play an essential role in this formation process [23]. Fibroblasts migrate from the nearby dermis to the wound site in response to cytokines and growth factors, such as PDGF, TGF‐β, and bFGF, produced by platelets and macrophages [24]. Once recruited to the wound, they proliferate and degrade the provisional fibrin clot using matrix metalloproteinases (MMPs) [25]. The signaling molecule TGF‐β acts as a potent stimulus for the activation and differentiation of fibroblasts into myofibroblasts. The acquisition of α‐SMA expression in these cells provides the necessary contractile force required for effective wound closure [26]. Furthermore, new collagen (mostly type III) and other ECM components (e.g., proteoglycans, hyaluronic acid, fibronectin, and collagen) are deposited to fill the wound gap for re‐epithelization [25]. Angiogenesis, the process of new blood vessel formation, occurs in response to tissue injury. It is stimulated by various growth factors and cytokines, including VEGF, angiopoietin, FGF, and TGF‐β, which collectively activate endothelial cells located at the tip of existing capillaries [27]. These mediators induce the activation of capillary tip endothelial cells. They subsequently secrete MMPs to degrade the basal lamina, facilitating invasive migration. Following degradation, endothelial cell proliferation and directed migration occur along the chemotactic gradient toward the source of the angiogenic signal [28]. The wound closure is achieved by the activation of keratinocytes, making a protective epithelial barrier. Keratinocytes at the leading edge of the defect area start migration and proliferate to cover the wound and rebuild the basal membrane through protein secretion [29]. Epithelial cell migration continues until the leading edges meet and establish a continuous barrier. Once this contact is achieved, the epithelial monolayer transitions to a proliferative state, subsequently rebuilding and maturing multiple layers of the epidermis to restore the barrier function of skin [15].

2.4. Remodeling

The final stage of healing involves the gradual remodeling of granulation tissue into scar tissue, a process that can last from several months to more than a year [30]. During this phase, the number of myofibroblasts and excess vascular cells are reduced by apoptosis [31]. The major component of the granulation tissue, collagen type III, is gradually replaced by collagen type I, and the orientation of the collagen fibrils becomes more organized [32]. This matrix remodeling consequently results in the increased tensile strength of the wound site.

3. Nanofiber as a Wound Dressing

Wound healing is a complex biological process that typically requires several weeks to months, depending on the severity and etiology of the injury. The rate and quality of tissue repair are influenced by numerous systemic and local factors, including vascular perfusion, microbial infection, and underlying pathological conditions [6]. Disruption of one or more phases of healing can result in delayed or impaired wound repair. Standard wound care strategies primarily focus on three key principles: (i) debridement, (ii) infection control, and (iii) moisture management [33]. Wound dressings play a central role in this process by protecting the wound from external contaminants, maintaining a moist environment, and minimizing mechanical trauma. An optimally moisturized wound environment promotes autolytic debridement of necrotic tissue, reduces patient discomfort, minimizes scar formation, and enhances collagen synthesis [34].

Conventional wound dressings, such as sterile gauze and bandages, have been widely used due to their low cost and ease of application. However, these traditional materials possess several inherent limitations. In wounds with high exudate levels, dressings often become saturated, leading to fluid accumulation and the need for frequent replacement. Conversely, when exudate production decreases, dressings may adhere to the wound surface, causing pain upon removal and potentially inducing secondary tissue damage [35]. As a result, conventional dressings are largely limited to passive moisture regulation and basic physical protection. To overcome these shortcomings, a variety of advanced wound dressing materials, including films, foams, hydrofibers, hydrogels, and hydrocolloids, have been developed. These materials are designed to maintain an optimal moisture balance while enabling the sustained release of antimicrobial agents such as silver and iodine, thereby providing localized infection control. Despite these advancements, persistent clinical challenges remain, such as the limited exudate absorption of films and hydrocolloids, the need for frequent dressing changes with foams, and the weak mechanical properties of hydrogels [36]. Consequently, there is a clear need for next‐generation bioactive dressings capable of actively promoting and accelerating the wound‐healing process.

Among emerging materials, nanofibers have attracted significant attention as promising candidates for next‐generation wound dressings. Their unique physicochemical properties make them particularly suitable for advanced wound management applications. Nanofibers, with diameters ranging from tens to hundreds of nanometers, closely mimic the nanoscale architecture of the ECM. This biomimetic structure resembles endogenous fibrous proteins such as collagen, fibronectin, and laminin, which exhibit hierarchical organization at the nanometer scale [37]. By providing critical topographical cues, nanofibrous scaffolds actively modulate key cellular behaviors, including adhesion, migration, proliferation, and lineage‐specific differentiation [38, 39]. In addition, their interconnected porous structure and high surface‐area‐to‐volume ratio support efficient gas exchange, moisture retention, and cellular respiration, while simultaneously forming an effective physical barrier against microbial infiltration [40, 41, 42]. Furthermore, nanofibrous wound dressings can be functionalized with a wide range of bioactive agents, such as growth factors, antioxidants, and antimicrobial compounds, enabling targeted and accelerated wound healing.

4. Techniques for Nanofiber Fabrication

4.1. Electrospinning

Electrospinning is the most prevalent method for fabricating nanofibers. This technique typically utilizes a high‐voltage power supply, a syringe pump, a spinneret, and a collector as conventional apparatus (Figure 2a). In this process, a polymer solution is extruded from the spinneret at a constant rate by the syringe pump, forming a pendant droplet via surface tension. The applied high voltage generates an electric field, and electrostatic forces separate positive and negative charges within the liquid. The solution with the same charge as the needle migrates to the surface of the droplet, producing excess charge. As the applied voltage gradually increases, the surface charge of the droplet accumulates, inducing the deformation of the droplet into a conical shape known as a Taylor cone. Consequently, the charged jet ejected from the droplet stretches toward the conductive collector in a whipping motion, and the solvent rapidly evaporates, forming thin fibers [43, 44].

FIGURE 2.

Schematic representation of nanofiber fabrication techniques: (a) electrospinning; (b) solution blowing; and (c) self‐assembly.

Electrospun nanofibers can be produced in various structural forms. Randomly oriented nanofibers deposited on a stationary collector, forming non‐woven fiber mats with a highly porous and interconnected structure resembling the natural ECM. Aligned nanofibers are fabricated by adjusting the electric field or using a rotating collector to orient fibers in a specific direction [45]. This orientation enhances the mechanical properties of the nanofibers and provides directional cues that can manipulate the migratory behaviors, ECM production, and differentiation of cells [46, 47]. For example, in a previous study, human adipose stem cells cultured on aligned nanofibers made of polycaprolactone/collagen exhibited elongated cell morphology, higher proliferation, a faster migration rate, and increased ECM production compared to cells cultured on randomly oriented nanofibers [48]. Another form of nanofibers is core–sheath nanofibers, produced through coaxial (two‐fluid) electrospinning. This process extrudes two different polymer solutions through coaxial capillaries to form fibers with a unique core–sheath structure. This distinctive structure is advantageous for protecting unstable components from degradation, enabling the sustained release of drugs, and enhancing mechanical properties and biocompatibility [49].

