Advanced regenerative solutions in diabetic foot ulcer therapy: delivery of mesenchymal stem cells in injectable hydrogels

Abstract

Diabetic foot ulcers (DFUs) are a severe complication of diabetes characterized by impaired healing driven by oxidative stress, chronic inflammation, reduced angiogenesis, and neuropathy, leading to high risks of infection and amputation. Current therapies remain insufficient, necessitating advanced regenerative approaches. Mesenchymal stem cells (MSCs) have demonstrated therapeutic potential through immunomodulation, angiogenesis, and extracellular matrix remodeling; however, their clinical application is limited by poor survival and retention, as well as potential safety concerns. Increasing evidence indicates that MSC-derived exosomes play a central role in mediating these therapeutic effects via paracrine signaling, delivering bioactive cargos such as microRNAs, proteins, and cytokines that regulate immune responses, angiogenesis, and tissue regeneration with lower immunogenicity and improved stability compared to cell-based therapies. Injectable hydrogels have emerged as biomimetic platforms that not only enhance MSC viability and retention but also provide an optimal delivery system for exosomes by protecting them from degradation and enabling sustained, localized release within the wound microenvironment. The integration of MSC-derived exosomes with hydrogel systems represents a synergistic strategy that simultaneously addresses multiple pathological barriers in DFUs, including inflammation, oxidative stress, and impaired vascularization. This review highlights recent advances in hydrogel-based delivery systems for MSCs and, importantly, MSC-derived exosomes, with a particular emphasis on their combined therapeutic potential in diabetic wound healing. Furthermore, emerging smart and stimuli-responsive hydrogels are discussed as next-generation platforms for optimizing exosome delivery and improving clinical outcomes in DFUs.

Introduction

Diabetic foot ulcers (DFUs) affect approximately 18.6 million people worldwide each year, including about 1.6 million in the United States. Despite treatment, only about 65% of patients heal within a year, while approximately 5–20% eventually require major amputation. DFUs precede 80% of diabetes-related lower limb amputations and are associated with significantly increased mortality risk [1, 2]. One of the key challenges in managing DFUs is their insidious onset. Underlying peripheral neuropathy and peripheral vascular disease in diabetic patients may cause these ulcers to develop without significant presentation in the early stages, leading to delays in detection and treatment. As the disease worsens, clinical presentations typically include chronic, non-healing ulcers, often complicated by infection and ischemia [3]. Current standard treatments include blood glucose control, infection management, surgical debridement, wound dressing, and angioplasty for peripheral artery disease (PAD). DFUs are often infected with polymicrobial organisms and frequently develop antibiotic resistance, contributing to persistent, non-healing ulcers. Surgical debridement promotes healing by accelerating granulation tissue formation and re-epithelialization and helps control infection by removing necrotic tissue. However, different monotherapies have shown inconsistent results, necessitating a new approach in DFU management [4–7].

Mesenchymal stem cells (MSCs), found in various tissues [8], support tissue repair through immune modulation, angiogenesis, and cell survival [9]. MSC-derived exosomes offer similar benefits with lower immunogenicity, making them a promising option for treating inflammatory and degenerative diseases [10]. However, the therapeutic potential of MSCs or their exosomes can sometimes be limited by poor cell survival, low retention at the injury site, or inadequate integration into the damaged tissue.

To overcome these challenges, hydrogels have been explored as supportive delivery systems. Hydrogels are three-dimensional, crosslinked polymer networks that can absorb large amounts of water. Their high water content, soft consistency, and porosity allow them to mimic the extracellular matrix, creating a proper environment for cell viability and wound healing [11]. Hydrogels are also responsive to external stimuli such as pH, temperature, and enzymatic activity, making them highly adaptable for biomedical applications by providing a suitable repair environment, reducing infection, and improving tissue regeneration [12, 13].

In this paper, we review the pathophysiology of DFUs and the underlying mechanisms of wound healing. We then highlight the therapeutic potential of various injectable hydrogels in regenerative medicine, particularly in the management of DFUs. Special emphasis is placed on the synergistic integration of MSCs with hydrogels, which enhances regenerative outcomes by combining MSCs’ biological activity with the structural and functional support of hydrogel matrices.

Pathophysiology of wound healing

Wound healing is a complex, dynamic process that occurs in four overlapping phases: hemostasis, inflammation, proliferation, and remodeling. In DFUs, this coordinated sequence is profoundly disrupted, resulting in prolonged inflammation and failure to progress toward healing.

Hemostasis

Immediately after injury, vasoconstriction occurs and platelets aggregate to form a fibrin clot. This clot stops bleeding and provides a provisional matrix for cell migration. In diabetes, advanced glycation end products (AGEs) impair platelet function and fibrin stability, delaying the initiation of repair [14, 15].

Inflammation

Platelet degranulation releases chemoattractants leading to initiating the inflammatory phase [16, 17]. Neutrophils are the first innate immune cells to infiltrate the wound site. They clear pathogens through phagocytosis and by releasing reactive oxygen species (ROS), antimicrobial peptides, and neutrophil extracellular traps (NETs) [18–21].

Macrophages infiltrate the wound site within 48 h after injury. Initially, Macrophages initially adopt a pro-inflammatory M1 phenotype and later transition to a pro-healing M2 phenotype, releasing growth factors that promote tissue repair. In DFUs, hyperglycemia and oxidative stress prolong neutrophil activity and persistently drive M1 macrophage dominance. This results in excessive production of pro-inflammatory cytokines (e.g., TNF-α, IL-1β, IL-6), sustained inflammation, and failure to resolve the inflammatory phase [17, 19, 22].

Proliferation

This phase involves fibroblast migration and extracellular matrix (ECM) deposition, keratinocyte proliferation and migration for re-epithelialization, and angiogenesis to restore blood supply. These processes are driven by key growth factors, including VEGF, PDGF, and TGF-β. In diabetic wounds, angiogenesis is markedly impaired due to reduced VEGF signaling and endothelial dysfunction, while fibroblast activity and keratinocyte migration are compromised, leading to poor granulation tissue formation and delayed wound closure [16, 20].

Remodeling and scar formation

During remodeling, the provisional ECM is reorganized, with type III collagen gradually replaced by stronger type I collagen to restore tissue tensile strength. In DFUs, an imbalance between matrix metalloproteinases (MMPs) and inhibitors of metalloproteinases (TIMPs) causes excessive ECM degradation, resulting in weak scar tissue and chronic ulceration [23, 24].

In summary, diabetic wounds are typically stalled in a prolonged inflammatory phase with impaired transition to proliferation and remodeling. The main underlying causes include persistent hyperglycemia, oxidative stress, neuropathy, peripheral arterial disease, and dysregulated immune responses. Understanding these disruptions is essential for developing targeted regenerative therapies such as mesenchymal stem cell (MSC)-laden injectable hydrogels.

Pathological barriers in diabetic wound healing

Persistent hyperglycemia and advanced glycation end product (AGE) formation

Chronic hyperglycemia activates multiple metabolic pathways, including the polyol, hexosamine, protein kinase C (PKC), and glycolytic pathways. These alterations increase oxidative stress, activate NF-κB signaling, promote cytokine release, and damage ECM components, ultimately leading to the excessive formation of AGEs [25]. AGEs are generated through the non-enzymatic glycation of proteins, lipids, and nucleic acids and accumulate in diabetic tissues, including the skin [26]. Their accumulation promotes a prothrombotic state, impairs fibrinolysis, and disrupts hemostasis through NF-κB–mediated tissue factor expression and glycation of fibrinogen and plasminogen .

At the cellular level, AGEs impair macrophage viability through the miR-361-3p/CSF1R axis and PI3K/AKT signaling pathway [27]. They also inhibit the transition from the M1 phenotype to the pro-regenerative M2 phenotype, thereby sustaining chronic inflammation in diabetic wounds [28]. Mechanistically, AGEs promote M1 polarization by activating RAGE/ROS/TLR4/STAT1 and MAPK signaling pathways, leading to increased expression of inflammatory mediators such as IL-1β, TNF-α, IL-6, NOS2, NLRP3, and SOCS3 [29]. Sustained activation of the NLRP3 inflammasome further enhances the release of IL-1β and IL-18 and promotes NET formation, thereby amplifying inflammation [30]. In addition, AGEs impair neutrophil migration and adhesion by downregulating PVR/CD155 and CTNND1 expression [31], reducing effective immune responses and delaying wound healing [32].

At the tissue level, AGE accumulation induces cellular senescence and apoptosis in keratinocytes and fibroblasts while disrupting ECM integrity through collagen and elastin glycation and increased MMP-9 activity [33]. Elevated MMP levels impair granulation tissue formation, angiogenesis, and oxygen delivery [34]. Furthermore, activation of the AGE–RAGE axis creates a self-perpetuating cycle of oxidative stress, inflammation, and matrix degradation. Inhibition of this pathway has been shown to improve neutrophil function and restore early inflammatory responses [35].

In addition, AGEs impair angiogenesis by reducing endothelial progenitor cell (EPC) function and VEGF expression. They also disrupt fibroblast–endothelial interactions through collagen glycation and contribute to vascular and neural dysfunction by reducing nitric oxide(NO) bioavailability.

Excessive oxidative stress

In diabetes, hyperglycemia leads to excessive ROS production via activation of enzymes such as NADPH oxidase, xanthine oxidase, and cyclooxygenase. At the same time, antioxidant defense systems, including glutathione peroxidase, catalase, and superoxide dismutase, are impaired, resulting in redox imbalance. Under physiological conditions, ROS play beneficial roles in wound healing by supporting antimicrobial activity, activating PDGF signaling, and promoting cell proliferation, migration, and angiogenesis [36]. However, sustained ROS overproduction in diabetic wounds is detrimental.

Mitochondrial dysfunction is a major source of ROS in hyperglycemic conditions. Excess ROS induces lipid peroxidation, protein misfolding, and DNA damage, leading to cellular senescence and apoptosis. These effects impair keratinocyte and fibroblast function and delay wound repair. In addition, oxidative stress disrupts peripheral nerve integrity by reducing blood flow and altering metabolic pathways [37].

At the molecular level, ROS activate transcription factors such as NF-κB and (Activator protein-1)AP-1, thereby increasing MMP expression and promoting ECM degradation. ROS also activate the NLRP3 inflammasome, sustaining pro-inflammatory macrophage polarization and maintaining a chronic inflammatory state in diabetic wounds [38].

Neuropathy

Diabetic neuropathy is a major contributor to the development and persistence of DFUs. It results in the loss of protective sensation, allowing minor injuries to go unnoticed and progress into chronic wounds. Motor neuropathy further contributes to foot deformities and abnormal pressure distribution, while autonomic neuropathy causes skin dryness and fissure formation, increasing susceptibility to infection. At the molecular level, neuropathy disrupts the production and release of key neuropeptides, including nerve growth factor (NGF), substance P, and calcitonin gene-related peptide (CGRP). These molecules play essential roles in immune cell recruitment, cytokine signaling, keratinocyte proliferation, and fibroblast activation [39]. Sensory nerve dysfunction also reduces leukocyte recruitment and impairs chemotactic signaling, delaying re-epithelialization and prolonging the inflammatory phase. Together, these effects significantly impair wound closure.

