3D Printing of Working Artificial Heart Tissue

Jetting-based bioprinting, or droplet-based bioprinting, uses a drop-on-demand strategy. It enables the precise, contactless deposition of cells and bioactive materials as droplets to fabricate 3D tissues or organs. It offers fine control over cell density by adjusting the number of cell-laden droplets deposited at specific locations [21]. Some commonly used examples of jetting-based bioprinting in the field of regenerative medicine include inkjet-based, laser-assisted and electrohydrodynamic jet bioprinting (Figure 1A). The following subsections introduce the major types of jetting-based bioprinting.

Stereolithography and digital light processing (DLP) bioprinting are both light-based fabrication techniques that rely on polymerizing photocrosslinkable bioinks to create 3D constructs. In stereolithography, a UV laser selectively cures a thin layer of photopolymer resin within a tank. The build platform lowers to form a thin film of liquid resin, which is then solidified layer by layer as the laser traces the desired pattern point by point (Figure 1B). A key advantage of stereolithography is that, by varying printing parameters such as laser attenuation, laser intensity, or scanning speed, it is possible to create graded crosslinking—i.e., some regions are more crosslinked while others are less. This mismatch generates internal stress that drives 4D morphing effects [29]. Using this strategy, Cui et al. [29] fabricated a cardiac patch that autonomously curves to match the heart's shape. With stereolithography, they also produced helically arranged myocardial fibers with precisely controlled angles to replicate the anisotropy of native myocardium. A limitation is speed: the laser must scan the entire surface point by point, which can be time-consuming. An innovative study by Miao et al. [30] integrated photolithography with stereolithography so that photolithography polymerizes an entire surface in seconds through a photomask, while leaving a ~10 µm liquid layer for fine fabrication to selectively cure the top film into microgrooves. This design completes prints within ~15 min while maintaining fidelity and patterning.

DLP uses a digital light projector and a digital micromirror device (DMD) to cure an entire layer of photocurable bioink simultaneously with UV light, then advances to the next layer (Figure 1B). This eliminates x–y mechanical scanning during curing and reduces print time [31]. Despite this advantage, axial resolution can be compromised due to discretization from mask changes between layers [32]. A notable advance is microscale continuous optical printing (µCOP), which fabricates complex microarchitectures continuously by moving the stage smoothly along the z-axis while dynamically updating the mask [33]. Liu et al. [33] also demonstrated multi-material printing—e.g., a cantilever for force analysis and a cardiac cell-laden bioink—using DLP without additional post-processing, enabling seamless transitions between materials.

Fewer studies encapsulate cells directly in the ink for stereolithography because photoinitiators and UV exposure pose cytotoxicity risks, including potential DNA damage [29,34]. As a result, many groups print acellular scaffolds and seed cells afterward, which can lead to non-uniform cell distribution [29]. DLP is generally associated with lower cytotoxicity risk [32,33], and several studies have printed cell-laden constructs with ~85–90% viability, likely due to lower local light intensity and shorter exposure times.

Extrusion-based bioprinting is a widely used technique in CTE that deposits bioinks—typically viscoelastic, cell-laden materials—through one or more nozzles in a computer-controlled, layer-by-layer process [5]. Pneumatic pressure or mechanical force (piston- or screw-driven) is applied to extrude the bioink as cylindrical filaments onto a build platform (Figure 1C). This approach is popular for cardiac tissues because shear stresses within the nozzle, along with extensional stresses during extrusion, can align microfibers in the bioink in a preferred direction [35,36,37], thereby guiding cardiomyocyte alignment. Recreating this anisotropy is a critical and desirable feature in cardiac constructs. Extrusion-based printing also allows straightforward parameter tuning—pressure, nozzle diameter and geometry, stand-off distance, and filament thickness—to mimic characteristics of native myocardium [11]. It is also relatively simple and cost-effective.

