3D Bioprinting in Cardiovascular Disease: Current Status and Future Directions

1. Introduction

Cardiovascular disease (CVD) is the leading cause of death and major contributor to disability worldwide, with a prevalence of 523 million people who suffer from CVD and an estimated 18.6 million deaths annually [1]. Clinical diagnosis and management of CVD largely relies on less invasive imaging modalities, such as use of computed tomography (CT), magnetic resonance imaging and echocardiography [2,3]. Despite being frequently used in daily practice, it is challenging to translate two-dimensional images of cardiovascular anatomy and pathology to three-dimensional (3D) structures based on these imaging modalities and this is especially obvious when visualising complex cardiovascular structures and pathologies, such as congenital heart disease which usually involves multiple cardiac structures with a spectrum of abnormalities. Three-dimensional (3D) printing technology has overcome these limitations by providing realistic models, thus offering superior advantages over current image visualisations [4,5,6,7,8,9,10].

Use of 3D printing in cardiovascular disease has been increasingly reported in the literature with applications mainly focusing on these areas: education value for medical students and health professionals, preoperative planning and simulation of complex cardiovascular procedures, development of medical devices for improvement of outcomes, optimising CT protocols using 3D phantoms to minimise radiation exposure to patients during routine CT scans for diagnosis and follow up, bioprinting of scaffolds or cardiovascular constructs [11,12,13,14,15,16,17,18]. Figure 1 summarises these current applications. 3D printed models serve as valuable tools for medical education by enhancing understanding of anatomy and pathology compared to current education approaches. Nearly 50% of the current applications of 3D printing lie in pre-surgical planning of cardiovascular procedures, with results from single and multi-centre studies showing changes in surgical decision-making with the use of 3D printed models [7,14,19,20,21,22,23,24,25,26,27]. 3D printed models are also used to simulate interventional cardiology or radiology procedures, such as endovascular aneurysm repair, increasing confidence through training and practice [28,29,30,31,32,33,34,35,36,37,38,39,40].

While there is sufficient evidence available to prove that 3D printed personalised heart and vascular models play an important role in clinical applications, 3D bioprinting represents a promising field that could revolutionise the future of treating cardiovascular disease. 3D bioprinting uses mostly the same additive manufacturing technologies but prints functional living structures from biological components such as living cells, biomaterials and growth factors [41,42]. This could eliminate the gap between the current animal study and human trials, and may in future contribute to repair or replace damage tissues or organ and optimise drug delivery techniques [43,44,45,46,47]. Although still in an early stage with many challenges to overcome, substantial progress has been made over the last decade in printing cardiovascular constructs as well ascardiovascular regeneration [48,49,50].

This review provides an overview of 3D bioprinting in cardiovascular disease, with a focus on the applications in cardiac tissues, vascular constructs and grafts, myocardium and heart valves.

2. 3D bioprinting: where are we now?

Heart transplant remains the gold standard treatment for selected people with end-stage heart failure [48]. However, the number of people requiring a heart transplant far outweighs the available donor hearts and the long-term outcome of heart transplantation remains uncertain, with acute rejection, cardiac allograft vasculopathy, malignancies, infection and the development of chronic kidney disease still significant complications [49,50].

Apart from organ transplantation, current treatment options for cardiac disease include cell therapy, implanted devices (stents or stent grafts), and bypass grafting [51,52,53]. Of these treatments, cell therapy using 3D bioprinting with patient-specific cells to repair damaged cardiac tissues or replace end-stage organ failure through tissue engineering is promising [48,54], and is theoretically able to bypass the limitations of poor biocompatibility, biofunctionality, immune rejection as well as the drastic shortage of organ donors [55,56,57,58].

Although recent advances in bioprinting methodologies and technologies have addressed limitations associated with bioprinting, it is still not clinically feasible to directly translate 3D printed cardiovascular tissues to patient therapy. The main limitation is the inability to construct thick tissues. Currently, only 1 mm tissues can be produced with incorporation of multiple cell types and induction of tissue differentiation of growth factors [59,60]. A further limitation is a lower density of cells compated to native tissues restricting the clinical value such as in myocardial ischemia [61].

