Decellularized cardiac ECM: A tissue-specific bioactive platform for drug release and cell delivery

Abstract

Decellularized extracellular matrix (dECM) materials derived from cardiac tissues have emerged as multifunctional biomaterials that recapitulate native microenvironmental cues while serving as controlled-release reservoirs of therapeutic agents and cells for cardiac repair. Recent advances have demonstrated that decellularized cardiac ECM (dcECM) hydrogels and patches can localize the release of angiogenic peptides, growth factors, or genes to promote neovascularization and mitigate adverse remodeling. Simultaneously, dcECM platforms have been applied to improve the survival and engraftment of stem cells, cardiomyocytes (CMs), and cardiac progenitor cells (CPCs) in infarcted myocardium, facilitating paracrine signaling and immune modulation. This review summarizes recent progress in the design and translational development of dcECM-based drug and cell delivery strategies for cardiac repair. We emphasized the interactions between matrix composition, its capacity to protect therapeutic factors and enhance cell retention, and therapeutic release profiles that collectively regulate cardiac tissue repair outcomes and highlighted emerging clinical applications and regulatory challenges for next-generation dcECM-based therapeutics.

Graphical Abstract

1. Introduction

Myocardial infarction (MI) is a critical pathological event, causing massive cardiomyocyte (CM) loss, fibroblast activation, and adverse left ventricular (LV) remodeling, ultimately leading to progressive cardiac dysfunction and heart failure[1–3]. Despite clinical treatments for MI such as pharmacological interventions, percutaneous coronary intervention (PCI), mechanical assistant devices, and coronary artery bypass grafting, may mitigate the MI symptoms, none of the current strategies effectively regenerate lost CMs and prevent adverse LV remodeling[4]. Heart transplantation remains the only curative option but is severely limited by donor shortage and immune rejection[5, 6].

Recently, cell, bioactive factors, and gene-based therapies have shown promises for cardiac repair [7–13]. However, direct intramyocardial injection of stem/progenitor cells results in extremely low retention and survival rate, rapid clearance, and inadequate targeting to the infarcted myocardium due to mechanical washout and the harmful inflammatory microenvironment of the infarct zone[14–16]. Likewise, soluble growth factors and cytokines including vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), stromal cell-derived factor 1 alpha (SDF-1α), and insulin-like growth factor 1 (IGF-1) exhibit short physiological half-lives and rapidly diffuse away from the injection site, limiting therapeutic efficacy [12, 17–20]. RNA- and gene-based therapies also face challenges such as nuclease degradation, immune activation, and inefficient intracellular uptake when delivered without protection[21]. These limitations highlight the need for biomaterial-based delivery systems capable of enhancing local retention, protecting therapeutics from degradation, enabling controlled release, and providing a reparative microenvironment that supports cell survival and positive tissue remodeling.

Over the past two decades, implantable biomaterials have emerged as promising therapeutic modalities for cardiac repair[22]. These include injectable hydrogels and biofunctionalized patch grafts that aim to deliver therapeutic drugs, bioactive factor, and cells to stabilize the infarct wall, promote angiogenesis, and enhance cardiac function[23–25]. Most synthetic or single-component natural scaffolds cannot fully replicate the complexity of the native multicomponent extracellular matrix (ECM), which provides not only structural support but also essential cues that guide cell adhesion, migration, differentiation, and paracrine signaling[5, 26, 27]. The cardiac ECM comprises a highly organized network of collagen, elastin, glycosaminoglycans (GAGs), and matrix proteins that orchestrates mechanotransduction, vascular remodeling, and immune regulation[28–31]. MI triggers extensive degradation and pathological remodeling of the cardiac ECM, leading to loss of structural integrity and disruption of matrix-dependent biochemical and mechanical cues that are essential for coordinated tissue repair, ultimately resulting in adverse LV remodeling[32]. Therefore, restoring or recapitulating the native cardiac ECM microenvironment has emerged as a critical design principle for effective myocardial repair.

Decellularized cardiac extracellular matrix (dcECM) derived from cardiac tissues is multicomponent and offers a unique combination of biocompatibility, bioactivity, and tissue-specific adaptability[27, 33–35]. Through removal of cellular components while preserving key structural and biochemical components, dcECM contains varieties of bioactive factors and integrin-binding peptides that promote cell recruitment and tissue remodeling[36]. The injectable dcECM hydrogels and acellular dcECM patches have demonstrated their abilities to attenuate fibrosis, enhance angiogenesis, and modulate immune responses for cardiac repair [37–42]. Importantly, dcECM also can be used as versatile carriers of therapeutic agents and cells, integrating mechanical support with localized biochemical signaling to strengthen therapeutic performance. Exosomes, angiogenic peptide, anti-inflammatory and antioxidant agents, and genes have been incorporated into dcECM biomaterials to increase neovascularization, reduce fibrosis, modulate immunoresponses, and improve cardiac function[38, 43–47]. In addition, stem cells or progenitor cells were integrated into dcECM scaffolds to provide both paracrine and structural cues for heart repair [48–50]. Co-delivery of cells with dcECM hydrogels or patches synergistically improved cell retention, angiogenesis, and LV function compared with either treatment alone[51–53].

These attractive studies highlight the dual functionalities of dcECM-based biomaterials that combine intrinsic bioactivity with localized, controlled delivery systems for drugs, bioactive factors, genes, and therapeutic cells. The inherent tissue specificity and tunable degradation of dcECM, when coupled with advances in nanotechnology and bioengineering, provide a versatile foundation for next-generation cardiovascular therapies. This review summarizes recent progress in dcECM-based biomaterials for cardiac repair, emphasizing the strategies for drug/bioactive factors release and cell delivery. We further discuss emerging design principles, translational challenges, and future directions toward clinically viable dcECM-based therapeutics.

2. Decellularized cardiac ECM

2.1. Decellularization techniques

Cardiac tissue decellularization approaches can be generally categorized into chemical, physical, and hybrid approaches (Table 1). Among these, chemical decellularization is the most widely used for preparing myocardial- or pericardium-derived ECM (Figure 1A and 1B). Common ionic detergents include sodium dodecyl sulfate (SDS) [54–57] and sodium deoxycholate (SDC) [58], which effectively lyse cells and solubilize cellular and nuclear components. However, prolonged exposure, particularly SDS, would cause damage to essential ECM proteins, thus optimization is requested to balance decellularization efficiency with matrix preservation[59]. The non-ionic chemical Triton X-100 primarily disrupts the bonds of lipid-lipid and lipid-protein and preserves protein-protein bonds, which results in milder ECM damage but reduces decellularization potency[60]. Physical factors such as temperature, pressure, and agitation can contribute to cardiac tissue decellularization. Repeated freeze-thaw cycles can induce cellular rupture via ice-crystal formation[42, 61]. The supercritical CO₂ (scCO₂)-assisted decellularization has been reported to increase removal of endotheliocytes, fibroblasts, and smooth myocytes while preserving the structure and reducing damage to mechanical properties [62], but this method requires specialized tissue suspension systems to ensure the optimal circulation of scCO₂ and other additives[63].

Table 1.

Decellularization methods used for dcECM preparation.

Mechanism	Common methods	Advantages	Disadvantages	Ref.
Chemical Methods	SDS, SDC/SD	Very efficient at lysing cells and solubilizing cellular/genetic material.	Prolonged exposure (especially SDS) damages essential ECM proteins.	[54–59, 74]
	Triton X-100	Disrupts lipid–lipid and lipid–protein interactions while preserving protein–protein bonds; widely used.	Potential concerns about environmental impact and biodegradability. Challenging characterization and quantification of residual detergent due to intrinsic molecular heterogeneity	[75–77]
Physical Methods	Freeze–thaw (thermal shock)	Simple, uses temperature cycling to disrupt cells via ice crystal formation.	May damage ECM and can be mitigated by cryoprotectants (e.g. trehalose).	[42, 61]
	Supercritical CO2 (scCO2)–assisted decellularization	Efficient removal of cell components; preserves structure and mechanical properties; faster than many unassisted methods (4–7 days).	Requires specialized tissue suspension/circulation systems.	[62]
Hybrid Methods	SDC + trypsin + peracetic acid (PAA)	Efficient removal of cell components; preserves structure and mechanical properties, especially in whole-heart decellularization; faster than many unassisted methods (4–7 days).	Requires specialized tissue suspension/circulation/perfusion systems.	[64]
	Tris/EDTA + SDS + DNase/RNase Trypsin/EDTA + Triton X-100 + SDC Heparinized PBS/adenosine + SDS + Triton X-100			[48] [34] [69]

Figure 1.

