Collagen, the most abundant protein in the ECM of animal tissues, plays a fundamental role in maintaining structural integrity and modulating cellular activities essential for tissue homeostasis and repair [24]. Of the 29 known collagen types, types I, II, and III account for over 90% of total collagen in the human body. These proteins provide mechanical strength and structural support to various tissues, including skin, bone, cartilage, and connective tissue [25]. Collagen’s structure is based on repeating Gly-X-Y amino acid sequences, where glycine occupies every third position to allow tight packing of the three $\alpha$-chains into a triple-helix configuration—critical for its biological function [26]. Owing to its biodegradability, low immunogenicity, and biocompatibility, collagen has been extensively used in biomedical applications such as wound healing, tissue scaffolds, and drug delivery systems [27]. Despite its advantages, animal-derived collagen presents challenges including batch-to-batch variability, potential for allergic reactions, and the risk of transmitting zoonotic pathogens. To overcome these limitations, recombinant collagen has emerged as a promising and safer alternative [28]. Advanced gene expression and protein engineering techniques enable the production of recombinant collagens that mimic the structure and function of native human collagen, including essential post-translational modifications such as hydroxylation and glycosylation [29]. Quantitatively, yeast-expressed collagens typically reach ~35–45% proline hydroxylation compared to ~45–50% in native human collagen, a difference that still allows formation of stable triple helices with melting temperatures (Tm) in the range of 38–40 °C, closely approaching the ~41–42 °C observed for tissue-derived collagen. This near-native hydroxylation correlates with comparable cell-adhesion capacity (integrin binding $>$80% relative to native collagen) and bioactivity in angiogenesis assays. Collagen types I, II, III, and V have been successfully expressed using recombinant systems, with the capacity to form homotrimers (e.g., types I–III), heterotrimers (e.g., type XI), or hybrid structures (e.g., type IX) depending on their $\alpha$-chain composition [30]. Among these, type III collagen plays a pivotal role in tissue remodeling, especially during early stages of healing and scar modulation. It constitutes about 8–11% of the dermal ECM and declines with age, making it a valuable component for skin regeneration therapies [31, 32]. Large-scale recombinant production typically employs microbial systems like E. coli and yeast. While bacteria offer cost-effective expression, they lack the cellular machinery for post-translational modifications. In contrast, yeast-based systems yield collagen with closer structural and biochemical resemblance to the native human form [33]. Recombinant collagen-based hydrogels are especially attractive for regenerative applications due to their structural similarity to the native ECM, excellent biocompatibility, and tunable mechanical properties [34]. These features support cell adhesion, infiltration, and nutrient diffusion—crucial parameters for promoting osteogenesis and angiogenesis in tissue repair settings. For instance, Rodríguez-Cabello et al. [35] functionalized a recombinant type I collagen protein (RCPhC1) with methacrylamide (RCPhC1-MA), norbornene (RCPhC1-NB), and thiol (RCPhC1-SH) groups, enabling high-resolution 3D printing via two-photon polymerization (2PP). The resulting microstructured scaffolds supported cell attachment and proliferation in vitro, showcasing the material’s potential for fabricating anatomically precise constructs in tissue engineering [35]. In another example, Xu et al. [36] integrated mesenchymal stem cell-derived exosomes (MSC-EVs) into recombinant type III collagen nanoparticles. These engineered hydrogels modulated the miR-223/pKNOX1 axis to polarize macrophages toward an M2 phenotype, thereby reducing inflammation and enhancing angiogenesis in cutaneous wound repair. In vitro and in vivo studies demonstrated reduced IL-6 