3. 3D bioprinting-bioinks

3D bioprinting involves deposition of living cells, bioactive molecules and biomaterials in a layered pattern to allow generation of 3D structures [62,63]. Different from the standard 3D printing process which uses various printing materials from rigid to soft and elastic depending on usefulness or application, bioinks serve as a medium to deliver living cells [64]. Bioinks possess viscoelastic properties with higher water content to protect the cells during the printing procedure from external risk factors. Bioinks used for cardiac tissue engineering mainly refer to the use of hydrogels which mimic the 3D extracellular matrix, with its: biocompatibility, biodegradability and mechanical support [65].

4. 3D bioprinting technologies

Bioprinting methods include inkjet-based, extrusion-based, and light-based printing to create geometrically complex and scalable tissues.

4.1. Inkjet-based bioprinting

Inkjet bioprinters store bioink in a cartridge and printer head; printer heads are deformed by thermal, piezoelectric, electrostatic or electrohydrodynamic actuators to eject droplets in a controlled fashion [63]. This is a low-cost method with high printing speed [64] and around 80–90% for fibroblast cell viability [65,66], but is limited by a low range of printable viscosities (<30 mPa.s) [67].

4.2. Extrusion-based bioprinting

Extrusion-based bioprinting is the most common technique used. Bioink is loaded into a syringe, then printed as continuous strands using mechanical or pneumatic forces [68]. This can print biomaterials with a wide range of viscosities (up to >6 x10^7 mPa/s) [69] and a high level of structural integrity [70]. However, cell viability decreases with increasing pressure and/or decreasing nozzle diameter [71]. Although extrusion-based bioprinting is a convenient and inexpensive platform with reasonable costs, it has limited resolution which is inferior to other bioprinting techniques.

4.3. Light-based bioprinting

Laser-assisted bioprinting projects a laser beam onto a thin layer of bioink causing a precise droplet to fall onto the printing surface [72]. Xiong et al [73] successfully printed both straight and Y-shaped vascular constructs; however, fibroblast cell viability was 68.1% and 70.8% for these constructs, respectively. The main advantages of laser-based bioprinting include high resolution, the ability to produce cardiac constructes with a high cell density (up to 10⁸ cells per mL), and bioprinting with low viscosity bioinks. The main disadvantages are complexity in controlling laser pulses, challenges in fabricating cell-embedding hydrogel constructs before printing and high cost of the laser system.

5. 3D bioprinting cardiac tissues

5.1. Human pluripotent stem cells and cardiac tissue engineering

Recent advancements in bioprinting of the cardiovascular system includes areas involving human pluripotent stem cells (hPSCs) and cardiac tissue engineering [74]. Promising benefits have been shown in using current hPSCs and cardiac tissue engineering technologies to construct biocompatible materials for the repair and regeneration of diseased cardiac tissue. Traditionally, cardiomyocytes are among those that are known to be difficult to expand and culture. As reprogramming of differentiated human cells back into a pluripotent state has been made possible in recent decades, the potential for creating patient and disease-specific stem cells in nearly any patient is becoming a possibility for the future [75,76]. Additionally, hPSCs technologies show great potential over conventional 2D monolayer culture systems for studying the development of the human heart, the heart’s response to therapeutic interventions in drug testing, and for disease modelling of acquired and inherited diseases [74,77,78]. Figure 2 provides a schematic view of the application of cardiac tissue engineering for disease modeling and drug screening in genetic heart diseases. Cardiovascular tissue engineering is a rapidly evolving field of medicine which involves multiple sophisticated components such as different cell types, biocompatible scaffold materials, cell differentiation factors and growth factors [79]. As this is an area of research that is still in its infancy, there are several issues which are yet to be solved.

5.2. Cellular maturity

hPSCs such as human-embryonic stem cells (hESCs) and human-induced pluripotent stem cells (hiPSCs) are promising approaches for heart regeneration and have the potential to be differentiated efficiently into cardiomyocytes, vascular endothelial cells, and vascular smooth muscle cells [74]. Although there is much international research in this space, there are concerns over the immature properties of the tissues created [75]. The maturity of current hPSC-derived cardiomyocytes resemble characteristics of fetal cardiomyocytes [80].

To promote maturation, many approaches have been proposed, including the use of biochemical stimuli (such as hormonal treatments), mechanical strain stimulation, electrical stimulation, topographic cues such as patterned surfaces and substrate stiffness, and interactions with other cells or extracellular matrix [77,78,81]. However, the optimal state of cellular maturity for hPSC cardiomyocytes is yet to be determined; immature cardiomyocytes may cause ventricular arrhythmias and other electrical disturbances, but have increased survival compared to adult cardiomyocytes upon transplantation into a host myocardium [74].