Decellularization workflows and compositional analysis of dcECM. (A) Schematic illustrations and representative images depicting the decellularization process for porcine myocardium tissue. (B) Schematic presentation of the decellularization process for human pericardium (Pc) and representative gross images of native Pc and decellularized pericardium (DPc) after decellularization. Modified and adapted with permission from Ref.[57]. © 2015 Wiley Periodicals, Inc. (C) Comparison of matrisome protein profiles between decellularized porcine myocardium and human pericardium. Modified and adapted with permission from Ref.[72] under the Creative Commons CC BY license. © 2018 The Author(s). (D) Quantitative analysis of total protein composition in injectable human dcECM using ECM-targeted quantitative QconCAT coupled with LC-SRM assay, revealing the relative contributions of fibrillar collagen, basement membrane proteins, matricellular components, and other ECM-associated proteins. Modified and adapted with permission from Ref. [73]. © 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim.

In addition, many decellularization protocols involve hybrid methods that combine detergents with enzymatic treatments to improve cellular removal. For example, trypsin is often used in combination with detergents, such as SDS, SDC, or with deoxycholic acid (DCA), to enhance decellularization efficiency [64–66]. In addition, ethylenediaminetetraacetic acid (EDTA) is commonly incorporated as a chelating and decalcifying agent. EDTA can assist decellularization by disrupting calcium-dependent cell–cell and cell–ECM adhesion and then improve cellular clearance. It can also prevent blood clotting and biofilm formation during perfusion-based or prolonged decellularization processes[67]. To further reduce immunogenicity, nucleases (DNase/RNase) are added to degrade residual nucleic acids and are often combined with detergent-based protocols for tissue decellularization [48, 68]. For whole-heart decellularization, perfusion-based methods are used to improve detergent diffusion and preserve native chamber geometry and vascular architecture[69–71]. The anticoagulant heparin and the vasodilator adenosine together with SDS, Triton X-100, and antibiotics have been shown to yield acellular heart architecture with intact chamber geometry and fine structure [69].

2.2. Components of dcECM

In the native cardiac tissue, the ECM consists of a highly complex components of structural and functional proteins, proteoglycans, glycoproteins, and soluble factors that collectively regulate myocardial architecture and cell behavior[50, 78]. Comparative proteomic analysis of the decellularized porcine myocardial and human pericardial matrices revealed a substantial overlap, with approximately 40% of matrisome proteins shared between the two tissues (Figure 1C)[72]. These shared components are dominated by multiple collagen subtypes that provide structural integrity and modulate cell differentiation, and basement membrane–associated components, including laminin and heparan sulfate proteoglycans, which regulate cell adhesion, matrix interactions, and CMs behavior[30, 79–81]. In addition, fibronectin, a multifunctional ECM glycoprotein, is present in both matrices and contributes to a broad range of cellular processes, including adhesion, survival, and differentiation[30]. Furthermore, quantitative proteomic analysis of injectable human decellularized myocardial matrices (Figure 1D) revealed a diverse composition of ECM proteins[73]. Fibrillar collagens comprise more than 70% of the total matrix, predominantly collagen types I and III, followed by basement membrane (~20%) and structural ECM proteins (~4%). The remaining minor fractions include matricellular and regulatory proteins that play essential roles in heart development and remodeling despite their low abundance[82].

2.3. Formats of dcECM

The dcECM is employed in two major formats as a therapeutic carrier: patches and injectable hydrogels. The dcECM patches are fabricated directly from decellularized myocardium slices or decellularized pericardium. They provide structural supports and tissue-specific biochemical cues that promote the stem cell attachment, migration, and lineage differentiation [83–85]. Beyond native dcECM patches, engineered dcECM-based patches have been developed by integrating dcECM with synthetic polymers, conductive components, or therapeutic agents using electrospinning or 3D printing techniques[53, 86–88]. These hybrid constructs increase mechanical strength, enable anisotropic architectures that better recapitulate native myocardial structure, or introduce additional bio-functionalities. These engineering designs expand the versatility of dcECM patches and further make dcECM as a multifunctional carrier for cardiac repair.

In contrast to patches, the injectable dcECM hydrogels offer a minimally invasive and conformal delivery strategy for MI treatment. They are typically generated by enzymatic solubilization of decellularized cardiac tissues and in situ self-assembly[54]. The dcECM hydrogels can be delivered via intramyocardial or catheter-based injection and rapidly adapt to the complex geometry of the infarcted myocardium[37, 54, 89]. The dcECM hydrogels are majorly used as injectable biomaterials, but they can also be fabricated into hydrogel patches that are placed on the epicardial surface for therapeutic delivery[88, 90]. In both formats, dcECM hydrogels retain key ECM components and biochemical cues that modulate cell behavior, inflammation, and tissue remodeling, while serving as local carriers for therapeutic cells, drugs, and bioactive factors[50, 91, 92]. Moreover, the physicochemical properties of dcECM hydrogels such as stiffness, degradation kinetics, and porosity, can be tailored through concentration control, chemical modification, or hybridization with synthetic polymers, nanoparticles (NPs), or bioactive molecules to enable spatiotemporally controlled release and improve tissue integration[44, 46, 55, 93–95]. Thus, the injectable dcECM hydrogels also provide a versatile platform that combines minimally invasive delivery with biomimetic regulation of MI repair.

2.4. Immunomodulatory roles of cardiac ECM and dcECM-based biomaterials

Dynamic alterations in ECM critically regulate post-MI inflammation, fibrosis, angiogenesis, and ventricular remodeling[32]. Beyond providing structural support, the cardiac ECM provides a dynamic bioactive microenvironment that actively regulates immune cell recruitment, activation, and phenotypic transition through biochemical and biomechanical signaling pathways[96, 97]. As a bioactive protein complex, ECM contains various immune-active components that contribute to immunomodulation during cardiac injury and repair[28, 39, 98]. Among immune cell populations, macrophages serve as critical responders to regulate cardiac remodeling and regeneration[99, 100]. Macrophage polarization is affected by multiple factors, including foreign pathogens and substances, intrinsic cellular debris, and ECM proteins[101]. The dcECM-based biomaterials facilitate interactions between immune cells and ECM components, such as collagens, hyaluronan, fibronectin, decorin, versican, and GAGs, through integrin- or non-integrin-mediated signaling pathway for immune modulations[28, 99, 102, 103]. For example, collagens can improve immune cell adhesion and integrin expression while suppressing inflammatory response[104–106]. The fibronectin fragments promote early pro-inflammatory macrophage activation that facilitates immune cell recruitment and initiation of tissue repair process[99, 107]. Certain GAGs, particularly hyaluronic acid-rich ECM microenvironments, have been reported to reduce fibrosis and promote macrophage polarization toward a pro-reparative M2 phenotype through modulation of inflammatory signaling pathways[108, 109]. In addition, non-structural ECM, known as matricellular proteins, include periostin (POSTN), osteoponin (OPN), thrombospondins (THBS), cellular communication network factors (CCNs), secreted protein acidic and rich in cysteine (SPARC), and others [110]. They can regulate immune responses by interacting with integrins, cytokines, and growth factor signaling pathways, and promote the transition from pro-inflammatory responses toward reparative remodeling and angiogenesis [110, 111]. Furthermore, the bioactive ECM components can modulate macrophage behavior through integrin-mediated cell–matrix interactions and mechanosensitive signaling pathways, including integrin-FAK and PI3K/Akt signaling[30, 112, 113]. These pathways suppress inflammation and promote a pro-regenerative microenvironment through TLR/NF-κB signaling modulation, resulting in increased expression of the anti-inflammatory cytokine IL-10 and the reparative cytokine TGF-β [114–116]. In addition to intact ECM components, matrix metalloproteinases (MMPs)-mediated cleavage of cardiac ECM proteins during heart injury generates biologically active matricryptic peptides [117, 118]. These bioactive fragments function as endogenous signaling molecules that regulate tissue repair. For example, collagen-derived matricryptic peptides could promote angiogenesis and regulate the scar formation through ECM remodeling [119]. These matricryptic peptides also promote M2 macrophage polarization and regulate inflammation and reparative processes [32, 120]. Consistent with these mechanisms, recent preclinical studies have demonstrated that dcECM-based biomaterials can promote macrophage polarization toward the reparative M2 phenotype, reduce inflammatory responses, and improve cardiac remodeling for MI treatment [98, 121–123]. In rat MI models, dcECM hydrogels have been shown to modulate post-MI repair by suppressing pro-inflammatory signaling and promoting pro-reparative macrophage polarization and Th2-associated immune responses[98]. Similarly, it also promoted M2 macrophage polarization through activation of JAK1/STAT6 and Erk1/2 signaling pathways and simultaneously increased anti-inflammatory and pro-angiogenic cytokine expression to support cardiac repair in a rat MI model[124]. Additionally, the injectable dcECM hydrogels could modulate the post-MI immune microenvironment by attenuating TGF-β-associated profibrotic signaling and preserving anti-inflammatory tissue-resident macrophages to promote pro-reparative immune remodeling [38, 98]. Despite increasing recognitions of the immunomodulatory properties of dcECM-based biomaterials, the precise molecular mechanisms underlying dcECM-mediated immune regulation remain to be fully elucidated because of the intricate interactions among multiple signaling networks, including NF-κB, JAK/STAT, PI3K/Akt, Hippo, and Notch pathways, as well as the dynamic cellular transitions during post-MI remodeling [125].

