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Article

Decellularized Extracellular Matrix/Gellan Gum Hydrogels Enriched with Spermine for Cardiac Models

by
Luca Di Nunno
1,2,
Marcin Wekwejt
2,3,
Francesco Copes
2,
Francesca Boccafoschi
1,* and
Diego Mantovani
2,*
1
Laboratory of Regenerative Anatomy, Department of Health Sciences, University of Eastern Piedmont, 28100 Novara, Italy
2
Laboratory for Biomaterials and Bioengineering (CRC-Tier I), Department of Mining, Metallurgical and Materials Engineering and Regenerative Medicine, CHU de Québec Research Center, Laval University, Québec City, QC G1V 0A6, Canada
3
Department of Biomaterials Technology, Faculty of Mechanical Engineering and Ship Technology, Gdańsk University of Technology, 80-222 Gdańsk, Poland
*
Authors to whom correspondence should be addressed.
Gels 2026, 12(2), 118; https://doi.org/10.3390/gels12020118
Submission received: 24 December 2025 / Revised: 26 January 2026 / Accepted: 27 January 2026 / Published: 28 January 2026
(This article belongs to the Special Issue Recent Advances in Novel Hydrogels and Aerogels)
Browse Figures
Versions Notes

Abstract

The physiological relevance of in vitro models is limited because conventional two-dimensional cell culture systems are unable to replicate the structural and functional complexity of native tissues. Extracellular matrix (ECM)-mimetic hydrogels have become important platforms for tissue engineering applications. This work developed hybrid hydrogels that mimic important biochemical and mechanical characteristics of cardiac tissue by combining decellularized bovine pericardium-derived (dBP) ECM, gellan gum (GG), and spermine (SPM). Although dBP offers tissue-specific biological cues, processing compromises its mechanical integrity. This limitation was overcome by adding GG, whose ionic gelation properties were optimized using DMEM and SPM. The hydrogels’ mechanical, biological, physicochemical, and structural characteristics were all evaluated. Under physiologically simulated conditions, the formulations showed quick gelation and long-term stability; scanning electron microscopy revealed an interconnected, ECM-like porous microarchitecture. While uniaxial compression testing provided Young’s modulus values comparable to native myocardium, rheological analysis revealed a concentration-dependent increase in storage modulus with increasing SPM content. H9C2 cardiomyoblasts were used in cytocompatibility studies to confirm that cell viability, morphology, and cytoskeletal organization were all preserved. All of these findings support the potential application of dBP−GG−SPM hydrogels in advanced in vitro cardiac models by showing that they successfully replicate important characteristics of cardiac ECM.