Two recent studies highlight the strategic use of pre-processing and auxiliary fabrication steps to enhance structural and functional outcomes. In one, melt-electrowriting (MEW) was used to fabricate hexagonal PCL microfiber scaffolds prior to extruding cell-laden bioinks [6] (Figure 2A). The resulting rigid, reinforcing scaffold provided mechanical support, promoted anisotropic cell alignment, and enabled integration of a patterned pre-vascular network—addressing limitations of soft hydrogel-only constructs and facilitating vascularization upon implantation. In another, the authors pre-formed anisotropic organ building blocks (aOBBs) on PDMS micropillar arrays before bioprinting, yielding tissues with elongated, microscale alignment and improved electromechanical maturation prior to encapsulation (Figure 2B); the additional shear and extensional stresses during extrusion further enhanced alignment [38]. As a result, aOBB-printed tissues exhibited significantly higher contractile force and more organized function while using fewer cells. Together, these studies also illustrate how integrated pre-fabrication strategies can streamline multicellular bioprinting workflows and improve tissue performance.

Volumetric bioprinting is an emerging light-based manufacturing technique that differs markedly from traditional layer-by-layer approaches. In this method, a cell-laden bioresin inside a rotating vial is exposed to a sequence of precisely controlled, spatially distributed light projections generated via computed tomographic reconstruction using a digital micromirror device (DMD) (Figure 4) [7]. These light patterns induce polymerization only within the voxels corresponding to the desired 3D structure, enabling the entire construct to form in a single, continuous step. This layerless process dramatically reduces fabrication time—often to seconds—and minimizes mechanical stresses on cells, preserving viability and delicate microarchitecture [49]. Because the build is not constrained by sequential deposition, it avoids issues such as interlayer delamination and prolonged light exposure that can harm embedded cells. Volumetric bioprinting also supports encapsulation at high cell densities.

Volumetric bioprinting is particularly well-suited for producing entire organoids. In a study by Bernal et al. [49], liver epithelial organoids were volumetrically bioprinted within a specially tuned gelatin bioresin in under 20 s. The process preserved organoid morphology and polarity and achieved higher post-printing cell viability than traditional extrusion methods. Volumetric bioprinting has also been explored for cardiac tissue models. Lian et al. [50] formulated a heart-derived decellularized extracellular matrix (dECM) bioresin laden with cardiomyocytes and a photoinitiator system; solidification was triggered by visible light at 525 nm. The resulting cardiac constructs were printed in ~30–45 s, formed complex 3D geometries, expressed sarcomeric α-actinin, and exhibited synchronized contractions.

In another study, Jones et al. [51] replaced computed tomographic reconstruction with filamented light (Flight), a volumetric approach that projects multidirectional, coherent UV light into cell-laden GelMA to photopolymerize the construct within a 3D volume. They exploited modulation instability, which occurs as UV light interacts with the hydrogel during photopolymerization, to generate internal microfilament structures that provide physical cues for cardiomyocyte alignment. Printing was completed in <30 s, encapsulating high cell densities (~15–60 million cells/mL). Notably, they fabricated torsionally contracting cardiac structures that mimic native heart mechanics.

Volumetric bioprinting allows ultrafast fabrication, polymerizing entire 3D constructs within seconds, whereas DLP or extrusion-based systems typically require minutes to hours. This capability presents substantial potential for scalable production of organoids and heart-implantable constructs. Because no nozzle extrusion or layer deposition is involved, mechanical stress on cells is greatly reduced. Furthermore, the technique can simultaneously print multiple constructs without stage movement, offering an attractive solution for high-throughput applications such as drug screening and robotic tissue actuators.

Researchers have expressed concerns about resolution and spatial control, particularly when printing large constructs, since limited local heterogeneity may impede features such as vascular path formation. An innovative integration of volumetric and extrusion-based printing was explored by Ribezzi et al. [52] to create multimaterial, multicellular tissue models. Using a new "plug-and-play" support bath based on low–molecular-weight gelatin, fine extruded filaments were held in place and later incorporated into the final structure via volumetric crosslinking. Two wavelengths were used—one for alignment and one for curing—combining the speed and design freedom of volumetric printing with the precise material placement of extrusion, enabling the fabrication of detailed constructs such as functional cardiac organoids.

Different bioprinting techniques each offer unique advantages for cardiac tissue engineering, yet their practicality and maturity vary. Jetting-based methods such as inkjet, LIFT, and EHD bioprinting provide high precision and contactless deposition. It is well-suited to patterning complex microstructures such as Purkinje-like networks and heterogeneous cell densities. However, they remain limited by small construct thickness, narrow bioink viscosity ranges, and unstandardized parameters. Light-based approaches, including SLA and DLP, achieve excellent geometric fidelity, gradient crosslinking, and even 4D morphing to replicate myocardial curvature and anisotropy, though challenges with UV-induced cytotoxicity, especially when long curing times are required, and non-uniform cell distribution remain. Volumetric bioprinting stands out for its ultrafast, layerless fabrication that minimizes mechanical stress on cells and supports high cell densities, making it promising for producing large organoids or heart models, but its spatial resolution and limited material selectivity remain challenges.