Various advanced technologies have been developed to overcome the low‐throughput limitations of conventional electrospinning methods. Multi‐needle electrospinning is a simple strategy that enables increased nanofiber production, higher material flow rates, the simultaneous processing of heterogeneous materials, and precise control of morphological properties. However, this method suffers from reduced efficiency due to electrostatic interference between needles and continuous needle clogging, which impacts fiber quality [50]. While conventional capillary‐based systems (both single‐needle and multi‐needle) suffer from yield limitations and maintenance difficulties, needleless (free‐surface) electrospinning is an alternative that offers high throughput. The needleless method, which induces spontaneous jetting directly from an open liquid reservoir, can avoid nozzle‐related instability. Nevertheless, achieving precise hydrodynamic control and a uniform jet distribution across the entire free surface remains a significant engineering challenge [51].

Although electrospinning is a facile method for nanofiber fabrication, the use of toxic solvents, such as chloroform, dichloromethane, and fluorinated alcohols, to dissolve the polymers is a remaining concern. The excessive exposure of these solvents can cause damage to multiple organs, including the liver, kidneys, nervous system, and reproductive system [52, 53, 54, 55]. Apart from the toxicity to humans, they are accumulated in the environment, contaminating the water, soil, and air. Thus, it was necessary to utilize the sustainable and environmental‐friendly electrospinning methods as alternatives: melt electrospinning, which extruding the molten polymer without the use of solvents, and aqueous systems to dissolve various hydrophilic polymers, thereby mitigating the reliance on hazardous organic solvents [56].

4.2. Near‐Field Electrospinning

Near‐field electrospinning (NFES) is a specialized additive manufacturing methodology for the deterministic fabrication of micro‐ and nano‐scale fibers in highly ordered configurations. Originally introduced by Sun et al. in 2006 [57], the technique utilizes a similar instrumental setup with conventional electrospinning but implements a significantly reduced working distance to facilitate aligned deposition. While traditional far‐field electrospinning typically yields stochastic, nonwoven architectures due electrohydrodynamic bending instability caused by Coulombic repulsion between the jets [58], the short nozzle‐to‐collector distance in NFES stabilizes the charged jet enabling the spatially precise deposition of individual or oriented fibers on a computer‐controlled substrate. The fiber diameter, shape, and spatial arrangement can be controlled by simultaneously controlling the feed rate, electric field strength, and solution properties [59]. Although early NFES platforms were limited by low throughput and limited uniaxial attenuation, recent advances in multi‐nozzle configurations and process optimization have significantly improved fiber uniformity and internal crystalizing [60].

4.3. Solution Blowing Spinning

Solution blowing spinning (SBS) is a simple method for manufacturing nanofibers that utilizes a polymer solution dissolved in a volatile solvent and a high‐pressure gas, such as nitrogen, argon, or air, to create fibers, as suggested by Medeiros et al. (Figure 2b) [61]. The formation of fibers in solution blowing based on the high‐speed stretching principle of airflow and the Bernoulli principle. As the high‐pressure gas exits the nozzle, it results in a drop in pressure in the center of the jet, which becomes the driving force accelerating the polymer solution. The gas flow induces shear force at the gas/solution interface in SBS and deforms the polymer solution into a conical shape. As the shear force overcomes the surface tension of the polymer solution, a jet is created, and the solvent rapidly evaporates during flight. Finally, thin polymer fibers are deposited on the collector [62, 63]. Although electrospinning provides precise control over nanofiber morphology, its reliance on high electric fields and low production rate limits its transition from lab‐scale to industrial application [64]. In contrast, SBS can deposit nanofibers on any substrate without the need for an electric field or a conductive collector [61], which enables in situ production for biomedical applications. Furthermore, its high polymer injection rate allows for scalable fiber production, making SBS more suitable for industrial applications.

4.4. Self‐Assembly

Self‐assembly is a promising method to generate stable and ordered nanostructures through non‐covalent interactions (van der Waals forces, electrostatic and hydrophobic interactions, hydrogen bonding, and coordination bonds) [65]. While this phenomenon is widely observed in nature and technology across various scientific fields, its complexity limits precise replication. This method predominantly allows for the formation of nanofibers using specific polymers, including diblock copolymers, triblock copolymers, peptide‐amphiphile triblocks, dendrimers, and bolaform glucosamine derivatives [66]. Among the many polymers, self‐assembling peptides are biocompatible materials that form hydrogels without the use of toxic chemicals and decompose into natural amino acids. The sol–gel transition of peptide hydrogels occurs under physiological conditions, and their hydration capacity and nanoscale fibrous network enable them to encapsulate cells or bioactive molecules for tissue engineering and drug delivery [67]. Ionic self‐complementary peptides form peptide nanofiber hydrogels through electrostatic interactions (Figure 2c). These peptides possess alternating positively and negatively charged amino acids. Charged amino acids on the hydrophilic surface determine the ionic complementarity pattern of self‐assembled peptides and categorize them into different moduli (modulus I, − + − + − + − +; modulus II, − + + − + +; modulus III, − − + + +; and modulus IV, − − + + + +) based on their charge sequence, which influences their self‐assembly and structural properties [67, 68]. Reversing the charge orientation of the peptides can generate completely different molecular behavior [68]. Ionic self‐complementary peptides readily form β‐sheet structures and nanofibers, producing stable hydrogels. Along with self‐assembling nanofiber hydrogels, peptide amphiphiles (PAs) are nanomaterials that can self‐assemble into well‐defined nanofibers. PA molecules consist of hydrophobic alkyl tails and short hydrophilic peptide sequences. They self‐assemble into nanofibers structurally similar to cylindrical micelles. The hydrophobic tails cluster together to form a fibrous core, while the hydrophilic peptide segments are exposed on the surface and interact with the surrounding environment [69].

5. Polymer Materials for Nanofibrous Dressing Formation

For applications in wound dressings or implantable scaffolds, nanofibers must exhibit biocompatibility, biodegradability, and the absence of cytotoxicity for both the fibrous structure and its degradation byproducts [70]. This section provides an overview of natural and synthetic polymers commonly utilized in the fabrication of nanofibers, highlighting the specific characteristics that render them suitable for wound‐healing applications.

5.1. Natural Polymers

Natural polymers, specifically proteins and polysaccharides derived from diverse biological origins, frequently exhibit superior biocompatibility. These materials undergo enzymatic degradation within physiological environments and subsequently absorbed. Furthermore, their capacity to mimic the cellular microenvironment of native tissues enables the provision of appropriate biological cues that modulate cellular process [71].

