Latest Advances in 3D Bioprinting of Cardiac Tissues

Abstract

Cardiovascular diseases (CVDs) are known as the major cause of death worldwide. In spite of tremendous advancements in medical therapy, the gold standard for CVD treatment is still transplantation. Tissue engineering, on the other hand, has emerged as a pioneering field of study with promising results in tissue regeneration using cells, biological cues, and scaffolds. Three-dimensional (3D) bioprinting is a rapidly growing technique in tissue engineering because of its ability to create complex scaffold structures, encapsulate cells, and perform these tasks with precision. More recently, 3D bioprinting has made its debut in cardiac tissue engineering, and scientists are investigating this technique for development of new strategies for cardiac tissue regeneration. In this review, the fundamentals of cardiac tissue biology, available 3D bioprinting techniques and bioinks, and cells implemented for cardiac regeneration are briefly summarized and presented. Afterwards, the pioneering and state-of-the-art works that have utilized 3D bioprinting for cardiac tissue engineering are thoroughly reviewed. Finally, regulatory pathways and their contemporary limitations and challenges for clinical translation are discussed.

1. Introduction

Cardiovascular diseases (CVDs), such as congenital heart disease, acute coronary syndrome, hypertension, and arrhythmias, are a major cause of death worldwide, currently accounting for more than 17.5 million deaths globally, and are expected to increase to 23.6 million by 2030.^[1] The progression of CVDs damages myocardium, heart valves, or vasculatures in the body and treatment requires an appropriate replacement of injured tissue.^[2] Currently accepted therapies are pharmacological, interventional, and surgical.^[3] Pharmacological therapy employs angiotensin receptor blockers, angiotensin-converting enzyme inhibitors, catecholamines, and aldosterone to promote systolic function during heart failure.^[4] Interventional therapy, on the other hand, utilizes the implantation of devices such as pacemakers to regulate electromechanical arrhythmia.^[5] Furthermore, grafting of tissues from the patient’s own body (autografts) such as coronary artery bypass grafting,^[6], from human donors (allografts) such as heart transplantation for treating heart failure,^[7] or from animals (xenografts) such as bovine or porcine heart valves^[8] have been used clinically.^[9] However, these medical practices have several limitations including the shortage of available donor organs, ^[10] a need to use anticoagulants,^[11] potential immune rejection,^[12] and low durability.^[13]

Cardiac tissue engineering aims to repair or regenerate the damaged heart tissue by using a combination of cells, biomaterials, and signaling molecules to form cardiac constructs.^[14] Several important considerations for engineering cardiac constructs are appropriateness of cell sources and biomaterial choices, structural properties such as oriented myofibers, perfusable and proper vascularization, comparable mechanical properties, electrical conductivity, and physiologically pertinent functionalities such as electromechanical coupling and synchronous beating.^[15] These constructs can significantly reduce immunogenicity and thrombogenicity and degrade within the body at a rate consistent with the rate of tissue regeneration.^[16] Natural or synthetic biomaterials support cell adhesion and proliferation by providing a porous and interconnected polymeric network as scaffold, which can be developed by several manufacturing techniques.^[17]

A novel approach for developing tissue-engineered scaffolds is three-dimensional (3D) bioprinting. By applying a precise layer-by-layer approach, 3D bioprinting techniques have several advantages over conventional scaffold fabrication technologies, including better control over multiple compositions, spatiotemporal distributions, and structural complexity.^[18] To mimic the complex architectural features of native tissue and create patient-specific 3D models, a wide variety of imaging technologies such as 3D scanners, computed tomography (CT), magnetic resonance imaging (MRI) systems, and computer-aided design (CAD) software have been used.^[19] Moreover, a broad range of biomaterials and cells has been implemented as bioinks for 3D bioprinting of cardiac constructs, including alginate,^[20] collagen,^[21] gelatin,^[22] hyaluronic acid,^[23] and decellularized extracellular matrix.^[24] Even though this technique is still in the early stages of investigation, it can lead to a better understanding of cardiac tissue abnormalities and a better directing of tissue regeneration than other fabrication procedures allow. This is accomplished through recapitulation of the compositional, structural, and physiological complexity of human heart tissue, which can result in the production of physiologically relevant cardiac models in vitro.^[25]

In this review, we first present cardiac tissue biology and available 3D bioprinting techniques and overview the cells and biomaterials available for 3D bioprinting of cardiac constructs. We then review the literature on the application of 3D bioprinting in vascularization, delivery systems, and regeneration of myocardium, heart chambers, heart valves, and coronary artery. Finally, we explain the application of bioreactors for cardiac tissue maturation, clinical translation of bioprinted structures and limitations, and future perspectives of contemporary 3D bioprinting techniques for cardiac tissue engineering.

2. Cardiac tissue biology

The heart muscle is an extremely vascularized and contractile network, surrounded by a double-layered protective membrane, the pericardium. The heart wall consists of three layers, epicardium, myocardium, and endocardium.^[26] The human myocardium is composed of at least four dominant cell types including cardiomyocytes (CMs) (75% by volume, 20–40% by number), cardiac fibroblasts (CFs) (60–80% by number), smooth muscle cells (SMCs) and endothelial cells (ECs).^[27] The myocardium can beat synchronously due to its unique architecture. It is composed of cardiac muscle fibers connected by intercalated discs that include desmosomes for mechanically anchoring the fibers and gap junctions for conducting action potentials between the fibers, allowing the heart to contract. Sarcomeres are the repeating units of myofibrils, the contractile machinery in CMs, and consist of actin and myosin fibers. Contraction occurs as actin fibers move along the fixed myosin fibers toward the inner area of the sarcomere.^[28] This uniform contractile activity of striated and rod-shaped CMs drives optimal cardiac pump function. The mechanism of contraction lies in a process called excitation-contraction coupling, a physiological process of converting an electrical stimulus into muscle contraction. The generation of action potential in the sinoatrial node (i.e., pacemaker cells) is conducted from the atrioventricular node to CMs in the atria and ventricles via gap junctions. Subsequently, the activation of calcium channels and cell depolarization occur by action potential passing through the sarcomeres. Furthermore, calcium binds to troponin-C, resulting in the movement of the troponin complex from the actin site, as actin binds to myosin and triggers contraction. In the repolarization step, calcium removal from the cell cytoplasm leads to the troponin complex retrieval to its original condition, thus concluding the contraction cycle.^[29]

Cardiac cells are enclosed within an extracellular matrix (ECM) network created by cardiac fibroblasts.^[30] Myocardial ECM provides structural integrity and mechanical support to the resident cells and promotes cell recruitment, attachment, orientation, maintenance, proliferation, differentiation, and maturation. Myocardial ECM contains structural components such as collagen, elastin, proteoglycans, glycosaminoglycans (GAGs), and functional components such as growth factors and cytokines.^[31] The main myocardial ECM proteins are collagen type I (almost 85%) and collagen type III (almost 11%), which also form the interstitial and perivascular matrix.^[32] The mechanical integrity of the heart is preserved by the unique assembly of helical structures of collagen type I, II, III, and V, while collagen type VI also has an important role in retaining tissue structure and collagen type IV is detected in the basement membrane and blood vessels. These aligned collagen fibers provide the anisotropic nature of the myocardium.^[33] Other proteins like laminin and fibronectin improve cell adhesion, spreading, and migration.^[34] Furthermore, GAGs like heparin promote certain viscoelastic properties of the heart as well as anchoring important growth factors such as vascular endothelial growth factor.^[35] Recapitulating these complex compositional and nanostructural features is a great challenge in cardiac tissue engineering since myocardial ECM is damaged upon heart failure, and the failed tissue should be regenerated using an appropriate tissue-engineered scaffold.^[36]

3. Available 3D bioprinting techniques

3D bioprinting is the most up-to-date approach available to scientists developing 3D cell-laden constructs for tissue engineering applications. 3D bioprinting has the advantage of in situ cell encapsulation, which makes it easier for fabricated scaffolds to develop into functional tissue constructs. Other major advantages of 3D bioprinting are high precision and flexibility. This technique is capable of placing multiple cell types and materials with exceptional precision to develop a hierarchical architecture that closely resembles the structure of the designated body tissue.^[37] To date, several 3D bioprinting strategies have been developed. These techniques can be classified based on the printing approach they use, as follows.

3.1. Extrusion-based bioprinting

The most common 3D bioprinting technique used in tissue engineering is extrusion-based bioprinting.^[38] To print a 3D construct using this method, the material, called the bioink, is extruded from a nozzle in a layer-by-layer manner, and the accumulation of the bioink creates the 3D construct.^[39] To extrude the filament, three different mechanisms can be used. These mechanisms are pneumatic-based, piston-based, or screw-based.^[39b] The pneumatic-based mechanism uses pressurized air to drive the material through the nozzle. The two other mechanisms mechanically push the bioink out of the printhead (nozzle) utilizing pistons or screws.^{[39a, 39b]} The piston-based mechanism allows for more control over bioink flow through the nozzle, and the screw-based system allows for better spatial control and better release of high viscosity bioinks.^[40] Multiple parameters can affect the printability of the bioink and the formation of fine structures: the nozzle moving speed, the nozzle diameter, and the type of pressure applied.^{[39c, 41]} For instance, some studies have shown that a larger nozzle diameter or the use of viscous bioinks can lead to thicker filaments and more stable structures.^[39a]

This method of operation allows extrusion-based bioprinters to print bioinks with a high cell density (up to 10⁹ cells/mL), which reduces the time needed for cell proliferation and maturation post-bioprinting.^[39b] This technique also allows utilization of a broader range of materials, as it can bioprint hydrogels, polymer solutions, cell aggregates, and decellularized ECMs. However, highly viscous fluids are a better choice for providing suitable mechanical support for the scaffold. Moreover, this mechanism can produce large scaffolds due to its ability to rapidly extrude the material. However, a drawback of extrusion-based bioprinting is lower cell viability due to stress induced by the nozzle during deposition. This stress is larger if smaller diameter nozzles are used or higher pressure is applied.^[39b] It is also important to note that a high shear force is needed to extrude highly viscous materials, which can be a negative factor for cell viability.^[41] Hence, to minimize this shear effect, a large nozzle diameter is usually selected (around 150–300 μm) to reduce cell damage and death.^{[9, 42]}