Among all methods, extrusion-based bioprinting remains the predominant technique in the field due to its simplicity, versatility, and compatibility with a broad range of bioinks, especially soft, natural hydrogels native to the heart. It enables directional cell alignment through shear forces and shows robust development in real-time feedback control systems. With continued improvement in scalability, cytocompatibility, and structural control, the other modalities could be more broadly implemented for cardiac tissue fabrication, enabling faster and more efficient construction of complex, functional tissues. Continued advancements in visible-light photochemistry, hybrid volumetric–extrusion strategies and closed-loop process monitoring are expected to further enhance the efficiency and fidelity of these approaches in fabricating cardiac constructs.

The biophysical properties of bioinks and biomaterials play a critical role in tissue engineering because they directly influence how cells interact with one another and with their surrounding matrix [55]. For cardiac applications, bioinks must be carefully tuned to replicate the mechanical environment of the native myocardium while still supporting cell survival and function. Several rheological and structural parameters are commonly used to evaluate hydrogel performance in recent studies in cardiac applications, including stiffness (elastic modulus), viscoelastic behavior (G′ and G″), print fidelity, swelling ratio, and degradation resistance. Each of these parameters contributes to the delicate balance between structural support and biological compatibility.

The elastic modulus (or compressive Young's modulus) describes material stiffness and is particularly important in cardiac tissues. If the bioink is too soft, it deforms easily and dissipates contractile forces rather than transmitting them between cardiomyocytes, which weakens coordinated and synchronous beating. Conversely, if the matrix is too stiff, it restricts cell spreading and migration, imposes mechanical load on cardiomyocytes and reduces their ability to shorten and contract effectively [8]. Hydrogel stiffness is also linked to a 'mechanical memory' effect, whereby cells "remember" their mechanical environment and alter their differentiation lineage accordingly; if the substrate is excessively stiff, this may activate fibrotic signaling. Therefore, the stiffness must be carefully tuned to support the desired long-term cell fate [55]. Complementing stiffness, most studies conduct the viscoelastic moduli (G′ and G″) measurements to assess how the material behaves as a solid (elastic response) versus a liquid (viscous response) [55]. For cardiac bioinks, post-crosslinking behavior should favor G′ > G″, indicating that the construct behaves more like a solid while retaining sufficient viscoelasticity for remodeling and mechanical flexibility.

Print fidelity, which is explicitly tested in extrusion-based bioprinting, ensures that the bioink maintains its intended geometry and resolution after printing; this depends on viscosity, shear-thinning behavior, and yield stress [55]. The swelling ratio determines how much water the hydrogel absorbs, affecting nutrient and oxygen transport as well as mechanical stability. Excessive swelling can weaken constructs or distort their shape, which becomes a significant parameter during implantation or in vivo testing [8]. Degradation resistance is needed so that the printed construct remains intact long enough to support cells through the maturation phase, before the cells themselves generate sufficient extracellular matrix [8]. These parameters can be tuned through polymer concentration, crosslinking chemistry and density, and the incorporation of additional molecules or secondary polymers.

Bioink or biomaterial design represents a balancing act between printability, mechanical stability, and cell viability. For example, in extrusion bioprinting, higher viscosity improves print fidelity but increases shear stress during extrusion, which can damage cells. Low stiffness favors survival but risks structural collapse during culture [8]. In light-based bioprinting, fast or aggressive crosslinking enhances construct stability but may expose cells to toxic conditions [29,34]. For this reason, optimizing bioink properties is an iterative process, where the rheology must be adjusted to achieve both structural fidelity and biological support during tissue maturation.