5.3. Microvasculature constructs

Creating functional microvasculature constructs is a challenging area in 3D bioprinting of complex tissues, such as the heart which is highly vascularised to allow for adequate gas exchange, nutrient diffusion and waste disposal [82]. As diffusion of oxygen through tissues is limited to approximately 100-200 µm, careful incorporation of microvasculature into bioprinted materials is a fundamental step for tissue survival [82]. Characteristics which are fundamental for proper microvascular perfusion to tissues include the presence of a hollow and endothelialised lumen, a hierarchy-based branched vascular network, and a complex signalling milieu [83]. Many biofabrication technologies such as extrusion-based 3D bioprinting, induced sprouting, electrospinning, and lithographic approaches have been explored to address these issues, each with their own advantages and disadvantages [83,84]. These techniques generally involve microfluidics-based molding with numerous other strategies such as the addition of proangiogenic factors (e.g vascular endothelial growth factor [VEGF]) within the biomaterial scaffolding [85]. Figure 3 is a schematic framework proposed by Seymour et al. to address the current issues in biofabricating microvasculature [83].

Despite rapid improvement in mimicking the vascular networks of the human cardiovascular system, there is yet to develop an approach which is able to address all the required characteristics of microvasculature in order to maintain perfusion and survival of tissues. Precisely, the spatiotemporal precision required for appropriate physiological function is still lacking [85]. Further exploration of biomaterials and bioprinting strategies for addressing the structural and physiological complexity of microvascular networks will be a key stepping-stone in the research field of 3D bioprinting of cardiovascular tissues.

5.4. Other issues

Other issues that are faced in 3D bioprinting of cardiovascular tissue include but are not limited to cellular heterogeneity in different anatomical regions of the heart [74], proliferation of primary cells in vitro entering senescence [81], foreign body reaction [86], and the risk of tumours such as teratomas in undifferentiated cell populations [87].

Although we are still far from a safe and functional 3D bioprinted cardiovascular construct, this field of research has been developing and improving rapidly. Future studies should focus on tackling the critical issues of cardiomyocyte maturation, improved microvasculature constructs as well as creating a safe microenvironment for cell culture and biocompatibility.

6. 3D bioprinting of vascular constructs and grafts

Vascular grafts are vital in surgical treatment of coronary artery disease and peripheral vascular disease; arterio-venous fistulae in haemodialysis; and repair of large-vessel aneurysms and congenital defects [88]. Autologous grafts remain the gold standard; however, 30% of patients lack suitable grafts [89,90]. Synthetic grafts suffer significantly lower patency than autologous equivalents [91] and are limited in <5mm-vessel repairs by high incidence of thrombosis [90]. 3D bioprinting to create tissue-engineered vascular grafts (TEVGs) offers a promising future alternative to produce individualised grafts.

6.1. Requirements of a TEVG

TEVGs must be able to generate functional endothelium; be non-immunogenic; and possess similar mechanical properties to human vessels [92]. Vascular grafts should be tested for several mechanical properties. Burst pressure is the maximum pressure a vessel can withstand, and is measured with burst pressure or burst probe testing [93]. Vessels must be compliant in order to accommodate pulsatile blood flow. Suture-ability of grafts is measured with suture retention testing, the maximum force required to cause the suture or graft wall to fail [94]. TEVGs also require an intact layer of endothelial cells (ECs), tested using cell viability studies [95].

6.2. Methods of 3D bioprinting of TEVGs

Biological grafts require complex structures, including branched and cellular vessels. Inkjet-based printing is either ‘vertical’ or ‘horizontal’: if the nozzle’s primary direction of movement is parallel to the vessel’s circumference, it is vertical; in horizontal printing, the primary direction of movement is parallel to the vessel’s longitudinal axis [96]. Christensen et al [65] successfully combined vertical and horizontal inkjet bioprinting to create vessels with multi-planar bifurcations. Fibroblast cell viability was >90% after 24-hour incubation. Cellular tubes (‘zigzag’ tubes), have also been successfully produced using inkjet techniques, with a cell viability of >82% after a 72-hour incubation [97].

