Natural polymer hydrogels are derived from biological macromolecules and typically resemble the composition of the native ECM. Examples include proteins (collagen, gelatin, fibrin and keratin) and polysaccharides [hyaluronic acid (HA), alginate, chitosan and agarose] (23-27). These materials intrinsically mimic the native matrix by providing ligands for cell adhesion and enzymatic degradation sites that facilitate tissue remodeling. Natural hydrogels are typically biocompatible and bioactive, meaning that cells recognize and interact with them favorably. However, they typically exhibit inferior mechanical strength (compared with synthetic hydrogels), batch-to-batch variability and a risk of immunogenicity if not adequately purified.

Collagen type I, the most abundant protein in the mammalian ECM, is found in skin, bone, tendon and myocardium and readily forms hydrogels through the self-assembly of its triple-helical monomers. Collagen gels exhibit intrinsic cell-binding motifs and undergo degradation by collagenase, which makes them cytocompatible. They have been used in tissue engineering and 3D cell culture (19). However, their limitations include low stiffness and the potential for xenogeneic immune responses unless properly processed (decellularization or pepsin digestion to decrease antigenicity) (19). Chemical cross-linkers (such as glutaraldehyde) can reinforce collagen gels, although this may compromise biocompatibility (5). Collagen hydrogels have demonstrated efficacy in cardiac repair. McLaughlin et al (28) demonstrated that collagen hydrogel treatment decreases methylglyoxal adduct formation within cardiomyocytes and enhances contractile function in the infarct border zone in a mouse model of MI (Fig. 2A-D). Histological analysis has revealed increased myocardial salvage and reduced fibrosis in the border zone (Fig. 2E-I) (28). Gelatin is a denatured form of collagen produced through partial hydrolysis. It retains numerous collagen bioactive sequences but is easier to handle because it forms gels upon cooling, a property known as thermal reversibility (13). Gelatin hydrogels are biocompatible and can be formulated to deliver growth factors. For example, heparin-binding growth factors can be incorporated to promote vascularization in engineered tissue (11). Because gelatin is derived from collagen, it represents a promising material for cardiac repair. Although its mechanical strength is limited, gelatin is typically chemically modified (for example, gelatin methacrylate) to enable cross-linking and enhance mechanical rigidity (13).

Fibrin is a natural clotting polymer formed through thrombin-mediated cleavage of fibrinogen. Fibrin gels polymerize rapidly at physiological temperature and can be obtained autologously from the patient blood plasma (5). They support cell infiltration and ECM deposition because fibrinolytic enzymes degrade the gel (29,30). Fibrin has clinical applications as a surgical sealant (11). Fibrin injection into infarcted rat hearts enhances cell engraftment and myocardial perfusion, resulting in smaller infarct size and improved cardiac output (31). However, fibrin has limitations, including rapid degradation and relatively weak mechanical strength; therefore, it may not provide sustained structural support unless reinforced or combined with other materials (30). Nonetheless, fibrin exhibits good biocompatibility, as it mimics the native wound-healing process. Fibrin-based heart valve scaffolds generated through injection molding of cell-laden fibrin have also been investigated (15).

HA is a glycosaminoglycan that is abundant in connective tissue and heart valves. HA hydrogels are enzymatically degradable by hyaluronidase and are non-immunogenic because HA is a natural component of the human ECM. However, unmodified HA is mechanically weak and fragile, necessitating chemical modification (such as methacrylation or thiol modification) or cross-linking to achieve sufficient mechanical stability for cardiovascular application (13). HA has been well-documented to contribute to wound healing and angiogenesis (14,18,32). Ding et al (32) developed an injectable hydrogel based on methacrylate-modified HA, which rapidly gels following ultraviolet irradiation via the Michael addition reaction (Fig. 3A). The aforementioned study demonstrated that this hydrogel suppresses cell apoptosis, increases the M2-to-M1 macrophage ratio, promotes angiogenesis, decreases infarct size and improves cardiac function (Fig. 3B) (32). In another study, application of a cross-linked HA patch to the heart increased ventricular wall thickness and vascular formation, highlighting the potential of HA hydrogels in cardiac regeneration (11).

Alginate, a polysaccharide derived from seaweed, forms hydrogels through ionic cross-linking with divalent cations (such as Ca²⁺). Alginate hydrogels are biocompatible, inexpensive and readily gel in situ by injection of a soluble alginate solution followed by a calcium-containing solution (33). They are non-immunogenic and non-thrombogenic, and their mild gelation conditions, along with the ease of encapsulating cells or drugs, make them suitable for cardiovascular application (33,34). Unmodified alginate is not enzymatically degraded in humans; the gels gradually dissociate through ion exchange of Ca²⁺ with the surrounding environment, which may compromise long-term stability. To address this limitation, alginates are partially oxidized to enable slow hydrolytic degradation or covalently cross-linked to enhance structural stability (34). Lee et al (35) presented findings from the first clinical trial evaluating the feasibility and safety of Algisyl-LVR™, a hydrogel composed of sodium and calcium alginate (LoneStar Heart Inc.), for the treatment of HF associated with dilated cardiomyopathy. During injection, mechanical stress disrupts the dynamic cross-links formed via Ca²⁺ coordination, causing the hydrogel to transition into a liquid state that facilitates catheter delivery. Once injected into the myocardium, the stress is relieved, and dynamic cross-linking is re-established (35).

Chitosan is derived from chitin, which is found in crustacean shells, through deacetylation. It is a cationic polysaccharide that forms hydrogels via ionically triggered gelation when mixed with β-glycerophosphate. Chitosan structure resembles that of glycosaminoglycans and it is biocompatible, with mild antibacterial properties. It dissolves in acidic solutions and forms gel as the pH or the temperature rises (36). For cardiac applications, thermosensitive chitosan hydrogels (liquid formulations that remain injectable at room temperature but solidify into a gel at body temperature) are particularly attractive because they can be delivered by catheter (36). Chitosan can be chemically modified [by conjugation with arginine-glycine-aspartic acid (RGD) peptides or growth factors] to enhance cell adhesion and therapeutic efficacy (12).

Several other natural materials have niche applications. Agarose, an algae-derived polysaccharide, forms gels upon cooling, and its pore size is altered by adjusting agarose concentration. Agarose is used as a cell-encapsulation matrix and for drug delivery (13). Kim et al (37) developed a low gel temperature agarose by incorporating β-cyclodextrin into ethylenediamine-modified agarose and demonstrated that this modified gel enables sustained release of a thermosensitive drug. Keratin-based hydrogels, derived from hair or wool proteins, leverage keratin self-assembling fibrous structure. They retain bioactive motifs associated with hair follicle development and demonstrate pro-angiogenic effects without eliciting inflammation. Matrigel, a complex ECM extract derived from tumor cells, has been used as an injectable scaffold for stem cell delivery to the heart (19). Studies (17,18) have shown that it preserves ventricular geometry and enhances vascularization. However, Matrigel non-human origin and batch-to-batch variability preclude its clinical use.

Overall, natural hydrogels typically exhibit excellent bioactivity and cytocompatibility because their biochemical components resemble those of native tissue (15). Cells typically attach, migrate and function more readily on natural matrices. For example, a collagen or fibrin patch applied to an infarct can recruit cells and support novel tissue formation more effectively than numerous synthetic materials (3). However, purely natural hydrogels often possess limited mechanical strength and structural stability. They tend to be soft and prone to rapid degradation. In a beating heart or a high-pressure artery, a mechanically weak hydrogel may not withstand physiological forces over time (5). Thus, although natural polymers establish a biologically favorable microenvironment, they typically require modification or reinforcement for load-bearing cardiovascular applications.

Synthetic hydrogels are formed from synthetic polymers such as poly(ethylene glycol) (PEG), poly(vinyl alcohol) (PVA) and polyacrylate (38-40). A key advantage of synthetic polymers is their tunability and reproducibility: Molecular weight, cross-link density and chemical functional groups can be precisely controlled, resulting in consistent batch-to-batch properties (5,15). Synthetic hydrogels can be engineered to exhibit a range of stiffness, degradation rate and network architectures to meet specific application requirements. They also tend to be chemically inert and non-immunogenic in their pure form. However, unlike natural polymers, most synthetic polymers lack inherent bioactive cues; cells do not readily recognize or adhere to them (13). Therefore, biofunctionalization is typically required. For example, peptides such as RGD can be grafted onto synthetic polymer backbones to promote cell adhesion (13).