3.2. Laser-based bioprinting

The laser-based bioprinting (LAB) technique is a less commonly used bioprinting technique due to its complexity.^[39c] It is a nozzle-free technique as it does not use a nozzle to deposit the material in the form of a filament.^{[38, 39c]} The LAB system includes a pulsed laser source, a donor slide, and a receiver (or collector) slide. The donor slide that is used to support and propel the printing material is made of thin laser-absorbing metal such as gold or titanium, and it is covered with the bioink.^{[38–39, 43]} For printing, LAB makes use of the laser-induced forward transfer (LIFT) technique. The laser beam is projected to the donor slide and is absorbed by the metal, causing a jet of the liquid biomaterial to be deposited from the donor slide onto the receiver counterpart.^{[39a, 43]}

To be properly bioprinted, the polymer solution used as the bioink must have a mid-range viscosity (1–400 mPa⋅s).^{[39a, 43–44]} For instance, bioinks that have been used with this technique include combinations of alginate and glycerol.^[43] Multiple parameters affect cell viability during the LAB process. It has been shown that an increase in laser fluence decreases cell viability after 24 hours.^[45] Moreover, printing speeds higher than 200 m/s can negatively impact cell viability. LAB resolution is dependent on bioink viscosity, bioink thickness, and the surface tension of the bioink.^{[39a, 43]}

A major advantage of LAB is that it results in high cell viability, ranging from 90% to 100% post bioprinting, and it allows the maintenance of the majority of cell functions and morphologies. In addition, LAB can provide a cell density of 10⁸ cells/mL.^{[39a, 43]} This technique also mitigates a major disadvantage of nozzle-based techniques—the problem of nozzles becoming clogged with cells or materials.^[39c] However, the drawbacks of LAB are its low availability, high cost, and complexity of the system.^[38–39]

3.3. Inkjet-based bioprinting

Inkjet-based bioprinting is another commonly used method, in which thermal or acoustic force is applied to release small droplets of bioink with controllable size on a collection plate.^{[39a, 39c, 46]} Thermal inkjet bioprinters use electricity to heat the printhead, and the resulting air pressure ejects bioink droplets from the nozzle. On the other hand, acoustic printers use a pulse created by the vibration of a piezoelectric crystal to cause the release of a small amount of the bioink out of the nozzle.^{[39a, 39c, 46]} It is possible to adjust bioink droplet size and deposition rate.^{[39c, 46]} Even though the use of inkjet printers for cardiac tissue engineering is still at an early stage, several studies have demonstrated the feasibility of using inkjet technology to incorporate vascularized structures.^[47] Moreover, this bioprinting method has been used for the fabrication of scaffolds for cartilage, bone, skin, and neural tissue engineering.^[38]

The main advantages of this technique are its high resolution and ability to form thin layers and patterned constructs.^[42] Inkjet printers can generate a high-resolution structure (around 20–100 μm) at a fast printing speed which can attain 1–1000 drops/s.^[39a] However, this method does have drawbacks, including nozzle clogging. Use of low-viscosity materials can mitigate this problem, though this limits the cell density that can be encapsulated for bioprinting.^{[39a, 46]}

3.4. Stereolithography

Another laser-based technique used for bioprinting tissue constructs is stereolithography (SLA). Like LAB, SLA is a nozzle-free technique. In SLA, photoirradiation is used to create crosslinks in the bioink, a photo-sensitive polymer that undergoes a photocuring process.^[38–39] Exposure of the polymer to the light source allows a defined pattern to be traced on the liquid polymer, which then crosslinks to form the final construct microstructure.^[39a] In most cases, the photocuring process can be facilitated by addition of a photoinitiator.^[38] Crosslinking is usually induced using a beam of visible, UV, or infrared light.^[39a] The complete SLA bioprinting system includes a light source, a polymer reservoir, and a stage that can move in three axes.^[39a]

Parameters that affect the resulting 3D construct include type and concentration of the photoinitiator, duration of light exposure, and type of bioink.^[38] Moreover, depending on the laser source, the 3D bioprinter can employ absorption of one or two photons, the latter allowing a better 3D bioprinting resolution since the photopolymerization occurs in a more precise region.^[39a] The advantages of SLA are fast printing and high cell viability since it does not introduce shear stress to cells.^{[39a, 48]} This mode of operation also allows a more realistic microstructure than those obtained by other bioprinting methods.^[39a] However, the use of laser and photo-dependent polymers can negatively influence cell viability.^[39a] To eliminate the negative impact of the beam on cell viability, visible light and a corresponding photoinitiator can be used in the bioprinting process.^[38]

4. Materials for bioinks

In a tissue-engineered scaffold, materials are used to form a temporary ECM to enhance cell recruitment, migration, adhesion, proliferation, differentiation, and maturation. This temporary ECM should also provide sufficient mechanical support for cells. In addition, it needs to be degraded over time and be replaced by the native ECM secreted by the resident cells. These characteristics should be taken into account when formulating bioinks for 3D bioprinting. So far, several types of materials have been investigated for development of bioinks.. Although, a thorough description is available in a number of relevant reviews, the following sections provide a brief overview of such materials.^[49]

4.1. Collagen

Collagen, a protein found in ECM, contributes significantly to the maintenance of the ECM’s structural and biological integrity.^[50] However, the mechanical properties and stability upon hydration of collagen are insufficient, and printing of collagen has been difficult due to the lack of control over its thermally driven gelation process as well as its low viscosity.^{[50b, 51]} To improve the mechanical properties and control of the gelation process of collagen, multiple methods such as physical or chemical intermolecular crosslinking, chemical modification, and adjustment of pH, temperature, and concentration can be exploited. Blending with other polymers can also improve collagen’s properties by modifying its degradation rate and viscosity.^{[50a, 51]} Therefore, as this biodegradable material presents low immunogenicity, excellent biocompatibility, and cell-binding sites, it is one of the most promising candidates for bioink formulation in 3D bioprinting.^{[50a, 51]}

4.2. Gelatin

Gelatin is a biodegradable protein that can be collected by hydrolysis of collagen.^[52] Gelatin can be obtained from different types of collagen through acid and base treatments. Depending on the collagen source, type, and method of hydrolysis, gelatin with specific characteristics, including amino acid composition, gel strength, isomeric point, and charge, can be achieved.^[53] Gelatin can form a thermo-reversible hydrogel and can be gelled at temperatures from 20 to 30 °C for mammalian gelatin and from 5 to 10 °C for fish gelatin.^{[53b, 54]} The mechanical properties of gelatin-based hydrogels depend on the concentration of the gelatin solution in the blend. However, physically crosslinked gelatin-based hydrogels exhibit weak mechanical properties, restricting their application in cardiac tissue engineering.^[53b] The blending of gelatin with different polymers such as polysaccharides is useful to improve cell adhesion, cell growth, vascularization, and stability of the gelatin-based hydrogel. Different components investigated with gelatin include cellulose, chitosan, alginate, and hyaluronic acid.^[53] In 3D bioprinting, gelatin-based hydrogels can either be used as sacrificial bioinks or printed as solid cell-laden constructs.^[53b] This makes gelatin an ideal material to bioprint porous structures or to be used as a sacrificial element for promoting angiogenesis and microvascular expansion into the constructs.

One of the best-known derivatives of gelatin is its methacrylated product (GelMA) which offers an improved viscosity to the gel. GelMA has been extensively used in studies for 3D printing of skin, bone, cartilage, and vasculature.^[53b] However, since GelMA exhibits low viscosity at room or elevated temperatures, printing this polymer has been challenging.^[55] Nonetheless, GelMA hydrogel shows promising characteristics for a bioink, having good biocompatibility, photo-crosslinking ability, and adjustable physicochemical properties.

4.3. Fibrin

Fibrin is a blood-derived protein that shows excellent biocompatibility. Fibrin is used in 3D bioprinting because of the cytocompatibility of fibrin-based hydrogels and their cell attachment properties.^{[51, 56]} Like collagen-based bioinks, fibrin-based bioinks show low viscosity, rapid gelation, and quick degradation. However, the drawbacks of using fibrin are poor mechanical properties and a fast gelation process that renders the printing process hard to control and can produce unstable constructs.^[51] To make fibrin suitable for 3D bioprinting, it has been used as an additive and is blended with other crosslinkable polymers by physical blending and chemical crosslinking.^{[51, 56]} An example of a fibrin-based bioink is the gelatin/alginate/fibrin hydrogel. The resultant hydrogel has shown excellent mechanical properties, cytocompatibility, and maintenance of physiological functions as well as strengthened 3D constructs.^[57]

4.4. Chitosan

Chitosan is a natural polysaccharide derived from chitin, an abundant component found in marine shells, and it is biocompatible, bioactive, and biodegradable.^{[51, 58]} Chitosan can be obtained from chitin using enzymatic or chemical processes.^{[51, 58a]} Chitosan is soluble in acidic environments, which allows it to be transformed into a hydrogel.^[58a] However, its weak mechanical properties and slow gelation rate make bioprinting chitosan difficult.^[51] It is possible to improve the mechanical properties of this material through physical blending and chemical crosslinking.^[51] As a result, in 3D bioprinting, chitosan is mostly used as a material blended with other biocompatible polymers. Recent bioinks that have been used are collagen/chitosan,^[59] alginate/chitosan/hydroxyapatite,^[60] and gelatin/sodium alginate/‌carbomethyl chitosan.^{[51, 61]}

4.5. Alginate

Alginate is an anionic polysaccharide and a natural biopolymer that can be found on the cell walls of brown algae and some bacteria.^[62] It is a non-toxic, biocompatible, non-immunogenic, and low-cost biomaterial.^[62a] It also possesses good water solubility and degradation in in vitro and in vivo environments.^{[62b, 63]} Alginate is used as a bioprinting material due to its ability to be ionically crosslinked with Ca²⁺.^[64] Other advantages of using alginate as a cell-laden material for 3D bioprinting are its abilities to support cell growth, be highly biocompatible, and not harm encapsulated cells during the bioprinting process.^[62a] However, the major drawbacks of alginate are its weak mechanical properties and slow degradation rate.^[62a] In addition, due to its lack of Arg-Gly-Asp (RGD) molecules (cell-adhesive ligands), it is a poor material for cell adhesion.^[62] It is possible to overcome the mechanical and cell attachment weaknesses through the use of crosslinking, surface modification, composites, and blends.^[64] For instance, to improve the mechanical properties of alginate-based bioinks, blends with different polymers have been investigated.^[62a] Also, the incorporation of growth factors as well as the addition of peptides favoring cell adhesion have been investigated for their biological enhancement. These modifications have shown an increase in cell adhesion in the resulting construct.^{[62a, 65]}