5.1.1. Collagen

As the most prevalent structural protein within the human ECM, collagen exhibits exceptional biocompatibility, low‐immunogenicity, and high amenability to chemical modification via its diverse functional groups. Numerous cell surface receptors, including integrin, discoidin domain receptors (DDRs), glycoprotein VI (GPVI), and leukocyte‐associated immunoglobulin‐like receptor 1 (LAIR‐1), specifically interact with collagen to activate intracellular signaling pathways that regulate fundamental cellular processes such as adhesion, migration, and differentiation [72]. It is produced by fibroblasts and contributes to the mechanical stability and elasticity of tissue. To date, 28 types of collagen have been identified; among these, collagen type I is the most abundant type, constituting 70%–80% of the dermis [73, 74]. In native tissue, the fibrous shape of collagen type I, ranging from 20 to 500 nm in diameter, contributes to cell attachment and proliferation [75]. For example, tilapia collagen nanofiber fabricated via electrospinning by Zhou et al. induced cell adhesion and proliferation of human keratinocytes [76]. The cells cultured on collagen nanofiber could firmly attach within 24 h and had 114% of proliferation rate after being cultured for 5 days. The collagen nanofiber could induce differentiation of keratinocyte, important for re‐epithelization, and it is speculated to be a synergistic effect of various amino acids, especially abundantly existing proline. Although collagen is known to be advantageous for cell adhesion, processing into nanofiber through electrospinning using fluoroalcohols, which is frequently used as a solvent for various polymers including collagen, denatures the collagen into gelatin losing it biological properties [77]. Rho et al. demonstrated that electrospun type I collagen nanofiber, which used 1,1,1,2,2,2‐hexafluoro‐2‐propanol (HFIP) as a solvent, exhibited less attachment of normal human keratinocyte than polystyrene surface [78]. The lower cell attachment was addressed by coating the nanofiber with type I collagen, which promoted cell spreading, proliferation, and migration through the nanofiber pores. Despite the advantages of collagen, S. aureus, a primary pathogen in chronic wound infections, can bind to collagen fibers to evade the host immune system [79]. Consequently, the application of collagen‐based dressings to severely infected wounds should be avoided unless combined with appropriate antimicrobial agents.

5.1.2. Gelatin

Gelatin is a derivative of collagen type I produced by the hydrolysis of collagen. During this processing, collagen undergoes thermal or chemical denaturation, resulting in the loss of its native triple‐helical conformation [80]. Due to its compositional homology with collagen, gelatin effectively recapitulates many of its biological functions, particularly in facilitating cellular development in vitro [81]. Gelatin's intrinsic properties, such as its low production cost, high biocompatibility, and tunable degradation profile, have facilitated extensive investigation into its use for various biomedical applications, including drug delivery, tissue engineering, and sophisticated wound dressings [82]. Furthermore, it contains the arginine‐glycine‐aspartic acid (RGD) sequence, which promotes cell adhesion through specific binding to integrin receptors expressed on the cell surface. This integrin‐mediated cell adhesion influences and regulates fundamental cellular processes [83]. While gelatin is more cost‐effective and exhibits lower immunogenicity than collagen, its inherent mechanical weakness and low thermal stability at physiological temperatures necessitate the use of cross‐linking methods. Rather than being a mere derivative of collagen, gelatin, particularly in its modified forms like gelatin methacryloyl (GelMA), is a highly tunable platform. These modifications enable precise control over 3D architecture and degradation kinetics, positioning gelatin as a versatile candidate for ‘smart’ bioactive dressing.

5.1.3. Chitosan

Chitosan is a linear polysaccharide composed of randomly distributed β‐(1→4)‐linked D‐glucosamine and N‐acetyl‐D‐glucosamine units [84]. It is primarily derived from the deacetylation of chitin, the second most abundant natural polymer after cellulose, which serves as a key structural element in the exoskeletons (shells) of crustaceans such as shrimp and crab [85]. The degree of deacetylation (DA) of chitosan significantly influences its solubility characteristics. Specifically, chitosan becomes soluble in an acidic environment due to the protonation of its free amine groups, which occurs when the DA is typically around 50% or higher, with an optimal solubility range in diluted acids at a pH below 6 [86]. The deacetylation process imparts significant cationicity to chitosan polymers, facilitating various chemical modifications and forming the basis for several desirable properties. Consequently, modified chitosan can exhibit valuable properties such as mucoadhesion, anti‐inflammatory effects, antioxidant activity, and wound healing, making it highly versatile for biomedical applications [87]. Among the various biochemical properties of chitosan, its antimicrobial activity in particular has attracted interest in the field of wound healing. For example, Arkoun et al. elucidated the underlying mechanisms of this activity by inoculating various Gram‐negative and Gram‐positive bacteria onto chitosan‐based nanofibers [88]. They observed a marked inhibition of bacterial proliferation, characterized by the extracellular leakage of intracellular proteins and DNA. This bactericidal effect is hypothesized to stem from the polycationic nature of the amine groups, which facilitate electrostatic interactions with the bacterial cell wall, resulting in membrane permeabilization and subsequent pore formation. Despite these advantages, the utility of chitosan is often limited by its poor solubility at physiological pH and the inherent brittleness of pure chitosan nanofibers, which can lead to dressing failure in mechanically active regions like joints. To address these constraints, researchers employ composite fiber fabrication or chemical quaternization (e.g., N, N, N‐trimethylchitosan), which enhances solubility and flexibility, ensuring durability under physical strain.

5.1.4. Hyaluronic Acid

Hyaluronic acid (HA) is a linear, anionic polysaccharide composed of repeating disaccharide units of D‐glucuronic acid and N‐acetyl‐D‐glucosamine. As a non‐sulfated glycosaminoglycan (GAG) with ubiquitous distribution throughout the body, HA exhibits inherent biodegradability and is subject to rapid in situ degradation by hyaluronidases. The highest concentrations of HA are found within the ECM of tissues such as the skin, synovial fluid, vitreous body of the eye, and the umbilical cord [89]. The abundant number of hydroxyl (─OH) and carboxyl (─COOH) groups in the HA structure facilitates extensive interactions with water molecules [90]. This molecular interaction renders HA highly hydrophilic and accounts for its remarkable ability to retain water up to several hundred times its own weight. Furthermore, those functional groups facilitate chemical modification to enhance their biological and physicochemical properties. HA is known to influence cell adhesion and migration through interaction with cellular receptors [91, 92]. Recent evidence suggested that HA serves as an immunomodulatory agent, facilitating the transition of macrophages from the M1 to the M2 phenotype, and this process is critically governed by the polymer's molecular weight [93]. Research by Rayahin et al. indicated that high molecular weight HA promotes a pro‐resolving macrophage phenotype characterized by the upregulation of anti‐inflammatory markers. In contrast, low molecular weight HA fragments, consisting of fewer than 12 disaccharide units, were found to trigger a pro‐inflammatory response, underscoring the molecular weight‐dependent nature of HA‐mediated immunomodulation

5.2. Synthetic Polymers

Synthetic polymers have been extensively investigated for applications in wound healing. These materials are characterized by their cost‐effectiveness, reliable material sourcing, and minimal lot‐to‐lot variability [94]. Furthermore, synthetic polymers enable precise modification over their intrinsic properties, offering consistent and homogeneous physicochemical characteristics and enhanced stability [95]. However, a primary limitation of most synthetic polymers is their inherent bio‐inertness. To enhance their biological functionality, various modification strategies are employed, including the direct chemical functionalization of polymer chains with biomolecules, surface modification of bulk substrates, and the incorporation of bioactive additives, such as collagen [96].