Naturally derived biomaterials/bioinks offer high biocompatibility and bioactivity; for example, decellularized extracellular matrix (dECM) retains key growth factors, extracellular-matrix (ECM) proteins, and glycosaminoglycans (GAGs) [56]. However, these materials are typically mechanically soft due to the extraction process and the breakdown of native crosslinks; therefore, they tend to collapse after several printed layers, making printability unstable and difficult to control [57]. In contrast, synthetic biomaterials provide improved mechanical integrity, tighter control over printability, and consistent batch-to-batch properties. Yet, manufacturing processes—such as UV-induced or high-energy crosslinking—can generate cytotoxic by-products, often limiting direct cell encapsulation before printing [36]. To minimize this trade-off between biocompatibility and printability, recent studies have explored hybrid formulations and combinatorial strategies that enhance construct stability while supporting cell growth in CTE applications, including dual-crosslinking mechanisms, interpenetrating networks (IPNs), and rheological modulation.

In Shin et al. [56], Laponite-XLG nanoclay and polyethylene glycol diacrylate (PEG-DA) were combined with porcine-derived ventricular dECM to develop a mechanically tunable and printable dECM bioink. Laponite-XLG is a disk-like charged nanoclay that forms an ionic gel network, acting as a rheology modulator, increasing shear storage modulus and raising viscosity at rest, thereby improving filament stability after extrusion. PEG-DA enabled rapid photocrosslinking (<1 min) to lock the construct's shape before sagging. By varying the concentration of PEG-DA, the authors achieved tunable stiffness values representing both healthy myocardium (≈13.4 kPa) and fibrotic myocardium (≈89 kPa), providing platforms for disease modeling and drug-screening applications [32].

Another study developed an IPN bioink comprising gelatin methacryloyl (GelMA) and methacrylated hyaluronic acid (MeHA) blended with dECM for cardiac tissue printing. GelMA provided cell-adhesive motifs and adjustable stiffness, whereas MeHA conferred slower degradation and greater rigidity [58]. A dual-crosslink method with UV light and microbial transglutaminase (mTGase) further enhanced stiffness and stability without excessive UV exposure. By varying GelMA/MeHA ratios, the team generated healthy and fibrotic myocardium analogs, showing that fibrotic tissues containing MeHA were ≈10× stiffer than those without, underscoring the influence of bioink composition on tissue mechanics [58].

Alginate remains a popular cardiac bioink due to its biocompatibility and instantaneous calcium-ion crosslinking capability [59,60]. However, similar to dECM and other naturally derived hydrogels, alginate-based systems exhibit low mechanical strength and poor shape retention. To improve these properties, recent formulations have integrated fibrinogen or secondary networks to enhance printability and post-printing stability. Budharaju et al. [5] incorporated fibrinogen into alginate to introduce cell-adhesion sites and generate a nanofibrous morphology after crosslinking, thereby better mimicking cardiac ECM. Fibrinogen also provided a secondary crosslinking mechanism through thrombin, which, together with calcium-mediated alginate gelation, enhanced construct stability. The study additionally pre-crosslinked alginate with low CaCl₂, increasing viscosity, yield stress, and shape fidelity during extrusion. Several groups have further coupled embedding (support-bath) printing with calcium-ion additives to stabilize filaments as they are extruded [43]. However, one study suggested that such external crosslinking can lead to uneven gelation due to nonuniform calcium diffusion, potentially limiting oxygen/nutrient transport [59].

To address this limitation, Stola et al. [59] developed an internally crosslinked alginate dialdehyde (ADA)/alginate/gelatin bioink. By embedding insoluble CaCO₃ and the pH-lowering agent glucono-δ-lactone (GDL) directly within the ink, gradual GDL hydrolysis releases protons that dissolve CaCO₃, yielding a controlled, homogeneous Ca²⁺ release and triggering internal crosslinking. Gelation rate was tuned via CaCO₃, GDL, and gelatin concentrations to achieve a usable printability window—stiffening late enough to permit extrusion but early enough to prevent sagging—eliminating the need for a support bath. Another recent report by Vettori et al. [8] added silk fibroin (SF) to alginate–gelatin bioinks and achieved hydrogel stiffness in the physiological range (≈38 kPa; native adult myocardium ≈10–50 kPa) with just 1% SF. SF introduced β-sheet and α-helix structures, providing additional intermolecular bonding that increased elasticity and enhanced contractile performance in the printed constructs.

Recent studies have incorporated conductive materials into bioinks for cardiac tissue printing to promote more synchronous and uniform cardiomyocyte contractions. Examples include carbon-based nanomaterials (graphene derivatives such as GO and rGO, and carbon nanotubes, CNTs), metallic particles (e.g., gold), MXenes, and conductive polymers such as poly(3,4-ethylenedioxythiophene) (PEDOT) [10,61]. The conductivities of these material-integrated cardiac constructs are summarized and compared with native adult myocardium in Table 2.