4.6. Agarose

Agarose is a biocompatible polysaccharide derived from marine red algae.^{[51, 66]}.It can be converted into a thermo-reversible gel and exhibits appropriate electrical conductivity for cells. The agarose-based hydrogel presents a structure similar to that of ECM and, with surface modifications, can be improved to support cell adhesion.^{[51, 66]} Since agarose promotes cell proliferation and matrix production, it is used in tissue engineering.^[66–67] However, the body lacks appropriate enzymes for agarose degradation.^[66] Moreover, agarose’s cell-adhesion capacity is weak, as agarose has no cell adhesion motifs. It is, therefore, usually used in 3D bioprinting in a blend or crosslinked with another polymer.^[51]

4.7. Hyaluronic acid

Hyaluronic acid (HA) is a material present in the ECM that plays a key role in regulating tissue functions and cell behaviors such as migration and proliferation.^[68] Hence, HA is an interesting choice for researchers to make scaffolds for tissue regeneration.^[68b] An advantage of HA is that it possesses bioactivity due to the presence of binding receptors such as CD44.^[69] However, pure HA is not suitable for 3D bioprinting since it is highly viscous and, as a consequence, is unstable during the bioprinting process, making its post-printing shape stability weak.^{[51, 68a]} To improve the mechanical characteristics of HA, HA-based hydrogels can be chemically crosslinked, physically modified, or blended. Hence, HA has mostly been studied as a component of multi-material bioinks.^{[51, 68a]} As a result, multiple HA blends have been made, such as HA/GelMA.^{[68a, 70]}

4.8. Cellulose

Cellulose is an abundantly found biopolymer and the main component of plant cell walls.^[71] This material possesses multiple derivatives such as esters, ethers, and microcrystalline cellulose.^[71a] The main advantages of cellulose fiber are its large availability, low cost, and high flexibility. Moreover, cellulose viscosity decreases when exposed to high shear rates during printing and quickly increases once the solution is printed. This property is advantageous for creating stable constructs.^[71a] Usually, unmodified cellulose is considered unsuitable for 3D bioprinting since it decomposes before it can be melted. However, nanocellulose hydrogels displaying shear-thinning abilities may be used as bioprinting material.^[71b] For instance, a decellularized ECM/cellulose nanoparticles ink was developed to recreate a physiologically relevant environment for gastric cancer cells. In the study, the incorporation of cellulose nanoparticles enhanced the mechanical properties of the printed constructs.^[72] Another study used a nanocellulose-alginate bioink with encapsulated chondrocytes to bioprint constructs showing cell viability of 86% after 7 days of culture.^[73]

4.9. Synthetic materials

While natural polymers have been more commonly used in 3D bioprinting, synthetic polymers have also been explored because of their capacity to provide good mechanical properties. To date, materials such as polycaprolactone (PCL), polylactic acid (PLA), polyglycolic acid (PGA), poly (glycerol sebacate) (PGS), and polyethylene glycol (PEG) have been used.^[74]

PCL is a linear hydrophobic polyester.^[74–75] Due to its excellent mechanical characteristics, this polymer has been extensively used in tissue engineering, including bone repair, vascular grafts, and valve tissue reconstruction.^[76] However, drawbacks of its application are its slow degradation rate under physiological conditions and the absence of biological properties.^[75–77] PLA is a biocompatible and biodegradable thermoplastic polymer extracted from renewable resources such as sugar, corn, and potatoes.^[78] It is used in tissue engineering for its excellent mechanical strength and biodegradability.^[77] It has been investigated for its use in cardiovascular scaffolds as a coating material for drug-eluting stents.^[74] However, PLA has poor cell adhesion and interaction as well as high hydrophobicity.^[77] Another synthetic polymer often blended with PLA is polyglycolic acid (PGA).^[74] PGA is used in tissue engineering because of its hydrophilicity and degradability properties.^[79] It is a biodegradable material that has been studied in cardiovascular scaffolds and heart valve applications.^[74] A copolymer of PLA and PGA, PLGA, has been approved by the Food and Drug Administration (FDA) in therapeutic devices. Constructs made from PLGA/GelMA, and other blends have shown strong mechanical properties, adjustable biodegradability, and in vivo biocompatibility.^{[53b, 80]} PGS is an FDA-approved biodegradable synthetic material that has been investigated for the tissue engineering of heart valve leaflets.^[81] PGS has shown excellent biocompatibility and degradation rates between four and six weeks.^{[74, 81]} However, it should be noted that none of the aforementioned synthetic polymers possess proper water-solubility. Consequently, they face critical limitations relating to combination with cells and development of bioinks.

Polyethylene glycol (PEG) is a water-soluble synthetic polymer, which has FDA approval for biomedical applications. It is also known as a biocompatible, non-toxic, and non-immunogenic material. PEG is considered a suitable polymer for 3D bioprinting due to its capacity to have tailorable properties that can facilitate the bioprinting process and improve proper shape fidelity.^{[80a, 82]} However, the bioinert properties of PEG prevents cell adhesion and due to its low viscosity, it cannot form a hydrogel independently. Consequently, PEG is considered to be inconvenient for extrusion-based bioprinting.^[80a] Fortunately, PEG can be modified by physical or covalent crosslinking to improve its properties.^[80a] Furthermore, to be used in bioprinting, it is usually blended with bioactive hydrogels or other components, such as the bioactive peptide RGD, to improve cell adhesion functions.^{[80a, 82]} Multiple composite PEG-based hydrogels have shown suitable degradation properties. For instance, after a few modifications, the reactive PEGX was created. It has been possible to combine this reactive group in order to form multiple bioink hydrogels: PEGX-PEG, PEGX-gelatin, PEGX-gelatin-fibrin, and PEGX-gelatin-atelocollagen.^[82–83] Another example is the blend of GelMA and PEG used for inkjet bioprinting.^{[80a, 84]} PEG can also be crosslinked by light irradiation when used in its photopolymerizable form, Poly(ethylene glycol) diacrylate (PEGDA). Another form of PEG, which has been studied by Skardal et al., is a four-armed PEG derivative blended with HA and gelatin. The study showed that this form allowed for better rheological properties and extrusion abilities than the PEGDA form.^{[82, 85]}

4.10. Composites

Composite bioinks are combinations of multiple components, such as blends of multiple polymers, nanoparticle-incorporated solutions, combinations of inorganic and organic biomaterials, etc. Composites attempt to improve the properties of the 3D printed construct by tuning the properties of the bioink by mixing of decisive elements.^{[74, 82]} Multiple composite bioinks for 3D bioprinting have been studied, including nanocomposite bioinks.^[86] These bioinks consist of the addition of nanoparticles (NPs) in solutions or hydrogels for 3D bioprinting to modify a variety of characteristics. NPs can alter the physical and chemical properties of the bioink by increasing stiffness, modifying shear-thinning properties, and modifying degradation rate in the physiological environment.^[86] Other bioink properties which can be altered by NPs are bioactivity, drug release, photoresponsivity, and conductivity.^[86–87]

Multiple studies have evaluated the addition of nanoparticles to create a composite bioink. For instance, Boularaoui et al. added gold nanoparticles (Au-NPs) to modify the rheological and printability properties of a GelMA-based bioink. In this study, they evaluated the addition of Ag-NPs or transition metal carbide sheets to enhance the printability, conductivity, and biological properties of a 5% GelMA bioink with encapsulated myoblasts. The addition of Au-NPs resulted in increased conductivity and rheological properties of the bioink, allowing for the printing of constructs with enhanced cell viability.^[88] Another use of composite bioink was demonstrated in a study that used Ag-NPs to bioprint a bionic ear. The alginate/Ag-NP bioink with encapsulated chondrocyte cells was formulated to take advantage of Ag-NPs’ conductivity and magnetic properties. The bioprinted construct was able to successfully receive radio frequency signals and allowed for the metabolic activity of cells.^{[82, 89]} Another interesting study sought to take advantage of the magnetic properties of nanoparticles to direct the architecture of printed constructs. In this study, streptavidin-coated iron nanoparticles were added to agarose and collagen bioinks. This addition resulted in the alignment of low-concentration collagen fibers by the application of a magnetic field during the printing process. This permitted the fabrication of constructs with alternating random and aligned fibers, which allowed for a higher expression of collagen II markers after 21 days in culture.^[90]

A composite scaffold can also be a combination of multiple polymers designed to modulate the biochemical, mechanical, or electrical characteristics of the printed scaffold.^[74] For instance, a study by López-Marcial et al. compared agarose-based and agarose/alginate blend hydrogels of different concentrations. They evaluated the rheological properties and printability of the hydrogels as well as the cell viability of the extrusion-based 3D printed constructs. They noticed that the shear-thinning and yield strength of the agarose/alginate composite hydrogels were similar to that of Pluronic, an existing model gel with known good printing capabilities. The agarose/alginate hydrogels also showed greater shape fidelity post-printing than the agarose-only hydrogels. Finally, they observed that the agarose/alginate constructs showed good cell viability for up to 21 days. They concluded that this 5% agarose/alginate hydrogel is potentially suitable for cartilage 3D bioprinting.^[67] Similarly, Noh et al. studied the development of a composite HA, hydroxyethyl acrylate, and GelMA (HA-g-p-HEA-gelatin) bioink for its potential application in tissue engineering. The developed hydrogel showed excellent physical properties in terms of swelling, rheology, gel morphology, and cytocompatibility. In their study, bone cells were viable when printed in lattice form, which demonstrated the potential of this hydrogel to be used as a bioink.^[68b] Another study by Yang et al. investigated a 3D printed scaffold made from a PGS/PCL hydrogel for use after myocardial infarction. They compared the PGS/PCL scaffold to separate PCL and PGS scaffolds. It was shown that the composite had better elasticity than PCL and better toughness than PGS. Moreover, the PGS/PCL scaffold showed good biodegradability in vitro, and suitable biocompatibility in a cell viability assay. After implantation in a rat model, they noticed that compared to the PGS- and PCL-only scaffolds, the PGS/PCL composite scaffold reduced the effect of myocardium apoptosis. These results indicated that this new blend allowed the preservation of heart function, promoted vascularization, induced tissue repair, and inhibited myocardial apoptosis.^[91]