PEG, also known as poly(ethylene oxide), is a hydrophilic polymer approved by the U.S Food and Drug Administration (FDA) for medical applications (41). It resists protein adsorption and is typically biologically inert, which decreases non-specific immune reactions. PEG hydrogels are typically formed by cross-linking multi-armed PEG molecules functionalized with acrylates, thiols or other reactive groups to generate a polymeric network. By adjusting PEG molecular weight or incorporating degradable linkers, such as ester bonds or peptide sequences, between PEG chains, the mechanical strength and degradation rate of the gel can be controlled (5). A notable advantage of PEG lies in its tunable architecture and well-defined chemical composition. Researchers engineer PEG hydrogels with specific pore sizes, stiffness and gelation kinetics. In its unmodified form, PEG does not support cell adhesion; cells typically remain rounded in pure PEG hydrogels. However, incorporation of ECM motifs (RGD peptides or heparin to facilitate growth factor binding) enables PEG hydrogels to become bioactive (5). In cardiac tissue engineering, PEG-based hydrogels have been investigated as scaffolds for engineered heart valves and cell delivery. Crocini et al (42) employed photo-clickable thiol-ene PEG hydrogels for 3D cell culture of adult mouse cardiomyocytes. These PEG hydrogels serve as versatile and biocompatible scaffolding materials, with precisely adjustable stiffness that enables them to mimic both physiological and pathological microenvironmental conditions. Compared with conventional cell culture systems, adult cardiomyocytes encapsulated within PEG hydrogels exhibit prolonged survival and preserved the integrity of their sarcomeric and T-tubular structures (42). The mechanical properties of PEG hydrogels are tunable, allowing formulations that are very soft for injectable applications or more rigid and elastomeric for load-bearing cardiac patches (43). Moreover, PEG can be combined with natural polymers (for example, through PEGylation of proteins such as fibrin or collagen) to create composite hydrogels that retain bioactivity while improving structural stability (5).

PVA is a synthetic, water-soluble polymer that forms hydrogels, typically through repeated freeze-thaw cycles or chemical cross-linking (40). It produces transparent, rubber-like hydrogels with good mechanical integrity and was among the earliest hydrogel materials used in medicine. PVA hydrogels are biologically inert, and few cells adhere to them without modification, however, they are non-toxic and chemically stable (22). PVA is used in cardiac applications, including the fabrication of heart valve leaflet models and vascular graft coatings, because of its flexibility and mechanical strength. Recent studies have developed nanocomposite PVA hydrogels reinforced with nanoparticles or nanofibers to enhance mechanical performance (41,44). Although unmodified PVA lacks cell-instructive cues, it can be functionalized (with gelatin or RGD peptides) to improve cell-material interactions (5). Mannarino et al (44) developed a composite hydrogel consisting of PVA coated with poly(acrylic acid) (PAA), which was subjected to heat treatment to form a physically cross-linked network. The mechanical properties and swelling behavior of the hydrogel are modulated by adjusting the heat input and duration of thermal exposure, thereby increasing the cross-link density of the matrix. In an in vitro study using a simulated aging model over 162.6 days, the composite hydrogel retained its mechanical properties and surface functionality following accelerated aging. When evaluated in an in vitro blood loop model, the PVA/PAA composite hydrogel exhibited a 97% decrease in platelet adhesion (44). Additionally, it demonstrated a slower thrombosis-induced occlusion rate at the catheter tip, in contrast to commercially available catheter products. Ultimately, this thromboresistant hydrogel shows potential as a biomaterial for vascular access application, with the goal of improving patient outcomes. PVA hydrogels can be loaded with drugs or growth factors and have been investigated as drug-eluting coatings for stents or cardiac implants (45,46). Hu et al (47) developed an injectable hydrogel composed of phenylboronic acid-grafted carboxymethyl cellulose and PVA for the localized delivery of curcumin and recombinant human collagen type III to the infarcted myocardium in a rat MI model. This hydrogel improves cardiac function, increases left ventricular (LV) wall thickness, decreases infarct size, attenuates cardiomyocyte apoptosis and decreases inflammation.

Hydrogels based on PAA or poly(HEMA) are classical synthetic materials that have been widely used in contact lenses and biomedical implants. PAA is hydrophilic and pH-responsive, exhibiting increased swelling at higher pH due to deprotonation of its carboxyl groups; this property has been exploited in drug delivery systems (36). Although these polymers are not commonly used alone in cardiac tissue engineering because of their limited bioactivity, they demonstrate the concept of stimuli-responsive synthetic hydrogels. By copolymerizing acrylic monomers with other functional monomers, responsiveness to pH or temperature can be induced (12). Poly(N-isopropylacrylamide) (PNIPAM) is a synthetic polymer that undergoes a thermal phase transition at ~32°C. It is soluble below this temperature and forms a gel above it (13). PNIPAM-based hydrogels are used to develop injectable systems that solidify at body temperature, conceptually similar to the chitosan/β-glycerophosphate system but fully synthetic. One limitation of PNIPAM is that its degradation products, if the polymer is not fully cross-linked, exhibit limited biocompatibility; therefore, PNIPAM is often combined with biodegradable segments or formulated as microgels to minimize potential accumulation (13).

Certain synthetic hydrogels bridge the distinction between natural and synthetic. For example, peptide-based hydrogels are fabricated from short synthetic peptides that spontaneously self-assemble into fibrous networks (such as the RADARADARADARADA peptide forms a β-sheet hydrogel) (13). Although these materials are synthetic in origin, they are biological in composition because they consist of amino acids. They have been investigated for myocardial repair because they present a nanofibrous architecture resembling collagen and incorporate bioactive sequences (5). Polyphosphazenes, a class of inorganic-organic polymers, can be engineered to be biodegradable and have been evaluated as hydrogel matrices for tissue engineering. Polyurethane and polyester urethane hybrids have also been formulated as hydrogels, referred to as elastic hydrogels, to provide rubber-like elasticity for cardiac patches. Synthetic systems offers advantages, such as electrical conductivity when doped with conductive polymers or high toughness in double-network hydrogels composed of two interpenetrating polymer networks (12).

In general, synthetic hydrogels provide greater control over material properties and typically exhibit superior mechanical robustness compared with natural hydrogels. They do not inherently support cell attachment or tissue integration, however, this limitation can be addressed through composite design (blending synthetic and natural polymers) or by conjugating bioactive molecules (5). Numerous contemporary hydrogel systems are hybrid in composition. For example, a PEG backbone can be functionalized with cell-adhesive peptides or combined with collagen or chondroitin sulfate (48). Such composite hydrogels aim to integrate the biofunctionality of the native ECM with the mechanical tunability of synthetic polymers. An example is PEGylated fibrin or PEG-collagen composites, in which the synthetic component slows degradation and increases mechanical strength, whereas the natural component preserves cytocompatibility (49,50). Table I summarizes key differences between natural and synthetic hydrogel materials in cardiovascular therapy.

A trend in hydrogel research is the development of stimuli-responsive (smart) hydrogels that undergo physicochemical changes in response to specific environmental triggers (5,13,16). In cardiovascular therapy, dynamic behavior is desirable because it enables minimally invasive delivery, on-demand drug release and adaptation to the changing post-infarction microenvironment (11). For example, a hydrogel may remain in a liquid state during injection but undergo gelation within the warm myocardium, or a drug-loaded matrix may release its therapeutic payload when it detects elevated levels of inflammatory cytokines at an injury site. By incorporating stimuli-responsive components, hydrogels enable on-demand therapeutic delivery and adapt to patient-specific conditions in real time (5).

Thermo-responsive hydrogels exploit the temperature difference between room temperature during injection and body temperature during gelation, enabling catheter-based delivery with rapid in situ gelation. This temperature-triggered transition allows therapeutic action to be aligned with the acute to subacute phases of cardiac injury. PNIPAM is a thermo-responsive polymer that exhibits solution-to-gel (sol-gel) transition at ~32°C (13). By copolymerizing PNIPAM with other monomers, the transition temperature is adjusted to physiological temperature. Other approaches include blending polymers that undergo temperature-induced assembly, such as chitosan combined with β-glycerophosphate, which gels upon warming, or Pluronic F127, a PEG-poly(propylene oxide) block copolymer that is liquid at low temperature and forms a gel at ~37°C. Duan et al (51) prepared a thermosensitive injectable hydrogel (PFSgel) synthesized from Pluronic F127 and sodium alginate (Fig. 4A). The poloxamer-alginate blend exists as a liquid solution) at room temperature and undergoes gelation in vivo (Fig. 4C). When loaded with PEDF and injected around an atherosclerotic arterial lesion, it forms a local depot that gradually releases the therapeutic agent and decreases plaque neovascularization and progression in mice by inhibiting MMP2/MMP9 and CD31 signaling pathways (Fig. 4B).

A thermo-responsive PLEL triblock copolymer hydrogel has also been developed that undergoes rapid sol-gel transition at body temperature. When sprayed or injected onto the injured epicardial surface, it forms a conformal and fouling-resistant barrier that prevents postoperative cardiac adhesion during the critical early healing phase (52). Another thermo-responsive system based on Pluronic F127 has been engineered to encapsulate HA-modified zeolitic imidazolate framework-8 nanoparticles loaded with shikonin/Cu²⁺ complexes, enabling sustained drug retention at the myocardial injection site following ischemia-reperfusion injury (53).