5.2.1. Polycaprolactone

Polycaprolactone (PCL) is one of the biodegradable polymers synthesized by ring‐opening reaction of ε‐caprolactone. It is a Food and Drug Administration (FDA) approved material, widely used for medical devices, implants, and tissue engineering showing a slow degradation rate due to its semi‐crystallinity and hydrophobic nature [97]. Despite its utility, PCL has limitations due to its inherent hydrophobicity and lack of a biomimetic niche for cell adhesion. A moderately hydrophobic surface can promote protein adsorption while maintaining its structural stability, but an excessively hydrophobic surface can lead to irreversible denaturation and structural rigidity of the adsorbed protein, thereby inhibiting its biological activity [98]. Therefore, surface modification techniques, including physical adsorption and chemical conjugation with bioactive molecules, are essential to modulate the surface properties of PCL and enhance its biological reactivity to impart biofunctionality. Among various surface modification techniques, hydrolysis and aminolysis are frequently employed to enhance the hydrophilicity of biodegradable polymers. Yaseri et al. demonstrated that these processes significantly increase the surface wettability of PCL nanofibers without altering the underlying semi‐crystalline structure of the bulk polymer [99]. Furthermore, a moderate degree of hydrolysis and aminolysis was found to improve the spreading, viability, and proliferation of cells cultured on the nanofiber scaffolds.

5.2.2. Poly(lactic‐co‐glycolic acid)

Poly(lactic‐co‐glycolic acid) (PLGA) is a representative biodegradable and biocompatible copolymer approved by the FDA for clinical applications, widely used in medical devices and drug delivery systems [100]. PLGA's utility stems from its ability to be broken down into natural metabolites, lactic acid and glycolic acid, which are metabolized through the Krebs cycle, minimizing systemic toxicity [101]. Despite the advantages of PLGA, the inherent hydrophobicity and the accumulation of acidic degradation products, which can lower localized pH and trigger inflammatory responses, is remaining drawbacks. Consequently, numerous strategies have been investigated to enhance its hydrophilicity and bioactivity to broaden its therapeutic application. Surface modification of PLGA nanofiber with bioactive material is one of the promising techniques improving the bioactivity of PLGA. For instance, Park et al. utilized plasma treatment followed by acrylic acid grafting to functionalize PLGA nanofibers. This modification significantly enhanced the nanofibers' hydrophilicity, leading to superior cell attachment and proliferation compared to the pristine PLGA counterparts [102].

5.2.3. Polyvinyl Alcohol

Polyvinyl alcohol (PVA) is a promising candidate for wound dressing applications due to its distinct physical and chemical properties, such as water solubility, crystallinity, mechanical strength, and biodegradability, depending on the degree of hydrolysis [103]. However, the inherent bio‐inertness and low elasticity of pure PVA continue to limit its application as a wound dressing [104]. To mitigate the shortcomings of pure PVA, a promising strategy involves the fabrication of hybrid polymeric scaffolds. Combining PVA with natural or synthetic polymers allows for tailored synergistic improvements in bioactivity, elasticity, and mechanical stability. Kang et al. demonstrated that chitosan coating to thermally crosslinked PVA nanofibers significantly enhances their tensile strength and structural integrity, while concurrently resulting in a moderate attenuation of the scaffold's inherent hydrophilicity [105].

6. Applications of Nanofibers for Wound Management

In recent years, nanofiber dressings have evolved into multifunctional scaffolds capable of integrating the diverse physiological requirements essential for effective wound repair. Beyond their morphological biomimicry of the ECM, the modulation of their mechanical and physicochemical properties can be achieved through chemical modifications to enhance cell–substrate interactions. In addition to these structural refinements, the incorporation of bioactive agents can endow these scaffolds with novel functionalities, such as immunomodulatory and antimicrobial properties. These advanced features are specifically designed to address the major pathological complications that typically characterize impaired wound healing. Table 1 summarizes the therapeutic applications of various natural and synthetic polymeric nanofibers, specifically those incorporating bioactive reagents to enhance clinical outcomes.

TABLE 1.

Bioactive properties of nanofibrous dressings for wound healing applications.

Basal materials	Bioactive reagents	Bioactive properties for wound healing	Reference
PCL (electrospinning)	Gelatin	Nanofiber expansion through gas‐forming and fast blood absorption	Chen et al. [108]
Cotton gauze	CNF, PTEF	Superhydrophobic surface preventing fibrin clot formation	Li et al. [109]
PCL (electrospinning)	PDA, thrombin	Increased stability of thrombin and enhanced adhesion to traumatic wounds	Cui et al. [110]
PLGA (solution blowing spinning)		Fast in situ formation of nanofiber and hemostasis	Behrens et al. [111]
PVA (sol–gel electrospinning)	SiO2, CaO, P2O4	High elasticity and adhesiveness with sustained release of ions for hemostasis.	Lu et al. [149]
CS/PEO (solution blowing)	HNT	Hemostatic and antibacterial effect	Huang et al. [150]
CS/casein (self‐assembly)		Rapid and robust blood clot formation	Mishra et al. [151]
PVP (electrospinning)	CuS	In situ nanofiber formation and antibacterial effect against superbacteria via photothermal effect	Liu et al. [152]
PVA (electrospinning)	PBChi	Antioxidant activity via ROS scavenging	Oh et al. [116]
SA/PVA (electrospinning)	Taxifolin	Enhanced wound healing through TLR4/NF‐κB/NLRP3 signaling pathway inhibition	Wang et al. [117]
PCL/gelatin (electrospinning)	OI	Antioxidant activity through activation of NRF2	He et al. [153]
HHA (electrospinning)		Macrophage phenotype change from M1 to M2	Liu et al. [118]
PVP/CS (electrospinning)	DHM	Anti‐inflammatory effect by TLR4/MyD88/NF‐κB signaling pathway suppression and autophagy‐protein regulation	Liu et al. [154]
SF/PVP (electrospinning)	Puerarin	Anti‐inflammatory effect by inhibiting TLR4/MyD88/NF‐κB and PI3K/AKT signaling pathway	Sun et al. [155]
PLLA (electrospinning)		Nanofiber topology induced immunomodulation suppressing M1 macrophage polarization	Xie et al. [156]
PCL (electrospinning)	Macrophage cell membrane	M2 macrophage cell membrane induced immunomodulatory effect	Nakkala et al. [157]
PLCL (electrospinning)	TA, poly‐L‐lysine	Antibacterial effect against S. aureus by membrane disruption	Xu et al. [125]
Chitosan/PEO (electrospinning)	AgNPs, ZnONPs	Antioxidant and antibacterial effect against S. aureus, E. coli, and P. aeruginosa	Bagheri et al. [126]
Silk fibroin (electrospinning)	Cys‐KR12	Antimicrobial activity against multiple pathogenic bacteria and prevention of biofilm formation	Song et al. [129]
Gelatin (electrospinning)	ε‐polylysine, PDA	Antimicrobial activity against P. aeruginosa and the fibroblast migration‐promoting effect	Mayandi et al. [158]
PVA (electrospinning)	SPC	Oxygen release and upregulation of HIF‐1α	Zehra et al. [138]
PLA (electrospinning)	CaO2 NP, MnO2 NS	Sustained oxygen generation through a catalytic reaction	Dos Santos et al. [139]
Ten‐2 (pro‐angigogenic peptide) (self‐assembly)		pH‐dependent self‐assembly and enhanced neovascularization	Chu et al. [140]
PLGA (electrospinning)	Gelatin, PR1P	PR1P‐mediated in situ recruitment of VEGF and immune modulation	Chen et al. [141]
Zein/PVP (electrospinning)	TA@CuNP	Upregulated VEGF and eNOS expression via PI3K‐aktsignaling pathway	Deng et al. [159]
PCL, Gelatin (electrospinning, electrostatic spray)		Angiogenesis promotion by the paracrine effect of ADSC on the nanofiber grid	Liu et al. [160]
PVA/Gelatin (co‐axial electrospinning)	A‐FRP	Angiogenesis stimulation via the release of PDGF‐AB and VEGF	Tavakoli et al. [161]
PVA/HA/chitosan (electrospinning)	Temproin RA	pH‐responsive antimicrobial peptide release	Cao et al. [145]
PU (electrospinning)	MnO2 NP, GSV	ROS‐scavenging and anti‐inflammatory effect through STING pathway inhibition	Ahmed et al. [146]
PLATMC (electrospinning)	GelMA, Epi‐1@CS NP	Temperature‐triggered scaffold contraction and antibacterial effect	Huang et al. [147]
GelMA, PLLA (co‐electrospinning)	PNVCL	Temperature‐responsive mechanotransduction induced fibroblast differentiation	Guo et al. [148]
PVB (electrospinning) polyvinylbutyral	GOx/CDs@MOF	pH‐responsive antibacterial effect and real‐time visual monitoring	Zhang et al. [162]
PAN (electrospinning)	Eu CPs	Real‐time visible monitoring of H2O2 level and angiogenesis stimulation	Wu et al. [163]