3. dcECM as a drug delivery platform

3.1. Rationale for drug incorporation

MI is accompanied by complex alterations in cellular behavior, enzymatic activity, and local biochemical signaling[126–129]. These changes progressively drive pathological processes, such as inflammation, oxidative stress, ischemia, and abnormal matrix remodeling, which can eventually lead to scar formation, loss of functional contractile or vascular tissue, and impaired perfusion. Such adverse remodeling hinders injured tissue repair and functional recovery. To regulate the local microenvironment and promote tissue repair, many therapeutic strategies have focused on the delivery of growth factors, cytokines, and small molecules[130–132].

Growth factors and cytokines play crucial roles in regulating angiogenesis, cell recruitment and survival, and resolution of inflammation[130, 133–135]. VEGF can promote angiogenesis in the ischemic myocardium to increase cardiac function recovery, and fibroblast growth factor 2 (FGF2) similarly induces the migration and proliferation of endothelial and smooth muscle cells to facilitate neovascularization and collateral vessel formation to improve blood perfusion and cardiac performance in ischemic regions[136–140]. Also, SDF-1α promotes the recruitment of stem cells and C-X-C chemokine receptor type 4 (CXCR4)-positive progenitor cells to the injured myocardium to enhance tissue repair [141–143]. However, the systemic administration of such protein-based agents faces several limitations. Most growth factors and cytokines are highly labile in the circulation, where they are rapidly degraded by proteases and cleared through renal filtration or hepatic metabolism, resulting in half-life in minutes [144, 145]. Repeated or high-dose injections are therefore required to maintain effective concentrations at the target site, which may trigger undesirable effects such as off-target angiogenesis, vascular leakage, or systemic inflammation[134, 146]. In addition, these soluble proteins exhibit poor localization and retention within the injured myocardium as the extracellular matrix undergoes extensive degradation and remodeling, which may affect the retention and stabilization of bioactive molecules[32]. On the other hand, small molecule drugs, such as TT-10, 5SM and colchicine, can support cardiac repair through multiple mechanisms, including antioxidant, anti-inflammatory, and pro-regenerative effects [147–150]. However, the systemic delivery of small molecule drugs has several challenges. Many of these drugs exhibit poor solubility in aqueous environments and are subject to rapid hepatic metabolism or renal excretion, which limits their bioavailability and duration of action [151]. Their small size and high diffusivity also result in widespread distribution across organs, sometimes causing off-target exposure and systemic side effects[134]. Moreover, under ischemic or inflammation conditions, the unusual pH, oxidative state, and enzymatic composition of the local tissue may accelerate the degradation or inactivation of these drugs, further reducing their therapeutic efficacy[152].

Besides protein therapeutics and small molecule drugs, nucleic acid–based drugs are emerging as promising therapies for cardiac repair. They can regulate gene expressions at different levels to improve cardiac function following myocardial injury[135, 153, 154]. MicroRNA (miRNA) mimics or inhibitors have been used to promote angiogenesis, reduce cardiomyocyte apoptosis, and modulate post-MI remodeling[155, 156]. Antisense oligonucleotides against maladaptive cardiac miRNAs, such as the clinical-stage inhibitor of miR-132 (CDR132L), have shown that selective silencing of pathogenic miRNAs can improve systolic and diastolic function in heart failure models and in early clinical studies[157, 158]. Messenger ribonucleic acid (mRNA)-based therapeutics provide an alternative strategy by tuning gene expression for the in-situ production of pro-angiogenic or pro-proliferative proteins to support cardiomyocyte survival, neovascularization, and cardiac repair for MI treatment [159]. Although nucleic acid drugs exhibit high therapeutic potential, their clinical efficacy is often limited by short half-life and limited drug retention due to nuclease degradation and rapid systemic clearance[160]. Thus, there is an urgent need for a delivery platform that can provide controlled release and protect therapeutic agents from rapid clearance or degradation, while adapting to the complex pathological environment of injured cardiac tissues.

3.2. Comparison of dcECM-based biomaterials with established biomaterial platforms

Varieties of local drug delivery platforms have been developed for cardiac repair, including synthetic and natural polymers [161]. The dcECM-based biomaterials possess several unique advantages for cardiac therapeutic delivery due to their tissue-specific biochemical composition and intrinsic bioactivity[89, 162, 163]. Synthetic polymers, such as poly(ethylene glycol) (PEG), poly(lactic-co-glycolic acid) (PLGA), poly(N-isopropylacrylamide) (PNIPAAm), and polyurethane (PU) provide highly tunable physicochemical properties, including crosslinking density, mechanical properties, and drug release profiles[164–169]. However, these synthetic materials generally lack bioactivity and often require biofunctionalization to support cardiac repair[170, 171]. In contrast, dcECM-based biomaterials can directly provide a more biomimetic microenvironment capable of regulating cell behavior, immune responses, angiogenesis, and tissue remodeling[39, 50, 163]. In addition, intrinsic affinity interactions between therapeutic molecules and ECM components could improve local therapeutic retention and bioactivity within the infarct microenvironment[93].

Single-component natural polymers/hydrogels, such as collagen, gelatin methacryloyl (GelMA), fibrin, and hyaluronic acid, exhibit favorable biocompatibility and biodegradability but cannot fully replicate the biochemical complexity and tissue specificity of the native cardiac extracellular matrix microenvironment[53, 172–176]. Compared with these conventional biomaterials, dcECM-based biomaterials preserve a tissue-specific composition by retaining multiple native ECM components and signaling molecules of cardiac tissue. In addition, dcECM contains various integrin-binding motifs that support cell–matrix interactions, while its proteoglycans and GAGs serve as growth factor reservoirs[177, 178]. However, dcECM-based biomaterials also face several important limitations, including batch-to-batch variability, heterogeneous composition, limited mechanical tunability, and challenges associated with scalable manufacturing, sterilization, and regulatory standardization, which are discussed in Section 6. Furthermore, precise control over degradation behavior and therapeutic release kinetics remains more difficult in dcECM-based biomaterials than in highly engineered synthetic biomaterials because of the susceptibility of dcECM to enzymatic degradation and its complex interactions with therapeutic agents. Strategies to improve release controllability and therapeutic retention in dcECM-based biomaterials are discussed in Section 3.4.