expression and increased levels of Ki67, CD31, and $\alpha$-SMA—markers indicative of proliferation and neovascularization—thus validating the material’s regenerative efficacy [36]. Munyemana et al. [37] synthesized porous hybrid nanoparticles using recombinant collagen and calcium carbonate. The collagen component modulated crystal growth through electrostatic interactions, enabling controlled morphology and high drug-loading efficiency. These nanostructures demonstrated pH-sensitive drug release and excellent cytocompatibility, making them suitable candidates for regenerative therapies involving localized drug delivery [37]. Similarly, Kong et al. [38] developed a recombinant type III collagen (rHCIII) hydrogel for delivering human adipose-derived stem cells (hADSCs) to diabetic wounds. The hydrogel preserved cell viability and function for over three weeks and significantly accelerated tissue repair in vivo, improving vascularization and re-epithelialization in treated mice [38]. In regenerative medicine, implantable scaffolds play a central role in guiding host tissue regeneration by creating a bioactive microenvironment that supports ECM deposition and cell integration. Both natural polymers (e.g., collagen, hyaluronic acid) and synthetic alternatives (e.g., PLA, PGA) have been explored, but recombinant humanized collagens stand out for their reproducible structure, customizability, and reduced immunogenicity. A recent study used a photoaging skin model to evaluate the regenerative efficacy of recombinant type III collagen (rhCol III). Treatment led to increased dermal collagen content, improved elasticity, and reduced epidermal thickening, confirming its therapeutic utility in skin regeneration and rejuvenation [39]. Table 1 (Ref. [35, 36, 37, 38, 39]) recaps the applications described.

Elastin is a crucial extracellular matrix protein that imparts elasticity and resilience to dynamic tissues such as blood vessels, lungs, and skin. Its structural architecture is composed of insoluble fibrillar networks stabilized by hydrophobic interactions, and is characterized by repetitive peptide sequences like elastin and elastin-like polypeptides (VPGG, VPGVG, and APGVGV). These properties not only contribute to the mechanical function of elastin but also support critical biological processes such as cell adhesion, proliferation, and migration—making it highly relevant for regenerative medicine applications [40]. However, the clinical translation of native elastin is hindered by several limitations: it is insoluble in aqueous media, potentially immunogenic when isolated from animal sources, and presents difficulties in processing and modification. Moreover, synthetic analogs are often expensive and inconsistent in structure [41]. A transformative advancement occurred when Urry and collaborators [42] identified VPGVG as a recurrent motif in natural elastin, leading to the development of elastin-like polypeptides (ELPs)—genetically engineered biopolymers mimicking elastin’s hydrophobic domain [43]. ELPs are typically produced via recombinant techniques, predominantly in Escherichia coli, due to its high-yield expression and scalability, although Pichia pastoris has also been utilized [44]. These recombinant polymers are generally composed of repeating pentapeptide units of the sequence Val-Pro-Gly-X-Gly (VPGXG), where X is any amino acid except proline. This modularity allows ELPs to replicate key elastin functions while enabling tunable mechanical and biochemical properties [45]. One of the most compelling features of ELPs in regenerative applications is their reversible phase transition behavior in aqueous solution. ELPs are soluble below a defined transition temperature (Tt), but undergo aggregation and phase separation above it. This thermoresponsive property is programmable through the choice and arrangement of guest residues, polymer length, and overall hydrophobicity, making ELPs ideal candidates for in situ forming biomaterials [46]. For tissue engineering purposes, researchers have developed multiblock ELP copolymers by linking hydrophobic and hydrophilic