5. Cells for cardiac tissue engineering

Adult cardiomyocytes are terminally differentiated cells and have minimal capability to self-regenerate. In fact, for a 25-year-old person, almost 1% of cardiac cells are replaced yearly by progenitor cells, and this amount will decrease over time.^[92] Upon myocardial infarction, recruiting cardiac progenitor cells is essential due to the huge loss of cardiomyocytes. Hence, a sufficient cell density must be supplied to support the electromechanical coupling of cardiomyocytes to the host tissue and improve functionality.^[93] Neonatal and fetal cardiomyocytes have adequate cell proliferation and functional beating behavior. However, several concerns restrict their use, including limited capacity for ex vivo expansion, immunogenicity because of their allogenic source, malignancy, limited access, limited cell survival in the hypoxic condition, and ethical issues.^[94]

Skeletal myoblasts were one of the first cell types studied as an alternative for cardiac tissue engineering. Skeletal myoblasts can be harvested from a patient’s own muscle tissue via a muscle biopsy and transplanted to heart tissue.^[95] They can be injected during coronary artery bypass graft surgery.^[96] The main advantages of using myoblasts are that they are resistant to hypoxic conditions, proliferate at a high rate in vitro, and avoid the need for immune suppression,^[97] while their main drawback is arrhythmia resulting from lack of electrophysiological coupling with the host cardiac cells.^{[15b, 98]} Furthermore, fibroblasts of mesenchymal origin can aid in preserving the tissue ECM and improving vascularization, which maintains the homeostasis of damaged tissue by inducing the proliferation of smooth muscle cells and fibroblasts.^[99] Fibroblasts can also be converted to cardiomyocytes to improve contractile properties^[100] but the formation of unpredictable cells and safety issues are still a challenge.^[101] In addition, myofibroblasts can help repair the damaged area by secreting fibronectin and releasing cytokines and growth factors.^[102]

To overcome these limitations, several naturally derived and engineered stem cells have been investigated for cardiogenic differentiation and formation of cardiomyocytes.^[103] This includes embryonic stem cells (ESCs),^[104] bone marrow derived stem cells (BMSCs),^[105] cardiac stem cells (CSCs),^[106] endothelial progenitor cells (EPCs),^[107] cardiac progenitor cells (CPCs),^[108] adipose-derived stem cells (ADSCs),^[109] and induced pluripotent stem cells (iPSCs).^[110]

ESCs derived from the inner cell mass of the blastocyst can be differentiated into a wide variety of cell types, including cardiomyocytes, by cultivating in a medium containing activin A and bone morphogenetic protein 4 (BMP4).^[94a] It has been shown that using human ESCs improved electromechanical coupling of cells with host cardiomyocytes, cell alignment, and restoration of heart function.^[111] Moreover, animal myocardial infarction (MI) models demonstrated integration with the host cardiac tissue and promotion of ventricular function.^[112] However, immunosuppressive therapy is essential for allogenic ESCs and many ethical and legal concerns accompany the usage of ESCs (for example, they can form teratomas in animal models).^[113]

Bone marrow is composed of several types of stem cells including hematopoietic stem cells, mesenchymal stem cells, peripheral stem cells, and endothelial stem cells.^[114] Hematopoietic and mesenchymal stem cells have angiogenic and anti-inflammatory effects and can be differentiated to cardiomyocytes.^[113] Endothelial progenitor cells can promote cardiac function. These cells can induce regeneration of cardiac cells along with micro vascularization and red blood cell perfusion, but their poor availability limits their application. This limitation also applies for cardiac stem cells and cardiac progenitor cells.^[113] On the other hand, the advantages of bone marrow-derived cells are their easy availability from the patient and their capability for transdifferentiation, but their effect on cardiac function is temporary.^[99b]

iPSCs are cells reprogrammed by augmentation of transcription factors Klf4, cMyc, Oct4, and Sox2 and can be differentiated to form cardiac progenitor cells from somatic cells.^[15b] For instance, different studies have been conducted using iPSCs derived from fibroblasts.^[115] Different methods have been studied for obtaining cardiomyocytes from iPSCs. In this regard, a two-step protocol has been examined in which cytokines such as BMP4, activin A, and Wnt3a were supplemented, and then BMP inhibitors and Wnt inhibitors were used to form cardiomyocytes.^[116] An important advantage of iPSCs is that the somatic cells can be derived from the patient, thereby reducing the probability of immunogenicity. This also removes the ethical concerns of using human ESCs. An additional benefit is easy availability of cells, due to their pluripotency and high rate of proliferation.^[117]

6. 3D bioprinting of cardiac tissues

The property of beating is one of the most important aspects of cardiac tissue, and the quality that distinguishes it from other organs. For an appropriate beating, it is important to have an ECM that supports sufficient electrical conductivity to enhance the electrical coupling of neighboring cells. To this end, Zhu et al. developed a new bioink for 3D printing purposes that contained gold nanorods (GNRs) in GelMA solution.^[118] The viscosity of the formulated bioink was low enough that it could support the incorporation of high cell numbers for bioprinting, while GNRs would reduce the electrical resistance of the GelMA polymer. For bioink preparation, they first coated commercially available GNRs with GelMA, with little change in their size and efficiency. Subsequently, coated GNRs were blended with 7% GelMA to form a base for the bioink. Conducting several tests on GNR-loaded GelMA hydrogels, they showed that incorporation of GelMA-coated GNRs increased Young’s modulus and surface roughness of the gels. However, at high concentrations, an adverse effect was observed due to UV reflection by the particles. Thus, GNRs at a concentration of 0.25 mg/mL gave the highest Young’s modulus, though for all gels the modulus was lower than 6 kPa, far below the modulus of native cardiac tissue. By culturing cardiomyocytes (CMs) on gels, they showed that GNR-loaded hydrogels possessed enhanced cell adhesion and retention in comparison with GelMA alone as well as higher expression of Cx43 and troponin I, two important factors in cardiac cells (Figure 1a). Moreover, the cells showed synchronous beating two days after incubation.

Figure 1.

(a) Fluorescence images showing higher expression of Cx43 on GNR-loaded hydrogels after one week of incubation, (b) printing of scaffolds with different grid sizes confirming the capability of bioinks for precise printing, and (c) printed spiral structures in the gelatin bath (top) and the scaffolds after removal from the gelatin bath showing intact structure. Reproduced with the permission from Ref.^[118]. Copyright 2017, John Wiley & Sons. (d) (i) Schematic representation of the patch, (ii) the printing technique and (iii) actual printed scaffold. (e) (i) In vivo implanted 3D bioprinted scaffolds in rat omentum, and (ii) immunostaining of explanted scaffolds after a week of implantation to observe nuclei (blue) and sarcomeric actinin (red). (f) The whole heart removed from bath showing the ventricles in blue and red; (g) confocal microscopy showing the survival of different cells one day after printing. Reproduced from Ref.^[119].

After showing enhanced properties for GNR-loaded GelMA, the GelMA/GNR composition for bioprinting was blended with 2% alginate, CMs, and cardiac fibroblasts (CFs) (with a concentration of 5 × 10⁶ cells/mL) and was printed by (1) core-shell needle with the bioink deposited in the core and CaCl₂ deposited from the shell, and (2) the same bioink printed in a gelatin bath containing CaCl₂. In the first approach, the incorporation of alginate into the solution mixture and the dual crosslinking strategy (i.e., ionic and photocrosslinking) could successfully result in the fabrication of GNR-loaded nanocomposite scaffolds with tailored porosity and grid size (Figure 1b). The second approach could also successfully result in the printing of spiral constructs as shown in Figure 1c. Interestingly, it was shown that GNRs are trapped within the bioprinted scaffold structure and would be released only by degradation of the scaffold. Co-encapsulation of CMs and CFs was also successfully achieved in bioprinted scaffolds, and due to gradual degradation of alginate, cells could spread within the construct 5 days after bioprinting, with continuous proliferation. Moreover, Zhu et al. observed the formation of a tissue layer with cells elongated rather than spherical, and with a higher expression of Cx43 (gap junction protein) after 12 and 14 days, respectively. With respect to cardiac beating, their scaffold showed improved synchronized contractile frequency. They also postulated that GelMA-coated GNRs prohibited excessive CF proliferation, which further enhanced the contractile behavior of scaffolds.

Although cardiac tissue bioprinting has evolved significantly in the last few years, there is a very limited amount of research investigating the bioprinting of a full and thick vascularized cardiovascular patch or heart-like structure. The work done by Noor and coworkers is one of these efforts.^[119] They used omental tissues to form a personalized collagen-based bioink that could match the immunological requirements of each patient. Using this material, they developed two models. In their first model, they used CT and computer-aided design to elaborate a 3D cardiac scaffold personalized to the patient’s left ventricle, with a specific patch size and large vessel geometry. To bioprint this patch, they used an extrusion-based bioprinter and mixed cells (either iPSC-derived CMs or neonatal cardiac cells) with their formulated bioink. To form the blood vessels, they prepared a sacrificial gelatin bioink loaded with ECs and fibroblasts (Figure 1d). Their results showed that in printed patches, ECs could successfully adhere to the edge of the formulated bioink and form a large (≈ 300 μm) vessel-like layer in between the CMs after a week of incubation. An in vivo examination of these millimeter-thick, vascularized patches confirmed their effectiveness as cells were elongated and aligned with contractile potency (Figure 1e).

In the second model (whole organ printing), to prevent the structure collapsing under its own weight, the gel was bioprinted within a support composed of alginate microparticles. They incorporated CMs and ECs in two separate bioinks and could successfully print a small heart with a height of 20 mm and a diameter of 14 mm. This bioprinted heart demonstrated suitable integrity of different compartments, which is an important finding for further use of the materials. It was also shown that cells are viable and well distributed after printing and cultivation for one day (Figure 1f and g).