3.3. Strategies for drug loading

The dcECM contains abundant structural proteins and bioactive macromolecules, including collagens, elastin, and GAGs, which provide natural binding sites for therapeutic agents. Drugs can be incorporated into dcECM by direct loading, and carrier-mediated loading [91]. These strategies enable localized, protected, and tunable delivery of proteins, small molecule drugs, and nucleic acid therapeutics within injured cardiac tissues (Table 2).

Table 2.

List of strategies for incorporating drugs/bioactive factors into dcECM.

Tissue resource	Drug	Loading method	Form	Release profile	Animal model	Ref.
Porcine myocardium	ASC-secreted factors	Direct loading	Injectable hydrogel	Sustained release over 5 days with factor-dependent kinetics	N/A	[45]
Porcine pericardium	bFGF	Direct loading	Injectable hydrogel	~28% released over 5 days	Rat I/R	[80]
Porcine pericardium	HGF-f	Direct loading	Injectable hydrogel	Slow sustained release over 5 days, with stronger retention compared to rapid burst from collagen gels.	Rat I/R	[46]
Porcine myocardium	miRNA and EVs	Direct loading	Injectable hydrogel	Prolonged and sustained release compared with rapid loss of free miRNA/EVs	N/A	[179]
Porcine myocardium	Human CSC–secreted factors	Carrier-mediated loading	Cardiac patch	Sustained release of VEGF, HGF and IGF for at least 4 days without burst behavior	Rat and Porcine MI	[180]
Porcine myocardium	Curcumin	Carrier-mediated loading	Injectable hydrogel	Slow sustained 7-day release, with only ~20–30% released in 24 h	Mouse MI	[43]
Porcine myocardium	SDF-1α	Carrier-mediated loading	Injectable hydrogel	Sustained SDF-1α release for about 4 weeks from the NP–dcECM composite	Rat MI	[55]
Porcine myocardium	Colchicine	Carrier-mediated loading	Injectable hydrogel	Colchicine released rapidly for 8 days, then continued with slower sustained release till day 14	Mouse MI	[38]

The first strategy, direct (physical) loading in dcECM, mainly relies on matrix affinity and physical entrapment of therapeutic agents within native ECM components such as collagen, elastin, and GAGs (Figure 2A) [45, 46, 80, 179]. The therapeutic agent is directly incorporated into the dcECM pre-gel solution before gelation, allowing the molecules to be entrapped within the matrix network during hydrogel formation [45, 46, 80]. The loaded agents may be further retained within the dcECM through non-covalent interactions, such as electrostatic interactions, hydrogen bonding, and hydrophobic interactions[91]. It can achieve effective loading without chemical modification, which can maintain the bioactivity of the incorporated agents and the structural integrity of the scaffold.

Figure 2.

Schematic illustration of drug incorporation strategies within dcECM-based delivery systems. (A) Direct incorporation of bioactive therapeutics, including proteins, small molecules, and nucleic acids by simple mixing within dcECM scaffolds. (B) Therapeutic drugs/factors are first loaded into micro- or nanoparticle carriers or exosomes, which are subsequently incorporated into dcECM scaffolds for carrier-mediated delivery.

The second strategy, carrier-mediated delivery, has attracted increasing attention in recent years. The carriers can be integrated with dcECM biomaterials to achieve more precise and tunable control of drug release. These particulate carriers, such as synthetic nano- or microparticles and naturally derived exosomes, can load growth factors, small molecules, or nucleic acids to protect them from enzymatic degradation and enable programmable release profiles through diffusion, carrier erosion, or environmental triggers (Figure 2B)[38, 43, 55, 180]. When drug-loaded carriers are embedded within the dcECM, a hierarchical system is formed. The dcECM provides a biomimetic microenvironment and tissue anchorage, and the carriers control the drug release kinetics. Most carriers are core-shell structures and encapsulate therapeutic agents within the inner core of NPs or exosomes, which provide enhanced protection from degradation and allow more sustained, diffusion-controlled release[38, 43, 55, 180]. The size, composition, and surface functionality of drug carriers can be tailored to achieve controlled or stimuli-responsive release to match the spatial and temporal heterogeneity of cardiac pathology. After MI, the ischemic core and border zone exhibit different oxygen levels, enzymatic activities, and inflammatory signals over time [129, 181, 182]. Engineering NPs to respond to pH, redox, or enzymatic changes in the infarct microenvironment may enable the therapeutic system to adapt to these dynamic conditions. Integrating responsive or multi-phase NPs into dcECM scaffolds may further enhance therapeutic efficacy.

3.4. Controlled release profiles

In the dcECM-based drug delivery systems, two characteristic release patterns are majorly observed: burst release and controlled (sustained) release. The difference between these two profiles mainly depends on how the therapeutic molecules are incorporated into the system, how strongly they interact with the scaffold, and how the dcECM structure regulates diffusion and degradation.

In systems that drugs or growth factors are physically mixed or weakly adsorbed onto the dcECM network, the drug-dcECM interactions are mainly dominated by weak electrostatic or van der Waals forces [46, 179]. Once the scaffold is exposed to an aqueous environment, loosely bound molecules rapidly diffuse from the scaffold, leading to an initial burst release. This phenomenon would be affected by the pore size and crosslinking density of dcECM[183]. The release kinetics are primarily regulated by the concentration of the therapeutic agents and the physicochemical properties of dcECM-based biomaterials [45]. It is notable that in vitro dcECM degradation is typically limited in PBS or other enzyme-free conditions[184], where in vitro release is controlled mainly by diffusion rather than matrix degradation. More precise control over therapeutic release can be achieved through strong matrix interactions or incorporation into a secondary carrier system, where matrix degradation further regulates therapeutic release in vivo.

Currently, dcECM-based systems enable temporal controls over drug release; however, precise spatiotemporal delivery in MI applications remains an ongoing challenge. Given the compositional complexity of dcECM-based biomaterials, therapeutic release is often modulated by multiple overlapping transport mechanisms, including diffusion-controlled transport, matrix degradation, matrix–therapeutic agent interactions, and environmental stimuli within the infarct zone, making precise control of release kinetics difficult compared with highly defined synthetic hydrogels[91, 185]. Drug release from hydrogel systems can be commonly classified into diffusion-, swelling-, degradation-, or erosion-controlled mathematical models, including the Higuchi, Korsmeyer–Peppas, Peppas-Sahlin, zero-order, and first-order equations [186, 187]. Among these models, semi-empirical (Korsmeyer–Peppas and Peppas-Sahlin) and reaction-diffusion models may be more applicable to naturally derived hydrogels because they can better describe the complex release behavior of these biomaterials[188–192]. However, there is currently no established mathematical model specifically for dcECM-based hydrogels. This limitation is largely attributed to the heterogeneous biochemical composition, dynamic enzymatic degradability, and matrix–therapeutic agent interactions inherent to dcECM biomaterials, which comprehensively result in complex release behaviors rather than a single dominant release mechanism. Moreover, accurate prediction of release kinetics in degradable biomaterials often requires simultaneous consideration of diffusion, network structure, polymer degradation, protein–matrix interactions, and enzymatic cleavage kinetics, particularly in bioactive matrices with drug-binding capabilities[185, 193].

In addition, degradation-mediated release represents an important characteristic of dcECM-based biomaterials because their structure and biochemical properties continuously change within the enzymatically active post-MI microenvironment. The dcECM scaffolds can be degraded in vivo by proteolytic enzymes, including MMPs that are highly upregulated after MI [177]. Implanted dcECM biomaterials also recruit immune cells that further contribute to scaffold degradation through the secretion of proteases and inflammatory mediators [194, 195]. Progressive degradation of the dcECM network increases matrix permeability and weakens matrix–therapeutic agent interactions to facilitate therapeutic release. Previous studies showed that collagenase treatment accelerated therapeutic release from dcECM-based biomaterials, but the dcECM still exhibited slower release kinetics and improved retention compared with single-component collagen hydrogels[80]. This indicates that the heterogeneous composition and structural complexity of dcECM may provide additional modulation of degradation-associated release behavior[80].