Following MI, sustained ischemia and hypoxia lead to a decrease in local pH, creating a spatially and temporally defined acidic microenvironment that evolves with disease progression (6). pH-responsive hydrogels exploit this dynamic change to achieve targeted, stage-specific therapeutic delivery (3) PAA hydrogels exhibit increased swelling at neutral-to-basic pH because deprotonated carboxylate groups repel one another. In CVD, this pH-responsive behavior can be harnessed because ischemic tissue typically becomes acidic (3). Alimirzaei et al (54) reported the fabrication of two pH-sensitive chitosan hydrogels and evaluated them as cell carriers and cardiac scaffold materials for infarcted myocardial regeneration. An injectable pH-responsive micellar system has been developed by encapsulating a small-molecule glycogen synthase kinase-3β inhibitor [(2'Z,3'E)-6-bromoindirubin-3'-oxime(BIO) within a PEG-polycaprolactone (PCL) matrix and integrating it with adipose-derived stem cells in Matrigel (55). Under acidic conditions simulating the post-MI microenvironment, micelles undergo accelerated degradation, triggering BIO release and mitigating reactive oxygen species (ROS)-induced injury (55). A pH-responsive conductive hydrogel composed of PVA, oxidized sodium alginate, borax, tannic acid and a vanadium-based MXene (a class of 2D transition metal carbides/nitrides) nanozyme has also been engineered. Under acidic conditions that mimic the ischemic microenvironment, the hydrogel exhibits enhanced nicotinamide adenine dinucleotide (NADH) oxidation activity, enabling sustained release of the nanozyme and restoration of the NAD⁺/NADH redox balance (56).

MMPs, particularly MMP-2 and MMP-9, are notably upregulated in the infarcted myocardium during the inflammatory and early remodeling phases. Their expression peaks within days after injury and declines as healing progresses. This positions MMPs as ideal endogenous triggers for on-demand therapeutic delivery aligned with the dynamic requirements of cardiac repair (57). Researchers have designed hydrogels containing cross-links cleavable by MMP-2 or MMP-9, which are abundant proteases in post-MI tissue (13,57,58). Such MMP-responsive hydrogels degrade in synchrony with the remodeling process, potentially releasing embedded therapeutics, such as growth factors or antifibrotic agents, at sites where MMP activity is elevated (57,58). Pereira et al (59) employed a biscysteine MMP-cleavable peptide as a cross-linking agent and fabricated a pectin-based hydrogel via a photo-click reaction: The fabricated hydrogel degraded in the presence of type II collagenase, suggesting its potential for in vivo enzymatic degradation (59). An MMP-2-responsive cardiac ECM (c-ECM) hydrogel has also been engineered to enable sequential release of pro-angiogenic and vascular-stabilizing factors (60). A collagen-binding vascular endothelial growth factor (VEGF) is incorporated directly into the c-ECM matrix, whereas an angiopoietin-1 (Ang-1) mimetic peptide is conjugated through an MMP-2-cleavable linker (60). As MMP-2 levels progressively increase, the enzyme cleaves the peptide linker, releasing the mimetic peptide to promote vascular maturation. For heart valve applications, an MMP-responsive hydrogel system encapsulating microRNA-93 (miR-93)-loaded nanoparticles has been developed. Following MMP-mediated degradation, the nanoparticles are released and internalized by infiltrating macrophages, where miR-93 induces M2 polarization and promotes constructive remodeling of decellularized valve scaffolds (61).

Because the heart is an electrically active organ, electro-responsive hydrogels that alter their properties in response to an applied electric field represent an emerging area of research (5). Although at an early stage of development, conductive or piezoelectric hydrogel composites incorporating graphene, gold nanowires or conductive polymers such as polypyrrole have been investigated for cardiac patch applications to enhance electrical signal propagation across scar tissue (41,43). Light-responsive hydrogels, which rely on photodegradable linkers or light-induced conformational changes, are being explored for externally regulated drug release (5,62). An injectable hydrogel containing light-sensitive bonds and photodegradable components, composed of PEG and heparin-derived polymers, effectively encapsulates fibroblast growth factor-2 (FGF-2) (62). The biological activity of FGF-2 following encapsulation is comparable to that pre-encapsulation and the hydrogel modulates its release profile. Magnetic field-responsive hydrogels typically incorporate magnetic nanoparticles and can be manipulated or heated under an external magnetic field; such systems may enable targeted thrombosis therapy or localized hyperthermia for the ablation of atherosclerotic plaque cells (12).

To harness the strengths of natural and synthetic components, numerous composite hydrogels have been developed (13,17). In composite systems, a natural polymer, living cells or nanoparticles, may be embedded within a synthetic polymer matrix, or a synthetic component may be incorporated into a natural matrix (5). This strategy can produce materials that balance biofunctionality with mechanical robustness (12). For example, combining type I collagen with glycosaminoglycans or a synthetic polymer such as PEG can enhance mechanical integrity while preserving cell-supportive surfaces (16). Composite hydrogels are applied in heart valve engineering, where a biological scaffold, such as a decellularized valve matrix, may be reinforced with a secondary synthetic polymer network to improve durability (48). In vascular graft design, the luminal surface of synthetic grafts [expanded polytetrafluoroethylene (ePTFE) tubes) is coated with a thin hydrogel layer to create a more biomimetic and endothelial cell-supportive interface (63). One study reported the bonding of a hydrogel lining to the interior of an ePTFE graft, which promotes endothelialization without compromising mechanical strength (63). Similarly, nanocomposite hydrogels, in which inorganic nanoparticles or nanofibers are dispersed within the polymer network, have attracted increasing attention (64-66). Incorporation of silica nanoparticles, clay nanosheets or carbon-based nanomaterials reinforce the hydrogel matrix and, in some cases, introduce additional functionality such as electrical conductivity or growth factor binding (13). For cardiac tissue engineering, embedding electrically conductive nanomaterials, including graphene, gold nanowires or MXene nanosheets, within a polymeric gel can generate electroconductive hydrogels that facilitate electrical signal propagation and may improve synchronization with myocardial contraction (64). Consequently, the incorporation of conductive biomaterials is key for enhancing the physiological relevance of in vitro bioengineered cardiac tissue models. These composite systems aim to mitigate post-MI arrhythmias by promoting integration with the cardiac electrical conduction network (64-66).

Additional stimuli, such as oxidative stress characterized by elevated ROS levels in infarcted regions and alterations in mechanical strain associated with changes in loading conditions, have also been explored (67). Overall, stimuli-responsive hydrogels introduce an additional functional dimension by enabling spatiotemporal control of therapeutic activity. By integrating multiple components or incorporating distinct responsive mechanisms, these hydrogels concurrently satisfy the mechanical requirements of cardiac tissue and the biological requirements necessary to support regeneration (3,11).

MI, commonly referred to as a heart attack, results from occlusion of a coronary artery, leading to ischemia and necrosis of cardiac muscle tissue (68,69). The primary goals of post-MI therapy are to preserve cardiac function, promote myocardial regeneration and prevent adverse ventricular remodeling (3,6). Standard post-MI management, including reperfusion therapy and pharmacological agents such as ACE inhibitors and β-blockers, aims to limit infarct size and decrease cardiac workload, however, these interventions do not restore lost myocardium (7,31). Functional hydrogels have been investigated as a multifaceted strategy for cardiac repair, serving as structural scaffolds, therapeutic delivery platforms and supportive biomaterials that actively promote myocardial regeneration (5,11).

One of the most direct approaches involves injection of a hydrogel or hydrogel precursor into the infarcted myocardium, typically via intracardiac administration, either surgically or through catheter-based delivery (33). Following administration, the hydrogel forms a bulking scaffold within the scarred region. This intervention confers an immediate mechanical benefit by increasing ventricular wall thickness and decreasing wall stress on the remaining myocardium. Finite element modeling and preclinical animal studies have demonstrated that intramyocardial injection of biomaterials attenuates ventricular dilation and preserves physiological geometry, thereby improving cardiac pump function (3). For example, an injectable HA-based hydrogel with optimized stiffness provides mechanical support and results in reduced infarct expansion in treated hearts compared with controls (70). Beyond passive mechanical reinforcement, injectable hydrogels serve as 3D scaffolds that permit cell infiltration (13). Ideally, the material undergoes gradual degradation and is replaced by newly formed tissue, such as vascularized myocardium, rather than persistent fibrotic scar tissue (16).