6.1. Myocardium

Proper incorporation of biological cues in scaffolds could significantly enhance the outcome of bioprinting for cardiac regeneration. The cardiac extracellular matrix hydrogel (cECM) is a great candidate for this purpose and has shown some promising results. In one work, Bejleri et al. combined cECM with GelMA and human cardiac progenitor cells (hCPCs) to form a bioink for bioprinting of cardiac patches.^[120] GelMA was added as it could adjust the bioink viscosity for proper printing, scaffold mechanical properties, and cell viability. The bioink was cooled to 10 °C before bioprinting commenced. In this way, GelMA formed physical gels to enhance the printability. Subsequently, they chemically crosslinked the scaffolds using visible light. The dispersion of cECM in the bioink was found to be homogenous, with formation of compact fibers at physiological pH and temperature.^[120]

The gels retained a good distribution of hCPCs within their structure, and high cell viability (near 80%) was observed up to 6 days after bioprinting. Gene expression analysis also revealed that cECM positively enhanced cardiac and endothelial differentiation of hCPCs. In addition, a sign of scaffold remodeling by hCPCs was observed, as stiffness was higher for cell-laden hydrogels incubated for 21 days compared with freshly bioprinted structures. The in vivo placement of patches on rat epicardial surfaces revealed that gels were capable of remaining on the surface and keeping their shape for the test time of two weeks while they integrated with the native tissue and formed a vascular network for nutrient support of implanted cells.

3D bioprinting of cardiac tissue can also be done without the use of biomaterials. Narutoshi Hibono’s group published a series of reports detailing the application of the following method for the formation of biomaterial-free 3D bioprinted cardiac constructs.^[121] First, they co-cultured human induced pluripotent stem cells (hiPSCs), human cardiac fibroblasts (HCFs), and human umbilical vein endothelial cells (HUVECs) for three days in ultra-low attachment plates and observed that spheroids were formed after 24 hours of co-culture, while beating was observed after 48 hours of incubation (Figure 2a).^[121a] For bioprinting, they took advantage of a needle array that gave high-precision positioning of spheroids (Figure 2b). Using this system, they were able to confirm that these multicellular spheroids can be bioprinted without losing their function, since the bioprinted tissues showed no distinguishable boundaries between individual spheroids while retaining their spontaneous beating. Further maturation of scaffolds after removal of the needle array showed the remodeling capacity of the structures, as holes created by removal of the needles were filled in (Figure 2c).

Figure 2.

(a) Spheroids of hiPSCs/FBs/HUVECs after 24 hours of incubation in plates. Reproduced with permission from Ref.^[121a]. Copyright 2017, MyJoVE Corporation. (b) Schematic representation of needle-based bioprinting used for the formation of tissues. Reproduced from Ref.^[121b]. (c) Bioprinted spheroids just after removal of needle array showing spheroids’ integrity (left) and remodeling of cardiac tissue to cover the holes upon further maturation of tissues for three days (right). Reproduced with permission from Ref.^[121a]. Copyright 2017, MyJoVE Corporation. (d) Isochronal activation maps for bioprinted structures with different cell ratios, (e) H&E and (f) immunohistological staining of the patches implanted on rat heart. Reproduced from Ref.^[121b]. (g) Confirmation of retention of the cells in rat 28 days after surgery (green shows troponin T and blue shows DAPI). Reproduced with permission from Ref.^[121c]. Copyright 2019, John Wiley & Sons.

Following this work, they utilized the same biomaterial-free bioprinting approach to form spheroids of hiPSCs/FBs/HUVECs with various ratios: 70:15:15, 70:0:30, and 45:40:15.^[121b] One of their interesting findings was that cardiospheres could be formed only if at least 15% of ECs or FBs were co-cultured within the spheroids. Furthermore, they noticed that only patches formed from 70:15:15 and 70:0:30 spheroids possessed appropriate electrical integration, while in the middle of the 45:40:15 patch, a functional electrical propagation blockage was observed (Figure 2d).

The in vitro evaluation of the reported optimum composition (70:15:15) confirmed high cell viability (near 95%) and functionality for all three cell types as well as the development of neovascular structures in the patches. Subsequently, an in vivo implantation of these patches in rats was conducted, which reconfirmed the vascularization of patches, viability of cells in the in vivo environment, and successful engraftment after one week (Figure 2e and f).

Finally, Yegung et al. conducted a more extensive in vivo examination using a rat myocardial infarction model.^[121c] They, too, found a 95% cell viability for biomaterial-free patches four weeks after implantation. More importantly, the patches increased the survival of rats (100% compared with 83.3% for control) four weeks post-surgery. Furthermore, cellular retention was confirmed even four weeks after implantation (Figure 2g). Consequently, it was observed that in patch-treated rats, the scar area was significantly smaller, with a higher degree of neovascularization at the infarction site. Together, these studies show the high potential of biomaterial-free 3D bioprinting for cardiac tissue regeneration. However, limitations such as the fragility of printing structures, sphere diameter, structural complexity requirements, and limitations for relatively large size printing should be addressed before further application.

Another approach for bioprinting heart tissues is to print electronic-filled cardiac patches. Asulin et al. developed a cardiac patch incorporating soft electronic components to stimulate cells and record heartbeat following implantation.^[122] To design this patch, they used 3 different bioinks; an ECM-based ink encapsulating neonatal ventricular cardiomyocytes, a PDMS ink containing graphite flakes to be used as the electrode, and another ink containing liquid PDMS used to passivate the electrode. They demonstrated that this patch could successfully record the tissue contraction and provide electrical stimulation, while being printed with homogeneously distributed cardiac cells. Finally, a cell viability assay was performed after 12 days of culture, depicting high cytocompatibility of the bioprinted constructs.

One of the most recent advances in 3D bioprinting of cardiac tissues is to develop a cardiac patch using an oxygenated bioink. Erdem et al. investigated the effect of oxygen release from a formulated GelMA calcium peroxide (CPO) bioink.^[123] They postulated that oxygen (O₂) could be released from the CPO particles by the production (and subsequent decomposition) of hydrogen peroxide (H₂O₂). Various concentrations of CPO were added to GelMA, and cell-laden hydrogels were carefully bioprinted at 4 °C. The addition of CPO had a significant effect on the extrudability, shape fidelity, and printing accuracy of the resulting structures. Moreover, incorporating CPO had a slight impact on the physical properties of the fabricated GelMA hydrogels as the swelling, pore size, and compression modulus were different when compated to CPO-free hydrogels. Specifically, CPO hindered UV irradiation and ultimately crosslinking, resulting in a looser structure. They also showed that the addition of CPO into the gels increased the viability of seeded CMs and fibroblasts when exposed to hypoxic conditions. Moreover, an evaluation of metabolic activity of cell-laden hydrogels cultured in hypoxic conditions showed similar outcomes to cells cultured in normoxic conditions. These important findings demonstrate the potential of advanced bioinks to enhance the survival of cardiac patches when exposed to ischemic conditions, such as myocardial infarction. However, they also showed that at high concentrations (> 1%), CPO could be toxic to cells. For instance, at 1% CPO, cell viability was reduced by nearly 60% at 4 and 7 days post-printing.

6.2. Heart chambers

Heart chambers serve as reservoirs for blood to be pumped through the human body. Consequently, appropriate modeling of the chambers is vital to a successfully tissue-engineered heart. To this end, Kupfer et al. sought to develop a bioink capable of supporting high hiPSC viability, along with promotion of their proliferation and differentiation into cardiomyocytes.^[124] The bioink contained GelMA, collagen methacrylate (ColMA), and other ECM proteins, i.e., fibronectin (FN) and laminin (LN), which not only provided cell-binding motifs and differentiation signaling to hiPSCs for an enhanced cellular response but also improved the printability of the prepared bioink. Optimization of the bioink composition resulted in a bioink that could support high cell density (1.5 × 10⁷ cells/mL), high cell viability, and colony formation (Figure 3a). In addition, based on fluorescence staining, optimum gel formulation (10 w/v% GelMA/0.25% ColMA/FN/LN) successfully supported differentiation of hiPSCs into cardiomyocytes after 32 days (Figure 3b). Furthermore, rheological assessment demonstrated that the optimum gel had suitable mechanical properties, as its stiffness was similar to the stiffness of the late embryonic heart.

Figure 3.

(a) Colony formation after two weeks of incubation; (b) successful differentiation of iPSCs toward cardiomyocytes in scaffolds; (c) digital design, bioprinted heart chambers, and geometrical deviation of the bioprinted hydrogel with the model design. (d) (i) Presence of cells at the outer boundary of printed structures (white oval shows the interior boundary) and (ii) higher magnification showing thick cell layers at the outer boundary and (iii) high magnification sarcomeric striations. Reproduced with permission from Ref.^[124]. Copyright 2020, American Heart Association, Inc. (e) Schematic representation of left ventricle tissue and image of printed ventricle structure. (f) (i) Immunofluorescent imaging showing the intact collagen layer in magenta and CMs in green, (ii) high magnification image showing high interconnectivity of CMs within the printed ventricle, and (g) calcium transient traces analysis showing spontaneous contractions of tissue and its response to 1 Hz and 2 Hz stimulation. Reproduced with permission from Ref.^[125]. Copyright 2019, American Association for the Advancement of Science.

This formulation was later used for extrusion bioprinting of a two-chamber structure with input and output vessels inside a gelatin support bath. Characterization of the bioprinted chambers using MRI scans revealed high precision in the printed structure (compared to the designed template) in both cross-section and overall structure (Figure 3c). Moreover, after bioprinting of cells in the structure and incubation for two weeks, they observed that around 90% of the structure was populated by cells, reconfirming the proliferative potential of the bioinks. Regarding cell differentiation in the bioinks, as confirmed with fluorescent staining shown in Figure 3d, cells were able to undergo differentiation to cardiomyocytes, smooth muscle cells, and ECs with high efficiency. However, cells were mostly present in the outer region of the bioprinted structures. Interestingly, they found that although the printed chambers were not perfused in a bioreactor or stimulated using mechanical or electrical cues, when incubated for maturation, constructs still showed vigorous expression of Cx43 between neighboring cardiomyocytes along with high levels of potassium ion channel expression. More detailed investigations of the heart chambers’ electrical functionality revealed that peak amplitude and beating frequency were not subject to change over time, up to six weeks after hiPSC differentiation. The fabricated tissues also responded properly to drug stimuli when calcium handling response was manipulated via drugs.