segments (e.g., [VPGIG]ₙ₁–[VPGSG]ₙ₂), or appending hydrophobic terminal blocks to facilitate self-assembly and mechanical reinforcement [47]. ELPs have also been functionalized with bioactive peptide sequences or engineered into fusion constructs to enhance cellular responses such as adhesion, angiogenesis, and matrix remodeling—key factors in tissue regeneration [48]. The injectable and thermoresponsive nature of ELPs makes them particularly useful in forming depot-like structures that provide controlled, localized, and sustained release of therapeutic agents. For instance, fusion of ELP to therapeutic proteins like interferon-alpha (IFN-$\alpha$) results in constructs that preserve biological activity while dramatically prolonging systemic half-life and enhancing tissue accumulation. Hu et al. [49] demonstrated that an IFN-$\alpha$–ELP fusion retained its biological efficacy, exhibited enhanced stability, and enabled sustained release at the target site in vivo. The same team later confirmed these benefits in glioblastoma models, where the ELP-fused cytokine formed an in situ depot that released the therapeutic in a zero-order kinetic profile, enhanced tissue penetration, and reduced tumor recurrence [50]. These design principles can be translated to regenerative contexts where extended presence of growth factors or cytokines is critical for orchestrating tissue repair over time. In vascular tissue engineering, ELPs have been employed to improve graft biocompatibility and function. For example, a nanofiber scaffold incorporating the REDV (arginine–glutamic acid–aspartic acid–valine) sequence demonstrated superior endothelial cell adhesion and proliferation, reduced platelet aggregation, and supported smooth muscle cell contractility—essential features for small-diameter vascular grafts [51]. Similarly, combining ELPs with other structural biomaterials has yielded promising results. Feng et al. [52] engineered a recombinant silk-elastin copolymer expressed in E. coli and integrated it with a nanocellulose layer to create a bilayer skin substitute. This construct showed exceptional mechanical strength, antibacterial properties, and regenerative efficacy in skin repair models [52]. Likewise, a recombinant scaffold composed of human collagen and elastin supported long-term membrane durability and facilitated dermal regeneration, offering a viable alternative to conventional skin substitutes [53]. Furthermore, ELPs’ amenability to chemical and genetic modification has enabled the development of multifunctional regenerative systems. For instance, elastin-based matrices have been tailored with pH-sensitive linkers for targeted drug release in acidic wound environments or fused with cell-penetrating peptides to enhance intracellular delivery of therapeutic cargos (e.g., Doxorubicin-Dox-) in tissue repair settings [54]. Li and Champion [55] further demonstrated how incorporating p-azidophenylalanine (pAzF) into the ELP backbone allows photo-induced crosslinking to stabilize nanovesicles, enhancing mechanical stability and control over drug release profiles. This strategy can be applied in regenerative contexts where spatiotemporal control over bioactive signaling is crucial [55]. Other ELP constructs have been designed to serve as immunomodulatory depots. For example, CpG oligodeoxynucleotides (known immune stimulants) were electrostatically bound to ELP-K12 (an ELP with 12 lysine residues) to form an injectable depot that released bioactive unmethylated cytosine-guanine (CpG) for up to three weeks. Although originally developed for immunotherapy, this sustained local immune activation concept could be leveraged in regenerative medicine to stimulate endogenous repair mechanisms and modulate inflammatory responses during wound healing [56]. ELPs can then be customized for scaffold formation, therapeutic delivery, immune modulation, and cellular microenvironment engineering—making them a powerful class of biomaterials for tissue regeneration. A summary of the discussed regenerative applications is presented in Table 2 (Ref. [49, 50, 51, 52, 53, 54, 55, 56]).