6.1. Hemostatic Nanofibrous Dressing

Excessive blood loss can lead to hemorrhagic shock, a life‐threatening physiological condition characterized by systemic hypoperfusion. The primary conventional strategy for controlling bleeding is applying continuous manual pressure to the injured area using sterile gauze. This method rapidly absorbs local fluids, and its high absorption capacity helps to concentrate blood cells, activate platelets, and consequently stop the bleeding [106]. However, the limitations of cotton gauze remain that it absorbs excessive blood for hemostasis and causes secondary damage upon removal. Additionally, other hemostatic materials possess several limitations, including causing allergic reactions, thrombosis, and heat generation [107]. To address the inconvenience of existing hemostatic materials, nanofibers had attention to be used as the next hemostatic material due to their high porosity, surface‐to‐volume ratio, and facile surface modification.

Chen et al. engineered a shape‐memorable nanofiber matrix via a gas‐forming expansion after electrospun PCL nanofiber [108]. This matrix was capable of expanding to over 30 times its initial volume, forming a highly porous scaffold with great water absorption capacity (Figure 3a). Functionalization with thrombin yielded a highly effective hemostatic material that significantly outperformed commercial alternatives, demonstrating accelerated clotting times independent of local pH changes. Li et al. focused on preventing the blood loss through hemostatic materials and developed a superhydrophobic hemostatic gauze by coating the carbon nanofiber on the cotton gauze [109]. The hydrophobic nature of carbon nanofiber promoted the formation of fibrin fibers when the droplet of blood is in contact with the material and facilitated the formation of fibrin fibers (Figure 3b). The superhydrophobic gauze resisted blood wetting, reducing the blood loss and allowed the facile detachment of clots due to the presence of air pockets at the blood–substrate interface.

FIGURE 3.

(a) Morphology of PCL nanofiber before and after gas‐forming expansion. Reproduced with permission [108]. Copyright 2018, Elsevier. (b) Scanning electron microscopy (SEM) image and fibrin fiber formation on of superhydrophobic fiber surface. Reproduced with permission [109]. Copyright 2019, Springer Nature. (c) PLGA nanofiber formation through the solution blowing technique using a commercial airbrush and an SEM image of PLGA nanofiber incubated with citrated‐treated human whole blood. Reproduced with permission [111]. Copyright 2014, American Chemical Society.

For rapid hemostasis, conjugating thrombin to the nanofiber surface was suggested by Cui et al. [110]. In this study, PCL nanofibers were surface‐modified with polydopamine (PDA) to facilitate the conjugation of thrombin. The PDA coating enhanced surface hydrophilicity and enabled a higher thrombin loading capacity via chemical conjugation compared to pristine PCL nanofibers. The higher degree of thrombin functionalization accelerated blood coagulation, which was comparable to commercial standards in both in vivo rat femoral artery and liver hemorrhage models.

In situ formation of a nanofiber is useful for the site‐specific application of hemostatic materials. Behrens et al. utilized a solution blow spinning technique, employing a commercial airbrush and compressed CO₂, to fabricate PLGA nanofibers (Figure 3c) [111]. This method facilitated the direct deposition of nanofibers onto various substrates, with the resulting scaffolds exhibiting mechanical properties comparable to several human soft tissues under optimized conditions. The hemostatic efficacy of the system was validated across multiple surgical models, where the nanofibers effectively achieved hemostasis and sealed air leakage within 60 s.

6.2. Anti‐Inflammatory Nanofibrous Dressing

Inflammation is an essential component of the wound healing cascade; its duration and intensity directly influence the quality of recovery outcomes [112]. An excessive inflammatory response significantly results in scar formation and impaired wound healing, which can develop an acute wound into a chronic wound [113]. Therefore, maintaining a proper homeostatic balance and facilitating a timely transition between the inflammatory and subsequent repair phases is critical for effective tissue regeneration and functional recovery. Accordingly, numerous approaches utilizing nanofibers have been suggested to regulate local inflammation.

As reactive oxygen species (ROS) are known participants in inflammatory responses that promote inflammation across a variety of diseases, ROS‐scavenging biomaterials have been designed specifically to mitigate this process [114, 115]. Oh et al. incorporated chitosan‐stabilized Prussian blue nanoparticles (PBChi) into PVA nanofibers to confer ROS scavenging activity (Figure 4a) [116]. The antioxidant efficacy correlates to chitosan molecular weight; specifically, low‐molecular‐weight chitosan (10 kDa) exhibited 60% antioxidant activity, which was approximately twofold higher than the high‐molecular‐weight (100 kDa) variant (Figure 4b). This enhanced performance was attributed to the increased accessibility of hydroxyl radicals toward hydroxyl and amine functional groups. Furthermore, the scaffolds demonstrated a dose‐dependent reduction in in vitro ROS levels, achieving a maximum attenuation of 50.7% (Figure 4c).

FIGURE 4.