In other work, Lee et al. utilized unmodified collagen as a bioink.^[125] This bioink had several interesting advantages including simple self-assembly procedure by pH adjustment, ability to be used at high concentrations, enhanced mechanical properties, and high printing fidelity for the fabrication of complex structures. Their bioprinting process was based on the freeform reversible embedding of suspended hydrogels (FRESH) technique with an improvement of the gelatin support bath they developed in their previous work.^[126] They adapted the gelatin particle synthesis to achieve microparticles with smaller size and narrower size distribution, uniform spherical shape, and adjustable mechanical properties of the bath, which all enhanced printing resolution. This modified printing procedure was further used to bioprint left ventricle tissue with a sandwich structure composed of a collagen bioink outer layer for the shell, and human embryonic stem cell-derived cardiomyocytes (hESC-CMs) and cardiac fibroblasts in the core (Figure 3e). The ventricles were then cultured in vitro for 28 days. They showed that this bioink and printing procedure resulted in high cell viability (≈ 96%) and formation of an interconnected dense layer with synchronized beating (Figure 3f). It was also confirmed that the observed contraction propagates within the whole printed tissue and that the printed structures had a spontaneous beat rate of 0.5 Hz which could be paced by 1 and 2 Hz stimulation as revealed by calcium transient traces analysis (Figure 3g).

6.3. Heart valves

Heart valves are mainly responsible for controlling the blood flow direction in the body, and any dysfunction of these valves can cause serious health problems. Therefore, attempts have been made in the field of bioprinting to develop functional heart valves. As early pioneers, the Butcher group was one of the first to start working on 3D bioprinting of heart valve conduits.^[127] In their first work, they used gelatin and alginate polymers for their bioink formulation and crosslinked printed scaffolds using CaCl₂. They found that in this system a high gelatin ratio hindered bioprinting while high alginate content resulted in material spreading. In addition, they observed that almost 80% of initial gelatin leached out over a week, resulting in a reduction of ultimate tensile strength of scaffolds over time. Nevertheless, when valvular interstitial cells (VICs) were encapsulated in the scaffolds, constant mechanical properties were observed over this period with high cell viability (Figure 4a). Further, they attempted to print the heart conduit by encapsulation of VICs in the heart valve structure and smooth muscle cells (SMCs) in the heart conduit. As confirmed by the images in Figure 4b and c, bioprinted scaffolds had high geometrical similarity with the native conduit. Moreover, the encapsulated cells were highly viable in the structure and expressed a higher amount of alpha-smooth muscle actin (α-SMA) and vimentin compared to cells encapsulated in hydrogel disks (Figure 4d).

Figure 4.

Mechanical response of (a) cell-free scaffolds and VIC-encapsulated scaffolds over one week of culture. (b) Model prepared from micro-CT images for bioprinting of heart conduit, (c) image of valve conduit after bioprinting, and (d) fluorescent image of first printed two layers (SMC green, VIC red) showing high cell viability. Reproduced with permission from Ref. ^[127]. Copyright 2012, John Wiley and Sons. (e) Live/dead assay showing the viability of cells within different GelMA/HAMA bioink formulations and depths, (f) live/dead staining showing high VIC viability in the depth of heart valve conduits, and (g) Masson’s trichrome staining showing newly deposited collagen in the heart valve conduit. Reproduced with permission from Ref.^[128]. Copyright 2014, Elsevier B.V.

In a subsequent study, they developed a new photocrosslinkable bioink composed of methacrylate hyaluronic acid (HAMA) and GelMA, which have been shown to support cells with stronger biological cues.^[128] They showed that by varying the concentration of GelMA and HAMA, bioink viscosity could be significantly altered to reach a formulation suitable for cell encapsulation and bioprinting. As in their previous work with simple shape bioprinting, with their new bioink and evaluation of cell viability, they confirmed high viability for cells (>90%) even at high depths (700 μm) (Figure 4e). Afterward, they bioprinted simplified trileaflet heart valve conduits using the bioink loaded with VICs. The bioprinted heart valve conduits showed promise as they had high fidelity and structural integrity one week after bioprinting. Furthermore, similar to the simple bioprinted structure, VICs could survive even several micrometers beneath the surface (Figure 4f). Also, based on histological evaluation, cells could deposit collagen and remodel the structure over time (Figure 4g).

Most recently, the same group studied the conjugation of TEMPO-modified nanocrystalline cellulose on GelMA (MNC-GelMA) and its application as a novel bioink for heart valve development.^[129] While investigating the properties of this new biomaterial, they found that MNC-GelMa hydrogels possess a nonlinear compressive behavior with increased stiffness, most likely due to MNC. Additionally, MNC-GelMA gels enhanced cell spreading and promoted GAG deposition. However, it was found that encapsulating human adipose-derived mesenchymal stem cells into the gels decreased α-smooth muscle actin (α-SMA) expression while increasing vimentin and aggrecan expression, suggesting a quiescent fibroblastic phenotype. Finally, as a proof of concept, they demonstrated that MNC-GelMA can be successuly bioprinted into a heart valve-like tissue with a multilayered configuration.

In a recent investigation, Maxson et al. bioprinted a simple disk shape from collagen/mesenchymal stem cell bioink as a heart valve scaffold.^[130] They used this structure as they only wanted to conduct an in vivo subcutaneous implantation to test the re-cellularization capability of scaffolds. Their examinations showed within a 12-week timeframe that scaffolds undergo resorption, ECM synthesis, stabilization, and remodeling stages. The resorption stage was observed at the beginning of the experiment and confirmed by the reduction of tensile strength and increase in inflammatory marker expression. In contrast, the other stages showed a reduction in inflammatory response and an increase in tensile modulus. Furthermore, cells began to infiltrate the structures and deposit collagen, elastin, vimentin, and α-SMA to strengthen the matrix. However, their scaffold still had a significantly lower strength than the aortic valve cusps.

Bioprinted heart valves can also be used as models for disease screening. Van der Valk et al. investigated calcific aortic valve disease.^[131] For this purpose, they first measured the mechanical properties of each of three aortic valve layers and then tried to develop a bioprinted aortic valve with biomechanics resembling leaflets. They also utilized a HAMA/GelMA mixture as their bioink and tuned its composition and UV crosslinking time to achieve the mechanical properties of the spongiosa and fibrosa layers. The bioprinting of VIC-encapsulated bioinks was then conducted inside a sacrificial Pluronic mold to form disk-shaped structures. Investigating the role of VIC encapsulation on hydrogel mechanical properties, they found that VIC encapsulation in spongiosa-like gels did not affect the mechanical properties of the hydrogels. However, for fibrosa-like gels, the presence of VICs resulted in a reduction of Young’s modulus, though they were still stiffer than the spongiosa-like layer. More importantly, they showed that VICs encapsulated in fibrosa-like hydrogels are capable of responding to osteogenic media by formation of microcalcification nodule density, while spongiosa-like hydrogel and 2D culture did not show any microcalcification. These results confirmed that the bioprinted models they developed have potential for high-throughput screening of calcific aortic valve disease in vitro.

6.4. Vascularization

Cells are completely dependent on receiving their required oxygen and nutrition supply through the blood for proper functioning, proliferation, and viability. Several works have stated that oxygen and nutrition diffusivity is only sufficient at small distances (≈ 200 μm). Hence, the vascular network is an essential component in tissue engineering.^[132] Skylar-Scott et al. investigated the performance of a novel bioprinting approach named “sacrificial writing into functional tissue (SWIFT)” for the fabrication of perfusable vascularized cardiac tissue.^[133] For SWIFT bioprinting, they first created iPSC spheroids by culturing cells in specialized plates. These spheroids were then blended with collagen/Matrigel as an ECM solution. Subsequently, using a sacrificial gelatin bioink, vascular channels were printed inside the structures as shown in Figure 5a. This approach had the significant advantage of using high cell density to resemble the in vivo condition more closely. The ECM solution could also mechanically support the cells for successful printing of vascular channels at low temperatures (0 to 4 °C), while upon heating to 37 °C, gelation of the ECM solution occurs and the sacrificial bioink is removed for immediate perfusion of the channels. They effectively utilized this technique to fabricate a vascular channel through a bed of iPSC-CM/cardiac fibroblast spheroids (Figure 5b). Furthermore, they investigated the capacity of this procedure to form a functional, perfusable cardiac construct composed of hiPSC-CM/stromal cell spheroids, ECM solution, and fibroblasts. The viability of cells within this perfusable tissue was confirmed after one day. Moreover, after 8 days, the tissue showed a pervasive sarcomeric architecture with a 20-fold improvement in contractility. It also showed enhancement in its beating properties, including synchronization, capacity, and response to paced stimulation. Finally, after confirming the functionality of vascular-channel-embedded cardiac tissue as proof of applicability, they showed that this technique has the potential to print complex structural geometries like the vascular network of the left anterior descending (LAD) artery. This pioneering work could confirm the capability of 3D bioprinting for vascular tissue engineering and the formation of complex vascular networks in 3D structures.

Figure 5.

(a) Detailed mechanisms for the fabrication of vascularized structures using the SWIFT technique. (b) (i) Cross-sectional image of cardiac spheroids and bioprinted vascular channel, and (ii, iii) the respective immunostained images. Reproduced from Ref.^[133]. (c) Immunofluorescence staining and cell orientation for iPSC-CM encapsulated hydrogel (left) and bioprinted scaffold (right). (d) Schematic representation of geometries used for bioprinting of vascularized scaffolds: Janus with HUVECs and iPSC-CMs printed in a single strand side-by-side, 4:2:4 with two layers of HUVECs sandwiched between layers of iPSC-CMs, and 2:2:2:2:2 prepared by altering two layers of iPSC-CMs and HUVECs and representative fluorescence images of each structure after a week of culture. Reproduced from Ref.^[134].