Silk-like polymers (SLPs) represent a prominent class of genetically engineered biomaterials with extensive applications in regenerative medicine. These polymers are composed of repeating sequences such as Gly–Ala–Gly–Ala–Gly–Ser, mimicking the structure of natural silk proteins produced by insects and spiders. Native silks, including those from Bombyx mori and Nephila clavipes, are well known for their impressive mechanical strength, elasticity, biocompatibility, and biodegradability [57, 58]. Among expression systems, Escherichia coli is frequently used for recombinant silk production due to its cost-effectiveness, rapid growth, and ease of genetic manipulation [59]. Genetically engineered SLPs often replicate silk motifs derived from the major ampullate gland proteins, with typical repetitive sequences such as [GGAGQGGYGGLGSQGAGRGGLGGQGGAG] and [GPGGYGGPGQQGPGGYAPGQQPSGPGS] from N. clavipes [60]. These sequences are frequently modified to regulate crystallinity or to introduce biologically active motifs such as RGD to enhance cell adhesion. Due to their low solubility and limited flexibility, silk domains are often copolymerized with elastin-like motifs, resulting in silk-elastin-like proteins (SELPs) that combine mechanical resilience with water solubility and flexibility [61]. This configuration is advantageous for producing injectable biomaterials and drug delivery vehicles with tunable degradation profiles [62, 63, 64]. For instance, Florczak et al. [65] demonstrated that MS1 silk gels functionalized with the H2.1 peptide and loaded with doxorubicin selectively targeted Her2-positive breast cancer cells in vivo, significantly reducing tumor volume with minimal systemic toxicity. Notably, unloaded particles showed no adverse effects in healthy mice, and detailed histology revealed no damage to major organs [65]. Herold et al. [66] further developed recombinant silk particles (eADF4(C16)) chemically linked with doxorubicin via a pH-sensitive hydrazine linker. These particles exhibited minimal drug release at physiological pH, while delivering the full drug payload under acidic conditions typical of intracellular vesicles, enabling precise intracellular drug release [66]. Similarly, Mulinti et al. [67] developed enzyme-responsive nanospheres using a recombinant spider silk copolymer with a thrombin-sensitive linker to release vancomycin in response to Staphylococcus aureus infection. These nanocarriers demonstrated strong antibacterial effects and potential for managing drug-resistant infections in tissue repair [67]. Lian et al. [68] created a nanofibrous membrane by electrospinning recombinant silk protein blended with sodium hydrogen sulfide (NaHS). The resulting scaffold sustained hydrogen sulfide (H₂S) release, supported endothelial progenitor cell (EPC) function, and significantly accelerated wound healing in a murine skin injury model [68]. Further regenerative applications have focused on hybrid constructs combining silk fibroin and recombinant spidroins. For example, composite hydrogels fabricated from silk fibroin and recombinant silk proteins functionalized with fibronectin motifs, growth factors (e.g., Fibroblast Growth Factor (FGF)), and antimicrobial peptides exhibited enhanced cell adhesion, antimicrobial activity, and stimulation of dermal cell proliferation—supporting the development of bilayered skin grafts [69, 70]. Other innovations include recombinant spidroins enriched with SIBLING-derived peptide tags (e.g., osteopontin and bone sialoprotein), which enhanced calcium phosphate mineralization and osteoblast adhesion—demonstrating strong potential for tendon–bone interface regeneration [71]. Additionally, eADF4(C16)-based scaffolds rich in carboxylic acid residues were shown to promote calcium ion binding and mineralization. Culturing mesenchymal stem cells (MSCs) on these hydrogels resulted in elevated alkaline phosphatase activity, highlighting their osteoinductive capacity [72]. Another engineered scaffold featured a chimeric fusion protein incorporating a spider silk-inspired sequence and a hyaluronic acid-binding motif (VTK), which guided osteogenic differentiation of bone marrow-derived human MSCs, underlining its applicability for bone tissue engineering [73]. Finally, a composite construct was developed by integrating decellularized extracellular matrix (dECM) derived from equine cartilage with a silk-based hydrogel. This composite served both as a structural scaffold and a bioadhesive to promote integration with host tissue. Mechanical and biological analyses confirmed its suitability for supporting chondral regeneration under load-bearing conditions, emphasizing the value of SELP-based hydrogels in developing advanced regenerative therapies [74]. Table 3 (Ref. [65, 66, 67, 68, 69, 70, 71, 72, 73, 74]) summarizes the described applications.