Peripheral arterial disease (PAD) and impaired angiogenesis

PAD is a major vascular complication of diabetes that reduces tissue perfusion and impairs wound healing. Chronic hyperglycemia induces vascular smooth muscle cell proliferation and intimal thickening, leading to reduced vessel lumen diameter and compromised oxygen and nutrient delivery to tissues. Approximately 43.87% of DFU cases are associated with PAD, which is strongly linked to delayed healing, increased infection risk, and higher amputation rates [40]. Angiogenesis, a critical process during the proliferative phase, is significantly impaired in diabetic wounds due to multiple factors. These include reduced expression of pro-angiogenic factors such as VEGF, increased levels of anti-angiogenic mediators, decreased mobilization of EPCs, and defective vascular maturation [41].

Hypoxia-inducible factor-1α (HIF-1α) is a key regulator of angiogenesis under hypoxic conditions. It controls genes involved in cell survival, metabolism, and neovascularization [42]. However, hyperglycemia suppresses HIF-1α activity, resulting in reduced EPC recruitment and impaired granulation tissue formation. Experimental studies have shown that restoring HIF-1α expression improves angiogenesis and accelerates wound healing [42].

Impaired barrier function and susceptibility to infections

Hyperglycemia compromises epidermal barrier integrity by disrupting tight junction proteins, increasing transepidermal water loss (TEWL), and reducing keratinocyte proliferation and differentiation [43]. These changes facilitate microbial invasion and weaken host defense mechanisms.

Antimicrobial peptides (AMPs), including β-defensins, LL-37, psoriasin, and RNase 7, play critical roles in innate immunity and wound healing. They promote keratinocyte activation, cell migration, angiogenesis, and immune modulation. However, their expression is significantly reduced in diabetic conditions, impairing both antimicrobial defense and tissue repair [44].

In addition, the skin microbiome is altered in diabetes, with increased colonization by pathogenic bacteria such as Staphylococcus aureus and Staphylococcus epidermidis, which increases infection risk [45]. Diabetic wounds often progress to chronic infection due to ischemia, neuropathy, and immune dysfunction. Impaired neutrophil function and reduced antimicrobial peptide production further contribute to persistent infection [46].

Approximately 60% of DFUs become infected, and 15–20% of these cases require amputation despite treatment [47]. Chronic infection also promotes biofilm formation, which inhibits keratinocyte and fibroblast migration and reduces the effectiveness of antibiotics and immune responses [48, 49].

Dysregulated immune responses

Immune cell dysfunction is one of the earliest and most critical barriers in DFU pathogenesis, creating a persistent pro-inflammatory microenvironment that stalls healing [50].

Neutrophils, the first responders, exhibit impaired chemotaxis, reduced phagocytosis, and, paradoxically, prolonged survival. In DFUs, they exhibit excessive formation of NETs driven by elevated ROS and PAD4 activation. While NETs have antimicrobial functions in normal healing, their persistence causes extracellular matrix degradation, endothelial damage, and amplified inflammation, significantly delaying wound resolution [21, 51].

Macrophages play a central role in the transition from inflammation to repair. In healthy wounds, they shift from M1 to M2 phenotype. However, in diabetic wounds, this transition is blocked by sustained activation of NF-κB, STAT1, MAPK, and the NLRP3 inflammasome. The resulting M1 dominance leads to excessive secretion of TNF-α, IL-1β, IL-6, and iNOS, while production of anti-inflammatory cytokines (IL-10, TGF-β) and growth factors (VEGF) is markedly reduced. This imbalance prevents resolution of inflammation and suppresses angiogenesis, fibroblast activation, and ECM deposition [22, 28].

T cell responses are also dysregulated, with increased pro-inflammatory Th1 and Th17 subsets, elevated IFN-γ and IL-17 levels, and reduced regulatory T cell (Treg) numbers and function. Aberrant expression of chemokine receptors such as CCR4 and elevated MIF and IP-10 further amplify immune cell recruitment and chronic inflammation [50].

Collectively, these immune cell-specific abnormalities create a self-sustaining inflammatory cycle that impairs keratinocyte migration, fibroblast proliferation, angiogenesis, and matrix remodeling.

Importantly, this multifaceted immune dysregulation greatly expands the therapeutic targets for MSCs and their exosomes. MSCs can simultaneously address multiple defects by reducing excessive neutrophil NET formation, promoting M1 to M2 macrophage polarization through PGE2, TGF-β, HGF, and IL-10 (via STAT3 and PI3K/AKT pathways), and restoring T cell balance by suppressing Th1/Th17 responses while enhancing Treg expansion via FOXP3, CTLA4, and IDO [52, 53]. Therefore, a detailed understanding of immune cell dysfunction is essential for optimizing MSC-based therapies and designing advanced hydrogel delivery systems capable of targeted immunoregulation in DFUs.

Therapeutic mechanisms and delivery challenges of MSCs in wound healing

MSCs offer a promising regenerative approach for DFUs by addressing the central pathological features of diabetic wounds, including persistent inflammation, oxidative stress, impaired angiogenesis, neuropathy, and abnormal extracellular matrix remodeling [54, 55].

In DFUs, MSCs primarily exert paracrine effects rather than long-term engraftment or direct differentiation into tissue cells. They release a broad spectrum of bioactive molecules, including growth factors, cytokines, and exosomes. These secreted factors help restore the disrupted healing process by modulating immune responses, promoting new blood vessel formation, supporting tissue regeneration, and creating a more favorable wound microenvironment [56, 57].

Numerous preclinical studies have shown that MSCs can accelerate wound closure, improve granulation tissue formation, enhance vascularization, and reduce scar formation in diabetic animal models. These beneficial effects make MSCs particularly attractive for treating chronic, non-healing DFUs that respond poorly to conventional therapies [54, 56].

Key signaling pathways mediated by MSCs and MSC-derived exosomes in wound healing

MSCs and their derived exosomes primarily exert therapeutic effects in DFUs through paracrine signaling rather than direct differentiation and long-term engraftment. They deliver bioactive molecules including growth factors, cytokines, miRNAs, lncRNAs, and proteins that activate multiple interconnected signaling pathways in target cells such as macrophages, fibroblasts, keratinocytes, and endothelial cells [58, 59].

Immunomodulation and macrophage polarization: MSCs and MSC-exosomes promote the shift from M1 to M2 macrophages mainly by secreting PGE2, TGF-β, IL-10, and HGF. These factors activate the STAT3 and PI3K/AKT pathways while suppressing NF-κB and NLRP3 inflammasome signaling. Consequently, production of pro-inflammatory cytokines (TNF-α, IL-1β, IL-6) is reduced, resolving the persistent inflammation typical of diabetic wounds [60, 61].

Angiogenesis and neovascularization: A major pathway is the HIF-1α/VEGF axis. MSCs and exosomes upregulate HIF-1α, which increases VEGF expression. This activates downstream PI3K/AKT/eNOS and MAPK/ERK signaling in endothelial cells, promoting proliferation, migration, tube formation, and NO production. Exosomal miRNAs such as miR-126 and miR-21 further enhance this axis by targeting negative regulators, leading to improved angiogenesis and granulation tissue formation in hypoxic DFU environments [59, 62].

Fibroblast activation, proliferation, migration, and ECM remodeling: MSCs stimulate fibroblasts via the TGF-β/Smad2/3, PI3K/AKT, and Wnt/β-catenin pathways. MSC-derived exosomes carrying miR-21-5p or lncRNA H19 inhibit excessive myofibroblast differentiation while promoting balanced collagen I/III deposition and MMP/TIMP activity. Notably, exosomal lncRNA H19 competitively binds miR-152-3p, thereby upregulating PTEN and activating the PI3K/AKT pathway. This reduces fibroblast apoptosis and inflammation while enhancing proliferation, migration, and ECM synthesis, ultimately limiting scar formation in diabetic wounds [63, 64].

Re-epithelialization and keratinocyte function: Paracrine factors from MSCs (EGF, KGF, HGF) and exosomes activate EGFR/ERK and PI3K/AKT signaling in keratinocytes, accelerating their migration and proliferation to restore the epidermal barrier [59].

Additional pathways modulated by MSC-exosomes include Nrf2 (to mitigate oxidative stress), JAK/STAT6 (to promote macrophage polarization), and SIRT3/SOD2 (to detoxify ROS). These coordinated cascades collectively restore the disrupted healing process in DFUs by dampening chronic inflammation, boosting angiogenesis under hypoxia, enhancing cellular survival/proliferation/migration, and optimizing extracellular matrix remodeling [60, 65].

Hydrogel-based delivery systems further amplify these signaling effects by protecting MSC viability and enabling sustained, localized release of exosomes and secretome components at the wound site.

Despite these advantages, the clinical application of MSCs remains limited by several delivery-related challenges. A major limitation is the poor survival and retention of MSCs after transplantation. Chronic wound environments are characterized by hypoxia, oxidative stress, and persistent inflammation, all of which reduce MSC viability and impair their therapeutic function. Furthermore, when MSCs are administered as free cell suspensions either locally or systemically, they are rapidly cleared from the target site, resulting in insufficient engraftment and reduced therapeutic efficacy [53]. To overcome these limitations, the use of biomaterial-based delivery systems has gained increasing attention. Hydrogels, scaffolds, and decellularized extracellular matrices (dECM) can provide a protective and bioactive niche that mimics the native tissue microenvironment. Among these, hydrogels are particularly advantageous due to their high water content, tunable mechanical properties, and ability to encapsulate cells and biologics. By embedding MSCs within hydrogels, it is possible to shield them from mechanical stress and immune attack, improve their retention and viability, and facilitate a sustained release of paracrine factors critical for regeneration.

Recent advances in hydrogel design include oxygen-releasing systems that alleviate hypoxia and improve MSC survival in ischemic wounds. Thermosensitive and pH-responsive hydrogels have also been developed to allow controlled delivery of cells and therapeutic molecules in response to physiological cues. Furthermore, hydrogels can be combined with MSC preconditioning strategies, such as hypoxic exposure, cytokine priming, or genetic modification, to enhance stress resistance, immunomodulatory capacity, and paracrine activity. These approaches collectively improve the therapeutic efficacy of MSCs in chronic and complex wound environments [46, 66].

In summary, MSCs play a central role in wound healing by regulating immune responses, promoting angiogenesis, and supporting tissue remodeling. However, their clinical success depends heavily on the development of advanced delivery platforms. Hydrogel-based systems provide a promising solution by enhancing MSC survival, retention, and functionality within the wound microenvironment.

Injectable hydrogels with inherent therapeutic functions as advanced delivery platforms for wound healing

An ideal wound dressing should create a favorable microenvironment by maintaining appropriate temperature, moisture, and pH, thereby promoting granulation tissue formation and re-epithelialization [67]. In addition, it should provide antimicrobial and anti-inflammatory properties to reduce the risk of infection during the healing process [68]. Because these materials are applied directly to exposed subcutaneous tissue, they must demonstrate high biocompatibility, biosafety, and minimal toxicity. Hydrogels with inherent therapeutic functions—such as antioxidant, pro-angiogenic, and antimicrobial properties—have gained increasing attention in wound care. Their unique physicochemical characteristics make them highly suitable as wound dressings for tissue regeneration and repair [11, 69]. These materials can maintain a moist environment, soften necrotic tissue, and allow atraumatic removal due to their non-adhesive nature, thereby improving patient comfort and clinical outcomes [48].