Graphical Abstract

1. Introduction

Cardiovascular diseases are the leading cause of death globally, claiming 17.9 million lives worldwide every year [1]. The complexity of these multifactorial conditions, influenced by both genetic and environmental factors, presents significant challenges for accurate modelling. Methods for assessing cardiac cytotoxicity and rhythm disturbances remain underdeveloped, hindering their ability to predict several adverse drug effects once tested in vivo [2]. This is because traditional models lack the biochemical, mechanical, and electrical signals within cell cultures needed to maintain the functional characteristics and differentiate into cardiomyocytes [3,4,5]. One of the critical aspects of cardiac tissue modelling is the mechanical mismatch that cells experience once cultured in a monolayer [6]. Therefore, there is a need for advanced biomaterials to improve model reliability.
Among the materials being used in cardiac modelling, natural polymer-based hydrogels have received considerable attention due to their structural similarity to the native extracellular matrix (ECM), their high water content, which facilitates nutrient diffusion, and their ability to guide cellular adhesion and differentiation [7,8,9]. By definition, the ECM is a sophisticated network of proteins, glycoproteins, glycosaminoglycans, and proteoglycans that supply structural and functional support to cells [10,11]. To replicate the structural complexity and mechanical stability of native tissues, diverse combinations of polymers and crosslinkers have been employed, while biochemical components are incorporated to provide appropriate cellular signals. Natural polymers offer advantages for developing advanced biomimetic engineered tissues and innovative regenerative approaches [12]. ECM-derived materials (e.g., collagen, fibrin, and hyaluronic acid) closely reproduce native biochemical signalling, whereas polysaccharides such as alginate and gums are primarily used to tailor mechanical properties and scaffold architecture. In addition, their degradation products are often non-toxic and can undergo enzymatic modifications, indicating that polymers are adequate for dynamic tissue remodelling [13].
The use of decellularized extracellular matrix to mimic a complex microenvironment for sophisticated in vitro models is a promising approach in tissue engineering [14,15]. Decellularization consists of removing the cellular component from tissues, leaving intact the biochemical composition of the native ECM. This process typically requires a combination of balanced physical, chemical, and enzymatic treatments to remove all traces of cells without damaging the integrity of the extracellular microenvironment [16]. These materials retain key components, including structural proteins, growth factors, and signalling molecules, which play a pivotal role in cell adhesion, proliferation, and differentiation. Furthermore, when mixed with other polymers, dECM can significantly enhance bioactivity, bridging the gap between bioinert scaffolds and biomimetic tissue environments. Bovine pericardium–derived extracellular matrix was chosen due to its high availability, established decellularization protocols, and documented clinical use in cardiovascular applications. The tissue is enriched in fibrillar collagens, predominantly type I, which are key structural components of cardiac ECM and contribute to relevant biological cues. Compared with myocardial tissue, pericardium provides higher material yield, reduced batch-to-batch variability, and improved process reproducibility, supporting its selection as a practical and translationally relevant ECM source. However, significant shortcomings of most dECM-based hydrogels include a lack of long-term stability and poor mechanical properties, which present challenges in mimicking a functional myocardium [17,18].
To address these limitations, gellan gum (GG), a microbial-derived anionic polysaccharide, has been extensively studied for biomedical applications [19,20]. In tissue engineering, temperature and ion-responsive gelation make GG highly tunable in controlling drug-release systems and scaffold development [21,22]. Due to its rapid and adjustable gelation mechanism, which can occur at physiological temperature and pH, it is a promising candidate for food and biomedical applications [23,24]. The carboxylic functional groups of GG offer a wide range of opportunities for chemical modification and crosslinking strategies, enabling fine control over its mechanical and physicochemical properties [25]. However, native GG does not induce cell adhesion and must be functionalized to induce cellular responses [23,26]. While GG exhibits greater intrinsic stability and processability than decellularized ECM, additional crosslinking strategies can be employed to further tailor its mechanical integrity and biological performance for tissue engineering applications. In this context, spermine-based crosslinking has been explored as an effective approach to modulate GG network formation while also introducing bioactive features.
Spermine (SPM), a natural polyamine, can bind electrostatically to the anionic groups of GG, strengthening the hydrogel network [26]. Beyond structural benefits, SPM is a key biological molecule in regulating oxidative stress and promoting cell survival [27,28]. Notably, in vivo studies have demonstrated that spermine induces autophagy by modulating the mTOR pathway, thereby contributing to cardio-protection, reducing apoptosis, and enhancing mitochondrial function [29,30]. These findings highlight the dual role as both a structural enhancer in scaffold fabrication and a bioactive molecule for biomedical applications [31,32]. However, concentration must be carefully evaluated because the cytotoxic effects of bioamine have been previously reported [33,34].
Beyond cytocompatibility considerations, mechanical properties are pivotal in controlling cellular activities, especially in tissues that undergo continuous mechanical stresses, such as the heart [35,36]. Thus, hydrogels used for cardiac applications not only need to be capable of resisting the mechanical loading typical of the natural heart, but they must also transmit these forces to the encapsulated cells [37]. This requires a delicate balance between viscoelastic compliance and stiffness, which must be achieved through careful control over composition and crosslinking density [26]. The balance between stability over time, mechanical properties, and bioactivity needs precise optimization. Hence, designing hybrid hydrogels based on cardiac dECM, GG, and SPM may hold potential for advancing cardiac tissue modelling [38,39]. Such biomimetic material provides new possibilities for disease modelling, drug screening, and regenerative therapy [40,41]. To date, only one group mixed cartilage dECM with GG using Ca2+ ions and glutaraldehyde as crosslinking agents, demonstrating the shear-thinning and bioprinting capability for effective cartilage repair [42].
Therefore, the aim of this research was to design, develop, and validate a hydrogel based on decellularized bovine pericardium (dBP) extracellular matrix, blended with gellan gum (GG) and enriched with spermine (SPM), to reproduce selected biochemical and mechanical features of cardiac extracellular matrix.

2. Results and Discussion

2.1. Microstructure

The dBP hydrogel exhibited a lamellar microstructure with a dense network of interconnected pores, indicating an anisotropic organization (Figure 1). In contrast, GG showed a more compact, sheet-like morphology. The blended formulations (dBP−GG and dBP−GG−SPM) displayed transitional features combining both lamellar and sheet-like domains, characteristic of hybrid architectures. No apparent morphological alterations were observed upon SPM incorporation.
Quantitative morphometric analysis, summarized in Table 1, supports qualitative observations. Mean pore diameter varied from 22.78 μm (dBP) to 43.57 μm (dBP−GG). The blend without SPM exhibited the largest maximum pore size (73.62 μm). Spermine addition induced a reduction in pore size without altering the pore distribution. Pore coverage analysis yielded percentages ranging from 45.33 to 55.45%, where dBP−GG−SPM1 had the largest area covered by pores. In general, no statistical differences were identified. dBP−GG hydrogels retained a porous structure, and the addition of SPM did not alter the structural integration among the polymers.

2.2. Chemical Characterization

FTIR spectroscopy was used to characterize the chemical structure of the dBP−GG hydrogels and to assess potential modifications following SPM incorporation (Figure 2A,B). Characteristic collagen absorption bands were identified in the dBP spectrum, including amide I at 1655 cm−1, amide II at 1546 cm−1, and amide III at 1184 cm−1, confirming the protein-rich composition of the decellularized pericardium (dBP) [43]. The GG spectrum displayed a broad O–H stretching band at 3300–3400 cm−1 and intense carbohydrate-related vibrations between 950 and 1150 cm−1, attributed to the glycosidic C–O–C linkage within the polysaccharide backbone [42]. In the dBP−GG blend, the principal spectral features deriving from both polymers were retained, demonstrating successful incorporation without evident phase separation or structural degradation. The symmetric and asymmetric stretching bands of COO− (~1358 and 1428 cm−1, respectively) appeared unmodified following SPM addition. This could also be due to overlapping with other functional groups present, especially from dBP. Finally, no specific peak shifts or new absorption bands were observed.