(a) Schematic representation of PBChi/PVA nanofiber preparation, exhibiting integrated reactive oxygen species (ROS) scavenging properties. (b) In situ antioxidant capacity of PBChi nanoparticles synthesized with varying chitosan molecular weights (10, 20, 50, and 100 kDa). (c) In vitro radical scavenging activity of PBChi/PVA nanofibers as a function of PBChi NP concentration (0, 1, and 10 µg/mL). Reproduced with permission [116]. Copyright 2022, Elsevier. (d) Photographs and SEM images illustrating the morphology of FHHA‐S/Fe nanofibrous hydrogels. (e) Comparative antioxidant capacity of HHA and HHA‐S at specified time intervals. (f) CLSM images of oxidative stress‐induced fibroblasts treated with diverse nanofibrous substrates, visualized via DCFH‐DA fluorescent staining. (g) The immunophenotypic profile of stimulated macrophages after treatment with nanofibers was analyzed by flow cytometry. (h) CLSM images representing macrophage polarization to M2 phenotype. Reproduced with permission [118]. Copyright 2020, John Wiley and Sons.

Wang et al. invented a nanofiber dressing made of sodium alginate, and poly(vinyl alcohol) loaded with taxifolin, an antioxidant flavonoid, for diabetic wound treatment [117]. Taxofolin exhibited rapid release characteristics for the first 4 h, followed by cumulative release thereafter. The nanofibrous dressing showed antioxidant, antibacterial, and promoted wound healing by taxifolin loading. Additionally, an anti‐inflammatory effect could be achieved by inhibiting the TLR4/NF‐κB/NLRP3 pathway and upregulating growth factor expression.

High molecular weight hyaluronic acid (HHA) has been shown to reduce inflammation‐inducing macrophage phenotype change. Liu et al. designed a rapidly degradable HHA‐based hydrogel nanofiber for chronic wound healing [118]. In this study, thioester groups were grafted onto the HHA backbone; following electrospinning, the resulting nanofibers were crosslinked via Fe³⁺ ions (Figure 4d). The scaffold demonstrated significant antioxidant efficacy, achieving a 29% reduction in ROS within six hours, which is attributed to the grafted thioester moieties (Figure 4e,f). Furthermore, the HHA component facilitated M1‐to‐M2 macrophage polarization, a process notably enhanced by the thioester modifications (Figure 4g,h). In a diabetic murine wound model, these nanofibers demonstrated effective immunomodulation and significantly accelerated the kinetics of wound closure

6.3. Anti‐Bacterial Nanofibrous Dressing

Bacterial infection at the wound site disrupts wound healing and causes chronic wounds. Sub‐infective bacterial levels may induce a degree of inflammation that contributes to the acceleration of wound healing and granulation tissue formation by increasing the blood flow and infiltration of granulocytes [119]. However, bacteria and their endotoxins that persist within tissues can induce tissue necrosis and increase the production of inflammatory cytokines, which can prolong the inflammatory state and eventually lead to chronic wounds [120]. In a state of chronic inflammation, the level of proteinase activity significantly increases compared to normal conditions, and most of the protease activity is attributable to MMPs, which can degrade ECM, growth factors, and their receptors essential for wound healing [121]. The extensive use of antibiotics to prevent wound infections has led to the emergence of antibiotic‐resistant bacteria, such as methicillin‐resistant S. aureus (MRSA) and vancomycin‐resistant S. aureus (VRSA). Therefore, these problems have necessitated the development of alternative antimicrobial agents, and these alternative materials, including antimicrobial polymers, antimicrobial peptides, nanoparticles, and metal oxides, have been applied to nanofibers to manufacture antimicrobial dressings [122, 123, 124].

Polycationic polymers are known to possess antimicrobial properties. Xu et al. engineered a poly‐L‐lysine and tannic acid incorporated poly(L‐lactide‐co‐ε‐caprolactone) (PLCL) nanofiber for chronic wound management (Figure 5a) [125]. This study demonstrated that the inherent nanofibrous architecture serves as an effective physical barrier against S. aureus infiltration. While nanofibers functionalized solely with tannic acid exhibited notable antibacterial activity, the co‐incorporation of poly‐L‐lysine (PLL) yielded a synergistic effect, significantly reducing bacterial viability (Figure 5b–d). The enhanced antimicrobial efficacy is attributed to the electrostatic interaction between the cationic PLL and the negatively charged bacterial cell membrane, which subsequently triggers membrane disruption and the intracellular accumulation of ROS.

FIGURE 5.

(a) Schematic illustration representing the sequential surface modification of PLCL nanofibers via the incorporation of tannic acid and L‐lysine. (b) Analysis of antibacterial effect of antibacterial nanofiber scaffolds against S. aureus. (c) Quantitative assessment of the bacterial isolation efficiency. (d) Statistical quantification of bacterial survival rates. Reproduced with permission [125]. Copyright 2025, Elsevier. (e) Schematically representing surface immobilization of Cys‐KR12 peptide onto the silk fibroin (SF) nanofiber. (f) SEM images of S. aureus and E. coli following 24 h of cultivation on pristine SF (left column) and K200 (right column). (g) Quantitative evaluation of the antimicrobial efficacy of Cys‐KR12 peptide‐functionalized SF nanofibers against Gram‐negative and Gram‐positive bacteria. Reproduced with permission [129]. Copyright 2016, Elsevier.

Metal‐based nanoparticles such as silver, zinc oxide, copper, and titanium oxide nanoparticles have been widely used for their antimicrobial properties [122]. Bagheri et al. developed an antibacterial nanofiber via blending silver and zinc oxide nanoparticles into chitosan/poly(ethylene oxide) polymer matrix [126]. The addition of metallic nanoparticles exhibited a synergistic antibacterial effect against both Gram‐positive (S. aureus) and Gram‐negative (P. aeruginosa and E. coli) bacteria. S. aureus was found to be more susceptible to the metallic nanoparticles than the Gram‐negative bacteria. The difference in antibacterial efficacy is attributed to the distinct cell wall compositions of Gram‐positive bacteria, specifically the absence of an outer membrane in Gram‐positive bacteria, which affects nanoparticle permeability.

Antimicrobial peptides (AMPs) are polypeptide molecules composed of 10 to 60 amino acid residues, found across all organisms [127]. They are considered a rising alternative to conventional antibiotics due to several advantages, including ease of synthesis, cost‐effectiveness, rapid bacterial killing, non‐selective killing effects against bacterial resistance phenotypes, and no disruption to the microbiota [128]. Song et al. engineered AMP‐immobilized silk fibroin nanofibers to attain antimicrobial activity at the wound site and to increase peptide stability using KR12, an antimicrobial motif derived from LL37 (Figure 5e) [129]. The antimicrobial activity of these modified nanofibers was confirmed against various pathogenic bacteria (S. aureus, S. epidermidis, E. coli, and P. aeruginosa), with no observation of biofilm formation (Figure 5f). The resulting antimicrobial efficacy was proportional to the density of AMPs immobilized on the nanofiber surface (Figure 5g).