4D bioinks—also called shape-morphing bioinks—enable printed constructs to undergo programmed structural changes over time in response to stimuli such as temperature, pH, magnetic fields, or electrical signals [31]. These transformations can include bending, folding, or reorientation, thereby mimicking the dynamic remodeling processes observed during cardiac development. This capability holds promise for replicating the native curved architecture of myocardium and for delivering constructs in compact forms that expand in situ to match patient-specific heart geometry, thereby improving fit, integration and functional performance [70].

Recent studies using 4D bioinks for cardiac patch fabrication highlight several design considerations compared to conventional bioinks [30,31].

First, printing modality is critical because certain techniques offer greater control over process parameters. For example, Miao et al. [30] used stereolithography and adjusted laser attenuation, intensity, and scanning speed to create crosslinking gradients. The resulting internal stress differentials drove 4D actuation, producing a curved cardiac patch [29,30].

Second, the stimulus-responsiveness mechanism must be carefully selected. Thermal triggers are among the most common: by defining a glass-transition temperature (Tg), the construct remains temporarily "fixed" below Tg and recovers its permanent shape above Tg—a phenomenon known as shape-memory behavior (SMP) [45]. For in vivo cardiac patches, Tg is typically tuned through crosslink density and network chemistry to operate near physiological temperature. Cui et al. [71] further demonstrated photothermal responsiveness by integrating graphene nanoplatelets within SMP matrices; near-infrared (NIR) light absorption generated localized heating, raising the temperature above Tg and triggering rapid shape recovery.

As 4D bioinks develop, several concerns have emerged. Biocompatibility and cytotoxicity must be carefully evaluated: although >60% cell survival was maintained after NIR-triggered shape recovery in one study, prolonged or high-intensity exposure may cause thermal damage [71]. Additionally, the constructs often exhibit MPa–GPa-scale stiffness—advantageous for maintaining programmed configurations but mechanically mismatched to the soft elasticity of native myocardium [31]. Because these patches are typically printed acellular and later post-seeded, achieving uniform and stable cell distribution before morphing is crucial. Finally, repeated folding/unfolding can introduce fatigue or microcracks; therefore, long-term mechanical durability under physiological loading must be thoroughly evaluated prior to clinical translation.

In volumetric bioprinting for cardiac tissue engineering, which relies primarily on light-based actuation, the bioink formulation must be carefully tuned to balance optical, chemical, and rheological properties.

First, optical transparency is critical: the material needs to allow sufficient and uniform penetration of visible or UV light throughout the rotating vial, as excessive absorbance or scattering from high solute concentrations can block light, limiting crosslinking depth and leading to poor resolution or incomplete structures [50]. Since volumetric printing enables very high cell densities, local variations in refractive index can also distort light paths; to address this problem, Jones et al. [51] added iodixanol to GelMA to match refractive indices throughout the vial and maintain uniform light propagation.

In parallel, the bioink must have adequate viscosity and stability during rotation to prevent sedimentation or diffusion that could blur the projected light patterns [50]. There is a trade-off between resolution and depth of crosslinking that needs to be carefully manipulated. As higher concentrations improve crosslink density and structural integrity, they also reduce light penetration, thereby lowering resolution [51].

Adult human primary cardiomyocytes (hPCMs), directly isolated from adult hearts, are considered more physiologically relevant as in vitro models with high drug-response sensitivity. Wang et al. [74] reported that preparation—from cell extraction to bioink mixing—can be completed in about one day, making the process relatively efficient [74]. However, hPCMs are limited by their non-proliferative nature, short culture window, and pronounced donor-to-donor variability [73]. Consequently, most 3D-bioprinting studies employing primary cardiomyocytes have instead used cells derived from neonatal rats rather than adult human hPCMs [74,75], thereby introducing a developmental gap that alters ion-channel expression and contractile force generation and limits patient-specific drug screening fidelity.