To further improve the controllability of therapeutic release and retention, engineering strategies have been explored for dcECM-based biomaterials. Increasing dcECM concentration or crosslinking density are commonly considered to reduce matrix pore size and diffusivity and then prolong therapeutic release and minimize burst release effects[196]. For example, secondary crosslinking approaches, such as genipin-mediated crosslinking, can enhance matrix stability and slow degradation-dependent release[50, 197, 198]. In addition, blending dcECM with natural polymers, synthetic polymers, bioactive peptides, or nanoparticle-based carriers, as well as chemical modification, has been investigated to further regulate degradation behavior, mechanical stability, and release profiles[38, 47, 55, 124, 199–201]. These modifications can alter crosslinking density, enzymatic susceptibility, and matrix–therapeutic agent interactions, enabling control over sustained therapeutic delivery within the dynamic post-MI microenvironment.

3.5. Preclinical studies & therapeutic outcomes

Preclinical investigations have consistently shown that dcECM-based drug delivery systems improved therapeutic efficacy in MI models and ischemia–reperfusion (I/R) injury models [38, 46, 55, 80, 202]. The dcECM acts as both a structural scaffold and a bioactive platform that prolongs local drug retention and supports tissue-specific repair.

The sustained release of bioactive factors or drugs within these matrices supports coordinated repair processes through stimulating angiogenesis, preserving viable myocardium, and mitigating fibrotic remodeling. The prolonged local retention could reduce off-target exposure and maintain bioactivity during the critical healing window. For example, a myocardium-derived dcECM hydrogel encapsulating SDF-1α–loaded PLGA NPs achieved prolonged and localized release of SDF-1α over 4 weeks in a rat MI model, which improved angiogenesis compared with dcECM alone (Figure 3A (i) and (ii)) [55]. Notably, the group treated with SDF-1α showed improved cardiac function at 2 weeks compared with other groups, but this effect diminished by week 8 as the SDF-1α release gradually decreased, which highlighted the need for sustained drug delivery (Figure 3A (iii), (iv), and (v)). Similarly, a myocardial-derived dcECM with curcumin-loaded exosomes markedly promoted vascularization compared with exosome-free curcumin delivery or dcECM alone in a rat MI model (Figure 3B (i), (ii), and (iii))[43]. In addition, effective curcumin delivery reduced fibrosis and preserved left ventricular wall thickness (Figure 3B (iv) and (v)). In a porcine MI model, a dcECM patch combined with synCSC-loaded PLGA microparticles (artCP) was implanted for cardiac repair (Figure 3C (i))[180]. The infarct percentages quantified across mid-ventricular sections (Slice 2) were significantly reduced in the artCP-treated hearts compared with controls (Figure 3C (ii) and (iii)). These preclinical findings exhibit the versatility of dcECM-based delivery platforms in integrating structural support with localized, sustained therapeutic signaling to promote cardiac repair.

Figure 3.

Preclinical evidence of therapeutic outcomes achieved by drug-releasing dcECM systems. (A) (i) Scheme of the fabrication and evaluation of the SDF-1α–loaded dcECM/PLGA nanoparticle composite hydrogel. (ii) Cumulative release of SDF-1α from dcECM hydrogels and dcECM/NP composite hydrogels at 37 °C. (iii) Immunofluorescent staining for α-SMA (green) and vWF (red) (Scale bar: 50 μm). (iv) Quantitatively analysis of total vessel density in the infarct region and (v) Left ventricular ejection fraction (LVEF) measurements (n=6, * p<0.05, **** p<0.0001). Modified and adapted with permission from Ref. [55], Copyright 2025, Elsevier. (B) (i) Illustration of the preparation of an injectable dcECM composite hydrogel incorporating curcumin loaded exosomes for myocardial infarction treatment. (ii) Immunofluorescence staining of CD31 (red) in myocardial sections from MI rats treated with myocardial-derived dcECM, curcumin-loaded exosomes, or exosome-loaded dcECM. (Scale bar: 50 μm). (iii) Quantitative analysis of CD31-positive vessel density for the corresponding treatment groups (n=3, *** p < 0.001, ^&&&p < 0.001). (iv) Masson’s trichrome staining of corresponding myocardial sections, highlighting collagen fibers (blue) and myocardium (red) (Scale bar: 1 mm). (v) Quantitative analysis of fibrosis area in myocardial sections from the different treatment groups (n=3, *** p< 0.001, ^& p < 0.05, ^&&& p< 0.001). Modified and adapted with permission from Ref. [43]. Copyright 2023, Elsevier. C (i) Scheme of the fabrication and application of the synCSCs-loaded dcECM artificial cardiac patch (artCP). (ii) Transverse heart sections (≈1 cm thick) from MI-only pigs and artCP -treated pigs, with infarct regions outlined in MI-only hearts (MI only) and artCP site marked in treated hearts (MI+ artCP); TTC-stained slices showing viable (red) and infarcted (white) myocardium, with patch location indicated by arrows. (iii) Quantification of infarct percentages from TTC-stained 2^nd ~4^th slices in MI-only and artCP-treated pigs, presented as mean ± SD with statistical comparisons indicated by lines and asterisks (n=3, * p < 0.05, *** p < 0.001). Modified and adapted from Ref. [180], licensed under Creative Commons Attribution-NonCommercial (CC BY-NC) (© The Authors).

4. dcECM for cell delivery

4.1. Challenges of cell therapy in cardiac repair

Cell therapy for MI and I/R diseases is hindered by the extremely low cell retention, engraftment, and survival within the hostile ischemic microenvironment in the infarcted myocardium. Fewer than 5% of transplanted cells typically remain at the target MI site after 48 h and only ~1% surviving beyond 4–6 weeks post-injection into the host myocardium[203]. The mechanical contraction of the myocardium and inflammatory environment rich in ROS, cytokines, and proteases further contribute to early cell loss and apoptosis[204, 205]. Clinical translation is also hampered by delivery route limitations. The intracoronary injections often lead to microvascular obstruction, and the intramyocardial injections are invasive and cause cardiac arrhythmias[206–208]. The harmful post-MI environment is characterized by excessive cytokines (IL-1β, IL-6), oxidative/inflammatory, and matrix degradation, which accelerate apoptosis and adverse fibrotic remodeling[51]. Thus, integrating therapeutic cells with a bioactive matrix such as the biomimetic cardiac ECM is essential to improve cellular retention and reparative efficacy.

4.2. dcECM as a cell delivery vehicle

The dcECM provides a native-like microenvironment that can substantially enhance cell survival, adhesion, and integration post-transplantation. The complex biochemical components of dcECM can interact with integrin and growth factor receptors to promote cell–matrix adhesion and mechanotransduction[31]. The dcECM hydrogels exhibit nanofibrous, porous structures that improve nutrient diffusion and support cell survival for loaded cardiomyocytes or progenitor cells[50]. The stem cell or progenitor cell loaded dcECM allows localized deposition, gradual degradation, and paracrine crosstalk with host cells, which can promote angiogenesis, reduce fibrosis, and improve cardiac function (Table 3).

Table 3.

Strategies of incorporating cells into dcECM for cardiac repair.