In another study by Maiullari et al., microfluidic-head extrusion bioprinting was utilized to bioprint alginate/PEG-fibrinogen bioinks and create vascularized cardiac tissues.^[134] Since alginate does not have cell-binding motifs, it was used in this study as a temporary material responsible for printability and precision improvement while PEG-fibrinogen provided cells with biological cues. In their investigations, they loaded iPSCs into their bioink and cultured the bioprinted scaffolds for two weeks. They found that compared with bulk hydrogels, colony size distribution was much more homogeneous in 3D bioprinted structures, as nutrients and oxygen could more easily diffuse to reach the iPSCs. Moreover, the gels could support keeping the stemness of the cells within this culture time. Similarly, encapsulation of iPSC-CMs within 3D bioprinted structures using modified bioink (supplied by PEGDA) resulted in the induction of cell maturation as well as higher cell alignment in the bioprinting direction in comparison with bulk hydrogels (Figure 5c). After confirming the good functionality of bioprinted iPSC-derived CMs, HUVECs were added to iPSC-CMs and bioprinted in three different geometries Janus, 2:2:2:2:2, and 4:2:4 as schematically represented in Figure 5d. Maiullari et al. demonstrated that Janus and 4:2:4 geometries have fewer cell-free spaces compared with 2:2:2:2:2. These structures also showed better HUVEC distribution and higher gene expression, while VEGF and cyclin D1 expression was highest in the Janus construct. However, iPSC-CMs did not form functional tissues in any of the structures. They attributed this observation to short culture time and medium composition. The in vivo evaluation also revealed better performance of bioprinted structures compared with bulk hydrogels. In this test, Janus geometry showed better formation of branched, well-developed blood vessels compared to the other geometries.

In another study, Mao, et al., used coaxial electrohydrodynamic bioprinting (EHD), a novel method to bioprint thick (≥ 3mm) pre-vascularized cell-laden constructs.^[135] This method is based on the application of an electric field around a coaxial printing nozzle, to act as the driving force for the deposition of the cell-laden hydrogel. Applying a low electric field minimizes cell damage and death during bioprinting. Alginate was used as a shell solution and a cell-laden collagen/calcium chloride solution was used as the core of the filaments. They also demonstrated that the perfusable constructs could be succesfuly fabricated by replacing the collagen/calcium chloride solution (core bioink) with calcium chloride. Additionally, this approach brought to light the potential of bioprinting pre-vascularized constructs by encapsulating viable and proliferative HUVECs to a collagen/calcium chloride solution.

7. Bioreactors for cardiac tissue maturation

Although similarity of both cell composition and composition and structure of the incorporated biomaterial to natural cardiac ECM are critical for the engineering of a successful cardiac tissue construct, they are not necessarily sufficient. Another important aspect that needs careful consideration is helping cells maturate to gain the desired cell phenotype, electrical impulses, and/or contractile force. An unmatured cardiac construct could not be used for appropriate drug screening and disease modeling.^[136] Bioreactors can be implemented to tackle this issue. Bioreactors are capable of simulating the in vivo environment in the in vitro culture, resulting in an improvement in cell and tissue maturation.^[137] Due to these advantages, bioreactors have been extensively investigated in tissue engineering.^[138] However, each tissue needs specific conditions to grow ex vivo which accentuates the need for engineering tissue-specific bioreactors.^[139]

Ronaldson-Bouchard et al. developed a new protocol to assist the maturation of hiPSCs toward adult-like CMs.^[140] The bioreactor design contained two PDMS pillars with heads designed for tissue alignment, a polycarbonate support that held the pillars, and a well that would hold the cell-laden scaffolds (Figure 6a and b). In their work, hiPSCs were differentiated toward CMs, and while cells still had developmental plasticity, were co-encapsulated in fibrin hydrogels with dermal fibroblasts and subjected to the bioreactor. The scaffolds were cultured in a serum-free medium supplemented with several required biological cues including fatty acids and hormones. After one week, when the tissues were formed, they were exposed to electromechanical stimulation to induce maturation. Instead of constant stimulation, they used what they called an “intensity training” regimen in which the cells were forced to adapt to increasing electrochemical signals in two weeks (from 2 to 6 Hz with a 0.33 Hz daily increase). Their extensive evaluations showed that the encapsulated cells possessed the main characteristics of adult-type CMs, including force-length and force-frequency relationships, designating their bioreactor and culturing process as yielding the most matured cardiac tissue phenotype developed so far.

Figure 6.

(a) A diagram showing the dimensions of pillars designed for the bioreactor and (b) the pillars attached to polycarbonate support. Reproduced with permission from Ref.^[140]. Copyright 2019, Springer Nature. (c) Schematic representation of the bioreactor chamber and (d) scaffold holder cross-section, (e) the optical image of the same parts presented by the schematic. Reproduced from Ref.^[141].

Visone et al. created the first bioreactor system to combine electrical stimulation, interstitial perfusion, and optical monitoring of the tissue within the culturing period.^[141] The bioreactor chamber was designed to be compatible with an oscillating perfusion bioreactor for bidirectional interstitial perfusion and to have a transparent part specifically designed for scaffold placement that included electrodes for electrical stimulation. The scaffold holder had male and female parts and pillars for control of scaffold positioning and its fixation (Figure 6c to e). The scaffold holder was designed to produce a uniform electrical field with fluid perfusion within the scaffold. Using this bioreactor with its online monitoring system to maturate neonatal rat CMs-loaded collagen scaffolds, they could prove that this system assures cell viability and enhances cell maturation, as CMs showed higher protein expression, beating properties, and contractility.

Recellularization of decellularized 3D structures is very challenging with conventional methods since cells may not reach distal points for complete recellularization. Recellularization challenges are more critical in heart valves; if cells do not infiltrate within the leaflets, they are vulnerable to detachment by blood shear after implantation. In a recent work, VeDepo and coworkers focused on solving this problem in heart valve leaflets using bioreactors.^[142] They used ovine aortic heart valves as their models and recellularized them with bone marrow mesenchymal stem cells, testing them within four different bioreactors: i) hypoxia with negative cyclic pressure (−20 mmHg), ii) hypoxia with high cyclic pressure (120 mmHg), iii) normoxia with negative cyclic pressure, and iv) normoxia with high cyclic pressure. For all four protocols, the seeded heart valves showed standard appearance and leaflet coaptation after two weeks while hypoxic conditioning resulted in higher cell infiltration into leaflet interstitium. On the other hand, protein expression and biochemical assay did not show any preference for one conditioning procedure over another, although all four secreted significant amounts of GAG and collagen compared with cryopreserved valves and decellularized valves, indicating their appropriate functionality.

8. Clinical translation of bioprinted structures

8.1. Regulatory pathways for translation

3D bioprinting has many potential and future clinical applications.^[143] However, as with all new biotechnologies, risks and benefits should be studied carefully. To this end, the ethical, safety, and regulatory challenges of translating 3D bioprinted tissues and organs should be addressed. Ethical issues are related to ownership, harvesting (cell type), and biomaterials in addition to commercialization pathways. Biocompatibility, aseptic conditions maintenance, and ex vivo manipulation are associated with safety concerns. From a regulatory perspective, regulations of bioprinted structures are not clear and a robust configuration is required. Despite numerous studies on 3D bioprinted tissues and organs, concerns associated with the translation aspects of this method have been neglected.^[144]

One of the ethical concerns of 3D bioprinting is ownership of the prototype.^[145] Often, patent protection for 3D bioprinted parts is complicated due to regulations surrounding non-patentable medical techniques. This means that tissues and human organs are not patentable entities, and in the best-case scenario, the bioprinting process or technique is patented, not the human product.^[146] Another consideration is that when a 3D bioprinted tissue is fabricated, ownership is not clearly defined, as the cell donor, clinician, or company may claim ownership.^[147] Living cells are an essential part of the creation of 3D bioprinted tissues and organs. Sometimes, using the patient’s own cells is not feasible, owing to issues such as poor regeneration capacity or an unknown level of cell differentiation. On the one hand, using already differentiated cells is difficult. On the other hand, adult stem cells may not be available, and sometimes their manipulation does not lead to a fully-functional tissue.^[148] Moreover, using embryonic stem cells entails many ethical and safety issues.

From the point of view of safety, the environment cells are grown in is important. Factors associated with the environment include (but are not limited to) culture media, sterilization, and growth factors. It is necessary to maintain an environment similar to the body’s conditions while not inducing any abnormalities. Cellular senescence has been reported due to ex vivo culture conditions and manipulation.^[149] Other studies have reported the transformation of human stem cells to malignant cells in culture.^{[149a, 149d, 149e, 150]} In addition to genotypic alterations, even the quality of regular tissue culture plastics available on the market can affect culture conditions.^[149b] During passaging, some of the cells in a given batch can lose their differentiation ability and alter their gene expression.^{[149c, 151]} Culture media is another game-changing factor that can cause toxicity and changes in pH levels. The composition of culture media is still a controversial topic. For example, fetal animal sera are typical components of culture media, but they can cause immune reactions.^[149c] Using human blood products to prepare the media is possible; however, the media may be affected by the state of the donor’s body.^[152] Sterilization to avoid contamination is another element that is essential in all steps from cell isolation to bioprinting. Due to the nature of the process and the number of steps needed before starting the bioprinting process, maintaining sterility is challenging. Many sterilization methods are available, but the effect of such techniques on cells is uncertain. Sterilization treatments can lead to insufficient growth of cells.^[153] Growth factors are commonly used to create cell-incorporated scaffolds. Nevertheless, their effect on cells is not clear,^[143] as growth factors can change the function of cells in an undesirable direction.^[154]

Regulation has been investigated, but still there is no unique framework to refer to as a guideline for bioprinting,^[155] and various concerns need to be addressed in this regard. At first glance, it is hard to classify a 3D bioprinted tissue or organ for the initial FDA investigation (e.g., medical device or drug) as there are different pathways to follow depending on the selected category. Nonetheless, the FDA has not mentioned the pathway for 3D bioprinting and generally refers to the need for additional manufacturing process considerations. Often, the approval of the Center for Biologics Evaluation and Research is needed as part of the FDA pathway, yet there are no specific provisions. Commonly, the regulation road map is as follows: the laboratory-based evaluation begins, followed by preclinical testing, then clinical trials in different phases are started, and after approval, post-marketing evaluation begins.

8.1.1. Worldwide regulations on 3D bioprinting

3D bioprinting is cutting-edge technology and different rules apply in different regions. In Europe, bioprinting is regulated as part of the tissue engineering provisions.^[156] In Asia, religious matters delay the approval process.^[157] There are no regulations for bioprinting worldwide and even the FDA is conducting more research on this topic to identify factors affecting safety and quality.^{[144, 158]} In Europe, 3D printing regulations are complicated and governed by the European Medical Devices Directive, the Active Implantable Medical Devices Directive, and the In Vitro Diagnostic Medical Devices Directive;^[159] the regulations depend on the type of device, 3D printing technology, software, and biomaterial. Australia has taken the initiative and updated their regulations by using the international definitions defined by the International Medical Device Regulations Forum, allowing low-risk devices to be mass-produced without any certificate, and regulating bioprinted tissues and organs as medical devices (not falling under the biologics category).^[160] Knowing this, different guidelines have been enforced in different parts of the world, and therefore the need for a global pathway is critical.