Beyond their intrinsic therapeutic effects, injectable hydrogels can function as multifunctional delivery platforms. Their hydrated polymer networks enable the encapsulation and sustained release of therapeutic agents, including drugs, growth factors, antimicrobial peptides, and stem cell–derived exosomes, directly at the wound site. This localized delivery enhances therapeutic efficacy while minimizing systemic side effects and reducing the frequency of administration. Another key advantage of hydrogels is their morphological adaptability, which allows them to conform to irregular wound geometries. Recently, carrier-free hydrogels have attracted considerable interest because they reduce issues such as batch-to-batch variability and potential toxicity associated with additional carrier materials [40, 41]. These hydrogels are typically fabricated from natural biopolymers (e.g., polysaccharides), synthetic polymers such as polyvinyl alcohol (PVA), or self-assembled biomaterials that closely mimic the ECM [43–45]. Importantly, many hydrogel systems are designed to be stimuli-responsive, enabling controlled release of therapeutic agents in response to environmental cues such as pH, temperature, or enzymatic activity. This allows the delivery profile to be aligned with the dynamic phases of wound healing.

In the following sections, recent advances in carrier-free hydrogels with intrinsic therapeutic properties are discussed, with particular emphasis on their role as injectable delivery systems for enhancing wound healing.

Formation mechanisms of injectable hydrogels

Injectable hydrogels for DFU therapy are formulated as low-viscosity precursor solutions that can be easily administered via syringe and rapidly gel in situ upon injection into the wound bed. This sol-to-gel transition enables minimally invasive delivery, precise filling of irregular and deep wound cavities, and effective encapsulation of MSCs or their exosomes while minimizing leakage and shear-induced cell damage. Gelation is typically triggered by physiological stimuli such as temperature, pH, ions, enzymes, or light, and is achieved through physical, chemical, or hybrid crosslinking strategies [1, 2, 70].

Physical crosslinking relies on reversible, non-covalent interactions that confer shear-thinning behavior (facilitating injection) and rapid self-recovery. Key mechanisms include:

a.Temperature-induced gelation, commonly observed in Pluronic F-127 (PF-127) and certain modified chitosan derivatives, driven by hydrophobic associations and micelle formation at body temperature (~ 37 °C) [3].

b.Ionic crosslinking, mediated by multivalent cations (e.g., Ca²⁺ or Zn²⁺) that bridge negatively or positively charged polymer chains.

c. Supramolecular interactions, including hydrogen bonding, π–π stacking, and host–guest assembly. These are widely used in self-assembled small-molecule and peptide-based hydrogels, providing excellent injectability, self-healing, and reversibility [4].

Chemical crosslinking, on the other hand, forms stable covalent bonds that enhance mechanical strength and long-term structural integrity. Prominent strategies include:

d.Schiff base (imine) formation between aldehyde and amine groups, frequently employed in oxidized polysaccharide–chitosan and catechol-modified systems [2, 71].

e.Enzyme-mediated crosslinking (e.g., horseradish peroxidase-catalyzed reactions in gelatin-hydroxyphenyl hydrogels), enabling rapid in situ solidification within minutes [58].

g.Photo-crosslinking of methacrylated polymers such as gelatin methacrylate (GelMA), initiated by UV or visible light after injection.

h.Michael-type addition and dynamic covalent chemistries (e.g., boronic ester bonds), which combine mechanical stability with self-healing properties [72].

Modern systems often employ hybrid physical–chemical crosslinking to achieve an optimal balance between rapid gelation (ideally within 5–10 min), injectability, mechanical robustness, and responsiveness to the dynamic diabetic wound microenvironment (changes in pH, glucose, ROS, or enzymatic activity) [73, 74]. Crucially, the selected gelation mechanism must be highly cytocompatible to preserve the viability and therapeutic functionality of encapsulated MSCs or exosomes by avoiding toxic crosslinkers, extreme pH shifts, or excessive shear stress during injection [58, 75].

In the context of DFU therapy, these formation mechanisms play a pivotal role in therapeutic success. Rapid in situ gelation shields MSCs and exosomes from the hostile wound milieu characterized by hypoxia, excessive oxidative stress, proteases, and chronic inflammation while enhancing cell retention and enabling sustained, localized release of paracrine factors. Furthermore, stimuli-responsive designs allow the hydrogel network to dynamically adapt its degradation rate and cargo release profile to the successive phases of diabetic wound healing [74, 76].

The following sections review specific classes of carrier-free injectable hydrogels (polysaccharide-based, synthetic polymer-based, self-assembled, and ECM-mimicking) and highlight how their distinct gelation mechanisms enhance MSC and exosome delivery in diabetic wound regeneration.

Comparative analysis of natural polysaccharide-based, synthetic polymer-based, and self-assembled carrier-free injectable hydrogels for DFU therapy

Injectable hydrogels for DFUs management can be broadly categorized into three classes according to their primary building blocks: natural polysaccharide-based, synthetic polymer-based, and self-assembled small-molecule or peptide-based systems. Each class presents distinct advantages and limitations with respect to biocompatibility, mechanical properties, intrinsic bioactivity, injectability, degradation kinetics, manufacturing reproducibility, and performance as delivery vehicles for MSCs or MSC-derived exosomes (Table 1).

Table 1.

Comparative overview of injectable hydrogel classes for DFU therapy and MSC/Exosome delivery

Hydrogel class	Key advantages	Main limitations	Best suited for in DFUs	Representative examples
Natural Polysaccharide-Based	High biocompatibility, biodegradability, intrinsic bioactivity (antimicrobial, immunomodulatory), ECM mimicry [75, 77]	Batch-to-batch variability, weak mechanical strength, rapid enzymatic degradation [75]	MSC/exosome support, moist healing, low immunogenicity	Chitosan derivatives, HA, alginate, fucoidan
Synthetic Polymer-Based,	Excellent mechanical strength, tunable degradation & release, high reproducibility[78]	Limited inherent bioactivity, potential cytotoxicity, poorer cell adhesion [78]	Mechanical stability, controlled pharmacokinetics	PVA, PCL, PEG-based systems
Self-Assembled Carrier-Free	Superior injectability, self-healing, intrinsic therapeutic molecules, low carrier toxicity [79]	Lower mechanical strength, production scalability issues, environmental sensitivity [79]	Smart/minimally invasive delivery, carrier-free systems	Baicalin-sanguinarine, glycyrrhizic acid, peptide hydrogels

Natural polysaccharide-based hydrogels (e.g., chitosan and its derivatives, hyaluronic acid (HA), alginate, cellulose, pectin, fucoidan) offer excellent biocompatibility, biodegradability, and intrinsic bioactivity. They closely mimic the native ECM, exhibit high water-absorption capacity, maintain optimal moisture balance, and often possess inherent antimicrobial, antioxidant, or immunomodulatory properties. These characteristics make them particularly advantageous for supporting MSC viability, enhancing exosome retention and sustained release, and promoting re-epithelialization and granulation tissue formation in the hostile DFU microenvironment [75, 77]. However, they often suffer from batch-to-batch variability, relatively weak and inconsistent mechanical strength, rapid enzymatic degradation in protease-rich chronic wounds, and limited precision in controlling gelation kinetics or long-term release profiles [75].

Synthetic polymer-based hydrogels (e.g., PVA, polycaprolactone (PCL), and polyethylene glycol (PEG) provide superior mechanical robustness, highly tunable degradation rates, excellent reproducibility, and precise control over porosity, swelling behavior, and crosslinking density. These features confer better structural stability in high-exudate or mechanically stressed DFUs and enable more predictable sustained release kinetics of bioactive cargos such as growth factors or exosomes [78]. Their primary limitations include limited inherent bioactivity, potential cytotoxicity from residual crosslinkers or degradation byproducts, poorer native cell adhesion (often requiring surface modification), and slower or incomplete biodegradability, which may provoke chronic foreign-body responses [78].

Self-assembled carrier-free hydrogels (formed through non-covalent interactions such as hydrogen bonding, π–π stacking, hydrophobic associations, or host–guest chemistry, using phytochemicals or peptide amphiphiles) combine outstanding injectability, self-healing and shear-thinning properties, and minimal use of additional carriers, thereby reducing toxicity risks and batch variability. Many incorporate intrinsic therapeutic small molecules with anti-inflammatory or antibacterial effects and can dynamically adapt to wound mechanics [79]. Their drawbacks typically include lower mechanical strength, challenges in large-scale production and standardization, and greater sensitivity to environmental factors, which can affect long-term stability in the variable DFU milieu [79].

In summary, natural hydrogels are often preferred when biological integration, low immunogenicity, and inherent regenerative cues are prioritized, making them highly suitable for MSC and exosome delivery in regenerative applications. Synthetic hydrogels excel in applications demanding mechanical durability and controlled pharmacokinetics. Self-assembled systems offer a promising “middle ground” for smart, minimally invasive, carrier-free delivery but generally require further optimization for clinical robustness. Hybrid strategies that integrate advantages from multiple classes are increasingly pursued to overcome individual limitations and create multifunctional platforms tailored to the complex pathophysiology of DFUs [80].

Chitosan and its derivatives for wound care

Chitosan is a cationic polysaccharide composed of β-(1→4)-linked N-acetyl-D-glucosamine units and is widely used in wound dressing applications due to its biocompatibility, biodegradability, and intrinsic antimicrobial properties [48]. Its strong water-absorption capacity and structural versatility enable the formation of hydrogels with favorable physicochemical properties for wound healing. The antimicrobial activity of chitosan is primarily attributed to electrostatic interactions between its positively charged amino groups and negatively charged bacterial cell membranes. This interaction disrupts membrane integrity and can also interfere with bacterial DNA, thereby inhibiting transcription and protein synthesis [48]. In addition, chitosan-based hydrogels have been shown to promote fibroblast proliferation, enhance ECM deposition, and reduce scar formation, thereby improving tissue regeneration [81]. The chemical structure of chitosan enables extensive functional modification via its hydroxyl and amino groups. This enables the development of various derivatives, including N, O-carboxymethyl chitosan, N-succinyl chitosan, O-succinyl chitosan, and N-acyl chitosan, which exhibit enhanced solubility, mechanical properties, and biological performance. From a drug-delivery perspective, chitosan hydrogels are particularly attractive due to their positive charge, which facilitates electrostatic interactions with negatively charged biomolecules, such as growth factors, nucleic acids, and proteins. This property enables efficient encapsulation and controlled release of therapeutic agents. For example, Lu et al. developed a bioadhesive hydrogel composed of carboxymethyl chitosan and allyl cellulose, which demonstrated excellent biocompatibility, strong adhesion, and antimicrobial activity [82]. Similarly, GelMA /sulfonated chitosan (GelMA/CS-SH) hydrogel exhibited high porosity, moisture retention, and mechanical stability, leading to enhanced wound closure and skin appendage regeneration, along with increased expression of vascular endothelial growth factor receptor (VEGFR) [47]. Some chitosan-based hydrogels have progressed to clinical evaluation. For instance, the MACF hydrogel sheet, composed of fluorinated methacrylamide chitosan, demonstrated improved healing efficiency and cost-effectiveness compared to conventional treatments [83]. These systems also offer opportunities to incorporate bioactive agents, enabling sustained, localized therapeutic delivery during the wound-healing process.