2.3. Degradation

The degradation profiles of the different hydrogel formulations over the 7-day period revealed distinct stability trends (Figure 3). During the first three days, no significant differences in mass loss were observed among the specimens, indicating comparable short-term stability. However, by day 7, the dBP hydrogel exhibited a pronounced degradation, with 56.6 ± 3.4% of mass loss, highlighting its limited structural stability in the absence of GG. In contrast, the GG control remained stable throughout the study, with negligible mass loss, underscoring its superior intrinsic stability under the tested conditions. Hydrogels containing both dBP and GG demonstrated a desired mass loss profile; dBP−GG degraded 10.6 ± 7.1%. No statistically significant differences were identified, adding spermine; in fact, these conditions resulted in a degradation of 17.6 ± 5.2%, 15.7 ± 4.0%, and 13.7 ± 2.6% for SPM1, SPM2, and SPM3, respectively.

2.4. Mechanical Characterization

The storage modulus of the dBP was 183.4 ± 23.5 Pa, as previously reported in our study evaluating batch-to-batch variability of dBP hydrogels [44]. All the blended specimens showed a statistically significant increase compared to the pristine dBP, with a corresponding increase in the concentration of SPM (Figure 4). The pristine GG possessed the highest values of 3174.1 ± 337.2 Pa, while the blends increased from 766.5 ± 34.5 to 1109.5 ± 94.6, 1683.2 ± 241.6, and 2165.7 ± 483.9 for dBP−GG specimens containing SPM1, SPM2, and SPM3, respectively, as illustrated in Figure 4. Gelation of dBP hydrogels occurs within 10–15 min, but G’ continues to increase slightly in the next hours, while dBP−GG samples take less than 1 min to start undergoing gelation and remain constant over time (data not included).

2.5. Uniaxial Compressive Test

Compressive testing showed that pristine dBP hydrogels were too soft and non–self-supporting to be evaluated under uniaxial load. In contrast, GG hydrogels exhibited a significantly higher stiffness, with a Young’s modulus of 43.3 ± 3.3 kPa. However, the addition of 0.083 µM of SPM (GG-SPM1) did not induce a significant increase in stiffness (45.2 ± 12.0 kPa), indicating that SPM had a minimal effect on the bulk mechanical reinforcement. When blended with dBP, the hydrogels exhibited a reduction in Young’s modulus compared to pristine GG, likely due to the softer nature of the collagen-rich matrix. The dBP-GG blend, with and without SPM, exhibited a modulus in the range of ~22.4–25.7 kPa. Only at the highest spermine concentration (SPM3) was a slight increase observed, although this was not statistically significant. Notably, enhanced variability was recorded for SPM3 samples. Overall, as shown in Figure 5, while spermine addition did not remarkably improve the Young’s modulus in either GG or dBP-GG formulations, dBP-GG hydrogels maintained stiffness within the physiological range of native myocardial tissue (i.e., 20 kPa) [26].

2.6. Biological Characterization

Cell viability on the hydrogels was evaluated for up to 7 days using the MTS assay, as shown in Figure 6. On day 1, samples containing GG exhibited significantly lower viability compared to CTRL. However, by day 7, a significant increase in metabolic activity was observed in all samples (p < 0.01), except for dBP-GG-SPM3.
Among the hydrogels, collagen and dBP showed the highest increase in enzymatic activity, corresponding to a 3.5 ± 0.1-fold and 2.4 ± 0.4-fold increase. The GG group displayed a metabolic activity at 1.5 ± 0.1-fold. Adding dBP into GG (GG-dBP) results in comparable metabolic activity at 1.7 ± 0.2-fold. This value remained constant for dBP-GG-SPM1 and dBP-GG-SPM2 at 1.7 ± 0.1-fold and 1.5 ± 0.3-fold, respectively. In contrast, higher spermine concentrations in dBP-GG-SPM3 resulted in toxicity, maintaining only 0.4 ± 0.1-fold of the original metabolic activity of CTRL on day 1.
Fluorescence staining with phalloidin (F-actin, green) and DAPI (nuclei, blue) at day 7 confirmed the MTS results, indicating adequate cell adhesion and spreading across all hydrogel formulations, except for the highest spermine concentration (GG-dBP-SPM3), which showed minimal cell presence. H9C2 cardiomyoblasts on dBP and GG-dBP samples showed a well-organized cytoskeletal network and an elongated morphology, consistent with active spreading and proliferation, resulting in a cell monolayer by day 7 (Figure 7).