8.1.2. Regulations in Canada

Medical device regulations are applied to 3D bioprinting in Canada.^[161] Health Canada reported the development of a guideline on how to fabricate 3D printed medical devices in 2018,^[162] and released a draft of this road map. This guidance clearly indicates the repeatability of the machine should be validated along with recording the machine maintenance and biomaterial. However, it does not include information about software, customized devices, and cell-incorporated tissues and organs.^[161]

8.1.3. Road map for clinical translation

In 2017, a guideline called “Technical Considerations for Additive Manufactured Devices” was published by the FDA to cover design, manufacturing considerations, and device testing matters.^[163] Due to inherent complexities, this guideline does not address all the regulatory necessities such as the use of cells or biologic products. As mentioned, the category of a bioprinted tissue or organ is not clear, and in addition to FDA instructions, engaging with the Center for Biologics Evaluation and Research (CBER) and the Center for Devices and Radiological Health (CDRH) is recommended. The biomaterial, design, bioprinting process, post-printing validation, characteristics/parameters of bioprinting, and mechanical/biological assessment are core elements that need to be well thought out. According to FDA guidelines, medical devices are categorized into groups of low (Class I), medium (Class II), and high (Class III) risk as per safety and effectiveness. For additive manufactured medical devices, there is a guideline, which includes design, software, and biomaterial fulfillments.^[163]

According to the guideline for additive manufacturing of medical devices,^[163] to fulfill the quality system requirements for different classes, specific design requirements must be met; these include 21 CFR 820.30 Design Controls and established procedures to monitor process parameters.^[164] To start, each step should be characterized; this means that the process parameters and the effect of every single parameter on others should be identified. This is not a simple task, owing to the effects of design, biomaterial, and fabrication-related parameters on each other.^[165] As such, identifying process parameters is critical because of the role this initial step has in finding the causes of failure during the final testing.

Considering the bioprinting process, file format conversion is an important part of the software workflow. Often, different software packages are used to do the design, meshing, rendering, and other related procedures, and every single software has its algorithm and coordinate system.^[158] Therefore, the geometrical dimensions are inevitably affected due to file conversion errors. To address this, standard protocols such as ISO/ASTM 52915 can be followed to have a robust configuration.^[163] Even after the design is ready, other steps are required to prepare it for printing in a layer-by-layer way; these steps include, but are not limited to, building volume placement (e.g., the orientation of filaments and packing density), adding support, slicing, and creating the build path as part of the software workflow.^[163]

Apart from the design and software workflow, it is important to take biomaterial into account. From a biomaterial perspective, additives and crosslinkers should be documented. The biomaterial itself should be identified in terms of supplier, common name, chemical name, and so on, in addition to referring to biomaterial standards and test methods. For example, polymers should be evaluated to find their composition, purity, molecular weight, and other specifications.^[163]

Overall, the whole process of 3D bioprinting should be validated. Sometimes, the quality is not the same for bioprinted organ/tissue using different machines with similar printing parameters, conditions, and biomaterials. Knowing this, it is very important to establish the relation between input parameters and processing steps from a part-quality perspective; this can happen using available protocols used for inspection.^[164] As part of the quality control, characterization methods are highlighted here. Biomaterial characterization is a common test, e.g., elastic modulus, strength (yield, ultimate), and viscoelasticity. While performing these tests, the effect of sterilization on the properties of biomaterials and cells should be considered. When performing mechanical testing, the anisotropic nature of the 3D bioprinting technique should be considered.^[163] The biomaterial may be crosslinked with various density gradients in different sections. Hence, it is essential to ensure that the biomaterial is fully cured. Especially for hydrogels, the water content should be investigated as the printing process may have an adverse effect on water uptake. Degradation tests should not be neglected. Biocompatibility is another critical factor to consider, and there is a protocol to biologically evaluate biocompatibility.^[163]

8.2. Challenges and limitations toward clinical application

It is vital to identify process parameters and obtain a reasonable understanding of 3D bioprinting machines.^[164] To gain this knowledge, test builds, processing evaluation and validation, and worst-case builds can be implemented. While manufacturing considerations were mentioned, these cannot be a comprehensive solution for the regulatory requirements, and it is essential to develop the process validation procedure further.^[164] To design a customized bioprinted tissue or organ for a specific patient, it is important to consider factors such as imaging and resolution, design manipulation software, data integrity (file conversion), and cybersecurity (personally identifiable information).^[164]

It should be noted that changing the machine or fabrication procedure causes deviations; as such, a revalidation is required.^[166] These changes can be as simple as changing the location of the machine or software changes and updates. The concept of repeatability is important because maintaining the consistency of bioprinted tissues and organs is challenging. Different FDA guidelines should be followed depending on the regulatory pathway.^[163]

Scientific gaps related to the bioprinting technique itself include structural integrity, mechanical stability, and pore interconnectivity.^[167] Other aspects are biological requirements such as cell loading, attachment, growth, and tissue formation. Additionally, manufacturability is a major factor, involving printability and fabrication-related matters. There is still a gap regarding the generation of new biomaterials. There is no robust protocol on either biomaterial formulation or living cell scale-up production.^[167]

One of the initial steps for clinical translation is producing cells commercially and cost-effectively in addition to complying with good manufacturing practice essentials. It is critical to establish guidelines for design and quality control of living cells and 3D bioprinted tissues and organs to address issues related to 3D bioprinting.^[167] Notably, it is different from traditional fabrication—because living cells are involved, the market potential has a limited lifetime. Cells might be damaged during the bioprinting process itself, which is another concern.^{[45, 168]} In terms of post-printing evaluation, various types of tools are required to access the function over time; there is no evaluation protocol to systematically access a 3D bioprinted tissue or organ.

Gaining knowledge on design for bioprinting is recommended; this includes investigating the effect of design and geometrical features on cell viability and function. Printability is a hot topic that maps the relationships between biomaterial, fabrication, and design. In addition, there is a demand for printing vascularized structures; thick tissues need a channel due to having living cells, and this is a challenge that needs to be addressed.^[167]

9. Conclusions and future perspectives

Over the past years, the effectiveness of tissue engineering to restore, maintain, or improve damaged tissues or organs has been demonstrated. The importance of cardiac tissue as well as the mortality related to CVDs highlights the attempts related to cardiac tissue engineering. However, the complexity of this specific tissue makes the formation of functional cardiac tissues quite challenging. The emergence of 3D bioprinting facilitated the design of viable cardiac constructs for various biomedical applications including tissue/cell transplantation, drug screening, and disease modeling. The fact that bioprinting allows the fabrication of sophisticated 3D structures with controlled layer deposition of materials and cells, enables this technique to create patient-specific microtissues with an unprecedented level of precision. Bioprinting can be categorized into several approaches based on the printing technique used. Although each method has its own pros and cons, they all allow the formation of high-resolution and complex bioprinted 3D structures. In addition, there is a vast variety of materials (polysaccharides, peptides, proteins, DNA, etc) that can be used, expanding the applicability and prominence of 3D bioprinting in tissue engineering and regenerative medicine.

Despite the promising results achieved in 3D bioprinted cardiac tissues, the field is still in its infancy. Further exploration into the following areas and addressing their issues are vital for reaching the ultimate goal of tissue engineering. First of all, considering bioinks, care should be taken to formulate bioinks that not only possess biological cues for cell attachment and proliferation, but also accommodate specific requirements, like electrical conductivity, that are necessary for proper specific cardiac tissue function. In addition, the mechanical properties of fabricated scaffolds need careful consideration if they are to match those of the native tissue. This could be achieved by formulation of new bioinks with more robust materials and/or modification of crosslinking density and procedures. Delivery of bioactive materials, including growth factors and pharmaceutical agents, is another aspect that needs to be considered more often in cardiac tissue engineering. 3D bioprinting can help to develop scaffolds containing these bioactive molecules with specially tailored temporal release profiles to significantly enhance tissue maturation, and cell differentiation. Vascularization is another critical aspect that needs more profound consideration when developing thick and large cardiac tissues. It is well-known that cardiomyocytes require various key nutrients and oxygens to function properly. Bioprinting methods such as sacrificial printing can significantly enhance the development of porous constructs with pre-defined or pre-formed vascular tubes inside. Another important challenge that needs careful attention is that following post-myocardial infarction, cells on the outer surface of the heart wall undergo epithelial-to-mesenchymal transition (EMT), ultimately forming a layer that acts as a barrier against therapeutics such as injected cells or implanted cardiac patches. The inclusion of growth factors and biomolecules (e.g., BMP-7) into 3D bioprinted constructs could be a potential solution to reverse the EMT process. Finally, translational regulations are still very controversial. Detailed and thorough regulations are critical to normalize procedures for generating bioprinted tissues and facilitating their clinical use. These regulations should be implemented by governmental agencies with the assistance of researchers.

Table 1.

Advantages and disadvantages of 3D bioprinting techniques.

Technique	Advantages	Disadvantages	Ref.
Extrusion	Fast printing speed A wide array of materials (e.g., hydrogels, cell aggregates, decellularized ECM) can be printed Can extrude materials with high cell density Good deposition control	Relatively low resolution (≈ 100 μm) Lower (than inkjet) cell survival after printing Bioink must tolerate shear thinning Only highly viscous bioinks	[9][38–39, 41]
Laser-based	Nozzle-free technique: no clogging High resolution (≈ 20 μm) Excellent cell viability Wide range of biomaterial viscosity (1–400 mPa⋅s)	High cost Complex laser-based system	[38–39, 39c]
Inkjet	High resolution (around 20–100 μm) Patterned structures High printing speed Low cost	Hydrogels limited to low viscosity materials Limited cell density Nozzle clogging	[9, 39a, 39c, 46]
Stereolithography (SLA)	High resolution (down to 20 μm) Fast printing speed Low cost High cell viability under visible light	Complex system Potential cytotoxicity caused by the laser system under UV light	[38–39, 48]