Other polysaccharide-based carrier-free hydrogels

In addition to chitosan, a wide range of natural polysaccharides, including alginate, HA, cellulose, pectin, starch, galactomannan, gellan gum, and fucoidan, have been extensively explored for hydrogel-based wound dressings. These materials are valued for their biocompatibility, biodegradability, low immunogenicity, and ease of processing [46, 84]. These biopolymers were among the earliest materials used for wound dressings [85], and hydrogels derived from them often replicate ECM-like characteristics that promote cell migration, proliferation, and tissue repair [86]. Furthermore, many of these matrices can be engineered for drug conjugation, encapsulation, and sustained release, enabling a synergistic combination of structural and therapeutic functions.

Cellulose, one of the most abundant natural biopolymers, consists of repeating cellobiose units linked by β-(1→4)-glycosidic bonds [87]. It supports wound healing by releasing growth factors such as bFGF, PGF, and EGF, which stimulate dermal fibroblast proliferation and inhibit bacterial growth [88]. Deng et al. prepared a cellulose–cucurbit gum composite hydrogel with porosity, thermal stability, and biocompatibility, which demonstrated effective hemostatic and healing properties in a mouse wound model [89]. Other cellulose-based systems, including cellulose–DOPA hydrogels and carboxymethyl cellulose–honey composites, have shown improved antibacterial activity and moisture retention, while also enabling sustained release of therapeutic agents [90, 91].

Pectin-based hydrogels, such as pectin–honey composites and polyphosphate–pectin conjugates, have demonstrated accelerated wound healing, improved moisture retention, and reduced bleeding time [92, 93]. These systems can also serve as carriers for antioxidants and anti-inflammatory compounds, contributing to enhanced tissue regeneration. Starch-derived hydrogels, particularly injectable adhesive systems such as starch–DOPA hydrogels, exhibit strong hemostatic properties, appropriate swelling behavior, and biodegradability. These features make them suitable for irregular wound geometries and enable controlled release of therapeutic molecules, including peptides and hemostatic agents [94].

Galactomannan-based hydrogels, derived from sources such as Cassia grandis, have been shown to maintain moisture and promote wound healing in preclinical models [95]. Similarly, gellan gum-based hydrogels, including methacrylated gellan gum systems and near-infrared (NIR)-responsive composites, exhibit injectability, antimicrobial activity, and controlled drug delivery.

Fucoidan-containing hydrogels, derived from brown algae, possess strong anti-inflammatory, antioxidant, and immunomodulatory properties. When incorporated into hydrogel matrices, fucoidan can promote macrophage polarization toward the M2 phenotype, enhance angiogenesis, and improve granulation tissue formation. In addition, these hydrogels can function as sustained-release platforms for growth factors and immune-modulating agents, further supporting wound healing.

Overall, polysaccharide-based carrier-free hydrogels provide a versatile platform that combines structural support with therapeutic functionality. Their ability to mimic ECM properties, deliver bioactive molecules, and respond to environmental stimuli makes them highly promising for advanced wound care applications.

Synthetic polymer-based carrier-free hydrogels

Synthetic polymers, such as PVA and PCL, are widely used to fabricate carrier-free hydrogels for wound-healing applications due to their tunable mechanical properties, structural stability, and reproducibility. These materials offer precise control over hydrogel architecture, including porosity, degradation rate, and mechanical strength, which are critical parameters for optimizing wound healing outcomes. However, unlike natural polymers, synthetic materials often lack intrinsic bioactivity and may require chemical modification to improve biocompatibility and cellular interactions. For example, Xie et al. developed a composite hydrogel composed of hydroxyapatite (HA), collagen, and PVA, in which collagen enhanced the overall biocompatibility of the system [96].

Similarly, Zhang et al. designed an (Aloe vera polysaccharide)AP/Honey@PVA hydrogel incorporating aloe vera polysaccharide and honey, which exhibited strong moisture retention and broad-spectrum antimicrobial activity against pathogens such as Staphylococcus aureus, Escherichia coli, and Candida albicans. This formulation significantly accelerated the healing of infected wounds [97].

Self-assembled carrier-free hydrogels

Although conventional polymer-based hydrogels, such as those derived from chitosan, cellulose, PVA, and PCL, exhibit beneficial properties, their in vivo therapeutic performance can be limited. As a result, increasing attention has been directed toward self-assembled small-molecule hydrogels, which offer enhanced functionality and tunable bioactivity [98]. Self-assembled hydrogels are formed through non-covalent interactions, including electrostatic forces, π–π stacking, and hydrogen bonding. These interactions enable the spontaneous organization of small molecules into three-dimensional networks, yielding injectable, structurally adaptable materials. For example, the BA-SAN hydrogel, composed of the phytochemicals baicalin (BA) and sanguinarine (SAN), self-assembles through multiple intermolecular interactions to form a malleable hydrogel. This system has been shown to modulate inflammatory mediators, inhibit bacterial virulence, and promote wound healing [40, 99]. Similarly, the PCEC-QAS hydrogel incorporates quaternary ammonium salt-functionalized nanoparticles with strong antimicrobial activity. This system is bioabsorbable and cytocompatible, and it forms a stable gel network upon thermal induction, making it suitable for treating methicillin-resistant Staphylococcus aureus (MRSA) infections [49]. Another example is the GN-Bn hydrogel, derived from glycyrrhizic acid, which demonstrated significant antimicrobial effects in MRSA-infected wound models. Its mechanism involves modulation of arginine metabolism and suppression of bacterial growth, highlighting its therapeutic potential [100].

From a delivery perspective, synthetic polymer hydrogels offer mechanical strength and tunable degradation rates that can be matched to the drug release profile.

They have been engineered for sustained release of antibiotics, anti-inflammatory drugs, and growth factors, as well as for responsive delivery triggered by environmental stimuli such as pH or temperature. For example, PVA-based injectable hydrogels have been combined with silver nanoparticles for prolonged antimicrobial delivery [101].

In drug delivery, self-assembled hydrogels offer the advantage of incorporating hydrophobic drugs into their nanofibrillar networks, enabling sustained release while maintaining local drug concentrations above therapeutic thresholds. Peptide-based self-assembled hydrogels have also been designed to release growth factors in a controlled fashion to accelerate chronic wound healing.

ECM-mimicking hydrogels for wound repair

Hydrogels, composed primarily of water and polymeric networks, can structurally and functionally resemble the native ECM, which plays a critical role in cell adhesion, intercellular communication, and cell recruitment during wound healing [29, 102, 103]. This ECM-mimicking capability allows hydrogels to regulate cellular behavior and promote tissue regeneration. For example, the GM-P@HA-P hydrogel, a glycopeptide–HA hybrid system, closely mimics ECM architecture while exhibiting strong antibacterial properties. This hydrogel has been shown to promote macrophage polarization toward the M2 phenotype and accelerate wound healing in both diabetic and burn wound models [14].

In addition, ECM-mimicking strategies have been developed using fucoidan, a sulfated polysaccharide derived from brown algae. When incorporated into collagen-based hydrogels, fucoidan enhances macrophage polarization toward the M2 phenotype, promotes angiogenesis, and accelerates wound closure in both diabetic and infected wounds. These systems also exhibit anti-inflammatory and immunomodulatory effects, contributing to improved tissue regeneration. Hydrogels with hierarchical fiber architectures further enhance ECM mimicry by reproducing the structural complexity and dynamic responsiveness of native tissues. These materials can respond to environmental cues and regulate multiple cellular processes, including proliferation, migration, and differentiation [104].

In drug delivery, ECM-mimicking hydrogels can act as bioinspired depots for controlled release of signaling molecules, stem cell secretomes, or extracellular vesicles, thereby guiding cell behavior during tissue repair. These systems can be functionalized with affinity ligands to selectively bind and release growth factors in response to changes in the wound microenvironment. Figure 1 presents a summary of these hydrogels.

Fig. 1.

Injectable carrier-free hydrogels and their therapeutic functions as delivery platforms for wound healing. the figure illustrates three primary categories of these hydrogels: polysaccharide-based, synthetic polymer-based, and self-assembled-based

Hydrogel-based exosome delivery systems

Although the primary emphasis of this review is the delivery of MSCs using injectable hydrogels, MSC-derived exosomes (MSC-exos) have emerged as a valuable cell-free adjunct that can largely recapitulate the paracrine therapeutic effects of MSCs [57].

MSC-exos are nanosized extracellular vesicles enriched with bioactive cargos, including microRNAs, proteins, cytokines, and lipids. These vesicles modulate key processes in diabetic wound healing, such as inflammation resolution, angiogenesis stimulation, fibroblast activation, and re-epithelialization, while offering distinct advantages, including lower immunogenicity, superior stability, and simplified storage compared to live cell therapy [105].

Recognizing the challenges of poor MSC survival and retention in the hostile DFU microenvironment, several studies have explored the integration of MSC-exos into injectable hydrogel platforms. Hydrogels serve as protective carriers that shield exosomes from proteolytic degradation and oxidative stress, while enabling controlled and sustained release directly at the wound site [13, 106].

Preclinical evidence supports the efficacy of this combined strategy. For example, GelMA hydrogels loaded with hypoxia-pretreated Adipose-derived stem cells (ADSC)-derived exosomes significantly enhanced angiogenesis and accelerated diabetic wound closure [106]. Similarly, thermosensitive Pluronic F-127 hydrogels incorporating hUC-MSC exosomes improved granulation tissue formation, collagen deposition, and skin appendage regeneration [98]. Other hydrogel systems, including chitosan- and ECM-based platforms, have also shown synergistic benefits when used for MSC-exosome delivery [13, 107].

In summary, hydrogel-mediated delivery of MSC-derived exosomes represents a promising complementary or alternative approach to whole-cell MSC therapy. This strategy leverages the structural and protective properties of injectable hydrogels to maximize the regenerative potential of MSC secretome in the treatment of diabetic foot ulcers.

Role of hydrogels in reversing pathological events and their biological properties

Hydrogels are highly versatile biomaterials capable of modulating pathological processes in chronic diabetic wounds and restoring the normal healing cascade. Their intrinsic ECM-like structure, tunable mechanical and chemical properties, and ability to be functionalized with bioactive molecules enable them to simultaneously target multiple biological pathways. By regulating the wound microenvironment, promoting cellular function, and delivering therapeutic agents in a sustained, localized manner, hydrogels address the multifactorial impairments in diabetic wound healing.

Controlling the hyperglycemic microenvironment

In diabetic wounds, persistent hyperglycemia disrupts normal cellular signaling, promotes oxidative stress, and delays tissue repair. Glucose-responsive hydrogels incorporate sensing components such as glucose oxidase (GOx), concanavalin A (Con A), and phenylboronic acid (PBA), enabling real-time detection of glucose levels and controlled insulin release. This localized regulation reduces glucose-induced tissue damage, improves fibroblast and keratinocyte migration, and supports granulation tissue formation in diabetic foot ulcers [83].