2.7. Discussion

The present study successfully developed and characterized novel hydrogels designed to mimic the biochemical and mechanical properties of cardiac tissue. The dBP-GG blend yields different tunable self-standing hydrogels with balanced properties, including fast gelation, reduced degradability, and a desired Young’s modulus, as well as cytocompatibility towards cardiac progenitor cells (H9C2), demonstrating potential for advanced cardiac tissue modelling.
The degradation studies revealed that the dBP and GG hydrogels exhibited different degradation rates. While the dBP sample lost 56.6% of its mass by day 7, dBP-GG hydrogels, regardless of SPM concentration, demonstrated gradual degradation, losing only about 15% of their mass compared to their initial mass. Previously reported results have shown that oxidized GGs with different molecular weights and blended with chondroitin sulphate undergo approximately 20% mass loss over 7 days in the presence of lysozyme and up to 30% with trypsin, indicating the influence of specific enzymatic activity on scaffold stability [45,46]. In contrast, many tissue engineering-oriented studies have examined GG degradation in ionic solutions, such as phosphate-buffered saline or culture medium, under physiologically relevant conditions (pH 7.4 and 37 °C). In these conditions, GG mass remains unaltered over extended periods, up to 168 days [47], which agrees with our findings (Figure 3) and with the intended in vitro application. Moreover, most in vitro-oriented studies typically report slow degradation rates limited to ion exchange or osmotic swelling effects [48]. For bioinert polymers (like GG), stability is often preferred over degradation, while a controlled degradation rate can be beneficial in supporting ECM cell remodelling while maintaining structural integrity. Recently, the first report of a decellularized ECM from cartilage and GG was reported; these hydrogels lose 30% of their mass in the first week and are fully degraded in 14 days, favouring cartilage repair and ECM components release [42]. Furthermore, most polysaccharides are notable for providing thickening properties, water solubility, and scalability, enabling the design of advanced, microstructured scaffolds [49,50].
The combination of dBP and GG into a hybrid network resulted in distinct microstructural variations, with dBP exhibiting a lamellar morphology and GG presenting a sheet-like structure. The blended formulations displayed an intermediate morphology characterized by both lamellae and sheets, effectively combining the structural features of each polymer. The pore size analysis indicated that while the dBP samples had a mean pore diameter of 22.78 µm, GG samples exhibited larger pores, averaging 28.34 µm. Those results are consistent with previous findings by Chen et al. for cell culture meat with realistic textures (based on GG and gelatin) [51,52] and Adel et al. for bioactive hydrogels for tendon tissue regeneration (based on GG and carrageenan) [53]. Porosity is one of the crucial features for cellular infiltration and nutrient diffusion, which are vital for tissue engineering [54].
The FTIR spectra of dBP, GG, and their blend exhibited characteristic absorption bands corresponding to the functional groups relevant to the individual components (Figure 2). Upon blending, no new absorption peaks were detected, suggesting that no covalent bond formation occurred [55]. Following SPM addition, no significant or systematic changes were observed in the COO−, NH3+, or O-H/N-H stretching regions. Although minor differences in band position or intensity could suggest ionic- or hydrogen-bonding interactions, these effects cannot be unambiguously attributed to specific SPM-related interactions. These may result partially from band overlapping, different hydration states, or be related to the high variability in the organic content of dBP, which has been previously shown to affect the precision of FTIR interpretation [43]. Comparable spectral features have been attributed in the literature to amine-driven collagen crosslinking [56] or to phenolic–polysaccharide interactions in GG systems [57], which could also be relevant in this case. Importantly, our observations are consistent with previously published reports. Torres-Figueroa et al. reported difficulties in the interpretation of ionic crosslinking due to peaks overlapping and slight changes in FTIR spectra following the addition of spermidine into comparable polymer-based systems (gellan gum, karaya gum, and humic acid), attributing this to the non-covalent nature of the interactions involved [58]. It should also be noted that the application of bioamine in hydrogel systems remains a relatively recent approach, which is reflected in the currently limited availability of detailed spectroscopic studies in the literature.
The mechanical performance of the tested hydrogel was used to determine whether amine-mediated crosslinking contributed to the mechanical reinforcement of the matrix. Measurements of the storage moduli (G′) reflect the hydrogel’s dynamic viscoelastic behaviour during small oscillations, which mimic what cells experience in the ECM. In fact, cardiomyocyte function is particularly responsive to dynamic stiffness, stress relaxation, and elastic recovery. A linear correlation between substrate stiffness and increased contractility and myofibrillar organization has been established [59]. This is because a stiffer substrate increases the amplitude of the calcium transient and force production.
The measured enhancement of G′ following crosslinking with SPM should enhance cardiac differentiation and effectively transmit synchronized contractions. Compressive measurements, in contrast, assess the quasi-static stiffness of the matrix and play a more structural role.