To overcome these limitations, researchers have increasingly adopted stem cell-derived sources such as human embryonic stem cells (hESCs) and induced pluripotent stem cell-derived cardiomyocytes (iPSC-CMs), both possessing self-renewal and cardiomyogenic differentiation potential. iPSC-CMs, in particular, are favored for patient-specific applications and avoid the ethical concerns associated with hESCs [76]. Accordingly, iPSC-CMs have become the predominant cell source in cardiac tissue engineering [77]. Choosing between printing immature cells and promoting their maturation post-printing versus using more mature cardiomyocytes depends on the intended application, bioprinting constraints, and desired maturation approach. Immature cardiomyocytes resemble fetal CMs—they tolerate mechanical stress more effectively, proliferate readily, and exhibit higher engraftment potential—making them particularly suitable for implanted cardiac patches [77]. They also respond well to 3D-bioprinting cues—anisotropic patterning, tuned stiffness, printed vascular architecture and mechanical/electrical stimulations, which collectively promote structural and functional maturation, rendering 3D bioprinting a practical platform for tissues derived from immature cardiomyocytes [78].

By contrast, encapsulating more mature iPSC-CMs within 3D environments can yield stronger contractile function, improved structural organization, and reduced arrhythmogenic activity [79]. Although fewer studies have bioprinted fully matured cardiomyocytes, casting-based models have demonstrated clear advantages over very early-stage cells: enhanced sarcomere organization, partial isoform shifts toward adult contractile proteins (e.g., increased β-MHC, cTnI, and MLC2v), improved calcium-transient amplitude and kinetics, and more stable electrophysiology (e.g., more hyperpolarized resting potentials and reduced spontaneous activity) [78,80,81]. These traits translate to stronger, more coordinated contractions once embedded into engineered heart tissues or hydrogels, as extended culture time facilitates structural and electrophysiological maturation prior to 3D encapsulation [82].

For a practical timing window for iPSC-CM incorporation, it is suggested to wait until the onset of spontaneous beating in 2D culture (~day 10–14) to ensure basic electrophysiological competence for subsequent 3D cues [83]; ~2–4 weeks of pre-culture in 2D before embedding is often optimal, allowing cells to develop early sarcomeric organization, partial isoform switching, and improved Ca²⁺ handling, which together enhance subsequent 3D tissue performance [82].

Some strategies instead print pluripotent stem cells (iPSCs) directly and induce differentiation in situ, using geometrical confinement to guide maturation within 3D constructs. Ellis et al. [84] showed that encapsulating iPSCs in defined 3D geometries can direct tissue-level organization without external stimulation, improving myofibrillar alignment, Z-line formation, and anisotropic contractility. Similarly, Kupfer et al. [85] found that in situ dense cardiomyocyte networks (~100 million cells/cm³) and thicker muscle walls (~500 µm), approaching native myocardial structure. Because this strategy avoids cell dissociation or reseeding, which can disrupt gap junctions and sarcomeres, cells form intercellular junctions during differentiation, more closely mimicking in vivo development [85,86]. However, non-uniform cellularization and phenotypic heterogeneity remain major challenges: proliferating and differentiated cells may accumulate near the construct peripheries, leaving core regions sparsely populated, and mixed lineage outcomes may arise—beneficial for multicellularity and co-culture, yet reducing control over specification [85].

Beyond single-cell-type approaches, co-culturing multiple cardiac cell populations has emerged as a powerful strategy to recapitulate the native heart's complex multicellular architecture. While cardiomyocytes (CMs) remain central to electrical conduction and force generation, non-myocyte cells—particularly cardiac fibroblasts (FBs) and vascular endothelial cells (ECs)—play equally critical and complementary roles [87].

Cardiac fibroblasts (FBs) provide structural support by secreting extracellular matrix (ECM) proteins and facilitate deposition of cell-derived ECM as biomaterial scaffolds degrade. They are also key modulators of electrical cross-talk. Khoury et al. [88] explicitly examined heterocellular gap-junction coupling (HC) between FBs and CMs, showing that FB–CM electrical coupling can modulate impulse propagation and influence arrhythmogenesis. Elevated fibroblast density is a hallmark of aging and post-myocardial infraction (MI) remodeling, making FB-enriched co-cultures valuable for drug screening in fibrosis-associated arrhythmias and conduction disturbances.