Controlling excessive oxidative stress

Excessive ROS accumulation in diabetic wounds leads to oxidative damage to DNA, lipids, and proteins, thereby sustaining inflammation and impairing angiogenesis. Antioxidant hydrogels incorporate enzymatic antioxidants such as superoxide dismutase, catalase, and glutathione peroxidase, as well as non-enzymatic antioxidants derived from natural sources (e.g., curcumin, polyphenols, DNA) or synthetic compounds (e.g., CORM-401, glycyrrhizic acid). These systems neutralize ROS and restore redox balance. Self-antioxidant hydrogels composed of antioxidant macromolecules, such as poly(ethylene glycol-co-citric acid-co-N-isopropylacrylamide), provide intrinsic antioxidant activity without requiring additional agents. Hybrid systems incorporating metal ions such as Zn²⁺ further enhance mechanical stability, promote cell proliferation, and reduce cytotoxicity. These antioxidant effects support endothelial cell survival, capillary formation, and collagen synthesis in the hypoxic diabetic wound environment [49].

Facilitating nerve regeneration

Diabetic neuropathy significantly impairs wound healing by disrupting sensory signaling and neurogenic inflammation. ECM-based conductive hydrogels, composed of materials such as gelatin methacryloyl and oxidized chondroitin sulfate combined with conductive nanoparticles, provide structural support and electrical conductivity to guide nerve regeneration. These hydrogels activate intracellular Ca²⁺ signaling and stimulate pathways such as PI3K/AKT and MEK/ERK, promoting axonal growth, reinnervation, and angiogenesis. They also enhance Schwann cell proliferation and myelination, which are essential for restoring nerve function. Improved neurovascular integration enhances tissue oxygenation and accelerates wound closure [108].

Facilitating angiogenesis

Impaired angiogenesis is a major feature of diabetic wounds, resulting in insufficient oxygen and nutrient supply. Hydrogels can function as delivery platforms for angiogenic growth factors (e.g., VEGF, FGF), chemokines (e.g., SDF-1), small molecules (e.g., desferrioxamine), and therapeutic cells (e.g., MSCs, ASCs). These systems protect bioactive molecules from degradation and enable sustained release. By enhancing HIF-1α signaling and promoting EPC mobilization, hydrogels support the formation and maturation of new blood vessels. Various strategies, including growth factor-loaded hydrogels, temperature-responsive systems, and cell-laden scaffolds, contribute to stable vascularization and improved wound healing outcomes [84].

Modulating inflammation and infection control

Persistent infection is a major barrier to wound healing in diabetic patients. Antibacterial hydrogels enable localized and sustained delivery of antimicrobial agents, including antibiotics (e.g., vancomycin, ciprofloxacin), metallic nanoparticles (e.g., silver, gold), natural compounds (e.g., honey, curcumin), and carbon-based nanomaterials (e.g., graphene oxide). Some hydrogels also possess intrinsic antimicrobial activity due to components such as chitosan or antimicrobial peptides. These systems reduce bacterial load, limit biofilm formation, and create a favorable environment for tissue regeneration [85].

Controlling inflammation

Chronic inflammation in diabetic wounds is characterized by prolonged activation of pro-inflammatory macrophages and excessive ECM degradation. Anti-inflammatory hydrogels modulate this response by promoting macrophage polarization toward the M2 phenotype, inhibiting MMP-9 activity, and delivering therapeutic agents such as siRNA, small-molecule inhibitors, or natural polyphenols. Advanced systems, including polydopamine-modified graphene oxide and tannic acid–iron complexes, suppress pro-inflammatory cytokine production while enhancing angiogenic signaling. ECM-mimicking glycopeptide hydrogels provide combined antibacterial, angiogenic, and immunomodulatory effects, thereby supporting tissue regeneration without the need for additional drugs [48, 109].

Biological properties of hydrogels in tissue engineering and regenerative medicine

Hydrogels possess a unique combination of properties, including ECM-like architecture, tunable mechanical characteristics, high biocompatibility, and responsiveness to external stimuli. These features make them highly suitable scaffolds for tissue engineering and regenerative medicine. Hydrogels provide structural support, regulate cellular behavior, and facilitate integration with host tissue by adapting to dynamic biological environments.

Cell adhesion and migration

Hydrogels play a critical role in promoting cell adhesion and migration by providing a biomimetic microenvironment that resembles the native ECM. These processes are essential for the formation of functional and well-integrated tissue constructs. Cell adhesion represents the initial attachment of cells to the hydrogel surface and is regulated by surface chemistry, ligand presentation, and topographical features [110]. Hydrogels inherently support cell adhesion through their functional groups and can further enhance this process when loaded with growth factors that activate intracellular signaling pathways. In addition, adhesion can be optimized through surface modification, biofunctionalization, and the incorporation of micro- and nanoscale structural features that regulate cell–material interactions.

Beyond adhesion, hydrogels facilitate cell migration within their three-dimensional network, enabling uniform cell distribution and tissue formation. This process is mediated by coordinated cell–cell and cell–matrix interactions and contributes to endothelialization and vascular network formation within engineered constructs [111]. Migration behavior can be further tuned by controlling ligand spacing, such as RGD nanospacing, which has been shown to enhance migration under moderate adhesion conditions [112].

Cell proliferation and differentiation

Hydrogels contain a wide range of functional groups and active sites that support cell adhesion and initiate biological processes such as proliferation, differentiation, and migration. These processes are essential for tissue repair and regeneration. By providing a biomimetic and supportive microenvironment, hydrogels create favorable conditions for the formation of new tissue, enabling coordinated proliferation and differentiation of resident or transplanted cells. The resulting engineered tissue can achieve functional and structural integration with host systems [113]. Cell proliferation, defined as the expansion of cell number through division, is a key requirement in tissue engineering. This process is influenced by hydrogel properties such as mechanical strength, surface chemistry, pore size, and porosity, which regulate nutrient diffusion, waste removal, and cell–material interactions [114]. Equally important is the ability of hydrogels to direct cell differentiation, ensuring the formation of tissues with appropriate structure and function. This can be achieved by modifying hydrogel composition, incorporating bioactive molecules such as growth factors, and engineering the physical architecture of the scaffold. These strategies allow precise control over cell fate and contribute to the development of functional tissue constructs [115].

Cell viability

Maintaining cell viability, particularly in thick or metabolically active constructs, is essential for successful tissue regeneration. Hydrogels with interconnected pore structures and controlled degradation profiles support the diffusion of oxygen and nutrients while facilitating the removal of metabolic waste. Stimuli-responsive hydrogels can further enhance cell survival by protecting embedded cells from oxidative and inflammatory stress. For example, hydrogels capable of delivering antioxidant molecules can reduce cellular damage and improve long-term viability [116, 117].

Angiogenesis and vascularization

Hydrogels have significant potential to promote angiogenesis and vascularization owing to their multifunctional properties and ECM-mimicking architecture. In tissue engineering, both angiogenesis and vascularization are critical for the development of viable, functional engineered tissues and organs, as they ensure adequate oxygenation, nutrient delivery, and waste removal. The intrinsic porosity of hydrogels facilitates metabolic exchange by allowing the diffusion of oxygen and nutrients while removing cellular waste products, thereby creating a microenvironment conducive to new vessel formation [118] .

Angiogenesis is particularly vital for the survival, integration, and performance of engineered constructs, especially under hypoxic conditions commonly encountered in early stages of tissue implantation. This process is driven by essential bioactive molecules such as VEGF and other angiogenic factors that stimulate endothelial cell proliferation, migration, and tubule formation. Successful vascularization ensures a continuous supply of oxygen and nutrients, both of which are indispensable for tissue regeneration and repair. Despite its importance, achieving stable vascularization remains a major challenge in tissue engineering. Hydrogels address this limitation by serving as sustained-release platforms for pro-angiogenic biomolecules, protecting them from rapid degradation and maintaining their bioactivity over extended periods. This controlled delivery approach can be tailored to match the temporal needs of the regenerating tissue, significantly enhancing neovascularization and overall construct viability [119].

Antibacterial and antiviral activities

Antibacterial and antiviral properties represent critical functional features of hydrogels, particularly in tissue engineering applications where preventing infection is essential for successful regeneration. Some hydrogels exhibit intrinsic antimicrobial activity due to their functional groups and structural characteristics. In addition, hydrogels can be engineered to incorporate antibacterial and antiviral agents, enabling sustained and localized release that provides long-term protection against microbial contamination [104].

Ensuring the safety and sterility of repaired or regenerated tissues is essential, as infections can significantly compromise healing outcomes and may lead to graft failure. The risk of infection can be reduced by embedding antimicrobial and antiviral agents directly within the hydrogel matrix. These agents include antimicrobial peptides, quaternary ammonium compounds, silver nanoparticles, and other bioactive molecules with broad-spectrum activity. Furthermore, hydrogel-based scaffolds or implant surfaces can be coated with antibacterial or antifungal compounds to prevent microbial adhesion and biofilm formation. This creates a protective interface that supports tissue integration while minimizing the risk of infection [114].

The multiplicative benefits of MSC–hydrogel systems

The integration of MSCs with hydrogel-based delivery systems has emerged as a promising strategy in regenerative medicine and wound healing, particularly in the treatment of chronic wounds such as diabetic ulcers. The synergistic interaction between MSCs and hydrogels provides three main multiplicative benefits. First, hydrogels enhance cell retention and protect MSCs from the harsh wound microenvironment. They act as three-dimensional biomimetic matrices that shield cells from inflammatory mediators, proteases, and oxidative stress, thereby improving cell survival at the injury site. Second, these systems enable the controlled release of growth factors, cytokines, and exosomes secreted by MSCs. This sustained-release formulation maintains therapeutic activity over extended periods and enables precise spatial and temporal delivery of regenerative signals. Third, hydrogels provide mechanical support together with bioactive cues, mimicking the native ECM both structurally and chemically. This promotes cell adhesion, proliferation, migration, angiogenesis, and tissue remodeling. The following sections describe how different MSC sources, when combined with hydrogels, leverage these advantages to improve wound-healing outcomes.

Bone marrow-derived mesenchymal stem cells (BMSCs)

Bone marrow stroma is a primary source of MSCs. Bone marrow-derived MSCs produce higher levels of collagen, fibroblast growth factor, and vascular endothelial growth factor compared to dermal fibroblasts. They also demonstrate enhanced effects on granulation tissue formation, epithelialization, and angiogenesis in vivo, supporting their application in wound healing [120]. However, no commercial therapies based on bone marrow-derived MSCs have yet been developed for diabetic foot ulcers.

Ahmed et al. investigated a NO releasing chitosan/PVA hydrogel to enhance the therapeutic efficacy of BMSCs in diabetic wound healing. This hydrogel promoted re-epithelialization, improved vascularization, and increased collagen deposition compared to controls. In addition, the combined treatment enhanced the expression of growth and cytoregulatory factors within 16 days, indicating a promising strategy for chronic wound management [121].