Bacterial cellulose (BC) is a polysaccharide macromolecule with the same chemical structure as plant-derived cellulose, yet it lacks associated plant components such as lignin, pectin, and hemicellulose that are commonly found in plant cell walls [77]. Structurally distinct due to its unique nanoscale fiber organization, BC exhibits superior mechanical strength, elevated crystallinity, and enhanced purity compared to its plant-based counterpart [78]. It is composed of $\beta$-1,4-linked D-glucopyranose units, stabilized by extensive intra- and intermolecular hydrogen bonding, and conforms to the molecular formula (C₆H₁₀O₅)ₙ [79]. The bacterium Acetobacter xylinum biosynthesizes two crystalline forms—cellulose I (ribbon-like) and cellulose II (thermodynamically stable)—by extruding glucan chains as protofibrils, which aggregate into nanofibrils and self-assemble into a web-like 3D network. The abundance of hydroxyl groups confers high hydrophilicity, chemical modifiability, and excellent biodegradability to BC, making it highly suitable for regenerative medicine applications [79]. Due to its intrinsic biocompatibility, porosity, mechanical tunability, and water retention, BC is an ideal scaffold material for the controlled release of therapeutic molecules and the engineering of regenerative constructs. In bone tissue engineering, three-dimensional (3D) scaffolds that integrate biomaterials with osteogenic cells and signaling agents are gaining traction for their ability to mimic native bone architecture and promote functional repair. Cao et al. [80] designed a nanocomposite scaffold by incorporating oxidized BC with chitosan (CS) and nano-hydroxyapatite (nHA). This formulation showed superior mechanical strength, favorable degradation kinetics, and enhanced water absorption compared to CS/nHA-only constructs. When tested with MC3T3-E1 preosteoblasts, the scaffold promoted cell proliferation and viability. In vivo studies using a rat calvarial defect model further confirmed enhanced new bone formation [80]. Zhu et al. [81] developed a hydrogel scaffold combining BC with CS and alginate (Alg), which formed a dense, fibrous network with desirable swelling behavior and degradation rate. This composite scaffold exhibited excellent apatite nucleation, cell compatibility, and protein adsorption capacity, while also enabling controlled drug release—key features for bone and cartilage regeneration [81]. In the context of cartilage tissue engineering, Li et al. [82] engineered a 3D hierarchical porous scaffold composed of BC and decellularized cartilage extracellular matrix (DCECM), crosslinked using N-hydroxysuccinimide (NHS) and N-[3-(dimethylamino)propyl]-N^′-ethyl carbodiimide hydrochloride (EDC). The scaffold facilitated robust adhesion and proliferation of rabbit-derived chondrocytes. In vivo implantation into a rabbit cartilage defect model demonstrated improved cartilage regeneration compared to native BC, attributed to its excellent elasticity, water retention, and hydrophilic properties—closely mimicking the biomechanical profile of native cartilage [82]. In ophthalmic regenerative medicine, Han et al. [83] fabricated a composite hydrogel using BC and poly (vinyl alcohol) (PVA) as a potential corneal stroma substitute. The BC/PVA scaffold exhibited increased optical transparency, enhanced water retention, and improved structural and surface morphology. In vitro cytocompatibility tests and in vivo rabbit implantation studies confirmed the scaffold’s ability to maintain corneal architecture, stability, and transparency more effectively than BC alone [83]. Other promising developments include the formulation of BC scaffolds functionalized with quince seed mucilage, a glucuronoxylan-based hydrogel. This modification improved swelling capacity and promoted enhanced fibroblast adhesion and proliferation, making it suitable for soft tissue engineering applications [84]. BC is also under active investigation as a component in advanced drug delivery platforms for regenerative therapies. Nanoparticles made from BC (BCNPs) have been developed as biocompatible carriers for sustained therapeutic release. These nanoparticles demonstrated increased crystallinity and size with prolonged culture, and remained thermally stable up to 90 °C. In vitro studies using bovine serum albumin (BSA) as a model drug confirmed successful loading and sustained release, supporting their use in nanomedicine for localized tissue regeneration [85]. Further advancements include the incorporation of poorly water-soluble glucocorticoids into BC-based hydrogels via microemulsion strategies. Zahel et al. [86] developed multiple formulations containing hydrocortisone or dexamethasone, achieving uniform integration into the BC matrix. Transmission electron microscopy confirmed the consistent dispersion of even water-in-oil (w/o) systems. Drug permeation across synthetic membranes was efficient and tunable, and anti-inflammatory efficacy was preserved post-release—underscoring the potential of BC composites in managing local inflammation during tissue repair [86]. Table 4 (Ref. [80, 81, 82, 83, 84, 85, 86]) summarizes the regenerative applications of BC-based systems discussed herein.