Tissue resource	Other components	Delivered cells or other factors	Form	Results	Ref.
Porcine myocardium	N/A	Human-derived CPC and MSCs, VEGF	3D printed hydrogel patch	Cardiac function↑, neo-muscle and capillary formation↑, cardiac hypertrophy and fibrosis↓	[88]
Porcine myocardium	GelMA	Human-derived CPCs	3D printed hydrogel patch	vascularization↑, Cardiac function and tissue remodeling↑	[53, 218]
Sheep myocardium	PPy	hESC-derived CPCs	Lyophilized hydrogel patch	Angiogenesis↑, Fibrosis↓, Cardiac function↑	[52]
Porcine myocardium	N/A	hiPSC-CPCs	Injectable hydrogel	Fibrosis↓, LV wall thickness↑, Angiogenesis↑, Cardiac function↑	[219]
Porcine whole heart	Poly(ester carbonate urethane) urea	Rat ADSCs	Hydrogel patch	Cardiac function↑, Fibrosis↓, Vascularization↑	[220]
Rabbit whole heart	N/A	Rabbit BMSCs, and bFGF	Injectable hydrogel	Fibrosis↓, cardiac function↑	[216]
Rat myocardium	N/A	Rat ADSCs	Injectable hydrogel	Fibrosis↓, cardiac function↑, Angiogenesis↑	[221]
Porcine myocardium	N/A	Mouse green fluorescent protein (GFP)-labeled MSCs	Injectable hydrogel	Cardiac function↑, Vascularization↑, Fibrosis↓	[215]
Porcine myocardium	N/A	Mouse primary CMs and SDF-1α	Injectable hydrogel	Angiogenesis↑, Fibrosis↓, Infarction size↓, Cardiac function↑	[44]
Porcine myocardium	N/A	Rat adipose-derived stem cells (ADSCs)	Injectable hydrogel	Angiogenesis↑, Fibrosis↓, Infarction size↓, Cardiac function↑	[51]
Porcine myocardium	N/A	Rat- and pig-derived ADSCs	Cardiac patch	Vascular formation↑	[83]
Rat whole heart	N/A	hiPSC-CMs	Cardiac Patch	Angiogenesis↑, Fibrosis↓, Cardiac function↑	[222]
Porcine myocardium	N/A	hiPSC-CMs and human MSCs	3D printed hydrogel patch	Angiogenesis↑, Cardiac function↑	[90]
Porcine myocardium	N/A	Mouse chemically induced cardiomyocytes (CiCMs)	Injectable hydrogel	Angiogenesis↑, Fibrosis↓, Cardiac function↑	[223]
Porcine myocardium	N/A	hiPSC-CMs and non-myocytes (ECs and CFs)	3D dcECM hydrogel (HEM)	Angiogenesis↑, Fibrosis↓, Cardiac function↑	[224]

Mechanistically, the dcECM matrix enhances cell adhesion through integrin-binding domains, provides biochemical guidance for lineage-specific differentiation, and reduces oxidative and inflammatory stress in the infarct region[209]. Moreover, the tissue-specific microenvironment of dcECM mimics native myocardium and has been shown in vitro to support electrical coupling and, in some studies, synchronous contraction of cardiomyocytes or progenitor cells[66, 210, 211]. Prevascularization strategies using endothelial or MSCs seeded with dcECM-based slices/patches or hydrogels also demonstrated enhanced angiogenesis capacity and vessel formation in vitro and in vivo[212–214].

4.3. Types of cells delivered with dcECM

4.3.1. Mesenchymal stem cells (MSCs)

MSC-laden dcECM scaffolds are the most investigated for post-MI repair. Intrapericardial delivery of mouse green fluorescent protein (GFP)-labeled MSCs loaded in an dcECM-based hydrogel has emerged as an effective strategy (Figure 4A). Compared with direct intramyocardial injection, this approach markedly increases acute cell retention, reduces cell washout, enhances angiogenesis, and further improves LV ejection fraction while limiting fibrosis in a mouse MI models[215]. In addition, the dcECM binding and retaining bFGF significantly enhanced the viability and differentiation of implanted rabbit BMSCs, resulting in augmented neovascularization, reduced infarct size, and enhanced cardiac function in a rabbit acute MI model[216]. Furthermore, the dcECM scaffolds supported efficient human umbilical cord blood-derived MSCs adhesion and proliferation with enhanced viability, highlighting the retention of native myocardial matrix bioactivity[217].

Figure 4.

Small and large animal MI models demonstrating dcECM scaffolds as delivery platforms for therapeutic cells in myocardial repair. (A) (i) Intrapericardial (IPC) injection of mouse-derived MSC-laden dcECM hydrogel in a porcine MI mode. (ii)Thoracoscopic images show catheter access to the pericardial space and delivery of MSCs (yellow arrows). (iii) Post-injection examination revealed no observable pericardial effusion. (iv) Ex vivo IVIS images at day 4 display the biodistribution and retention of NIR680-labeled MSCs within myocardial tissue. Modified and adapted with permission from Ref. [215]. Copyright 2021, Elsevier. (B) Human iPSC-CPCs encapsulated in dcECM hydrogel for MI therapy in rats. (i) Schematic illustration of intrapericardial (iPC) injection and in situ formation of a cardiac patch. (ii) Representative Masson’s trichrome–stained heart sections (red = viable myocardium; blue = scar) collected 4 weeks after treatment (Scale bar: 60 μm). (iii) Infarct size and LV wall thickness were quantified from the stained sections. Left ventricular ejection fraction (LVEF) and left ventricular fractional shortening (LVFS) were assessed by echocardiography. Adapted with permission from Ref.[219] under Creative Commons CC BY license. Copyright © 2021, The Author(s). (C) (i) Pre-vascularized stem-cell cardiac patch fabricated using multiple bioinks containing human-derived CPCs, MSCs, and angiogenic cues. Representative image of the implanted patch (ii) and whole-heart Masson’s trichrome staining from MI, CPC, mix C/M, and patterned C/M groups (iii). (iv) Quantification of fibrosis at 8 weeks demonstrates enhanced structural and functional repair with spatially organized bioinks (n = 3, * p < 0.05, ^# p < 0.05, ^‡ p < 0.05). Modified and adapted with permission from Ref.[88]. Copyright 2016, Elsevier. (D) (i) Schematic of engineered human cardiac tissue fabrication using three cell types, heart ECM (HEM) hydrogel, and a microfluidic platform. (ii) The constructs, containing hiPSC-CMs, were transplanted into a rat MI model. (iii) Cardiac function was assessed by quantifying LVEF and LVFS. (iv) Representative Masson’s trichrome staining at 4 weeks post-transplantation with corresponding fibrosis quantification. (v) Representative CD31-stained images (green) show enhanced capillary formation in the infarct (IZ) and border (BZ) zones, with capillary density quantified at 4 weeks post-transplantation (scale bar = 200 μm). Modified and adapted from Ref.[224] under Creative Commons CC BY license. Copyright © 2024, The Author(s).

4.3.2. Adipose-derived stem cells (ADSCs)

The combination of ADSCs with dcECM scaffolds enables synergistic therapeutic effects, as the dcECM provides a protective microenvironment that protects transplanted cells from the hostile environment in the infarcted myocardium. In a rat MI model, porcine dcECM hydrogels with rat-derived ADSCs reduced infarct size by > 50 % and increased LVEF by ~22 % compared to the dcECM hydrogel alone, attributed to elevated VEGF, EGF, and Ang-1 expression and enhanced capillary density[51]. In addition, a combination of porcine dcECM and poly(ester carbonate urethane) urea with the rat-derived ADSCs formed a cellularized biodegradable, elastic cardiac patch, which significantly improved the engraftment of ADSCs, promoted the neovascularization and enhanced the cardiac function in a rat myocardial ischemic model [220].

4.3.3. Cardiac progenitor cells (CPCs)

CPCs exhibit cardiomyogenic potential but suffer from poor survival after transplantation. Integration within dcECM scaffolds significantly enhances their reparative function. A polypyrrole-loaded cardiogel patch strengthened electrical conductivity and improved human embryonic stem cell (hESC)-derived CPC retention, leading to reduced fibrosis and increased neovascularization in a rat MI model [52]. Similarly, 3D printed myocardial dcECM/GelMA patch loaded with human-derived CPCs enhanced vascularization and supported CPC survival and differentiation following epicardial implantation on rat hearts [53]. Intrapericardial (iPC) delivery of human iPSC-derived CPCs encapsulated within a dcECM hydrogel (Figure 4B) formed an in situ cardiac patch-like structure that significantly enhanced therapeutic retention and attenuated immune response, promoted angiogenesis, reduced infarct size, and ultimately improved LVEF and LVFS in a rat MI model [219]. In addition, a 3D-printed pre-vascularized cardiac hydrogel patch fabricated using stem cell–laden dcECM bioinks demonstrated improved cell survival, vascularization, and structural integration after implantation in a rat MI model (Figure 4C)[88]. The patterned human-derived CPC/MSC within a porcine dcECM patch reduced fibrosis, preserved ventricular wall thickness, and enhanced perfusion after MI, leading to greater functional recovery compared with non-patterned or non-dcECM constructs [88].