Ravari et al. conducted a case series of patients with chronic wounds resistant to conventional treatments. BMSCs were combined with platelets, fibrin glue, and a bone marrow-impregnated collagen matrix. The results showed significant improvement in wound healing, reduction in wound size, and complete closure in 3 out of 8 patients within four weeks [122]. Yao et al. developed a gelatin-hydroxyphenyl hydrogel that was cross-linked in situ within five minutes using a dual-enzyme system. When loaded with BMSCs, this hydrogel accelerated re-epithelialization, enhanced granulation tissue formation, and increased collagen deposition compared to control groups [123].

Martin et al. designed a hydrolytic hydrogel containing degradable ester linkages, which enabled controlled degradation and prolonged MSC retention in the wound environment. This approach improved localized immune modulation and enhanced wound healing in diabetic mice [124]. In another study, a PEGDA hydrogel co-encapsulating human BM-MSCs and rat insulin-secreting cells improved cell viability and function, resulting in wound repair rates nearly three times faster than controls. However, limitations such as invasiveness, low engraftment efficiency, and extended in vitro expansion remain challenges for clinical application [125].

Adipose-derived stem cells (ADSCs)

ADSCs have been identified as an important source of MSCs. They promote wound healing through the secretion of growth factors such as vascular endothelial growth factor, hepatocyte growth factor, transforming growth factor-β1, insulin-like growth factor-1, epidermal growth factor, and keratinocyte growth factor, which support angiogenesis, collagen synthesis, and epithelialization [126–129]. ADSCs also exhibit anti-inflammatory effects by inhibiting T-cell activation and modulating immune signaling pathways, leading to reduced levels of inflammatory cytokines such as tumor necrosis factor-α and interferon-γ [126, 130].

Da Silva et al. designed ADSC-laden HA spongy hydrogels to enhance diabetic wound healing. In vitro studies demonstrated that the hydrogel effectively retained MSCs and facilitated cellular interactions, with findings indicating high MSC viability and proliferation over 2 weeks, particularly when hASCs were conditioned in neurogenic media. Four weeks post-transplantation of hydrogel into diabetic mouse wounds, in vivo studies demonstrated that MSC-loaded HA hydrogels exhibited enhanced collagen deposition, a more structured extracellular matrix, and improved angiogenesis compared to the control group. Moreover, the hydrogel sheet facilitated the polarization of M1 macrophages toward the M2 phenotype and enhanced effective neoinnervation, establishing this hydrogel as an effective microenvironment for accelerated diabetic wound healing [131].

Xia et al. designed a curcumin-incorporated gelatin methacryloyl hydrogel using 3D bioprinting to deliver ADSCs. This system reduced reactive oxygen species levels and ADSC apoptosis, leading to increased collagen deposition and enhanced angiogenesis in diabetic wounds [132].

Eke et al. developed a hydrogel composed of HA, gelatin, and PEGDA that could be cross-linked in situ under ultraviolet irradiation. ADSCs within this hydrogel promoted endothelial cell migration and vascular formation, resulting in a threefold increase in angiogenesis compared to cell-free hydrogels [133].

Chen et al. examined the efficacy of delivering MSCs via a hydrogel derived from a human decellularized adipose matrix (DAM) to enhance the healing of chronic wounds. The DAM-based hydrogel was synthesized via pepsin digestion and pH neutralization to create a supportive microenvironment that enhances MSC stability, viability, and functionality at the wound site. The study used in vitro and in vivo analyses to demonstrate that the DAM hydrogel markedly enhanced MSC adhesion, ADSC paracrine activity, and the secretion of hepatocyte growth factors, which are essential for neovascularization. In a diabetic mouse wound model, the application of DAM hydrogel-embedded MSCs resulted in enhanced collagen deposition, revascularization, and cellular infiltration, which facilitated quicker wound closure and improved tissue regeneration relative to the untreated control, local ADSC injection, and acellular hydrogel groups [134].

Moon et al. reported that the allogenic AMSC-hydrogel complex improved healing outcomes in a diabetic mouse model by enhancing cell viability, wound closure rates, increasing angiogenesis, and reducing inflammation in the treated group compared to controls [128]. An injectable, non-crosslinked HA gel containing human adipose-derived stem cell (hASC) spheroids was developed by Feng et al. to enhance cell delivery and tissue repair. In the absence of further cross-linking, its gel-like consistency supports cell viability and functionality, as well as regenerative applications after injection. In vitro, the spheroids demonstrated improved cell-cell interactions and increased growth factor secretion relative to conventional monolayer cultures, indicating enhanced regenerative potential. In vivo results from a mouse diabetic wound model showed significant improvements in wound epithelial growth, increased dermal thickness, and enhanced angiogenesis compared to control groups [135].

Umbilical cord-derived mesenchymal stem cells (UCMSC)

UCMSCs were first isolated in 2013 from Wharton’s jelly, a gelatinous tissue found within the umbilical cord blood. Due to their unique multi-lineage differentiation property, low immune resistance, easy and noninvasive access to the umbilical cord, and lack of controversial ethical problems, UCMSCs are expected to partially replace BMMSCs in some applications [136, 137].

Xu et al. investigate a hybrid hydrogel that encapsulates human UCMSCs within an injectable hydrogel system composed of GelMA and chitosan-catechol, aimed at improving diabetic wound healing. Chitosan-catechol exhibits a significant hemostatic effect, and the incorporation of zinc ions into the hydrogel enhances angiogenesis. The combined treatment of UCMSCs and hydrogel facilitated DW healing by suppressing the inflammatory factors TNF-α and IL-1β in vivo, achieving a wound closure rate of 92.2% within 14 days. On day 7, the UCMSCs-hydrogel-treated group exhibited a marked increase in collagen deposition. The refined 10% GelMA hydrogel facilitates superior stem cell adhesion and function, promoting faster wound closure, reduced inflammation, enhanced vascular regeneration, and greater collagen accumulation in diabetic mice [138].

An adaptable hydrogel system employing gelatin microspheres (GMS) that encapsulate UCMSCs was created by Shi et al. to improve the delivery and integration of these cells in diabetic wound settings. This strategy in a diabetic mouse model demonstrated enhanced angiogenesis, collagen deposition, and facilitated extracellular matrix remodeling, ultimately expediting wound healing and skin regeneration. Moreover, the MSCs within the GMS were found to stimulate anti-inflammatory responses, resulting in minimizing local inflammation, thereby accelerating wound closure relative to the control group [139].

Other mesenchymal stem cells

Although BM-MSCs were among the first MSCs to be isolated, the invasive extraction process used for their isolation has shifted attention toward placenta-derived mesenchymal stem cells. Similar to the umbilical cord, the placenta provides a rich alternative source of perinatal MSCs with lower immunogenicity and fewer ethical concerns compared to BM-MSCs [140–142].

Zeng et al. reported a clinical case of a 57-year-old female with type 2 diabetes mellitus and a 20-day DFUs (Wagner classification III) that did not respond to conventional treatments. Topical application of placenta-derived MSC hydrogels reduced wound size after 3 weeks, resulting in shorter healing time and the formation of dense granulation tissue that facilitated wound recovery. Furthermore, PD-MSCs release paracrine factors that promote vascular development and regulate the immune system, thereby enhancing wound healing [143].

Recent clinical studies have provided stronger evidence supporting the use of placenta-derived MSCs in DFUs. Meamar et al. (2021) conducted a randomized pilot trial using human placenta-derived mesenchymal stem cells delivered in gelatin electrospun nanofibrous scaffolds combined with platelet-rich plasma gel, demonstrating significantly improved wound-healing rates, granulation tissue formation, and clinical outcomes compared with standard care. More recently, a Phase 2 multi-center, randomized, double-blind, placebo-controlled trial (Pollak et al. 2025) evaluated placenta-derived cells (PDA-002) in DFU patients with and without peripheral artery disease. The study confirmed safety across dose levels and showed promising efficacy signals for wound healing in this difficult-to-treat population. These findings suggest that placenta-derived MSCs, when delivered via hydrogel or scaffold systems, offer a clinically feasible and effective option for refractory DFUs [144, 145].

Chen et al. developed a nanoscale hydrogel-based oxygen-releasing stem cell transplantation device to enhance stem cell survival, retention, and therapeutic efficacy in the hypoxic environment of diabetic foot ulcers. This temperature- and pH-sensitive hydrogel is composed of cardiosphere-derived stem cells and oxygen-releasing microspheres. The system maintains intracellular oxygen levels for up to three weeks, accelerating wound contraction, enhancing skin vascularization, improving re-epithelialization and tissue repair, and reducing inflammatory responses [146]. Although clinical data for cardiosphere-derived cells in DFUs remain limited, their strong paracrine and pro-angiogenic properties, combined with oxygen-releasing hydrogel delivery, make them a promising candidate for ischemic diabetic wounds where hypoxia is a major barrier.

While the multi-differentiation capability of mesenchymal stem cells opens up numerous applications, it also highlights the risk of tumorigenicity. Cell-free treatments, including exosomes and artificial cell products derived from the MSC secretome, have recently attracted attention as potential solutions because they preserve the paracrine factors found in stem cells.

While the multipotent differentiation capacity of mesenchymal stem cells enables numerous applications, it also raises concerns regarding tumorigenicity. Cell-free therapies, including exosomes and artificial cell products derived from the MSC secretome, have attracted attention as potential alternatives because they preserve stem cell paracrine factors. Exosomes, which are extracellular vesicles characterized by a nanoscale phospholipid bilayer, enhance cell–cell communication by delivering biomolecules such as RNA and proteins to recipient cells [147, 148]. They facilitate wound healing by promoting cell migration, proliferation, vascularization, collagen deposition, re-epithelialization, and macrophage polarization [149–151].

Xu et al. developed a multifunctional injectable hydrogel incorporating foreskin-derived mesenchymal extracellular vesicles, demonstrating effectiveness in chronic wound repair and hair growth in a diabetic wound model. Using polyvinylpyrrolidone and silicotungstic acid, a dual-physical crosslinked hydrogel was created to encapsulate FM-EVs, exhibiting self-healing, adhesive, and antibacterial properties. Both in vivo and in vitro results showed that the combination of EVs and hydrogel significantly enhances diabetic wound repair through M1 to M2 macrophage polarization, increased angiogenesis, and improved microenvironment regulation [152].

Long et al. formulated an MSC-EV-7 A composite hydrogel to enhance diabetic wound healing through the synergistic effects of mesenchymal stem cell-derived extracellular vesicles (MSC-EVs) and the 7 A peptide. This hydrogel addresses key challenges in diabetic wound healing, including excessive inflammation and impaired angiogenesis. Treated wounds showed significantly reduced inflammatory markers, increased capillary density, and accelerated wound closure compared with untreated controls [153].

In a study by Geng et al., a multifunctional hydrogel composed of MSC-exosome-loaded carboxyethyl chitosan and dialdehyde carboxymethyl cellulose was developed to enhance diabetic wound healing. This system combines antibacterial and self-healing properties while delivering BMSC-derived exosomes. These exosomes provide bioactive factors that promote cell migration, angiogenesis, and macrophage polarization toward the M2 phenotype, thereby reducing inflammation. The hydrogel structure is formed via Schiff base reactions, which contribute to its stability and functionality [107].