Hyaluronic acid (HA) is a high molecular weight glycosaminoglycan extensively used in regenerative medicine due to its unique physicochemical and biological properties. While HA is naturally present in the ECM of vertebrate tissues, its biotechnological production via microbial fermentation—particularly from Streptococcus equi subsp. zooepidemicus and recombinant platforms such as Lactococcus lactis, Bacillus subtilis, Corynebacterium glutamicum, Escherichia coli, and Pichia pastoris—has become the preferred method for obtaining pharmaceutical-grade HA. This shift ensures enhanced biosafety, scalability, and product standardization while eliminating risks associated with animal-derived HA, such as immunogenicity and pathogen contamination [87, 88]. Structurally, HA consists of repeating disaccharide units of D-glucuronic acid and N-acetyl-D-glucosamine linked via $\beta$(1$\to$3) and $\beta$(1$\to$4) glycosidic bonds. Its linear, non-sulfated configuration and ionized carboxylic groups enable strong hydration through hydrogen bonding and interaction with cations, supporting tissue integrity and cell signaling. These characteristics confer viscoelasticity and pseudoplastic rheological behavior—essential properties for its role as a dynamic ECM component and injectable scaffold material [89]. In regenerative medicine, bacterial-origin HA has emerged as a bioactive scaffold and signaling molecule with applications across various tissue types. The biological and rheological properties of HA strongly depend on its molecular weight. Low-molecular-weight (LMW) HA is typically defined as 250 kDa, intermediate as 250–1000 kDa, and high-molecular-weight (HMW) HA as $>$1000 kDa. LMW HA exhibits lower viscosity and reduced viscoelastic moduli (G^′ 50 Pa), which facilitates rapid diffusion and degradation, making it suitable for applications requiring fast turnover or enhanced cell signaling. Conversely, HMW HA forms highly viscous hydrogels with storage moduli (G^′) exceeding 200–500 Pa, improved shear-thinning behavior, and slower degradation rates (half-life ranging from 2–4 weeks in vivo). Then, high molecular weight (HMW) HA supports structural integrity, dampens mechanical stress, and exhibits anti-inflammatory, anti-angiogenic, and immunosuppressive properties—making it suitable for maintaining tissue homeostasis and modulating immune responses. In contrast, low molecular weight (LMW) HA penetrates tissues more efficiently and acts as a pro-regenerative agent by stimulating angiogenesis, inflammation, and cell proliferation. Intermediate-weight HA is involved in wound healing and embryonic development and demonstrates good transdermal permeability [90]. HA’s regenerative effects are largely mediated by its interaction with specific cellular receptors, particularly CD44 and RHAMM (CD168), which regulate cell proliferation, motility, and differentiation. These interactions enable HA to influence a broad range of regenerative processes, including angiogenesis, inflammation resolution, ECM remodeling, and cell recruitment [91]. In tissue engineering, HA serves as a temporary ECM substitute and delivery platform. Bacterial HA-based hydrogels have been engineered for the regeneration of bone, nerve, cartilage, muscle, and skin tissues. Chemically modified HA derivatives or polyelectrolyte complexes with other biopolymers (e.g., chitosan) enhance mechanical stability, tune degradation rates, and allow for controlled drug and cell release [92]. For example, HA-based hydrogels have been loaded with stem cells, bioactive peptides, or growth factors to promote localized differentiation and tissue repair in musculoskeletal and neural contexts [93]. Advanced fabrication technologies further expand the versatility of bacterial HA. Electrohydrodynamic (EHD) techniques like electrospinning and electrospraying allow the production of HA-based fibrous matrices and microparticles with defined porosity and drug release profiles, ideal for guiding cell behavior and controlled healing [94]. In parallel, 3D bioprinting using bacterial HA-derived bioinks enables the precise construction of biomimetic scaffolds with tunable viscoelastic properties, supporting spatially organized tissue development and facilitating