A variety of materials have been evaluated as injectable cardiac hydrogels. Natural polymers, including collagen, fibrin and alginate, were among the earliest candidates investigated (11,17). Collagen injections in animal models improves cardiac function, potentially by providing a provisional matrix that facilitates cardiomyocyte alignment and limits increases in ventricular wall stress (28). Leveraging its capacity to form stable gels through ionic cross-linking, calcium-cross-linked alginate injected into rat and porcine MI models increases wall thickness and improves ejection fraction (EF) (3,34). These effects are attributed to a combination of mechanical reinforcement and pro-angiogenic bioactivity. This strategy has progressed to clinical evaluation as the IK-5001 hydrogel (BioLineRx) (33). In a first-in-human study, intracoronary administration of this alginate hydrogel shortly following MI was demonstrated to be feasible and safe (33). Another injectable hydrogel, VentriGel (Ventrix, Inc), is derived from decellularized porcine c-ECM (71). It is supplied as a lyophilized cardiac matrix powder that is reconstituted and administered as a liquid, followed by in situ gelation within the myocardium. VentriGel retains native cardiac ECM biochemical cues and, in preclinical studies, recruits endogenous cells and promotes novel myocardial tissue formation (19,71). A Phase I clinical trial evaluating catheter-based delivery of VentriGel in patients following MI demonstrated safety and preliminary signs of functional improvement, including increased exercise capacity and trends toward improved cardiac function, particularly in patients treated >1 year after MI (71). These findings represent important translational milestones, demonstrating that catheter-based hydrogel therapy is feasible and safe in human patients.

A notable application of hydrogels in MI therapy is their use as cell delivery vehicles. Numerous cell types, including mesenchymal stem cells (MSCs), adipose-derived SCs, bone marrow-derived mononuclear cells, induced pluripotent SC (iPSC)-derived cardiomyocytes and cardiac progenitor cells, have been investigated for cardiac regeneration (3). Delivery of these cells within a hydrogel enhances their retention and survival in the myocardium compared with injection in saline alone (20). The hostile post-MI microenvironment, characterized by inflammation, oxidative stress, and ECM disruption, causes the majority of injected cells to undergo apoptosis or be rapidly cleared (31). Hydrogels provide a protective and adhesive microenvironment that promotes cell retention and supports cell viability. For example, a chitosan-based hydrogel conjugated with survival-promoting peptides increases the engraftment of transplanted cardiomyocytes and enhances neovascularization in the infarcted region (72). Similarly, fibrin-based patches loaded with SCs demonstrated higher cell survival rates and greater improvements in cardiac function compared with cell transplantation alone (24). Hydrogels can be pre-seeded with cells and implanted as cardiac patches. In this approach, cells are cultured within or on the hydrogel scaffold in vitro, typically under dynamic mechanical stimulation to promote tissue maturation, and the resulting cell-laden construct is surgically applied to the infarcted myocardium. Such tissue-engineered cardiac patches, fabricated using scaffolds composed of collagen, fibrin or synthetic polymers, have demonstrated partial remuscularization of injured myocardium in animal models (64).

A study (72) reported the development of a self-healing, electrically conductive injectable hydrogel based on chitosan-graft-aniline tetramer for cardiac therapy. Owing to its decreased viscosity under shear stress, allowing easy injection) and self-healing properties, this hydrogel enables minimally invasive administration and provides a protective niche for transplanted cells. In a rat model, it exhibits favorable injectability, controllable cell release and complete biodegradation within 45 days, underscoring its potential as a cell delivery platform for cardiovascular therapy (72).

Hydrogels serve as versatile drug delivery platforms in the infarcted myocardium. Instead of systemic administration, therapeutic agents are incorporated into a hydrogel and delivered locally to the injury site, thereby achieving high local concentrations while minimizing systemic adverse effects (3,5). A range of cargos has been investigated, including small-molecule agents such as anti-inflammatory and antioxidant compounds, cytokines and growth factors, nucleic acids and extracellular vesicles (EVs) (11). The structural characteristics of the hydrogel regulate release kinetics, either through diffusion-driven release mediated by water uptake and swelling or through degradation-triggered release mechanisms (57).

Delivery of pro-angiogenic factors to promote neovascularization within the infarcted region is key for supporting regenerating myocardium. VEGF and basic FGF (bFGF) are well-established angiogenic proteins with short biological half-lives when administered in free form. Encapsulation within a hydrogel enables sustained release over days to weeks (73). In a representative study, a fibrin-based hydrogel patch releasing VEGF was applied to infarcted rat hearts (51). Over a 4 week period, it increases capillary density in the infarct border zone and improves cardiac function compared with controls. The hydrogel ensures prolonged local retention of VEGF, allowing it to exert therapeutic effects during the key healing phase. Similarly, dual delivery of an angiogenic peptide, N-acetyl-serylaspartyl-lysyl-proline (Ac-SDKP) and the chemokine stromal cell-derived factor-1α (SDF-1α) from an injectable hydrogel enhances angiogenesis and promotes SC recruitment to the infarcted myocardium, resulting in decreased infarct size and improved LV function in a mouse MI model (73,74).

Another strategy involves modulation of inflammation and prevention of cardiomyocyte death during the acute phase of MI. Immediately following infarction, waves of inflammatory cells, including neutrophils and macrophages, infiltrate the injured myocardium and may cause collateral tissue damage through oxidative stress and protease release (64). Hydrogels have been loaded with anti-inflammatory agents, such as IL-10, glucocorticoids or non-steroidal anti-inflammatory drugs (NSAIDs), to locally attenuate this inflammatory response. For example, an injectable peptide hydrogel carrying an NSAID and iron oxide nanoparticles has been evaluated for targeting macrophages in atherosclerotic cardiac tissue, demonstrating decreased inflammatory markers and promotion of a reparative macrophage phenotype (75). In MI, a hydrogel-based platform may similarly deliver immunomodulatory agents to promote macrophage polarization from a pro-inflammatory M1 to a pro-reparative M2 phenotype, thereby limiting secondary tissue injury. One approach incorporates microparticles releasing IL-10 within a cardiac patch, resulting in decreased fibrosis and improved cardiac function in a rat MI model (34).

Hydrogels encapsulate antifibrotic or anti-remodeling agents. During post-infarction healing, excessive fibrosis and ventricular dilation contribute to progressive cardiac dysfunction (7). The peptide Ac-SDKP inhibits cardiac fibroblast activation. Firoozi et al (76) synthesized a RADA-SDKP hydrogel composed of a self-assembling gel-forming core sequence (RADA) and a bioactive motif (SDKP), providing structural flexibility through dynamic self-assembly to facilitate injectability. Following transplantation of the hydrogel, LVEF is improved. This functional improvement is associated with a marked reduction in fibrotic tissue deposition, enhanced microvascular density and attenuation of the inflammatory response within the infarcted region (76). Similarly, tissue inhibitors of metalloproteinases have been delivered via MMP-responsive hydrogels to locally counteract MMP-mediated ECM degradation that contributes to ventricular wall thinning (77).

Timing of therapeutic delivery can be modulated through hydrogel design. Physiological requirements of the heart evolve over time following MI, progressing through inflammatory, proliferative and remodeling phases (64). Hydrogels provide the capacity for temporally controlled release profiles (for example, immediate release of an anti-apoptotic agent to preserve jeopardized cardiomyocytes, followed by delayed release of a growth factor to stimulate regenerative processes). Multiphase hydrogel systems have been engineered to achieve sequential delivery (62). For example, a two-component hydrogel system has been developed in which one component rapidly releases SDF-1α to recruit progenitor cells, while the second component enables sustained release of VEGF to promote neovascularization (73). This staged delivery approach enhances cardiac repair in a mouse MI model.

Beyond proteins and cells, hydrogels are explored as platforms for localized gene therapy in the myocardium (78). Plasmid DNA, mRNA or viral vectors encoding regenerative factors can be incorporated into hydrogel matrices to enable sustained and site-specific gene delivery (21,78). An injectable biocompatible hydrogel capable of facilitating efficient delivery of a nanocomplex composed of graphene oxide (GO) and the pro-angiogenic VEGF-165 gene demonstrates therapeutic potential for myocardial repair (Fig. 5A) (78). Exosomes (nanoscale EVs secreted by cells that contain miRs and proteins) have attracted attention as cell-free therapeutic agents (79-81). Exosomes attenuate apoptosis and promote cardiac repair in a manner comparable with their parent cells (79-81). However, freely injected exosomes are rapidly cleared from the myocardium. Encapsulation within hydrogels prolongs their retention at the injury site (82). Improved retention of exosomes in injured myocardium has been demonstrated using an injectable shear-thinning gel (STG) loaded with EVs (82) (Fig. 5B-D). The STG shows even distribution of EVs within the gel matrix (Fig. 5D) and sustained EV release over 21 days (Fig. 5E), which collectively enhances myocardial function, improves hemodynamic parameters and increases neovascularization (82). Furthermore, exosomes derived from human adipose-derived SCs, gelatin and laponite are combined to formulate a shear-thinning nanocomposite hydrogel (nSi Gel) as an injectable carrier of the stem cell secretome, which increases vascular density surrounding the myocardium, improves cardiac function and decreases scar formation (83). These findings demonstrate how hydrogels function as reservoirs for advanced biological agents, such as nucleic acids or EVs, ensuring localized therapeutic activity rather than rapid systemic dispersion.