Guo et al. developed a multifunctional hydrogel system incorporating ADSC-exos and thermosensitive components, which accelerates wound healing through antioxidative stress reduction, immunomodulation, and pro-angiogenic effects under mild heat stimulation. The thermosensitive components enable controlled release of exosomes at target sites, enhancing healing without invasive intervention [154]. As illustrated in Fig. 2, integrating MSCs with hydrogel-based delivery systems plays a significant role in promoting and regulating wound healing.

Fig. 2.

The integration of mesenchymal stem cells (MSCs) With hydrogel-based delivery systems. The figure illustrates various types of MSCs, their multiplicative benefits, and how hydrogels contribute to reversing pathological events and enhancing biological properties

However, current hydrogels still have limitations as ideal cell carriers. Their slow solidification rate may result in cell extrusion during muscle contraction or unintended solidification within blood vessels, potentially causing tissue damage by obstructing veins in the lower extremities. Table 2 provides a comprehensive overview of various hydrogel-based systems and their associated bioactive components used in diabetic wound healing.

Table 2.

Hydrogel-based systems and bioactive components for diabetic wound healing

Hydrogel system	Bioactive component	Key properties / functions	In vivo / clinical outcomes	Reference
GelMA-HExo	Anoxically pretreated ADSC-exosomes	Enhanced degradation resistance, controlled swelling, high biocompatibility	Accelerated diabetic ulcer healing via EC proliferation, migration, and neovascularization	[106]
ADM/GelMA composite hydrogel	HUVEC-exosomes	Antioxidant, antimicrobial, pro-regenerative; encapsulated in acellular dermal matrix	Inhibited bacterial colonization, reduced ROS, promoted collagen deposition and angiogenesis in diabetic wound models	[155]
ECM hydrogel with ADSC-exosomes	ADSC-exosomes	Sustained exosome release at wound site	Enhanced proliferation, migration, angiogenesis, collagen deposition; reduced inflammation; rapid wound closure; restored skin architecture	[13]
PF-127 thermosensitive hydrogel	hUC-MSC-exosomes	Continuous local exosome release	Improved wound closure, enhanced skin appendage regeneration, increased CD31 & Ki67; promoted granulation tissue; upregulated VEGF and TGF-β1	[98]
Chitosan/PVA NO-releasing hydrogel	BMSCs	Nitric oxide release, improved BMSC therapeutic efficacy	Accelerated re-epithelialization, improved vascularization, increased collagen deposition; enhanced growth factor expression in diabetic rabbit wounds	[121]
Bone marrow + platelets + fibrin glue + collagen matrix	BMSCs	Multi-component scaffold for chronic wounds	Enhanced healing in refractory chronic wounds; wound size reduction; complete closure in 3/8 patients in 4 weeks	[122]
Gelatin-hydroxyphenyl hydrogel (dual-enzyme crosslinked)	BMSCs	Rapid in situ crosslinking, improved cell survival	Accelerated re-epithelialization, enhanced granulation and collagen deposition, faster wound closure	[123]
Hydrolytic ester-linked hydrogel	MSCs	Controlled degradation, prolonged MSC persistence	Enhanced localized immune modulation; promoted wound healing in diabetic mice	[124]
PEGDA hydrogel	Human BM-MSCs + rat insulin-secreting cells	Co-encapsulation, improved cell viability	Wound repair nearly 3x faster than controls; limited by invasiveness and low MSC engraftment	[125]
HA spongy hydrogel	ADSCs	MSC retention, supports proliferation, promotes anti-inflammatory M2 polarization	Enhanced collagen deposition, structured ECM, improved angiogenesis and neoinnervation in diabetic mouse wounds	[131]
Curcumin-incorporated GelMA hydrogel (3D bioprinted)	ADSCs	Inhibits ROS, reduces ADSC apoptosis	Increased collagen deposition, enhanced angiogenesis, improved diabetic wound healing in nude mice	[132]
HA/gelatin/PEGDA hydrogel (UV crosslinked)	ADSCs	Rapid in situ gelation, supports endothelial migration and duct formation	3-fold increase in angiogenesis; enhanced wound healing compared to cell-free hydrogels	[133]
Decellularized adipose matrix (DAM) hydrogel	MSCs	Supports MSC adhesion, paracrine activity, growth factor secretion	Enhanced collagen deposition, revascularization, cellular infiltration; accelerated wound closure and tissue regeneration	[134]
Allogeneic ADSC-hydrogel complex	ADSCs	Injectable hydrogel scaffold	Improved cell viability, wound closure, angiogenesis; reduced inflammation in diabetic mice	[128]
Non-crosslinked HA gel with hASC spheroids	hASCs	Gel-like, supports cell-cell interactions and growth factor secretion	Enhanced epithelial growth, increased dermal thickness, improved angiogenesis in diabetic mouse wounds	[135]
GelMA + chitosan-catechol hybrid hydrogel	hUCMSCs	Hemostatic, zinc ions enhance angiogenesis	Reduced TNF-α & IL-1β, 92.2% wound closure in 14 days, increased collagen deposition, enhanced vascularization	[138]
Gelatin microspheres (GMS) encapsulating UCMSCs	UCMSCs	Adaptable delivery system, anti-inflammatory	Enhanced angiogenesis, collagen deposition, ECM remodeling, accelerated wound closure	[139]
Placenta-derived MSC hydrogel	PD-MSCs	Paracrine factor release, immunomodulation	Clinical case: reduced wound size, dense granulation tissue formation, decreased healing time in a 57-year-old diabetic patient	[143]
Oxygen-releasing hydrogel	Cardiosphere-derived stem cells	Temperature- and pH-sensitive, maintains intracellular O₂	Accelerated wound contraction, increased vascularization, improved re-epithelialization and tissue repair, inhibited inflammation	[146]
Hydrogel + foreskin MSC-derived extracellular vesicles (FM-EV)	FM-EVs	Dual-physical crosslinked, self-healing, adhesive, antibacterial	Enhanced diabetic wound repair, hair growth, M1to M2 macrophage reprogramming, angiogenesis, improved microenvironment	[152]
MSC-EV-7 A composite hydrogel	MSC-EVs + 7 A peptide	Targets inflammation and angiogenesis	Reduced inflammatory markers, increased capillary density, accelerated wound closure in diabetic wounds	[153]
BMSC-exo-loaded CEC-DACMC hydrogel	BMSC-derived exosomes	Antibacterial, self-healing	Enhanced migration, angiogenesis, M1 to M2 macrophage conversion, mitigated inflammation	[107]
ADSC-exosome + thermosensitive hydrogel	ADSC-derived exosomes	Accelerated the healing process	Antioxidative stress, immunomodulation, and proangiogenesis with mild heat stimulation	[154]

Clinical translation of MSC and MSC-derived exosome therapies

The clinical application of MSCs and their derivatives in DFU therapy has advanced from preclinical models to early-phase human trials. These approaches target the multifactorial barriers in diabetic wound healing, including chronic inflammation, impaired angiogenesis, oxidative stress, and poor tissue regeneration. MSCs from bone marrow (BM-MSCs), adipose tissue (ADSCs/ASCs), umbilical cord (UC-MSCs/hUC-MSCs), and placenta have been evaluated in case reports, pilot studies, and randomized controlled trials (RCTs). They often demonstrate accelerated wound closure, enhanced granulation tissue formation, improved vascularization, and reduced amputation risk compared to standard care [54, 56].

Notable examples include allogeneic ADSC-hydrogel sheets (e.g., ALLO-ASC-DFU). In a Phase 2 randomized trial, this approach achieved 82% complete wound closure at 12 weeks versus 53% in controls, with good tolerability and no major adverse events. Placenta-derived MSCs (e.g., PDA-002/cenplacel) have shown encouraging results in Phase 1/2 trials for DFUs, with or without PAD. Intramuscular administration proved safe and produced signals of improved healing and durable wound closure in PAD-complicated cases—a high-risk subgroup with limited options. A randomized pilot trial using placenta-derived MSCs in gelatin electrospun scaffolds ± platelet-rich plasma reported significantly better healing rates and granulation tissue formation than standard care [54, 145, 156].

Meta-analyses and systematic reviews of MSC therapies (primarily BM-MSCs, ADSCs, and UC-MSCs) confirm overall improvements in ulcer healing rates, reduced amputation risk, better perfusion (e.g., ankle-brachial index), and enhanced pain-free walking distance, with a favorable safety profile. Delivery methods vary (intramuscular, topical, or scaffold-assisted), but hydrogel or biomaterial-assisted approaches consistently improve retention and outcomes over direct injection [157, 158].

Despite these advances, MSC therapies face limitations: poor cell survival and retention in the hostile DFU microenvironment (hypoxia, oxidative stress, inflammation), variability in cell sourcing/expansion/dosing, potential immunogenicity or (rare) tumorigenicity risks, and challenges in scalability and standardization [54].

MSC-derived exosomes (MSC-exos) offer a compelling cell-free alternative. These nanoscale extracellular vesicles carry bioactive cargos (miRNAs, proteins, cytokines, lncRNAs) that mediate paracrine effects on immunomodulation (e.g., the M1-to-M2 macrophage shift), angiogenesis (via the HIF-1α/VEGF axis), fibroblast/keratinocyte activation, and ECM remodeling. Compared with live cells, exosomes exhibit lower immunogenicity, better stability during storage/handling, and reduced safety concerns [159, 160].

Preclinical models robustly show that MSC-exos (from ADSCs, BMSCs, UC-MSCs, etc.) accelerate DFU closure, collagen deposition, re-epithelialization, and vascularization. Early clinical translation includes trials of Wharton’s jelly MSC-conditioned media/exosomes, purified exosome products (PEP), and adipose-derived exosome dressings, with promising safety and healing signals in chronic ulcers. Ongoing Phase 1/2 trials evaluate topical or injected MSC-exos for DFUs (e.g., NCT06319287 for PEP-TISSEEL) [160].

Hydrogel-based delivery systems synergize with both MSCs and exosomes. They enhance retention, provide sustained/controlled release, protect payloads from degradation, and create a biomimetic niche that modulates the wound microenvironment. Examples include GelMA, chitosan, HA, and thermosensitive (e.g., Pluronic F-127) hydrogels loaded with MSCs or exosomes, which have shown superior angiogenesis, reduced inflammation, and faster closure in preclinical and early clinical settings compared to free administration [65].

Challenges for clinical translation remain. These include optimizing large-scale GMP manufacturing and standardization (especially for exosomes), dosage regimens, long-term safety monitoring, regulatory classification of combination products (cell/exosome + biomaterial), and demonstrating cost-effectiveness. Most evidence derives from small Phase 1/2 trials; larger, multicenter RCTs with standardized endpoints (e.g., complete closure at 12–24 weeks, amputation-free survival, quality of life) and extended follow-up are essential [54].

In summary, MSC-based therapies have demonstrated clinical feasibility and efficacy signals in DFU management, while MSC-derived exosomes represent a safer, more scalable next-generation approach. Integration with advanced injectable or smart hydrogels addresses key delivery barriers and simultaneously targets multiple DFU pathophysiologies. With ongoing optimization and rigorous trials, these regenerative platforms hold strong potential to reduce healing times, lower amputation rates, and improve outcomes in this challenging patient population [161].