The comparable values of compressive Young’s moduli indicate stiffness within a physiologically acceptable range, which would otherwise cause problems for cardiomyocyte spreading and contractions. Surprisingly, the uniaxial compressive test revealed no statistically significant increase in Young’s modulus for dBP-GG hydrogels upon the addition of SPM, probably due to competition with ions and dBP functional δ+ groups. This is, in contrast to previous studies in GG and spermine hydrogel crosslinking with SPM [60]. Notably, all the blended dBP-GG hydrogels exhibited stiffness values in the range of native myocardial tissue (~20 kPa) [23,61]. Furthermore, the Young’s modulus in our results (Figure 5) aligns with GG hydrogels previously selected as candidates to mimic cardiac mechanical properties [23]. To ensure consistency, the uniaxial compressive parameters, hydrogel volume, and geometry were kept identical to reduce as much as possible methodological variability [40,41].
In contrast, the oscillatory test (Figure 4) revealed a statistically significant concentration-dependent effect of SPM. The storage modulus (G′) of pristine dBP hydrogels was found to be consistent with previous findings [44]. Upon incorporation into dBP-GG, G′ increased from 766.5 ± 34.5 Pa (medium-only crosslinking) to 1109.5 ± 94.6 Pa, 1683.2 ± 241.6 Pa, and 2165.7 ± 484.0 Pa for SPM1, SPM2, and SPM3, respectively. These results suggest that SPM participates more actively in reinforcing the viscoelastic (frequency-dependent) network rather than the bulk compressive properties. Comparable storage modules were reported in the GG–collagen interpenetrating network for burn wound therapy [62]. Moreover, previous studies have suggested that only hydrogels with a storage modulus greater than 1 kPa support cells’ mechanotransduction [38]. These results collectively show that SPM modulates dynamic viscoelastic properties, which are particularly important for mimicking native cardiac cells’ dynamics without excessively stiffening the construct. Therefore, the SPM1 and SPM2 addition into our hydrogels was selected as the optimal concentration after an extensive polymer concentration optimization process [GG: 0.20–0.75% (w/v); dBP: 2–5 mg/mL; and SPM: 0.083–0.500 (µM)]. The use of HEPES/sucrose buffer was important to maintain GG liquid at 37 °C, as previously optimized by Kellomäki et al. [23].
GG degradation products are considered non-toxic; therefore, they are generally recognized as safe (GRAS) by the Food and Drug Administration (FDA) for food manufacturing and biomedical applications. The main drawback of employing GG alone is that it does not support cell adhesion and differentiation [22,63]. Our group previously demonstrated that dBP hydrogel can provide adhesion sites, promote viability, and induce myogenic differentiation markers like α-smooth muscle actin, myogenin, and myosin heavy chain [64]. In addition to biochemical cues, combining mechanical stimulation with biomaterials can further improve cardiac cells’ maturation and cytoskeletal organization using cyclic substrate deformation (5%) at a frequency of 1 Hz for 7 days [65]. In this study, we combined SPM because increasing Ca2+ or Mg2+ ions for additional crosslinking (a common strategy for alginate and other gums) can inhibit cell function and induce cell death [66]. Therefore, the concentration of these endogenous molecules is crucial for optimal cellular response.
Our results highlight a dose-dependent relationship between SPM concentration and cell viability for the tested hydrogels. In detail, low-to-intermediate SPM levels in the gel (0.083–0.167 µM) maintained cellular migration, viability, and adhesion (Figure 6 and Figure 7). In contrast, 0.335 µM of SPM in the hydrogel compromised cell metabolic activity and morphology, ultimately leading to cell death. These concentrations are higher compared to those tolerated in a 2D monolayer of cells; the literature reports a critical range above ~10 µM free spermine as detrimental to intestinal cells [34]. This higher tolerance can be attributed to an effective lower concentration of free SPM due to ionic interactions with dBP-GG hydrogels. Moreover, similar concentrations were used to design biomaterials for multiple biomedical applications [40,67,68]. Interestingly, in vivo research (8-month-old mice with type 2 diabetes) of bioamine treatment showed anti-ageing and positive cardiac effects up to 10 mg/kg/d [32,69]. Hence, this result suggests not only higher in vivo tolerance but also that the addition of SPM into biomaterials can further improve bioactivity and tissue remodelling, and it can open new possibilities for bioactive materials for regenerative medicine strategies.
Our study also had some limitations: for instance, FTIR spectroscopy did not allow for the unquestionable identification of interactions related to SPM addition. Therefore, future studies should employ complementary techniques, such as Raman spectroscopy, to provide additional molecular-level insight into the proposed system. Moreover, degradation and mechanical tests were performed in static conditions. This may underestimate the actual degradation rate and mechanical fatigue expected in a dynamic setting. In addition, a more comprehensive evaluation of spermine release kinetics and functional tests on cells should be considered to better define long-term biological responses. Lastly, although H9C2 cardiac progenitor cells are widely used for cardiac cytocompatibility testing, these cells do not fully represent the contractile and electrophysiological properties of mature cardiomyocytes. However, based on the results, we selected the dBP-GG formulation containing less than 0.167 µM as the most promising. Further studies will also consider using primary cardiac cells and functional assays (e.g., calcium transients, contraction force, or connexin-43 expression) to evaluate dBP-GG-SPM bioactive potential in vitro.