In addition to injectable systems, hydrogels have been developed in patch form for application to the epicardium, the outer surface of the heart, overlying the infarcted region (84). These patches may be acellular or seeded with therapeutic cells. The underlying concept is to provide a scaffold that covers the damaged myocardium, promotes tissue regeneration from the epicardial surface inward and serves as a mechanical restraint to limit adverse ventricular dilation (5). Epicardial patches have been fabricated from synthetic polymers, such as degradable polyester meshes embedded within hydrogel matrices, as well as from naturally derived ECM sheets (17,64). Patch-based approaches cover a broader area and accommodate larger numbers of cells or higher drug loads, however, they typically require surgical implantation. Consequently, minimally invasive strategies, including thoracoscopic placement or catheter-based attachment techniques, are under investigation (5). In addition, epicardial priming through placement of bioactive materials on the cardiac surface has been explored to stimulate recruitment of endogenous progenitor cells from the pericardial space (64). Hydrogels that can be applied by painting or spraying onto the epicardium are in development, with some formulations incorporating mucoadhesive properties to enhance adhesion to the moist cardiac surface (52). Furthermore, hydrogel-based patches can be prevascularized or engineered with microchannel architectures to facilitate oxygen and nutrient diffusion, supporting cell survival following implantation. A 3D bioprinted cardiac patch composed of a PEG hydrogel has been fabricated with an internal microchannel network (64). When loaded with MSCs and implanted in a mouse MI model, the microchannel-containing patch results in greater vascular ingrowth and superior functional recovery compared with a solid patch, underscoring the importance of internal architecture (34). Wang et al (85) established a rabbit MI model and administered an epicardial injection of a hydrogel incorporating bone marrow-derived mesenchymal stem cells (BMSCs). Compared with control groups, retention of BMSC grafts within the infarcted myocardium is prolonged, an effect attributed to the properties of the α-cyclodextrin (CD)/PEG-b-polycaprolactone-(dodecanedioic acid)-polycaprolactone-PEG (MPEG-PCL-MPEG) hydrogel system. The transplanted BMSCs contribute to restoration of cardiac function, as evidenced by increased vascular density in the infarct border zone (85). Although the long-term therapeutic efficacy of this approach requires further investigation, the α-CD/MPEG-PCL-MPEG hydrogel, characterized by its cytoprotective properties, improves LV function and attenuates adverse LV remodeling (85).

Vascular diseases, particularly atherosclerosis affecting arterial vessels, contribute to life-threatening conditions such as coronary artery and peripheral arterial disease and stroke (1,2). Conventional treatments include bypass graft surgery using either autologous vessels or synthetic grafts and percutaneous angioplasty with stent implantation to restore luminal patency in occluded arteries (5). Although these interventions are typically effective in the short term, they present notable limitations. Synthetic small-diameter grafts (for arteries <6 mm) exhibit high failure rates due to thrombosis or intimal hyperplasia, defined as excessive neointimal tissue formation (18). Stents may undergo restenosis or induce chronic inflammatory responses. Functional hydrogels offer strategies for engineering artificial blood vessels, enhancing integration and healing of vascular implants and treating vascular pathologies such as atherosclerosis (45,63).

The development of TEVGs has long been an objective, particularly for small-diameter arteries (≤6 mm), such as coronary or peripheral vessels (86). Conventional synthetic grafts, typically composed of PET or ePTFE, perform adequately in large-diameter applications but demonstrate high failure rates in small-caliber vessels due to thrombosis and intimal hyperplasia (86). Clinical outcomes for large-diameter grafts (>6 mm) have been satisfactory, with reported patency rates of up to 95% at 5 years (87). By contrast, the aforementioned study reported a patency rate of 30% for small-diameter grafts (<6 mm) (87). Additional investigations have documented patency rates ranging from 0 to 25% in canine and rabbit models after implantation periods of weeks to months (88-90).

Hydrogels serve as scaffold materials capable of forming the structural matrix of bioengineered vascular grafts (16). One strategy involves fabrication of a tubular hydrogel scaffold (optionally reinforced with supportive materials) that can be seeded with vascular cells and matured into a functional vessel. Natural hydrogels, such as collagen or fibrin, have been molded into tubular configurations for this purpose (24,46). These materials provide a favorable microenvironment for smooth muscle and endothelial cells to deposit ECM components and organize into a vessel wall. However, hydrogel-only tubes typically exhibit insufficient mechanical strength, particularly in terms of burst pressure, during the early implantation phase (63). To address these limitations, strategies have been developed in cardiovascular tissue engineering, including the use of diverse scaffold types, such as natural and synthetic polymers (91-93), as well as decellularized xenografts and homografts (94-96), in combination with different cell sources, including valvular interstitial cells (97,98), BMSCs (99) and progenitor cells obtained from peripheral blood or amniotic fluid (100-102). In addition, 3D bioprinting techniques have been employed to fabricate vessel-like constructs using hydrogel-based bioinks laden with cells, enabling the creation of defined multilayered wall architectures (45). Another strategy involves cell sheet engineering, in which cell monolayers cultured on temperature-responsive hydrogel surfaces are detached as intact sheets and wrapped into tubular structures. Although the final graft does not contain a hydrogel component, this approach relies on PNIPAM, a thermoresponsive hydrogel, to enable enzyme-free harvesting of cell sheets (103). These cell sheet-based grafts have demonstrated promising outcomes in small animal models (45,85,103).

Hydrogels serve a more direct role in fully synthetic TEVGs through the development of hybrid graft materials. One design involves electrospinning a blend of biodegradable polymer fibers with an incorporated hydrogel component to enhance graft compliance and facilitate cell infiltration (104). Another approach involves chemically bonding a hydrogel layer to the luminal surface of a porous synthetic graft, thereby creating a bioactive interface that promotes endothelialization (63). This inner hydrogel layer can be functionalized with bioactive agents such as heparin to reduce thrombogenicity, or VEGF to stimulate endothelial cell recruitment. In a study by Chen et al (105), a small-diameter polyethylene terephthalate graft was coated on its luminal surface with a heparin-containing hydrogel. In a rabbit carotid artery implantation model, this modification results in decreased thrombosis and more rapid formation of a functional endothelial lining compared with uncoated grafts (105). A key objective in TEVG development is to promote host-mediated remodeling and repopulation of the graft with autologous cells, effectively transforming the construct into a living vascular tissue (86). Hydrogels, owing to their cytocompatible and bioactive properties, are well-suited to support this process. For example, a bioinspired small-diameter vascular graft incorporated a hydrogel derived from decellularized arterial matrix on the luminal surface, providing biochemical cues that accelerate endothelial coverage following implantation (106). Establishment of a confluent endothelium is key for long-term graft patency, as it suppresses thrombosis and pathological smooth muscle cell proliferation (86).

In vascular intervention, such as percutaneous angioplasty with stent implantation, a challenge is the host response to the implanted stent (107). Drug-eluting stents incorporate antiproliferative agents (sirolimus or paclitaxel) within a polymer coating to suppress excessive smooth muscle cell proliferation and thereby prevent restenosis, however, these agents may also delay endothelial healing (107). Ideally, a stent should inhibit neointimal hyperplasia and rapidly achieve coverage by a functional endothelial layer, thereby mimicking the behavior of a native artery and decreasing the risk of thrombosis. Hydrogel-based coatings are being investigated to create pro-healing stent surfaces. One approach involves nitric oxide (NO)-releasing hydrogel coatings applied to stents (105). A tough NO-eluting (NOE) hydrogel coating has been designed for sustained NO release (Fig. 6A). Following stent deployment in porcine coronary arteries (Fig. 6B), digital subtraction angiography shows that the NOE coating prevents severe restenosis observed in control stents (Fig. 6C; yellow arrow). Histological analysis (Fig. 6D) and luminal surface imaging at 2 weeks and 3 months (Fig. 6E) confirms that the NOE coating decreases neointimal hyperplasia and thrombosis. Confocal microscopy (Fig. 6F) reveals rapid and complete endothelialization (CD31⁺ cells) on the NOE-coated stents (105). Another strategy employs antibody-functionalized hydrogels, such as those conjugated with anti-CD34 antibodies, to capture circulating endothelial progenitor cells and accelerate endothelialization (106). Zwitterionic hydrogels, characterized by resistance to non-specific protein adsorption, have also been utilized as coatings to reduce platelet adhesion while selectively promoting endothelial cell attachment (104). In one study, a zwitterionic hydrogel loaded with VEGF was applied to a heart valve prosthesis to enhance endothelialization and decrease calcification; this may also be applicable to vascular stents (108). For bioresorbable stents composed of biodegradable metals, such as magnesium, or biodegradable polymers, hydrogel coatings modulate degradation kinetics and improve the tissue-device interface. In one study, magnesium alloy stents were coated with a fucoidan-containing hydrogel, which enhanced hemocompatibility and promoted endothelial cell colonization during gradual metal degradation (109). Fucoidan, a sulfated glycosaminoglycan with anticoagulant properties, contributes to localized inhibition of thrombosis and inflammation (109).