Challenges and limitations

Injectable hydrogels have shown promise as delivery systems for MSCs in DFUs therapy; however, several challenges limit their translation. Natural hydrogels such as collagen, gelatin, and HA provide inherent bioactivity and enzymatically tunable biodegradability, but they suffer from batch-to-batch variability, limited mechanical properties, and difficulties in handling during wound care. In contrast, synthetic hydrogels offer precise control over structure and function, but many are not inherently biodegradable or biocompatible and require additional chemical modification and extensive purification, raising concerns about cytotoxicity. From a clinical perspective, the weak mechanical strength of hydrogels limits their stability in load-bearing areas and in wounds with high exudate, where they may deform or be washed out, leading to poor MSC retention. Their non-adhesive nature often necessitates secondary dressings, and repeated applications increase both cost and infection risk. Sterilization introduces additional complexity, as conventional methods such as heat or irradiation may alter hydrogel structure, degrade encapsulated factors, or compromise MSC viability; therefore, aseptic manufacturing becomes essential but costly [162].

Controlling degradation rates also remains challenging, as DFUs present variable enzyme levels, pH, and oxidative stress that can accelerate or delay hydrogel breakdown. Rapid degradation can result in premature loss of cells and bioactive molecules, whereas slower degradation can limit nutrient diffusion and hinder tissue remodeling [163]. The choice of crosslinking strategy further affects release kinetics, mechanical stability, and cell compatibility; chemical crosslinkers may leave toxic residues, and overly dense networks restrict MSC migration and factor diffusion, whereas insufficiently crosslinked gels risk burst release or rapid erosion [164].

In addition to material limitations, reproducibility and manufacturing remain major barriers. Subtle differences in polymer characteristics, water content, and cell handling (source, passage number, cryopreservation, and shear stress during injection) can influence MSC viability and potency. Finally, the chronic DFU microenvironment—characterized by infection, biofilm formation, protease activity, hypoxia, and oxidative stress can further impair MSC function and accelerate hydrogel degradation, thereby reducing therapeutic efficacy. Collectively, these challenges emphasize that, although hydrogels are promising carriers for MSC-based therapy, further optimization of material design, sterilization processes, degradation control, and clinical handling protocols is required before these systems can be reliably implemented in DFU treatment.

Future perspectives and conclusions

Comparative analysis of injectable versus conventional hydrogel dressings for DFU therapy

Conventional preformed hydrogel dressings, usually supplied as sheets, films, or amorphous gels (e.g., alginate-, PVA-, or gelatin-based), have been a standard option in DFU management. Their primary benefits include maintaining a moist environment, absorbing moderate exudate, promoting autolytic debridement, and reducing pain on dressing change. Meta-analyses indicate that these conventional hydrogels improve healing rates (OR 4.09, 95% CI 2.83–5.91) and reduce mean healing time by approximately 11 days compared with basic wound contact dressings [165, 166]. However, they have notable limitations: poor conformability to irregular, deep, or tunneling wounds (creating dead space), limited capacity for sustained or stimuli-responsive release of biologics, inability to effectively deliver MSCs or exosomes, and the frequent need for dressing changes, which increases infection risk and healthcare costs [167].

In comparison, injectable in situ-forming hydrogels offer several important advantages. Administered as low-viscosity solutions that rapidly gel at the wound site through physical, chemical, or hybrid crosslinking, they provide excellent conformability to complex DFU geometries, complete filling of wound cavities without dead space, and minimally invasive application [168]. These systems enable superior retention and controlled, stimuli-responsive release of MSCs, exosomes, growth factors, and antimicrobial agents, aligning delivery with the dynamic phases of diabetic wound healing. Many injectable formulations also incorporate self-healing, antioxidant, oxygen-releasing, or smart-responsive properties that simultaneously target multiple pathological barriers in DFUs (inflammation, hypoxia, infection, and oxidative stress) features that are difficult to achieve with static preformed sheets [73, 75].

Preclinical studies consistently demonstrate that injectable MSC- or exosome-loaded hydrogels achieve faster wound closure, significantly higher angiogenesis (up to 3-fold increase), better collagen organization, and stronger anti-inflammatory effects than conventional cell-free or preformed hydrogel dressings in diabetic models [168, 169]. Nevertheless, injectable hydrogels also present challenges, including more complex manufacturing and sterilization requirements, potential variability in gelation kinetics in the harsh DFU microenvironment, and higher regulatory hurdles for cell-laden products.

Overall, while conventional preformed hydrogel dressings remain simple and suitable for basic moist wound care, injectable hydrogels combined with MSCs or MSC-derived exosomes represent a clear advancement by offering superior conformability, better cell/exosome retention, sustained bioactive release, and simultaneous targeting of multiple DFU pathologies (chronic inflammation, hypoxia, and impaired angiogenesis) [73, 75]. These features have translated into faster wound closure, enhanced angiogenesis, and improved tissue quality in preclinical diabetic models compared to traditional therapies [106].

Looking forward, the most promising directions include the development of smart, multi-stimuli-responsive hydrogels that dynamically respond to glucose, pH, ROS, or enzymes in the wound bed [74]. The design of hybrid systems that combine the biocompatibility of natural polymers with the mechanical strength and tunable release of synthetic materials, and the use of personalized approaches using patient-derived MSCs or exosomes. Large-scale, head-to-head randomized controlled trials are now essential to confirm clinical efficacy, quantify the reduction in amputation risk, and evaluate cost-effectiveness relative to standard care [165, 167].

With continued multidisciplinary collaboration and optimization, injectable MSC- and exosome-laden hydrogels have strong potential to become a transformative regenerative platform that significantly improves healing outcomes and reduces the burden of diabetic foot ulcers.

Abbreviations

ADM

Acellular dermal matrix

ADSC(s)

Adipose-derived (mesenchymal) stem cell(s)

AGE(s)

Advanced glycation end product(s)

AMP(s)

Antimicrobial peptide(s)

AP

Aloe vera polysaccharide

AP-1

Activator protein-1 (transcription factor)

BMSC(s) / BM-MSC(s)

Bone marrow–derived mesenchymal stem cell(s)

bFGF

Basic fibroblast growth factor (FGF2)

CGRP

Calcitonin gene-related peptide

CMC

Carboxymethyl cellulose (oxidized CMC also appears)

Con A

Concanavalin A

CORM-401

Carbon monoxide–releasing molecule-401

CTLA4

Cytotoxic T-lymphocyte–associated protein 4

CTNND1

Catenin delta-1

DAM

Decellularized adipose matrix

dECM

Decellularized extracellular matrix

DFU(s)

Diabetic foot ulcer(s)

DOPA

3,4-dihydroxy-L-phenylalanine (catechol; adhesive chemistry)

DNA

Deoxyribonucleic acid

EC(s)

Endothelial cell(s)

ECM

Extracellular matrix

EDA

Ethylenediamine (appears in “GG-EDA@pDA”)

EGF

Epidermal growth factor

EV(s)

Extracellular vesicle(s)

FGF2

Fibroblast growth factor-2 (basic FGF)

FM-EV

Foreskin mesenchymal-derived extracellular vesicles

FOXM1

Forkhead box protein M1

FOXP3

Forkhead box P3 (Treg marker)

GelMA

Gelatin methacrylate

GG

Gellan gum

GITR

Glucocorticoid-induced TNFR-related protein

GOx

Glucose oxidase

HA

Hydroxyapatite (materials context) and Hyaluronic acid (biopolymer context)

HGF

Hepatocyte growth factor

HIF-1α

Hypoxia-inducible factor-1 alpha

HUVEC

Human umbilical vein endothelial cell

IDO

Indoleamine 2,3-dioxygenase

IFN-γ

Interferon-gamma

IP-10

Interferon-γ–induced protein-10 (CXCL10)

Ki67

Proliferation marker protein

LL-37

Human cathelicidin antimicrobial peptide

MACF

(Fluorinated) methacrylamide chitosan (hydrogel sheet)

MAPK

Mitogen-activated protein kinase (pathway)

MCP-1

Monocyte chemoattractant protein-1 (CCL2)

MEK/ERK

MEK/ERK signaling pathway

MIF

Macrophage migration inhibitory factor

MMP(s)

Matrix metalloproteinase(s)

MRSA

Methicillin-resistant Staphylococcus aureus

MSC(s)

Mesenchymal stem cell(s)

MSC-exos / MSC-EV(s)

MSC-derived exosomes / extracellular vesicles

MSC-EV-7A

MSC-EVs combined with “7A” peptide (study-specific)

M1 / M2

Macrophage phenotypes (pro-inflammatory / pro-regenerative)

NF-κB

Nuclear factor kappa-B

NO

Nitric oxide

NOS2 (iNOS)

Nitric oxide synthase 2 (inducible)

NLRP3

NOD-like receptor family pyrin domain–containing 3 (inflammasome)

NGF

Nerve growth factor

PAD

Peripheral artery/arterial disease

PBA

Phenylboronic acid

PD-MSC(s)

Placenta-derived mesenchymal stem cell(s)

PDGF

Platelet-derived growth factor

PEDF

Pigment epithelium–derived factor

PEGDA

Poly(ethylene glycol) diacrylate

PF-127

Pluronic F-127 (Poloxamer 407)

PGE2

Prostaglandin E2

PI3K/AKT

Phosphoinositide 3-kinase / AKT signaling pathway

PVA

Polyvinyl alcohol

PVR/CD155

Poliovirus receptor (CD155)

QAS

Quaternary ammonium salt

RAGE

Receptor for advanced glycation end products

RGD

Arg-Gly-Asp (cell-adhesion motif)

RNase

Ribonuclease (as in RNase 7)

ROS

Reactive oxygen species

SC(s)

Schwann cell(s)

SDF-1

Stromal cell-derived factor-1 (CXCL12)

siRNA

Small interfering RNA

SOCS3

Suppressor of cytokine signaling 3

STAT1

Signal transducer and activator of transcription 1

TEWL

Transepidermal water loss

TGF-β / TGF-β1

Transforming growth factor-beta / beta-1

Th2

T helper type 2

TIMPs

Tissue inhibitors of metalloproteinases

TLR4

Toll-like receptor 4

TNF-α

Tumor necrosis factor-alpha

Treg(s)

Regulatory T cell(s)

UCMSC(s) / hUC-MSC(s)

(Human) umbilical cord mesenchymal stem cell(s)

vSMC(s)

Vascular smooth muscle cell(s)

vWF

von Willebrand factor

VEGF / VEGFA / VEGFR

Vascular endothelial growth factor / A isoform / receptor

Author contributions

F.S. designed the review study. S.A.S., M.R., M.J.E., Y.K., and M.L. contributed to collecting the data and writing the manuscript. M.S. contributed to editing the text. F.S., S.A.S and M.R. contributed to reviewing the manuscript. All authors have confirmed the final version of the manuscript and are accountable for the contents of all parts of the work. All the authors have read and approved the final manuscript.

Funding

This study was not financially supported by any institution or funding agency.

Data availability

All data are available in this article.

Declarations

Ethics approval and consent to participate

Not applicable.

Consent for publication

Not applicable.

Competing interests

The authors declare no competing interests.