personalized regenerative therapies [95]. In wound healing, HA-based gels and dressings maintain a moist microenvironment, enhance epithelialization, and promote angiogenesis, with demonstrated efficacy in both acute injuries (e.g., burns) and chronic lesions (e.g., diabetic ulcers) [96]. In cartilage repair, injectable HA hydrogels—especially those modified with thermo-responsive polymers like N-isopropylacrylamide (NIPAM)—facilitate minimally invasive interventions by gelling in situ at body temperature and improving chondrocyte viability and matrix synthesis, thus showing significant promise in osteoarthritis management [97]. Bacterial HA also plays a central role in joint regeneration. As a physiological constituent of synovial fluid, it is a key agent in viscosupplementation therapies aimed at restoring lubrication, reducing joint friction, and stimulating cartilage regeneration. Clinical studies have reported significant improvements in pain relief and joint function following intra-articular injection of bacterial HA in osteoarthritic patients [98]. In bone tissue engineering, HA is combined with inorganic materials such as hydroxyapatite, calcium phosphates, and osteogenic factors (e.g., Bone Morphogenetic Protein BMP-2) to fabricate hybrid scaffolds that provide both mechanical support and biochemical cues for osteogenesis and neovascularization [99]. Similarly, in myocardial regeneration, injectable bacterial HA-based hydrogels loaded with mesenchymal stem cells enhance cell retention and promote in situ differentiation, aiding in the repair of ischemic cardiac tissue [100]. Nerve tissue repair also benefits from HA’s bioactivity. HA-based conduits and hydrogels have been shown to support axonal regrowth, myelination, and functional recovery in models of spinal cord injury, owing to their permissive microenvironment and compatibility with neurotrophic agents [101]. The regenerative medicine applications of bacterial hyaluronic acid are summarized in Table 5 (Ref. [93, 97, 98, 99, 100, 101]).

Alginate is a naturally occurring polysaccharide predominantly found in brown algae (Phaeophyceae), where it serves as a key structural component of cell walls, providing both mechanical stability and osmotic regulation [102]. Although bacterial production is also possible, most commercially available alginate is extracted from algal sources such as Laminaria, Macrocystis, Ascophyllum, and Sargassum. Its high viscosity and gelling capacity make alginate a valuable material, particularly in the food and pharmaceutical industries [103]. Due to its sustainable origin, biocompatibility, and biodegradability, alginate has gained considerable attention in biotechnology and biomedicine, where it is used in drug delivery systems, scaffolds for tissue regeneration, and 3D bioprinting applications [104]. Alginate is recognized as safe by the FDA, further supporting its role in regenerative medicine [105]. Traditional industrial alginate extraction from seaweed involves multiple steps: raw material pretreatment, alkaline solubilization, precipitation with acids or alcohols, followed by drying and milling. However, environmental concerns related to seaweed harvesting and variability in the alginate’s composition—affected by seasonal and geographic factors—complicate standardization, especially in pharmaceutical contexts. As an alternative, bacterial systems such as Pseudomonas and Azotobacter species have been explored for more controlled alginate biosynthesis [106]. Among these, Azotobacter vinelandii stands out for its capacity to produce alginate with high molecular weight and rheological stability. This non-pathogenic bacterium naturally synthesizes alginate as both a carbon reserve and a nitrogenase protective barrier. The complete genome of A. vinelandii facilitates genetic modification, enabling strains that produce alginates with customized properties. For instance, oxygen tension and growth rate during fermentation significantly affect polymer molecular weight, and limiting oxygen availability can promote synthesis of highly homogeneous polymers [106, 107]. Chemically, alginate is a linear anionic copolymer composed of $\beta$-D-mannuronic acid (M) and $\alpha$-L-guluronic acid (G), connected via $\beta$-(1$\to$4) bonds. These monomers form homopolymeric (MM, GG) or heteropolymeric (MG) blocks, with their distribution depending on species, growth stage, and environment. The physical and biological properties of alginate are closely linked to this structural configuration. Due to its carboxyl groups, alginate interacts electrostatically with proteins and cationic polymers, enabling development of controlled-release systems for cationic drugs [108]. In the presence of divalent cations (e.g., Ca²⁺), alginate forms hydrogels at room temperature following the “egg-box” model, where G-blocks facilitate crosslinking. High G-content results in rigid gels, whereas M-rich alginate yields softer gels [109]. Gelation kinetics can be modulated by temperature, concentration, and calcium ion availability—slower gelation methods using phosphate buffers, insoluble calcium salts, or low temperature improve structural homogeneity [110, 111]. These properties make alginate hydrogels ideal for bioactive molecule immobilization and for applications in cancer therapy, gene delivery, and stem cell encapsulation. Alginate has also demonstrated immunomodulatory effects, including macrophage modulation and anti-inflammatory activity, making it relevant for regenerative contexts [112]. Despite its advantages, alginate exhibits limitations such as low cell adhesion, variable degradation rates, and poor mechanical strength, restricting its use in advanced biomedical applications. These drawbacks are addressed through chemical modifications aimed at enhancing native properties or introducing new functionalities. For example, enzymatic modification using epimerase enzymes can convert $\beta$-D-mannuronic acid residues into $\alpha$-L-guluronic acid, significantly increasing the rigidity and mechanical strength of the resulting gels. This makes them crucial for bone and cartilage regeneration, as exemplified by injectable alginate-hydroxyapatite hydrogels enriched with gelatin microspheres, which have shown improved stability and osteointegration for repairing large segmental bone defects [113]. Similarly, an alginate hydrogel enriched with platelets and integrated into a 3D porous scaffold of chitosan, chondroitin sulfate, and silk fibroin effectively mimics native cartilage ECM, enhancing metabolic activity and promoting type II collagen expression for cartilage regeneration [114]. Oxidative modification using sodium periodate introduces reactive aldehyde groups, significantly enhancing alginate’s ability to bind proteins, peptides, and bioactive molecules, while also conferring improved antimicrobial properties—a critical advantage for wound healing. For example, biodegradable alginate–chitosan hydrogels loaded with vitamin E, known for their high porosity, have demonstrated enhanced regenerative properties for skin tissue engineering [115]. Fibrous scaffolds of calcium alginate, fabricated via microfluidic spinning, offer excellent biocompatibility and an ECM-mimicking architecture, making them ideal for chronic wound healing [116]. Sulfation of alginate through the addition of sulfate groups imparts anticoagulant activity comparable to heparin, broadening its applicability in cardiovascular settings. This modification improves cell adhesion and reduces clot formation, inflammation, and undesired immune responses, as seen in alginate-based polyurethane elastomers that exhibit excellent elongation capacity, self-healing properties, and enhanced endothelial cell adhesion [117]. For cardiac regeneration following myocardial infarction, incorporating fullerenol nanoparticles into alginate hydrogels creates an injectable antioxidant system that scavenges reactive oxygen species, supporting functional cardiac recovery [118]. In neurological applications, alginate hydrogels with anisotropic capillary structures have been investigated for axonal growth and spinal cord injury repair. These scaffolds promote linear, oriented axonal regrowth and neuronal reinnervation, integrating well in vivo without inducing inflammatory responses, thereby positioning them as a promising strategy for post-injury spinal cord regeneration [119, 120]. Table 6 (Ref. [113, 114, 115, 116, 117, 118, 119, 120]) recaps the described applications.