Atherosclerosis, characterized by the accumulation of lipid-rich plaques within arterial walls, underlies numerous cardiovascular events, including MI and stroke (1,2). Although risk factor modification and systemic pharmacotherapy, such as statins and anti-inflammatory agents, remain the cornerstone of management, attention has been directed toward localized therapeutic strategies aimed at stabilizing or regressing high-risk plaques (75). Hydrogels serve as promising platforms in this context, particularly as injectable depots that can be delivered to vulnerable plaque sites for localized drug release (110).

One strategy targets intraplaque neovascularization, a hallmark of advanced atherosclerotic lesions. Fragile microvessels that develop within plaques are prone to rupture, leading to intraplaque hemorrhage and contributing to plaque instability. In a recent study, a thermosensitive hydrogel composed of poloxamer 407 and alginate was employed to deliver PEDF, an endogenous inhibitor of pathological angiogenesis, to atherosclerotic plaques (51). The formulation remained in a liquid state during catheter-based administration and gelled at physiological temperature, forming a localized drug reservoir at the plaque site. Sustained release of PEDF decreases intraplaque neovessel density and plaque progression in a mouse model of atherosclerosis (51). Treated plaques exhibited smaller lipid cores and thicker, more stable fibrous caps, indicating enhanced lesion stabilization. This approach demonstrates how hydrogel-based delivery systems can localize therapeutics that would lack plaque specificity if administered systemically (110).

Hydrogels have been investigated as vehicles for delivering agents that modulate plaque-associated inflammation (75). For example, an injectable peptide hydrogel incorporating an anti-inflammatory agent (naproxen) and a ferrofluid for magnetic targeting was developed to treat atherosclerotic plaques by promoting macrophage polarization toward a reparative phenotype. In principle, a hydrogel could be directed to the lesion site using an external magnetic field via embedded Fe₃O₄ nanoparticles and provide sustained local drug release (75). Similarly, another group developed a filamentous nanofiber hydrogel depot for prolonged delivery of anti-inflammatory nanoparticles, resulting in sustained suppression of inflammatory markers in an experimental model of atherosclerosis (110). Gene and RNA-based therapeutic delivery to plaques using hydrogel systems has also been explored (66,78). For example, administration of small interfering RNA (siRNA) or miR targeting pro-atherogenic genes in macrophages or vascular smooth muscle cells may modulate plaque progression (78). Hydrogels shield such nucleic acids from enzymatic degradation and enhance their local retention within the plaque microenvironment (62,78). Preliminary studies employing DNA-loaded hydrogels to modulate macrophage signaling pathways in arterial tissue have demonstrated decreased pro-inflammatory cytokine expression (78,110,111). Although these strategies remain largely at the preclinical stage, they represent a paradigm shift toward localized resolution of plaque pathology. While intramural hydrogel injection into the arterial wall may appear invasive, such delivery could be integrated into existing angioplasty procedures or implemented via specialized catheter systems. This localized therapeutic approach has the potential to complement systemic pharmacotherapy by targeting high-risk plaques individually (107,110).

Heart valve disease typically necessitates valve replacement (112), with conventional options including mechanical and bioprosthetic valves (Fig. 7A and B). To the best of our knowledge, however, all currently available prosthetic devices are associated with limitations that increase the risk of morbidity and mortality. For example, mechanical valves are associated with thromboembolic events and hemorrhagic complications, thereby requiring lifelong anticoagulation therapy (113). Bioprosthetic valves demonstrate limited long-term durability due to structural degeneration, calcification and fibrotic remodeling and they may also provoke immunogenic responses (114-116). To the best of our knowledge, no existing valve replacement strategy fully overcomes these limitations, rendering tissue-engineered heart valves (TEHVs) (Fig. 7C) an appealing therapeutic alternative for affected patients (117-121).

Hydrogels are attractive materials in heart valve engineering owing to their high water content, biocompatibility and ability to mimic the compliant connective tissue of native valve leaflets, and they can be seeded with cells capable of remodeling the construct (122-124). Native valve leaflets exhibit a trilayered architecture, comprising collagen fibers that provide tensile strength, proteoglycans that confer shock absorption and elastin that enables elastic recoil (115). Pure hydrogels, such as collagen or fibrin, replicate certain characteristics (for example, softness and pliability) but may lack sufficient mechanical strength to withstand repetitive cyclic loading during valve opening and closing (24). Consequently, hydrogel application in TEHVs typically involves composite designs or the use of hydrogels as cell carriers within a mechanically robust framework (15). One strategy employs a decellularized valve matrix as the primary scaffold, which preserves native architecture and mechanical integrity, followed by infusion or coating with a cell-or growth factor-laden hydrogel (5). The hydrogel facilitates reseeding of the decellularized scaffold with autologous endothelial and valvular interstitial cells and can fill microstructural defects. Hydrogels such as gelatin methacrylate or HA-based formulations have been used to impregnate decellularized porcine valves, generating hybrid constructs that support cellular infiltration while maintaining cusp flexibility (25). Another research direction focuses on fabricating synthetic or hybrid valve scaffolds incorporating hydrogels. Li et al (125) developed a PEG/poly(sulfobetaine methacrylate)] dual-network hydrogel coating on decellularized heart valves (Fig. 8A). Following implantation, the hydrogel coating provides sufficient flexural stiffness and fatigue resistance to withstand mechanical cycling, while facilitating endothelial cell coverage and ECM remodeling (Fig. 8B) (125). Hydrogel coatings mitigate calcification of bioprosthetic valves. For example, a zwitterionic hydrogel coating loaded with heparin and VEGF demonstrates reduced calcific deposition and enhanced endothelialization in a rabbit animal model (106).

Four-dimensional (4D) bioprinting, where printed constructs are designed to undergo controlled structural changes over time, may enable the fabrication of hydrogel-based valves that mature in response to physiological forces (123,124). A valve could be printed in a predefined geometry with cells embedded within a hydrogel matrix; as the cells synthesize ECM components and the hydrogel degrades, the construct may transition into a mechanically functional configuration shaped by bioreactor-mediated conditioning (121). Conditioning engineered valves in a pulsatile flow bioreactor enhances collagen alignment and improves mechanical strength (15). Hydrogels serve a key role in these strategies by providing an initial structural template that is progressively replaced by newly formed tissue.

Hydrogels promote vascularization through coordinated biochemical signaling and mechanotransductive pathways (126,127). VEGF-loaded HA nanobubble/Pluronic F127 diacrylate) hydrogels activate PI3K/Akt signaling, thereby enhancing endothelial cell migration and proliferation (126). Spatially heterogeneous platelet-rich plasma (PRP) fibrin hydrogels direct 3D vasculogenesis via the integrin β1/phosphorylated focal adhesion kinase/phosphorylated myosin light chain pathway and stabilize vascular endothelial-cadherin/β-catenin junctions (127). UCL-TRO-1938, a PI3Kα activator delivered from alginate hydrogels, stimulates angiogenesis in a temporally controlled manner during the acute ischemic phase (128). Moreover, c-ECM hydrogels enable sequential release of VEGF and Ang-1 mimetic peptides in response to MMP-2, thereby promoting the formation of mature and functional blood vessels (60).

Hydrogels reshape the cardiac immune microenvironment through modulation of conserved anti-inflammatory and antioxidant signaling pathways. MXene-PVA hydrogels activate AMPK/Nrf2/Keap1 signaling to enhance ROS clearance and suppress NF-κB activity, while promoting macrophage polarization toward a reparative M2 phenotype (129). Similarly, ROS-responsive adhesive hydrogels releasing andrographolide derivatives, NLRP3 inhibitor-loaded folic acid-HA-mesoporous silica nanoparticles and gelatin methacrylate/gelatin norbornene hydrogel patches promote M2 polarization independent of exogenous cytokine supplementation (130,131). MMP-responsive hydrogels enable on-demand IL-4 plasmid delivery to suppress post-MI inflammation, whereas p38-responsive supramolecular hydrogels inhibit p38 mitogen-activated protein kinase signaling to attenuate oxidative stress and inflammatory responses (132,133). Furthermore, thermosensitive hydrogels delivering thiamet-G enhance M2 polarization via signal transducer and activator of transcription 6 O-GlcNAcylation (134). Annexin A1-loaded alginate hydrogels modulate macrophage polarization through the AMPK-mTOR axis (135), while exosome-laden hydrogels regulate PI3K/Akt/mTOR, Hippo, TGF-β, HIF-1 and FoxO signaling pathways to facilitate cardiac repair (136). Collectively, these findings underscore the capacity of hydrogels to modulate the cardiac immune microenvironment through multiple convergent molecular mechanisms.