3. Conclusions

In summary, a gellan gum with decellularized bovine pericardium hydrogel crosslinked with spermine presents a promising and biomimetic platform for modelling cardiac tissue. By combining the biochemical cues of dBP with the tunable gelation and mechanical properties of GG, these systems form stable, self-supporting hydrogels that mimic the native cardiac microenvironment. Scanning electron microscopy confirmed a hybrid microarchitecture, where the fibrous and porous structure of dBP is preserved and complemented by the compact, dense matrix of GG. Further, the GG and dBP hydrogel with SPM showed significantly enhanced stiffness and stability over time, reaching values of Young’s modulus within the physiological range of cardiac tissue. Importantly, in vitro studies have verified the cytocompatibility of the materials, demonstrating sustained viability and the typical morphology of cardiac progenitor cells on the hydrogels. Collectively, these findings highlight the potential of dBP-GG-SPM hydrogels to serve as a highly tunable system, modulated by SPM concentration (found at an appropriate threshold), capable of providing an optimized and biologically supportive microenvironment, thus offering a reliable substrate for the development of physiologically relevant in vitro cardiac models.

4. Materials and Methods

4.1. Materials

dBP hydrogel was provided by Tissuegraft srl (Novara, Italy; Italian patent number 102020000007567, patented on 29 April 2022; international patent number PCT/IB2021/052779 pending).
Gellan gum (Gelzan™), low acyl, MW 1 kg/mol; spermine tetrahydrochloride (SPM); 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES); and sucrose were all high laboratory-grade purity and were obtained from Sigma Aldrich (Milan, Italy).
In brief, a 10% (w/w) sucrose solution in deionized water was used as a solvent for GG and SPM. GG was dissolved at a concentration of 10 mg/mL by heating to 85 °C for 1 h. SPM was prepared at three concentrations: 0.175, 0.35, and 0.70 mg/mL in 10 mL (SPM1, SPM2, and SPM3, respectively). All solutions were sterile filtered with 0.2 μm from Sartorius Minisart™ (Göttingen., Germany), adjusted to pH 7.4 with 25 mM HEPES buffer, and stored at 4 °C. dBP-GG blends were formed by mixing GG and dBP solutions (1:1 v/v) at room temperature for 30 min, followed by the addition of SPM solutions (1:5 v/v). The pre-gels were cast and incubated at 37 °C for 30 min before cell seeding. Cell suspensions in DMEM were gently overlaid onto the hydrogel. Final concentrations of SPM are listed in Table 2, following the same rationale as previously reported by Kellomäki et al. [60].
As controls for materials characterization, dBP and GG were also included in the study at the final concentration of 5 mg/mL in PBS or blended, without the addition of SPM. For cell culture experiments, CTRL represents a standard plastic-treated culture plate (Carlo Erba, Milan, Italy) as a standard control used to normalize the data, while type I collagen extracted from rat tail tendons solubilized in 0.02 N acetic acid was selected as a standard polymer at a final concentration of 4 mg/mL [70].

4.2. Microstructure

Hydrogel morphology and porosity were examined using scanning electron microscopy (SEM, JSM-IT500 InTouchScope™ series, JEOL, Akishima, Japan) at 15 kV voltage. Samples were mounted on aluminum stubs and gold-coated (5 nm gold layer, Smart Coater, JEOL). Briefly, the samples were lyophilized, soaked in liquid nitrogen, and cut transversally with a bistoury. Images were acquired at 50×,100×, and 150× magnification. To measure porosity, particle function analysis was selected using Mountains10 (Digital Surf, France) software. Brightness and contrast were adjusted to allow consistency across the three pictures for each sample. The threshold was kept constant at 52% of the material ratio.

4.3. Chemical Characterization

In order to characterize the chemical composition of hydrogels (after their lyophilization), Fourier-transform infrared spectroscopy with attenuated total reflectance (Agilent Cary 660 FTIR, Agilent Technologies, CA, USA), equipped with a deuterated L-alanine-doped triglycine sulphate (DLa-TGS) detector and a Ge-coated KBr beam splitter, was used. Spectra were recorded in absorbance mode, with 64 scans taken between a wavenumber range of 500 and 4000 cm−1 at a spectral resolution of 4 cm−1. Baseline correction and normalization were applied, ensuring consistent comparison between samples. Spectra were then shown with a vertical offset and overlapped. Normalization was performed using the amide I band (1650 cm−1) as a reference for collagen-rich samples or the peak at 1040 cm−1 (C–O–C stretch in GG) for polysaccharide-rich samples. Raw data and spectra were plotted using Origin 2025 PRO.

4.4. Degradation

To measure hydrolytic degradation, 1 mL of hydrogel was cast in a 24-well multiplate. The initial mass of (Mi) was recorded after the specimens were soaked in 2 mL of standard culture medium (DMEM) at pH 7.4 and incubated at 37 °C. Following this, the excess of DMEM was gently blotted, and the mass of hydrogels (Mt) was measured on days 1, 3, and 7; they were then re-immersed. The % mass lost over time was calculated as follows [71]:
D e g r a d a t i o n   ( % ) = M i − M t M i × 100

4.5. Mechanical Characterization

Oscillatory mechanical characterization was performed using the ElastoSens Bio (Rheolution Inc., Canada) to characterize the hydrogel’s viscoelasticity. Storage modulus (G’) was recorded on 5 mL of each specimen at 37 °C. The mean G’ was used to describe the stiffness of the hydrogels once the stability was reached after 24 h. Three specimens per condition were tested in soft mode for the pristine dBP hydrogels. While GG-containing samples were tested in hard mode. Data were provided by Soft Matter Analytics (Rheolution Inc., Canada).

4.6. Uniaxial Compressive Test

To evaluate the bulk equilibrium elastic modulus of the dBP hydrogels, stress-relaxation unconfined compression tests were performed using the MACH-1 mechanical testing system (Biomomentum Inc., Laval, QC, Canada). In order to ensure complete gelation before compression, the specimens were stored at a constant temperature of 37 °C overnight in DMEM. The applied parameters were previously reported by Gering et al. [40]: ramp velocity (10 mm/s) applied after 5% initial compression until 65% of the thickness to reach the point of failure.