Endothelial cells (ECs) are essential for promoting vascularization and facilitating nutrient and oxygen transport within engineered constructs [89]. Fabres et al. [90] used EC co-culture to recreate the in vivo CM–endothelium interface; under flow, ECs self-organized into a monolayer shell surrounding the tissue, which shielded CMs from shear stress and regulated drug diffusion kinetics, thereby better mimicking native tissue physiology. In cardiac patches, ECs can pre-form vascular networks that anastomose with host vasculature upon implantation, enhancing long-term perfusion and survival [91].

Other supporting lineages including mesenchymal stromal cells (MSCs) and smooth muscle cells (SMCs) are also studied and incorporated into the bioprinting process. MSCs contribute stromal and angiogenic support, recapitulating mesenchymal roles in vessel maturation [87]. In Brady et al. [91], MSCs promoted EC proliferation, and a subset differentiated into α-SMA⁺ pericyte-like cells that wrapped endothelial lumens, stabilizing and maturing microvessels. Smooth muscle cells (SMCs) serve as mural partners that stabilize EC organization and support capillary-like networks; Fabres et al. [90] reported that removing SMCs resulted in reduced EC coverage and compromised vessel integrity.

3D bioprinting offers unique advantages for co-culture systems, as it preserves multicellular organization (including spheroids) while defining precise geometries with high reproducibility [92]. It enables patterning strategies that control cell–cell proximity and spatial contact at micrometer scales—critical parameters for coordinated vascularization and regulated electrical conduction—thereby promoting hierarchical vessel formation over the disordered microvasculature typically seen in bulk hydrogels [35,91]. In aligned, micro-architected constructs, co-culture has been associated with enhanced synchronous contraction and lower excitation thresholds compared to CM-only tissues, reflecting improved electromechanical integration [90]. For example, Brady et al. [91] reported that patterned patches guided linear "trunk" vessels (~20–40 µm diameter) with branching microvessels (~5–15 µm), whereas non-patterned patches generated smaller, scattered vessels (<10 µm).

There are several key design considerations for establishing successful co-culture systems in 3D bioprinting, including medium compatibility, passage control, and cell-loading strategy. The culture medium must be optimized to support all cellular linages equally; for instance, researchers have employed a blended "Spheroid Medium" composed of a 2:1:1 ratio of CM-maintenance, endothelial, and fibroblast media, a 1:1 mixture of CM differentiation medium with EGM-2 for CM–EC co-cultures, or serum-free CM conditions supplemented with lineage-specific growth factors such as VEGF and bFGF to sustain non-myocytes [90]. Passage numbers—particularly for ECs and FBs—should be carefully controlled to minimize phenotypic drift and senescence-related changes that could impair function [88,93]. Finally, the sequence of cell deposition plays an important role: sequential or staged printing can outperform simultaneous loading by reducing competition for limited micro-niches and preserving cardiomyocyte alignment and functional coupling within the construct [92].

Cardiac patches have long served as temporary mechanical support systems following myocardial infarction (MI), aiming to restore cardiac function in damaged myocardium. However, several translational challenges remain—most notably insufficient vascularization, limited electrical and mechanical coupling with the host myocardium, incomplete innervation for physiological regulation, and only modest improvements in overall cardiac function and infarct size reduction after implantation.

Recent 3D bioprinting strategies have begun addressing these challenges by integrating advanced fabrication approaches to produce patches with geometrically optimized architectures, incorporating diverse cellular populations to promote vascular and stromal support, and functionalizing bioinks with conductive materials, bioactive molecules, or imaging and sensing components to enhance integration, monitoring, and therapeutic efficacy.

3D-bioprinted cardiac tissues can serve as high-throughput drug-testing systems, preserving cell–cell and cell–matrix interactions as well as native-like architecture. To ensure these systems efficiently and robustly model pathological conditions or screening scenarios, the printed tissues must be sufficiently mature. Across the field, parameters used to enhance maturation in engineered heart tissues (EHTs) include refined structural patterning, vascularization, multimodal stimulation and prolonged 3D culture. An overview of recent bioprinted tissues and their maturation strategies is summarized in Table 4. Additionally, 3D bioprinting can directly recapitulate disease stages by incorporating pathological biomaterial cues into the printed constructs.

3D bioprinting is especially valuable for biomanufacturing irregular, large, and complex geometries such as the heart. As research progresses from small-scale cardiac tissues to organ-scale, chambered constructs, investigators are leveraging innovations including embedded printing for continuous builds, modular assembly, and integrated perfusion-flow conditioning to achieve target shape and function.