Hydrogels mitigate maladaptive cardiac fibrosis through multiple complementary mechanisms. Co-delivery of the MMP-2 inhibitor CTT (a cyclic peptide) and bFGF inhibits fibroblast-to-myofibroblast differentiation (137). ROS-responsive hydrogels releasing anti-IL-11 antibodies reduce scar thickness and border zone expansion by suppressing fibroblast activation (138). Furthermore, injectable alginate composite hydrogels enabling pH-responsive release of bone morphogenetic protein-9 effectively suppress TGF-β mediated fibroblast activation and myocardial fibrosis by antagonizing maladaptive ECM deposition (128). Similarly, a ROS-responsive hydrogel system (superoxide inhibitor, and FT011/liposome-quercetin-thioketal ROS-responsive linker; S1&FT/Lipo-QCFT) sequentially releases the superoxide production inhibitor S1QEL1.1 and tannic acid to attenuate early oxidative injury and promote M2 macrophage polarization, followed by sustained release of the anti-fibrotic agent FT011 to counteract TGF-β1-induced fibrosis and prevent adverse ventricular remodeling (139). Notably, injectable conductive hydrogel patches (GRAXe, a gelatin-base MXene-containing ROS-responsive conductive patch with asymmetric adhesion) integrate ROS-responsive degradation, MXene-mediated electrical conductivity and asymmetric tissue adhesion to restore electromechanical coupling while decreasing fibrotic remodeling (140).

Conductive hydrogels restore cardiac electrical signal propagation by re-establishing electrophysiological continuity, modulating connexin 43 (Cx43)-dependent gap junctions, stabilizing ion channel function and decreasing electrical heterogeneity across fibrotic infarcted tissue (5). Studies have demonstrated that polypyrrole-chitosan conductive hydrogels bridge electrically isolated cardiomyocyte clusters, synchronize calcium transients and improve conduction velocity across cryoinfarcted myocardium (141-147). In vivo administration shortens QRS duration and restores longitudinal conduction, whereas ex vivo scar tissue exhibits a three- to four-fold increase in electrical signal amplitude compared with controls, confirming restoration of conductive pathways (141). Injectable conductive hydrogels composed of polypyrrole/PAA and loaded with cardiomyocyte-derived exosomes enhance electrical coupling by upregulating Cx43 expression and improving sarcomeric organization, while concurrently modulating SIRT1/p53 signaling to attenuate oxidative stress and electrical instability (142). In vivo, these hydrogels normalize action potential and calcium transient duration and decrease arrhythmia inducibility and electrical resistivity within infarcted tissue. MXene-based conductive hydrogel patches with intrinsic ROS-scavenging capacity and electroconductive properties increase Cx43 expression and enhance electrical signal amplitude in the infarct zone, thereby synchronizing cardiomyocyte contraction through restoration of gap junction communication (143). Their conductivity approximates that of native myocardium, enabling physiological electroconductive restoration. Injectable GO-based conductive hydrogels capable of sustained NO release also provide tissue-matched electrical conductivity. These systems preserve electrical signal propagation, enhance endothelial network formation and maintain electrophysiological stability by reducing ROS-mediated electrical disruption (144). Printable PVA-glutaraldehyde/PEDOT [poly(3,4-ethylenedioxythiophene) (PEDOT) conductive patches efficiently transmit electrical signals to adult cardiomyocytes without introducing notable conduction resistance, demonstrating safe in vivo performance and preservation of electrocardiographic parameters (145). Ionic-conductive PEDOT/poly(ionic liquid) hydrogels enhance Cx43 and α-actinin expression, thereby restoring intercellular electrical coupling and decreasing infarct-associated electrical heterogeneity (146). Aniline tetramer-grafted conductive supramolecular hydrogels preserve Cx43 expression in the infarct border zone, maintaining gap junction integrity, improving electrical communication and attenuating adverse ventricular remodeling (147).

ECM-derived hydrogels represent one of the most promising platforms owing to their intrinsic bioactivity. A principal advantage of ECM-derived hydrogels is their capacity to modulate the host inflammatory response and promote tissue regeneration, rather than merely providing passive mechanical support (71). However, their limitations include batch-to-batch variability inherent to biologically sourced materials, low mechanical strength compared with synthetic counterparts and regulatory complexity associated with xenogeneic products (60,71). Alginate-based hydrogels exhibit good biocompatibility and have been extensively investigated for mechanical reinforcement of the infarcted ventricle (3,8). Preclinical studies have demonstrated increased wall thickness, decreased wall stress, and improved cardiac function (33,35). Nevertheless, a major limitation of alginate hydrogels is their lack of intrinsic bioactivity. As non-degradable or slowly degrading materials, they primarily provide mechanical support without actively stimulating tissue regeneration (33,35). Fibrin-based hydrogels have been employed as cell delivery vehicles because of their rapid gelation, biocompatibility and capacity to support cell infiltration and ECM deposition (24). Preclinical studies have demonstrated enhanced cell engraftment and angiogenesis (24,30,31). However, rapid degradation and limited mechanical strength restrict their application as standalone scaffolds in load-bearing cardiac environments (31). Collagen and gelatin-based hydrogels support cardiomyocyte function and have demonstrated efficacy in decreasing infarct size and improving contractility in preclinical models (11,28,70). However, their low stiffness and rapid enzymatic degradation necessitate chemical crosslinking or composite formulation to achieve adequate mechanical stability, which may compromise biocompatibility (5,28). HA hydrogels exhibit immunomodulatory properties, including promotion of M2 macrophage polarization and angiogenesis (32). Although promising for myocardial repair, unmodified HA is mechanically weak and structurally fragile, requiring chemical modification to enhance mechanical stability, thereby increasing manufacturing complexity (32). Synthetic hydrogels offer tunability and reproducibility, with precisely controllable mechanical properties, degradation kinetics and network architectures. Despite these advantages, most synthetic hydrogel systems remain confined to preclinical development (13). Their intrinsic bioinertness prevents effective cell recognition and adhesion, thereby limiting tissue integration and remodeling unless biofunctionalized with adhesive ligands, such as RGD peptides, or incorporated growth factors (41,45). Conductive composite hydrogels incorporating graphene, gold nanowires or MXene nanosheets represent an emerging frontier in cardiac applications, with the objective of enhancing electrical signal propagation across scar tissue and decreasing arrhythmogenic susceptibility (141,143). Preclinical studies have demonstrated improved electrical integration and synchronized myocardial contraction (141,143). However, concerns regarding nanomaterial-associated cytotoxicity, long-term biocompatibility and fabrication complexity constrain their clinical translation.

The divergent outcomes observed across hydrogel platforms suggest that optimal cardiovascular therapy may require integration of multiple functional attributes. This informs future research priorities and underscores that no single platform currently satisfies all therapeutic requirements in cardiovascular medicine. To facilitate direct quantitative and qualitative comparison, the key efficacy endpoints and safety outcomes of major hydrogel platforms are summarized in Table II.

Discrepancies between preclinical and clinical outcomes arise from multiple contributing factors. In addition to efficacy considerations, several critical challenges must be addressed to ensure safe and effective clinical translation, including immune response, biodegradation control, long-term mechanical stability, thrombogenicity and scalability (5,12,148).

Comprehensive biocompatibility evaluation in animal models and in vitro systems is essential prior to clinical application (148-150). Hydrogels derived from xenogeneic sources carry a theoretical risk of immunogenicity. Decellularization decreases this risk by removing cell components and debris, thereby promoting constructive tissue remodeling rather than immune rejection (151). VentriGel demonstrates no serious immune-associated adverse events, supporting the biocompatibility of decellularized ECM products &1). Nevertheless, each newly developed material must undergo rigorous immunotoxicological assessment. Strategies include incorporation of immunomodulatory cues within hydrogel matrices to promote immune tolerance, as well as the use of autologous or allogeneic sources to eliminate xenogeneic-associated risks (132,148). Synthetic hydrogels may elicit inflammatory responses if degradation yields irritant byproducts, particularly in the case of certain polyester-based systems [such as poly(lactic acid), poly(glycolic acid) and poly(lactic-co-glycolic acid)] (13). Ensuring degradation products are biologically inert and readily metabolized or excreted represents a key design consideration. This challenge may be mitigated by optimizing degradation kinetics and incorporating buffering components, however, the balance between material persistence and degradation must be tailored to each specific application.

Control of biodegradation is key to balance mechanical support and tissue remodeling in cardio-vascular hydrogels. Premature degradation leads to early loss of structural support before myocardial repair is complete, whereas slow degradation impedes native tissue ingrowth, disrupts electromechanical integration and increases the risk of chronic inflammation or foreign-body reactions. Therefore, degradation kinetics must be tailored to align with the repair timeline of infarcted myocardium (13,16). Recent advances (128,152) have enabled programmable degradation through modulation of crosslinking density, incorporation of stimuli-responsive cleavable bonds, or design of composite networks with controlled erosion profiles. Increased cross-linking density typically slows degradation by limiting water penetration and enzymatic accessibility, thereby extending hydrogel persistence (152). Degradation byproducts must be chemically inert, non-toxic and readily metabolized or excreted to prevent chronic inflammation, local acidosis or tissue injury. Consequently, contemporary cardiac hydrogels increasingly favor natural polymers or synthetic systems engineered to yield neutral, biocompatible metabolites (13). Recent studies (138,139) emphasize the importance of dynamic mechanical matching during degradation. Gradual decrease in hydrogel stiffness should parallel recovery of native tissue mechanics to prevent pathological loading or functional decline. Collectively, rational control of degradation rate, mechanism and byproduct biocompatibility is key for the safe and effective application of hydrogel-based cardiac therapies (12,152).