6.4. Pro‐Angiogenic Nanofibrous Dressing

During the early stages of wound healing, the wound site becomes hypoxic due to disturbed blood supply and increased oxygen consumption by infiltrated inflammatory cells involved in the healing process [130]. Hypoxia, which refers to a reduced level of tissue oxygenation, promotes wound healing by angiogenesis, cellular proliferation, and immune modulation through activation of hypoxia‐inducible factor‐1 alpha (HIF‐1α) in acute wounds [131]. Despite the advantages of hypoxia in wound healing, prolonged hypoxia in chronic wounds impairs wound healing by altering the pH of the wound environment, increasing oxidative stress, hindering angiogenesis, and impeding immune cell functions [132]. Insufficient local vascularization limits oxygen supply to the wound site, contributing to the poor or inconsistent therapeutic effects of systemic and topical therapies widely reported in clinical trials to promote chronic wound healing [133, 134, 135].

In previous studies, oxygen‐generating dressing materials have suggested to overcome the limitations of systemic oxygen supply by increasing the localized oxygen concentration at the wound site. The generation of molecular oxygen typically leverages the aqueous hydrolysis of metal peroxides. This chemical pathway is characterized by the initial formation of hydrogen peroxide as a transient precursor, which subsequently undergoes disproportionation into water and oxygen [136]. However, a primary limitation associated with current solid peroxides is the burst release of hydrogen peroxide (H₂O₂) and the potential formation of cytotoxic byproducts. To mitigate these challenges, researchers have explored strategies such as encapsulating solid peroxides within hydrophobic materials to control the hydrolysis rate of the peroxides by limiting their exposure to water molecules [137].

For instance, Lai et al. invented an oxygen‐releasing nanofiber by incorporating sodium percarbonate (SPC), which has low water solubility, to a PCL nanofiber matrix for diabetic wound healing [138]. This nanofiber demonstrated sustained oxygen release for 10 days under physiological conditions, which was attributed to the slow degradation of the PCL matrix. The oxygen‐releasing scaffolds exhibited significant angiogenic potential, as evidenced by the upregulation of HIF‐1α expression in human skin fibroblasts and a greater density of newly formed blood vessels observed in chorioallantoic membrane (CAM) assays compared to the neat PCL controls.

Alternatively, Santos et al. proposed the use of manganese dioxide (MnO₂) as a stable catalyst for the decomposition of H₂O₂, offering a more robust alternative to catalase, which is susceptible to denaturation [139]. They engineered oxygen‐releasing nanofibers designed to modulate the local partial pressure of oxygen by incorporating calcium peroxide (CaO₂) nanoparticles and MnO₂ nanosheets within a polylactic acid (PLA) matrix. In this biphasic system, CaO₂ generates H₂O₂ upon hydration, while the MnO₂ nanosheets function as nanozymes to facilitate the catalytic conversion of H₂O₂ into molecular oxygen. This composite scaffold demonstrated sustained oxygen release for over seven days, with the release kinetics appearing directly proportional to the concentration of incorporated CaO₂ nanoparticles.

In another approach, directly promoting angiogenesis using bioactive reagents, such as peptides and growth factors, has been suggested. Chu et al. developed a nanofibrous hydrogel composed of the proangiogenic peptide GEETEVTVEGLEPG (Ten‐2) [140]. This peptide was designed with alternating hydrophilic and hydrophobic residues, facilitating self‐assembly into nanofibrous architectures via the formation of β‐sheet secondary structures, which subsequently induced gelation under acidic conditions (Figure 6a,b). The resulting peptide hydrogel significantly enhanced the adhesion, proliferation, and migration of human umbilical vein endothelial cells (HUVECs), thereby promoting angiogenesis in a dose‐dependent manner (Figure 6c,d). In a mouse subcutaneous implantation model, the hydrogel facilitated the infiltration of inflammatory and endothelial cells, orchestrating robust neovascularization over a 14‐day period concurrent with the progressive biodegradation of the peptide scaffold.

FIGURE 6.

(a) Chemical structures and molecular design of self‐assembling peptides. (b) Optical micrographs and SEM images illustrating the morphology of the self‐assembled peptide hydrogels. (c) Evaluation of the in vitro angiogenic potential of self‐assembled peptide hydrogels. (d) Dose‐response kinetics and morphological quantitative analysis of self‐assembled peptide hydrogels. Reproduced with permission [140]. Copyright 2021, American Chemical Society. (e) Schematic representation and (f) quantitative assessment of VEGF recruitment by nanofibrous dressings. Evaluation of angiogenic potential mediated by dressing‐recruited VEGF as a function of PR1P concentration (0.02, 0.2, and 2 mg). (g) Assessment of tubular network formation in human umbilical vein endothelial cells (HUVECs). (h) In ovo chorioallantoic membrane (CAM) assay evaluating the neovascularization capacity of nanofibrous scaffolds. Reproduced with permission [141]. Copyright 2023, Springer Nature.

Addressing the inherent instability of VEGF, Chen et al. developed an in situ VEGF‐recruiting dressing designed to facilitate localized angiogenesis and immunomodulation [141]. The nanofibrous dressing was fabricated via the electrospinning of a PLGA/gelatin blend solution incorporated with the VEGF‐binding domain of the Prominin‐1‐derived peptide (PR1P). The encapsulated PR1P exhibited a sustained release profile over a 72‐hour period, effectively recruiting VEGF to promote the migration of HUVECs and the formation of tubular networks in both in vitro models and CAM assays in a dose‐dependent manner (Figure 6e–h). Furthermore, the nanofibrous dressing accelerated wound closure in splinted excisional rat models by enhancing neovascularization and inducing macrophage polarization toward the pro‐healing M2 phenotype, irrespective of diabetic status.

6.5. Smart Nanofibrous Dressing

Smart wound dressings indicate a sophisticated theranostic wound management platform that integrates autonomous sensing, diagnostic feedback, and programmable response capabilities to regulate the regenerative process [142, 143]. These systems are designed to continuously monitor physiological parameters, including pH, temperature, mechanical force, and specific biomolecules [144]. Smart dressings, utilizing stimulus‐responsive architectures, effectively optimize the wound microenvironment by responding to external or internal stimuli to deliver targeted therapy and spatiotemporally controlled delivery of bioactive agents as needed. Ultimately, the integrated platform delivers significantly superior therapeutic outcomes than passive approaches, accelerating tissue remodeling and improving clinical outcomes.

Among various external stimuli, pH remains the most extensively investigated parameter for the design of ‘smart’ wound dressings. Cao et al. engineered a pH‐responsive drug delivery scaffold by electrospinning a ternary mixture of chitosan, hyaluronic acid, and PVA (Figure 7a) [145]. Subsequently, the antimicrobial peptide Temporin‐Ra (FP‐14) was immobilized onto the nanofiber surface via electrostatic interactions. Optimal peptide loading was achieved in an acidic environment (pH 5), attributed to the high isoelectric point (pI 10) of FP‐14 (Figure 7b). This electrostatic tethering facilitated the sustained release of the peptide under alkaline conditions (Figure 7c). In a murine excisional wound splinting model, the wound microenvironment underwent a physiological pH shift from 5 to 9; this alkaline transition triggered the accelerated release of FP‐14, enhancing the wound closure rate to 60%, significantly outperforming the non‐treated control group, which displayed a closure rate of only 12% (Figure 7d,e).