4.3.4. Cardiomyocytes (CMs)

Cardiomyocyte-based therapies have attracted significant attention for cardiac repair because transplanted CMs can partially remuscularize the infarcted myocardium and restore contractile function following myocardial injury [225]. Decellularized natural heart ECM has been used as a matrix to deliver CMs in vivo [44, 222]. When transplanted onto an infarcted rat heart, hiPSC-derived cardiac cells seeded dcECM patch showed improved LVEF, reduced infarct size, and increased wall thickness compared with untreated or dcECM patch-only groups at 4 weeks [222]. Moreover, the dcECM hydrogels enhanced mouse-derived chemically induced cardiomyocyte-like cell (CiCM) maturation by improving sarcomere organization, calcium-handling properties, and electrophysiological features in vitro. It is notable that the implantation of CiCMs with the dcECM hydrogel improved ventricular function and reduced adverse remodeling compared with CiCMs seeded on 2D Matrigel in a rat MI model[223]. In addition, a recent perfusable 3D dcECM hydrogel was used to generate human iPSC-derived cardiac tissues for MI treatment (Figure 4D). Mixed populations of hiPSC-derived cardiomyocytes, cardiac fibroblasts, and endothelial cells were embedded in the porcine dcECM hydrogel and matured under microfluidic perfusion[224]. When these engineered human cardiac tissues were transplanted into a rat MI model, they resulted in higher LVEF and LVFS, reduced ventricular dilation, and increased capillary density compared with control groups. These findings demonstrated that the use of dcECM as a supportive matrix for hiPSC-derived cardiac cell delivery is promising for post-MI repair.

4.4. Dynamic interactions between dcECM remodeling and cell fate

Beyond serving as a structural scaffold, the dcECM can function as a dynamic microenvironment that actively regulates cellular behaviors during tissue repair. The therapeutic efficacy of dcECM-based delivery systems is closely linked to its degradation kinetics, which temporally coordinate cell retention, migration, and integration with host tissues. In the early phase following implantation, dcECM provides a biomimetic niche that increases cell retention and survival through its nanofibrous architecture and tissue-specific adhesion ligands [226]. This initial stabilization is critical in the infarcted myocardium, where dynamically mechanical stretching and hostile ischemia and inflammatory microenvironment severely limit cell engraftment [227]. The injectable dcECM hydrogels have significantly improved the retention and survival of implanted cardiomyocytes [44]. As remodeling progresses, the dcECM undergoes cell-mediated enzymatic degradation which is primarily driven by inflammatory cells and matrix metalloproteinases. This progressively reduces matrix stiffness and increases matrix porosity, which facilitates the transition from initial cell retention to cell migration and host integration [177]. Importantly, this degradation process is not simply associated with matrix breakdown, but also generates bioactive fragments, commonly referred to as matrikines, which can modulate cell behaviors and the local immune response [78]. These degradation products also could act as chemoattractants to recruit progenitor cells, regulate macrophage polarization, and promote angiogenesis, amplifying endogenous repair mechanisms[228]. Meanwhile, cell–matrix interactions influence paracrine signaling dynamics because the biochemical and mechanical properties of dcECM could modulate the secretion profile and therapeutic activity of encapsulated cells[229, 230]. This paracrine activity is important to coordinate angiogenesis, immunomodulation, and tissue remodeling over time[231]. Together, these dynamic cell–matrix interactions facilitate dcECM-mediated cardiac repair.

4.5. Synergistic effects with drug release

A growing number of hybrids dcECM systems integrate controlled drug or factor release with cell therapy to further enhance therapeutic efficacy. SDF-1α-loaded dcECM hydrogels have been shown to recruit endogenous progenitor cells and promote angiogenesis, resulting in synergistic functional recovery when combined with mouse primary cardiomyocytes [44]. Similarly, a bFGF-binding dcECM hydrogel significantly enhanced the therapeutic efficacy of rabbit BMSCs and improved ventricular function in a rabbit MI model by promoting angiogenesis and attenuating fibrosis through the sustained activation of the PI3K/Akt and ERK 1/2 pathways [216]. In another study, 3D-printed constructs composed of stem cell-laden dcECM bioinks with VEGFs provided structural and biochemical cues that increased human-derived CPCs and MSCs survival and tissue organization in vivo, and improved myocardial repair and neovascularization after implantation in a rat MI model [88]. These dual-delivery approaches of combining structural and biochemical support from dcECM with localized, sustained therapeutic release represent a convergent strategy to maximize cell retention, angiogenesis, and functional recovery for MI treatment.

5. Translational and clinical progress

Over the past decade, decellularized heart ECM-based biomaterials have been investigated for restoring ischemic or infarcted myocardium tissues [37, 39, 40, 89]. Their intrinsic bioactivity, native structural hierarchy, and immunomodulatory properties confer advantages over synthetic polymers or single-component natural hydrogels, enabling more physiologically relevant repair and regeneration. These advantages have positioned dcECM hydrogels, patches, and hybrid constructs as promising candidates for myocardial repair. To date, substantial progress has been made from rodent proof-of-concept studies to large-animal validation and early-phase clinical evaluation. In particular, the safety and therapeutic efficacy of the injectable dcECM hydrogels have been proven in small and large animal MI models, as well as in clinal trials[89, 163, 232].

Several dcECM–based biomaterials have entered human investigation for cardiac repair in clinical use or in clinical trials listed in Table 4. The VentriGel, from porcine myocardial matrix hydrogel, represents the first-in-class biologic intended for minimally invasive catheter-based myocardial restoration (Figure 5 A and B). Early large-animal studies demonstrated VentriGel’s safety and functional efficacy. Trans-endocardial injection of the solubilized dcECM hydrogel into infarcted pig hearts significantly reduced infarct expansion, improved regional wall thickness, and enhanced ejection fraction without arrhythmogenic events[162, 232]. These data directly supported the Phase I clinical trial (NCT02305602). In the Phase I VentriGel trial, Traverse et al. conducted a prospective, multicenter, open-label study enrolling 15 post-MI patients (2–36 months after infarction)[89]. VentriGel was delivered via NOGA-guided intramyocardial injection using a percutaneous catheter system and demonstrated favorable safety, with no serious device-related adverse events or ventricular arrhythmias observed. Cardiac MRI at 6 months revealed improvements in left-ventricular end-systolic volume (LVESV) and regional strain, particularly in patients treated ≥12 months after MI, indicating reverse remodeling potential. Importantly, this trial demonstrated the feasibility of catheter-based delivery of an acellular ECM hydrogel for MI treatment. After demonstrating safety in Phase I clinical trials, VentriGel was approved by the U.S. Food and Drug Administration to proceed to Phase II clinical studies.

Table 4.

Commercially approved heart-derived decellularized ECM products and undergoing clinical trials for cardiac repair.

Product name	Source	Application	Clinical trials	Ref.
VentriGel	Decellularized porcine myocardium	Injectable MI treatment	NCT02305602, Phase1completed	[89]
CardioCel	Decellularized bovine pericardium	Pediatric cardiac repair	FDA-approved, commercialized	[233]
PeriCord	Decellularized human pericardium	Cardiac/vascular repair	NCT03798353, Phase1	[234]

Figure 5.

An injectable dcECM hydrogel currently under clinical evaluation (VentriGel). (A) Fabrication workflow of VentriGel from porcine ventricular myocardium. Fresh myocardium is minced, decellularized, lyophilized, milled into a fine powder, and enzymatically digested to yield a solubilized dcECM solution suitable for injection through a standard syringe and 27-gauge needle. Representative image shows left-ventricular pacing one-week post-injection in a rat model for arrhythmogenesis assessment. Modified and adapted with permission from Ref. [162], Copyright 2012, Elsevier. (B) First-in-human clinical evaluation of porcine dcECM hydrogel (VentriGel) in patients with early (60 days–1 year) or late (1–3 years) post-MI injury. VentriGel was delivered via NOGA-guided transendocardial catheter injection into the left ventricle. Clinical follow-up at 3 and 6 months demonstrated safety, feasibility, improved 6-min walk distance, reduced NYHA class, and attenuated LV remodeling in late-MI patients. Modified and adapted with permission from Ref. [89], Copyright 2019, Elsevier.