Fang et al. [44] introduced a novel sequential embedded bioprinting approach (SPIRIT) to print a ventricular model with open, perfusable blood vessels. They used a microgel-based biphasic (MB) bioink, capable of serving both as a printable bioink and a suspension medium, providing structural support while allowing subsequent printing of secondary inks. In the first stage, the external ventricle geometry was printed within a support bath. Without crosslinking the construct, they then performed a second embedded printing step in which sacrificial endothelial cell (HUVEC)-laden gelatin ink was written directly inside the freshly printed ventricle to generate freeform vascular channels. Upon removal of the sacrificial ink, these channels were perfusable and became uniformly lined with endothelial cells, forming a confluent endothelium by day 7 and even showing endothelial invasion into the surrounding matrix. The printed ventricle constructs achieved high shape fidelity (>95% match to CAD), hierarchical vascular trees, and—when loaded with cardiomyocytes—enhanced viability, spontaneous calcium transients, and sarcomeric α-actinin and connexin-43 expression, confirming preliminary functional maturation.

A recent innovation seeks to recreate the myocardial fiber orientation, as native left ventricular function depends on transmural fiber helices that generate left-ventricular twist (LVT) during contraction and relaxation [113]. A recent study used a bioprinting-assisted tissue assembly (BATA) approach to build a modular ventricle. It first printed uniaxially aligned EHT modules containing iPSC-derived cardiomyocytes and cardiac fibroblasts in hdECM. These were printed onto PEVA post frames to induce tension, compaction, and alignment. These modules were then assembled like "LEGO" blocks to form multi-layer structures with programmable fiber directions (Figure 10). Individual modules showed spontaneous and paced contraction, optical mapping signals, and drug responses (isoproterenol and omecamtiv increased inotropy; nifedipine decreased it). After assembly, modules fused structurally through fibroblast-mediated ECM remodeling and formed connexin-43 junctions without losing structure. The final "MoCha" chamber exhibited synchronized beating, calcium wave velocity of ~1.4 cm/s, and a net left-ventricular twist of ~0.68°. Although this is lower than the native 12–15°, it demonstrates successful fiber programming toward anisotropic contraction [114].

Beyond the single-chamber construct, researchers are now creating multi-chamber heart models with perfusable geometries. Kupfer et al. [85] developed a 3D bioprinted, chambered human cardiac muscle pump (hChaMP) by printing hiPSCs in an ECM-based bioink and differentiating in situ. The constructs formed muscular walls of 100–500 µm, with ~88% cardiomyocytes, and exhibited synchronous beating at ~40–80 BPM for at least 6 weeks. They generated pressure oscillations of ~0.3–0.5 mmHg, pressure–volume loops, and stroke work of ~14.5 nJ with a maximum ejection fraction of 6.5%. The constructs also exhibited robust electrical coupling and dose-dependent pharmacological responses.

Zeng et al. [114] further expanded this concept by designing two- and four-chamber heart models using DLP printing with a conductive PGI hydrogel ink. They embedded helical micro-grooves in ventricular walls and circumferential grooves in atrial walls and integrated a hollow auxetic re-entrant layer to reproduce anisotropic stiffness. They further included vessel inlets and tri-leaflet valves to achieve unidirectional flow. After NaOH strengthening and ECM coating, the scaffolds were seeded with hiPSC-derived cardiomyocytes, which aligned along the patterned grooves, formed anisotropic tissues, and established gap junction coupling. The two-chamber models achieved valve-mediated one-way pumping, while the four-chamber model displayed synchronized contractions with ventricular-like torsional twist (~7°), sustained over 10 million contraction cycles without structural failure. These results demonstrate that structurally encoded scaffolds can guide organ-level organization and function.

Despite rapid progress toward organ-level constructs, several limitations persist. Viable tissue thickness remains below native myocardium, and uniform cell infiltration is challenging, with viability often concentrated at the periphery [85]. Immaturity of cardiomyocytes further limits function [114]. Ventricular twist angles also fall short of physiological ranges—likely due in part to coarse, layer-by-layer fiber orientations; recent work suggests that imposing a gradual transmural rotation in fiber angle, rather than abrupt changes between layers, can better recapitulate native mechanics [113,114]. Even so, it is noteworthy that complex, chambered geometries produced by 3D additive manufacturing can remain structurally stable in vitro for over a month.