The beating heart and pulsatile vasculature impose cyclic mechanical loading that may result in material fatigue, fracture or functional deterioration. Hydrogels must therefore exhibit not only appropriate initial stiffness but also sufficient fatigue resistance and viscoelastic properties that approximate those of native tissue (5,64). Stiffness matching represents a foundational strategy: Hydrogels can be engineered to approximate native myocardial stiffness, thereby minimizing abnormal stress transmission to the surrounding myocardium (153). Beyond static stiffness, viscoelastic design is key. Hydrogels exhibiting stress relaxation that mimics the time-dependent mechanical behavior of native tissue promote cell spreading, proliferation and ECM deposition, which are key for functional integration and repair (154). In addition to intrinsic material durability, a hydrogel patch or injectable formulation must integrate with the myocardial wall and deform synchronously during each cardiac cycle. If the material is excessively stiff, the adjacent myocardium may experience abnormal mechanical loading; conversely, if it is overly compliant, it may fail to provide adequate structural support. Achieving mechanical integration with dynamic cardiac tissue remains challenging (64,67). Early-generation polymeric restraint devices, for example, restricted diastolic filling. Although hydrogels are inherently soft and deformable, offering an advantage in this context, their mechanical contribution evolves as degradation progresses. Designers must therefore anticipate shifts in myocardial workload as the hydrogel resorbs. Premature degradation may lead to loss of structural support before tissue healing is complete, whereas excessively delayed degradation may impede physiological remodeling. Alignment of degradation half-life with the myocardial recovery timeline is key (155); for example, an ideal hydrogel for MI repair may degrade over 3-6 months, coinciding with scar maturation (31). Electrical compatibility is also critical. An implanted hydrogel should not interfere with cardiac electrical conduction or contractile synchrony. Although a highly insulating hydrogel may induce a conduction block if it forms a substantial mass, clinical injections are typically small and distributed within native tissue and to the best of our knowledge, no increase in arrhythmic events has been reported to date (156). Nevertheless, studies (7,156) monitor patients for potential proarrhythmic effects [using magnetic resonance imaging (MRI) and electroanatomic mapping]. Conductive hydrogels incorporating conductive nanomaterials (such as graphene, gold nanowires or MXene) or electrically conductive polymers have been developed to restore electrical coupling across scar tissue (141,143). Preclinical studies (141-143) have demonstrated enhanced Cx43 expression, decreased arrhythmogenic susceptibility and improved synchronization of cardiomyocyte contraction. Collectively, achieving long-term mechanical and electrical stability under dynamic cardiovascular conditions requires an integrated design strategy that balances stiffness, viscoelasticity, degradation kinetics and electrical compatibility. By tailoring these properties to the specific biomechanical and electrophysiological demands of the target tissue, as well as the temporal course of healing, functional hydrogels provide durable support that promotes native cardiac repair.

Thrombogenicity is a key concern for any intravascular application of hydrogels or for materials that come into direct contact with blood. Although numerous hydrogels (alginate or PEG) are considered relatively hemocompatible, clinical experience from the Humacyte human acellular vessel trial demonstrated that even human ECM-based grafts may thrombose when flow dynamics are disturbed (157). In the Symvess trauma study, peri-implantation anticoagulation was administered to mitigate thrombotic risk (158). For future hydrogel-based vascular devices, surface modification strategies may be necessary to maintain long-term patency. Heparin or heparin-mimetic coatings decrease platelet adhesion and activation, as demonstrated in heparin-coated ePTFE grafts and stent coatings (159). NO-eluting hydrogels and mechanically robust NO-releasing hydrogel coatings applied to stents attenuate neointimal hyperplasia and thrombosis in porcine models by leveraging the antiplatelet and anti-smooth muscle proliferative effects of NO (105). Zwitterionic hydrogel coatings exhibit enhanced antifouling properties and decrease thrombogenicity of decellularized swim bladder matrices used for bioprosthetic heart valves (160). Hydrogels incorporating fucoidan, a sulfated glycosaminoglycan with anticoagulant properties, are applied as coatings for magnesium alloy stents, improving hemocompatibility and promoting endothelial cell colonization (109). In cardiac applications such as injectable hydrogels, delivery techniques must avoid the introduction of air bubbles or particulate contaminants that could result in embolization. Clinical protocols developed to date, including hydrogel filtration and meticulous catheter-based delivery techniques, have prevented embolic complications (33,71), however, these procedures require standardization as therapies advance toward broader clinical implementation.

Complex hydrogel systems, particularly those incorporating biological components, present challenges for consistent large-scale production. Batch-to-batch variability in biologically derived materials, such as differences in ECM composition among donor lots, directly influence therapeutic performance. For scale-up of composite and multi-component hydrogels, advances in continuous manufacturing technology, including microfluidic systems, automated mixing and dispensing platforms and 3D bioprinting, are enabling reproducible large-scale production with improved quality control (13,19). These platforms permit precise regulation of component ratios, crosslinking kinetics and final product homogeneity, thereby mitigating the inherent variability associated with multi-component formulations (161-163). Sterilization constitutes a notable translational barrier for cardiovascular hydrogels because of their high water content, thermosensitivity and susceptibility to structural degradation under conventional sterilization conditions. Most hydrogels cannot tolerate autoclaving or dry-heat sterilization, as elevated temperatures may induce premature gelation, network disruption, polymer hydrolysis or loss of thermoresponsive properties (164). Alternative terminal sterilization strategies are required but must be rigorously validated to preserve mechanical integrity, biocompatibility and drug release characteristics (150,164). Gamma irradiation and electron beam irradiation have been investigated for heat-sensitive hydrogels, however, radiolysis of water generates ROS, which can cause polymer chain scission, crosslinking imbalance, decreased mechanical strength and accelerated degradation (164). Ethylene oxide sterilization enables low-temperature processing; however, it requires prolonged degassing to eliminate cytotoxic residues and is incompatible with highly hydrated formulations. Hydrogen peroxide plasma sterilization decreases residual toxicity but may induce oxidative damage to polymer networks. Supercritical carbon dioxide sterilization has emerged as a promising low-temperature, residue-free alternative that preserves hydrogel structure, mechanical properties and biocompatibility while achieving high sterility assurance levels (164). For injectable and in situ-forming hydrogels, sterile filtration of prepolymer solutions prior to crosslinking is a practical and widely implemented approach, provided that all components are amenable to filtration and aseptic processing conditions are maintained. Lyophilized (freeze-dried) hydrogel formulations enable terminal sterilization of dry solids before reconstitution, thereby decreasing moisture-mediated degradation during sterilization and storage (164). No universal sterilization strategy exists; selection must be tailored to polymer composition, crosslinking chemistry, water content and therapeutic payload, with rigorous validation of sterility, structural integrity, mechanical performance and biocompatibility. Achieving good manufacturing practice compliance is also essential. Regulatory approval requires controlled cleanroom environments, validated cleaning protocols and comprehensive quality control of composition, mechanical properties, gelation behavior and sterility. Taken together, these considerations underscore that scalability is an integral component of hydrogel design, requiring the integration of robust material engineering with validated manufacturing processes to enable reproducible, off-the-shelf production of functional hydrogels for cardiovascular applications.

Any novel therapy must complement, rather than replace, established treatments unless it demonstrates a clear clinical advantage. For patients with MI, reperfusion therapy (percutaneous coronary intervention) and contemporary guideline-directed medical therapy are life-saving interventions and the cornerstone of management (6,7); hydrogel injection may serve as an adjunct strategy aimed at improving ventricular remodeling. Accordingly, clinical trials must demonstrate incremental benefit beyond optimal standard-of-care therapy (8,31,37). This necessitates sufficiently powered studies to detect, for example, additional improvements in LVEF or decreased HF-associated hospitalizations. Similarly, in patients with HF, hydrogel-based therapy may be administered in combination with established pharmacological and device-based treatments (34). The multimodal nature of HF management complicates attribution of therapeutic effects, making it challenging to isolate the independent contribution of the hydrogel intervention unless trials are rigorously designed and appropriately controlled (5,34).

Hydrogel-based therapies typically straddle the boundary between medical devices and biologics, rendering regulatory classification complex (150). FDA has categorized certain formulations (hydrogels incorporating cells such as mesenchymal stem cells or hydrogels containing bioactive factors like growth factors) as combination products, particularly when they incorporate cells or bioactive factors (150). Comprehensive safety and efficacy data are required as these therapies represent relatively new treatment modalities. Clinicians may also exercise caution in adopting such technologies until robust evidence from Phase III clinical trials demonstrates clear clinical benefit (71). Practical considerations influence adoption. Interventional cardiologists may require specialized training to deliver hydrogels safely via catheter-based techniques and cardiac surgeons must become familiar with handling hydrogel materials during operative procedures. In addition, reimbursement pathways must be defined, as these therapies may initially incur notable costs due to manufacturing complexity (5,150).