4.7. Biological Characterization

Rat embryonic cardiomyoblasts (H9C2) (ATCC, Mannassas, VA, USA, CRL-1446) were cultured under standard conditions (37 °C, humidified atmosphere with 5% CO2) and employed for cytocompatibility and morphological evaluations. Cells were expanded in 75 cm2 flasks using Dulbecco’s modified Eagle medium with high glucose (DMEM; ECM0749L, Euroclone, Milan, Italy), supplemented with 10% (v/v) fetal bovine serum (FBS; Gibco, Milan, Italy), 5 mM L-glutamine (Sigma-Aldrich 1294808, Milan, Italy), penicillin (100 U/mL), streptomycin (0.1 mg/mL), and amphotericin (0.25 µg/mL); all products were from Euroclone, Milan, Italy. Cells were subcultured upon reaching 70–80% confluence.
Hydrogel cytocompatibility toward cardiac progenitor cells was assessed using the MTS viability assay (Promega Italia Srl, Milan, Italy). Briefly, 200 μL of hydrogel was cast in 24-well plates and incubated for 30 min at 37 °C. H9C2 were subsequently seeded onto the hydrogels at a concentration of 3000 cells/cm2. Cells cultured on a standard plastic-treated culture plate and collagen type I (4 mg/mL) were used as controls. At 1 and 7 days post-seeding, culture medium was removed and replaced with 250 µL of MTS working solution, prepared by mixing phenol red–free DMEM (Fisher Scientific, Milan, Italy) and MTS reagent in a 4:1 ratio, following the manufacturer’s instructions. Absorbance was recorded at 490 nm using a Victor 4X Multilabel plate reader (PerkinElmer, Milan, Italy), and experiments were performed in triplicate.
Cell morphology was analyzed by phalloidin/DAPI fluorescence staining. Samples were incubated with phalloidin and labelled with fluorescein isothiocyanate (Sigma-Aldrich, P5282, Milan, Italy) at 37 °C for 45 min to label F-actin, which was followed by nuclear counterstaining with 1 μg/mL DAPI (Sigma-Aldrich, Milan, Italy) for 1 min. A mounting medium consisting of 60% glycerol in PBS was applied prior to imaging. Fluorescent images were acquired using a Leica DM2500 microscope (Leica Microsystems, Wetzlar, Germany) at 40× magnification and processed with LAS V4.7 software (Leica Microsystems, Wetzlar, Germany).

4.8. Statistical Analysis

All the experiments were performed in triplicate. Raw data were collected in Excel (Microsoft, Redmond, WA, USA), and analyses were performed using Prism 8 (Graphpad Software Inc., San Diego, CA, USA). The results are expressed as mean ± standard deviation. Two-way ANOVA and a post-Bonferroni test were used to compare more than two groups.

Author Contributions

L.D.N.: Conceptualization, Data curation, Formal analysis, Investigation, Methodology, Software, Validation, Visualization, Project administration, Writing—original draft, Writing—review and editing. M.W.: Data curation, Formal analysis, Investigation, Methodology, Validation, Writing—review and editing. F.C.: Conceptualization, Supervision, Validation, Writing—review and editing. F.B.: Conceptualization, Funding acquisition, Methodology, Resources, Supervision, Validation, Writing—review and editing. D.M.: Conceptualization, Funding acquisition, Resources, Supervision, Validation, Writing—review and editing. All authors have read and agreed to the published version of the manuscript.

Funding

This research was carried out with the support of the Natural Science and Engineering Research Council of Canada (Alliance and Discovery) and PRIMA (Quebec Ministry for Economy and Innovation). D.M. holds a Canada Research Chair Tier I (2012–2026).

Institutional Review Board Statement

All cell-handling procedures were carried out in accordance with institutional biosafety regulations.

Informed Consent Statement

Not applicable.

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.

Acknowledgments

The authors gratefully acknowledge the use of the facilities at CHU de Québec Research Centre (Quebec City, QC, Canada) and UPO (Novara, Italy).

Conflicts of Interest

The authors declare no conflicts of interest.

Abbreviations

The following abbreviations are used in this manuscript:
ECMExtracellular matrix
dBPDecellularized bovine pericardium
GGGellan gum
SPMSpermine
FTIRFourier-Transform Infrared Spectroscopy
SEMScanning Electron Microscopy
GRASgenerally recognized as safe
FDAFood and Drug Administration
MiInitial mass
MtHydrogel mass
G’Storage modulus
MTS3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium salt
PBSPhosphate-buffered saline
DMEMDulbecco’s Modified Eagle Medium
CTRLPlastic Treated Culture Plate