FIGURE 7.

(a) Schematic illustration of nanofiber fabrication and pH‐dependent peptide release. (b) Quantitative evaluation of loading efficiency and (c) release kinetics of the peptide within the nanofibrous matrix under varying pH conditions (pH 5, 7.4, and 10). (d) Representative images of the in vivo wound healing process in various nanofiber treatment groups on days 0 and 6. (e) Wound closure effect of the nanofiber substrates quantified by the reduction in residual wound area from day 0 to day 6. Reproduced with permission [145]. Copyright 2023, American Chemical Society. (f) SEM images of the thermo‐responsive nanofibers, accompanied by corresponding schematic representations and fiber orientation (angle) distribution diagrams. (g) Representative image showing the shrinkage of thermo‐responsive nanofibers after incubation at 37.8°C for 90 min. (h) Mechanical properties of thermo‐responsive nanofibers. Reproduced with permission [147]. Copyright 2024, Elsevier.

Increased levels of ROS and the subsequent activation of the cGAS‐STING pathway significantly impede the healing of diabetic wounds. To address this, Ahmed et al. engineered ROS‐responsive, diselenium‐containing polyurethane nanofibers (PUF) functionalized with MnO₂ nanoparticles and gelsevirine (GSV) for targeted ROS scavenging and immunomodulation [146]. Under simulated oxidative stress (H₂O₂ exposure), the cleavage of diselenium bonds within the nanofibrous matrix facilitated the accelerated release of GSV, while a synergistic ROS‐scavenging effect between MnO₂ and the PUF scaffold was demonstrated. Consequently, a significant reduction in the pro‐inflammatory cytokines was observed in vitro, primarily attributed to GSV‐mediated inhibition of the STING pathway. In a diabetic murine model, the MnO₂/GSV/PUF‐encapsulated nanofibers achieved a superior wound closure rate of 92% by day 14, driven by the attenuation of inflammation by GSV and oxidative stress through the integrated scavenging activity of MnO₂ and selenium.

For enhanced wound healing, not only the delivery of bioactive agents, but also mechanical stimulation serves as one of the factors. Huang et al. developed a bi‐layered nanofiber/hydrogel composite dressing using temperature‐responsive Poly(lactic acid‐co‐trimethylene carbonate) (PLATMC) nanofibers and gelatin hydrogel on the other side (Figure 7f) [147]. The nanofiber exhibited different directional contraction at 37°C corresponding to the different axial orientation of fiber alignment, which affected the mechanical strength of the nanofiber (Figure 7g,h). The bi‐axial orientation increased the mechanical strength of the nanofiber, which can provide mechanical support withstanding external stress. The antimicrobial nanoparticle, which is composed of epinecidin‐1 and chitosan, within a gelatin layer suppressed colony formation against S. aureus, E. coli, and MRSA in a dose‐dependent manner in vitro, and enhanced the wound healing of a bacterial‐infected diabetic mouse wound model.

Similarly, Guo et al. targeted modulation of fibroblast behavior providing temperature‐responsive mechanical stimulation to promote wound healing [148]. A thermo‐responsive nanofibrous hydrogel was synthesized via the co‐electrospinning of poly‐L‐lactide (PLLA) and GelMA. Subsequently, poly(N‐vinylcaprolactam) (PNVCL) was grafted onto the GelMA nanofiber layer to impart reversible mechanical responsiveness. Fibroblasts subjected to dynamic mechanical stimulation exhibited significantly higher rates of spreading and migration compared to those in a static environment. This mechanical signaling upregulated the expression of ECM‐related genes and induced the differentiation of fibroblasts into myofibroblasts. In a murine skin‐defect model, it was confirmed that dynamic mechanical stimulation promoted wound closure, and this result primarily attributed to the fibroblast polarization to the myofibroblast phenotype.

7. Conclusion and Perspectives

According to the increase of population, chronic wounds have emerged as a major concern, and the expense for impaired wound care has become an economic burden. The rising incidence of chronic wounds necessitates specialized therapeutic strategies to mitigate the complications associated with impaired healing. This demand has inspired the development of functional dressings characterized by antioxidant, antimicrobial, immunomodulatory, and pro‐angiogenic properties. Current research focuses on integrating these synergistic functions into a multifunctional platform to overcome the limitations of existing therapies and accelerate regenerative outcomes. Consequently, the unique structural characteristics of nanofibers, combined with the incorporation of bioactive agents, have established these scaffolds as prominent candidates for advanced wound management.

Through these advancements, the commercialization of nanofiber‐based wound dressings is led by innovative firms that leverage biomimetic structures to enhance tissue repair. Nanomedic Technologies has pioneered the SpinCare system, a portable device that generates a bespoke, on‐the‐spot nanofiber layer directly onto the patient's skin to treat burns and chronic wounds. In parallel, Nanofiber Solutions (RenovoDerm) offers the Phoenix Wound Matrix, an FDA‐cleared synthetic scaffold that mimics the natural extracellular matrix to facilitate cellular infiltration, while the Stellenbosch Nanofiber Company produces Nanotrix, a high‐porosity dressing designed to optimize the microenvironment for rapid re‐epithelialization. Collectively, these products represent a significant shift toward precision regenerative medicine for chronic wound management by providing breathable, biocompatible, and structurally advanced alternatives to traditional dressings. However, despite their therapeutic potential, critical obstacles remain to clinical application. The fabrication and functionalization of nanofibers often rely on toxic solvents and cross‐linking agents, necessitating more rigorous investigations regarding long‐term safety as well as environmental contamination. Second, large‐scale industrial manufacturing remains a major hurdle; while modern techniques have improved throughput compared to conventional methods, they have not yet achieved the cost‐effectiveness required for commercialization. To address these limitations, melt electrospinning and solution blow spinning are emerging as alternatives for the development of efficient, sustainable, and multifunctional nanofiber dressings.

This review details the four stages of the wound healing process and highlights key factors that regulate each stage. The mechanisms of electrospinning within the context of nanofiber fabrication for wound healing were examined, alongside a comprehensive analysis of the diverse methodologies currently employed. A comprehensive overview of natural and synthetic polymers employed for nanofiber manufacturing is described, based on their characteristic and biofunctionality for wound healing. A significant advantage of nanofibrous scaffolds lies in their capacity for the incorporation of therapeutic reagents, enabling the programmed modulation of the regenerative microenvironment through the spatiotemporal control of bioactive agent release tailored to specific healing phases. Furthermore, the relevant literature suggests that bioactive nanofiber dressings provide superior therapeutic efficacy compared to conventional treatments, providing spatiotemporally tailored intervention at each distinct phase of the regenerative process.

The authors declare no conflicts of interest.

Data Availability Statement

The data that support the findings of this study are available in the supplementary material of this article.