Other dcECM-based scaffolds have also completed clinical trials or are currently approaching early-phase clinical trials. CardioCel^® derived from decellularized bovine pericardium, has been approved by FDA for congenital and valvular reconstruction, establishing a clinical benchmark for dcECM processing, sterilization, and long-term biocompatibility in humans[233]. PeriCord is a human decellularized pericardial matrix engineered as an epicardial scaffold for early intervention after MI. In a first-in-human Phase I clinical study (NCT03798353), PeriCord was implanted surgically onto the epicardial surface of the infarcted heart to evaluate safety, feasibility, and preliminary efficacy. Early outcomes demonstrated favorable tolerability, with no device-related serious adverse events. Imaging and biomarker analyses showed reduced ventricular dilation, attenuated fibrosis, and improved regional contractility in treated patients. These results support the translational potential of human-derived pericardial ECM as a mechanically robust, biologically active scaffold capable of modulating post-MI remodeling and promoting reparative neovascularization[234].

To date, VentriGel remains the only clinically tested dcECM hydrogel in human subjects, yet its success provides a template for the field. The transition from animal models to clinical feasibility demonstrates that naturally derived ECM biomaterials can be produced reproducibly, delivered safely, and elicit measurable functional improvement without immune rejection. As regulatory familiarity increases, these products are expected to join the growing class of bioactive combination products for ischemic heart disease and heart failure management.

6. Challenges and future perspectives

6.1. Balancing biological activity, immunogenicity, and manufacture

A central challenge in dcECM-based biomaterials lies in preserving bioactivity while minimizing its immunogenic risk and ensuring manufacturing consistency. Incomplete removal of cellular debris or xenogeneic epitopes may elicit chronic inflammation or fibrosis, whereas excessive processing can denature structural proteins and eliminate essential growth-factor binding domains. Optimization of decellularization and post-processing strategies is critical to preserve ECM composition and bioactivity. Compliance with residual DNA criteria is also essential for clinical translation.

Beyond biological considerations, large-scale translation of dcECM biomaterials is constrained by manufacturing heterogeneity and limited process control. Source tissues inherently vary in donor age, anatomical region, and disease status and minor deviations in decellularization efficiency or matrix composition can significantly alter biochemical, mechanical, and rheological properties, leading to batch-dependent therapeutic performance[235]. Reliable clinical translation of dcECM technologies requires scalable, GMP-compliant manufacturing pipelines. Batch-to-batch variability necessitates rigorous donor selection, standardized decellularization methods, and release criteria, including DNA, collagen, and GAGs measurements[236]. Advanced proteomic fingerprinting, immune cell co-culture assays, and rheological “bio-signatures” are increasingly recognized as additional quality-control tools for predicting host response and identifying out-of-specification lots[224, 237–239].

The dcECM biomaterials are biodegradable and are not intended for permanent implantation. Therefore, their long-term safety risks are generally lower than those of permanent implantable materials. Numerous preclinical studies in rats and pigs, as well as early clinical trials, have demonstrated the safety of these dcECM materials[89, 162, 163, 232]. To date, injectable dcECM biomaterials have generally been evaluated following a single administration, and there are no reports about repeated dosing in preclinical studies. Thus, the in vivo safety and performance for repeated dosing remain unknown and should be carefully evaluated if the repeated dosing is executed. In addition, future studies are needed to assess long-term functional benefits and integration of dcECM biomaterials with the host myocardium [89, 232].

6.2. Sterilization and storage stability

Sterilization and long-term preservation remain particularly challenges for dcECM-based biomaterials, as ECM bioactivity is highly sensitive to irradiation and chemical residues. Ethylene oxide and carefully validated low-dose γ-irradiation are acceptable when shown not to compromise collagen integrity or gelation behavior [240]. However, these methods still need to be further validation [241]. Beyond sterilization, preservation methods critically influence scaffold stability and clinical usability. For preservation, lyophilization is the predominant method for dcECM products, as it effectively stabilizes matrix composition and bioactivity while enabling extended shelf life. Lyophilized dcECM powders have been shown to maintain structural integrity and functional performance for long-term storage when stored at −20 °C or −80 °C, whereas rehydrated formulations require frequent potency verification to ensure batch consistency and clinical reliability[163, 242].

6.3. Achieving spatiotemporal control of therapeutic release

Most dcECM hydrogels exhibit diffusion-controlled release of encapsulated bioactive factors and drugs, and their degradation further contributes to therapeutic release in vivo[183, 195]. Achieving precise spatiotemporal control of therapeutic delivery is essential for maximizing efficacy and minimizing off-target effects [243, 244]. Integration of factor-binding domains and micro/nanocarriers within dcECM matrices enables controllable and programmable release kinetics [55, 245]. Further tuning crosslinking density and degradation kinetics may alter temporal control of therapeutic release [246]. Such designs are particularly promising for sustainable delivery of growth factors, extracellular vesicles, or small-molecule drugs during cardiac remodeling and repair.

Despite these advances, precisely controlling therapeutic release with post-MI healing remains challenging. Ideally, the therapeutic release should be aligned with the temporal progression of cardiac repair. After MI occurrence, the healing process progresses from an inflammatory phase involving immune cell recruitment and debris clearance, to a proliferative phase characterized by angiogenesis and matrix deposition, and subsequently to a remodeling phase associated with scar formation and cardiac tissue remodeling[247, 248]. However, these phases overlap substantially during cardiac repair[248, 249], thus it is hard for the biomaterials to execute sequential therapeutic release to specific stages of healing.

6.4. Integration with emerging technologies

In recent years, advances in biomaterials science and biomanufacturing have expanded the functional versatility of dcECMs. For example, 3D bioprinting with dcECM inks provides structural precision and patient-specific geometry, enabling the creation of vascularized myocardial or pre-vascularized constructs with recapitulated native tissue anisotropy[250, 251]. Furthermore, micromanufacturing strategies, including microfluidic fabrication, emulsion templating, and electrospray-based techniques, enable dcECM to be engineered into microgels and then granular scaffolds with well-defined size, porosity, and release profiles, offering new opportunities for injectable, minimally invasive, and modular therapeutic delivery[252–255]. Besides, smart or stimuli-responsive systems, such as pH, enzyme, or ROS-triggered hydrogels, allowing on-demand drug release in response to the dynamic post-MI microenvironment[256–258].

6.5. Toward tunable and precision regenerative cardiology

The dcECM technologies have potentials to contribute to precision regenerative cardiology. However, the clinical feasibility of generating patient-specific myocardial dcECM remains limited [50]. Current approaches primarily are based on well-characterized allogeneic or xenogeneic dcECM [259]. Meanwhile, engineered dcECM biomaterials are being developed to mimic the patient’s myocardial microenvironment [50, 260]. Integration of patient-derived omics data with advanced manufacturing techniques, such as 3D bioprinting and microfluidic systems, further enable fabrication of customized cardiac constructs with spatially controlled composition, anisotropic architecture, and tunable mechanical and biochemical properties [261–263]. Such platforms may be able to engineer dcECM-based biomaterials for specific pathological conditions, including infarct stiffness, remodeling stage, and inflammatory status, and then strengthen therapeutic relevance and efficacy without requiring fully autologous ECM sources.

Combining with stem cells, growth factors, genes, or RNA therapeutics, dcECM-based biomaterials would further improve cardiac repair through providing a bioactive microenvironment that supports cell survival, integration, and functional recovery. Importantly, dcECM hydrogels and patches can serve not only as structural bioactive scaffolds but also as localized delivery vehicles. Promising preclinical and early-phase clinical results demonstrated the feasibility and safety of injectable dcECM hydrogels as well as their translational potential. Advances in dcECM processing, controlled-release engineering, integrated therapeutic delivery, and scalable manufacturing will be critical for further clinical translation. Rather than fully patient-specific materials, future strategies would focus on adaptable dcECM platforms engineered to meet patient-specific pathological and biomechanical needs. Integration of dcECM with smart-release systems and advanced biofabrication technologies may enable precision-engineered regenerative therapies. In short, the tissue-specific complex dcECM is a promising bioactive material for drug or cell carriers for cardiac repair.

Data availability

No data was used for this review paper.