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Gels 12 00118 g001
Figure 1. SEM images of the lyophilized hydrogels at 100× magnification (representative for 3 specimens, scale bar: 100 μm).
Figure 1. SEM images of the lyophilized hydrogels at 100× magnification (representative for 3 specimens, scale bar: 100 μm).
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Figure 2. FTIR spectra of the lyophilized hydrogels: (A) spectra with vertical offset show representative absorption peaks of dBP, GG, dBP−GG, dBP−GG−SPM1; and (B) overlapped representation of the same peaks and samples. Reference lines representing the chemical species for each overlapping peak analysis.
Figure 2. FTIR spectra of the lyophilized hydrogels: (A) spectra with vertical offset show representative absorption peaks of dBP, GG, dBP−GG, dBP−GG−SPM1; and (B) overlapped representation of the same peaks and samples. Reference lines representing the chemical species for each overlapping peak analysis.
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Figure 3. Degradation behaviour estimated by percentage wet mass variation over time of 5 mg/mL dBP−GG hydrogels incubated in DMEM at 37 °C (n = 5; * p < 0.001).
Figure 3. Degradation behaviour estimated by percentage wet mass variation over time of 5 mg/mL dBP−GG hydrogels incubated in DMEM at 37 °C (n = 5; * p < 0.001).
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Figure 4. Stabilized storage modulus (G’) values of hydrogels after 24 h of incubation in DMEM at 37 °C (* p-value < 0.05, ** p-value < 0.01, *** p-value < 0.005).
Figure 4. Stabilized storage modulus (G’) values of hydrogels after 24 h of incubation in DMEM at 37 °C (* p-value < 0.05, ** p-value < 0.01, *** p-value < 0.005).
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Figure 5. Young’s modulus obtained from compression of tested hydrogels (n = 3; red line as reference value of myocardial tissue; * p-value < 0.05).
Figure 5. Young’s modulus obtained from compression of tested hydrogels (n = 3; red line as reference value of myocardial tissue; * p-value < 0.05).
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Figure 6. dBP-GG hydrogel cytocompatibility (n = 3). MTS assay performed on H9C2 cardiomyoblasts cultured onto a hydrogel surface at day 1 and day 7. Viability is normalized to CTRL at day 1. Significant increases were observed in all groups except GG-dBP-SPM3. (* p-value < 0.05, ** p-value < 0.01 vs. CTRL at day 1).
Figure 6. dBP-GG hydrogel cytocompatibility (n = 3). MTS assay performed on H9C2 cardiomyoblasts cultured onto a hydrogel surface at day 1 and day 7. Viability is normalized to CTRL at day 1. Significant increases were observed in all groups except GG-dBP-SPM3. (* p-value < 0.05, ** p-value < 0.01 vs. CTRL at day 1).
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Figure 7. Representative fluorescence images of H9C2 cardiomyoblasts cultured on various hydrogel formulations for 7 days. F-actin was stained with phalloidin (green), and nuclei with DAPI (blue). CTRL is a plastic culture plate.
Figure 7. Representative fluorescence images of H9C2 cardiomyoblasts cultured on various hydrogel formulations for 7 days. F-actin was stained with phalloidin (green), and nuclei with DAPI (blue). CTRL is a plastic culture plate.
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Table 1. Pore characteristics, including pore size and surface coverage (%), were quantified from SEM micrographs (n = 5).
Table 1. Pore characteristics, including pore size and surface coverage (%), were quantified from SEM micrographs (n = 5).
Hydrogel
Nomenclature		Mean Diameter (μm)Maximum Diameter (μm)Minimum Diameter (μm)Coverage (%)
dBP22.78 ± 2.4642.04 ± 1.5314.93 ± 4.4649.09 ± 1.97
GG28.34 ± 4.9652.47 ± 3.8416.65 ± 6.9845.33 ± 7.49
dBP-GG43.57 ± 6.6573.62 ± 4.5225.88 ± 1.0351.14 ± 0.85
dBP−GG−SPM133.75 ± 9.1258.46 ± 6.6916.31 ± 1.2455.45 ± 5.80
dBP−GG−SPM230.60 ± 1.0357.61 ± 2.1219.43 ± 1.2546.94 ± 3.92
dBP−GG−SPM329.32 ± 4.1752.84 ± 2.9817.26 ± 6.6151.00 ± 4.86
Table 2. Hydrogel compositions used in this study, the calculated details of SPM per hydrogel, and GG in the specified concentrations.
Table 2. Hydrogel compositions used in this study, the calculated details of SPM per hydrogel, and GG in the specified concentrations.
Hydrogel
Nomenclature		SPM Solution (mg/mL)SPM Solution (µM)SPM Mass/dBP-GG Mass (w%)SPM/Total Mass (w%)SPM in Hydrogel (µM)SPM µMoles/GG (g)Positive Charge/GG (g)
dBP-GG-SPM10.175502.60.3490.05383.7720.1080.40
dBP-GG-SPM20.3501005.00.6950.106167.5040.20160.80
dBP-GG-SPM30.7002010.01.3810.212335.1080.40321.60
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Di Nunno, L.; Wekwejt, M.; Copes, F.; Boccafoschi, F.; Mantovani, D. Decellularized Extracellular Matrix/Gellan Gum Hydrogels Enriched with Spermine for Cardiac Models. Gels 2026, 12, 118. https://doi.org/10.3390/gels12020118

AMA Style

Di Nunno L, Wekwejt M, Copes F, Boccafoschi F, Mantovani D. Decellularized Extracellular Matrix/Gellan Gum Hydrogels Enriched with Spermine for Cardiac Models. Gels. 2026; 12(2):118. https://doi.org/10.3390/gels12020118

Chicago/Turabian Style

Di Nunno, Luca, Marcin Wekwejt, Francesco Copes, Francesca Boccafoschi, and Diego Mantovani. 2026. "Decellularized Extracellular Matrix/Gellan Gum Hydrogels Enriched with Spermine for Cardiac Models" Gels 12, no. 2: 118. https://doi.org/10.3390/gels12020118

APA Style

Di Nunno, L., Wekwejt, M., Copes, F., Boccafoschi, F., & Mantovani, D. (2026). Decellularized Extracellular Matrix/Gellan Gum Hydrogels Enriched with Spermine for Cardiac Models. Gels, 12(2), 118. https://doi.org/10.3390/gels12020118

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