Despite these advances in inkjet bioprinting, its use for TEVGs has been limited. Inkjet printer head orifices can be clogged by bioinks of higher viscosity; the use of lower viscosity bioinks is unfavourable for mechanical strength [98], a quality essential for vascular grafts.

Hinton et al [99] used extrusion-based bioprinting with a ‘freeform reversible embedding of suspended hydrogels’ (FRESH) support bath technique to create part of the right coronary artery, with accurate anatomical and mechanical fidelity to the 3D model. The support bath behaved as a rigid body at low shear stresses, but as a fluid at higher shear stresses, allowing for low resistance to the printer nozzle and high resistance to movement of the printed construct (Figure 4).

Co-axial extrusion bioprinting utilises multiple distinct nozzles arranged coaxially, in order to print several concentric layers of biomaterial [100]. Hong et al [101] bioprinted vascular-like structures using a rapid-gelling gelatin-PEG-tyramine polymer with a dual coaxial nozzle. Viability of printed cells (fibroblast and endothelial cells) was 80–90%, and patency was confirmed with flushing of trypan-blue solution through the tube.

Composite bioinks can improve stability [102]; alginate and carboxymethylcellulose were combined by Milojević et al [103] to print vessels with good biocompatibility and sufficient mechanical stability (Figure 5). Wang et al [104] and Liu et al [105] (Figure 6) added gelatin methacryloyl (GelMA) to alginate, producing vessels with enhanced EC proliferation, adhesion, and spreading; the alginate sheath provided mechanical strength. Jia et al [106] used a novel bioink of GelMA, sodium alginate, and 4-arm poly(ethylene glycol)-tetra-acrylate. The combination of three bioinks stabilised the vessel and supported the maturation of vascular cells; their triaxial printing technique was able to achieve a wide range of diameters and wall thicknesses; and the printed vessels maintained high fidelity to their models, even when printing tortuous structures.

Mandrel-based printing uses a rotating rod as the support structure for the surrounding printed vessel. Gao et al [107] used coaxial nozzles to print hollow alginate filaments in a spiral configuration along the rotating support rod, creating a macro-channel (equivalent to the vessel lumen) and microchannels within the vessel wall. This approach produced biocompatible blood vessels with relatively strong mechanical strength. In a recent study, Jin et al [108] used a novel approach: nanofibre electrospinning to create an inner EC tube, and mandrel-based extrusion printing to create an outer layer of smooth muscle (SM) cells (Figure 7). This process yielded >90% cell viability, tensile strain sufficient to withstand blood pulsation, and suture retention strength greater than the common carotid artery. The combination of biocompatibility and mechanical strength render these processes promising for clinical application.

Xu et al [109] created a vascular construct with both good biocompatibility and mechanical strength using a novel integrated tissue-organ printer [110]. Two syringes of bioink were used, one for the inner layer of ECs and the other for the outer SM cells. The porosity of the GelMA bioink was varied, creating different-sized pores for each layer, essential for the different sizes of endothelial and SM cells. Gold et al [111] disposed with a supporting sacrificial structure, and used direct extrusion printing to fabricate vessels with a novel bioink: nanoengineered extracellular matrix (nECM). Their nECM comprised GelMA, poly (ethylene glycol) diacrylate, and two dimensions nanosilicates. This nECM had high printability and fidelity, cell viability >80%, and mimicked the thromboinflammatory response of native vessels.

Stereolithographic bioprinting uses patterned UV light to photocrosslink entire layers of bioink at once. In contrast to extrusion or inkjet printing, it does not print in points; as such, it produces larger-scale vascular structures more easily [112]. Krishnamoorthy et al [113] used this layer-by-layer approach to print vascular constructs using fibroblasts encapsulated within GelMA bioink. Cell viability after 48-hour incubation was 80%. They identified an ideal cure depth (photocrosslinking thickness of an individual layer) range of 200 µm for this printing process, which gave the best shape fidelity and resolution. Thomas et al [114] trialled multiple photoink formulations to find the ideal sacrificial photoink in which to embed ECs. 1.5% methacrylated hyaluronic acid (HAMA) was superior due to its rapid digestion time, resulting in a faster release of ECs. They created a sophisticated bifurcating vessel with sufficient EC proliferation (Figure 8).

6.3. Summary

3D bioprinting represents an important potential solution to the shortage of biocompabile vascular grafts. Multiple techniques have been trialled in vitro, with extrusion-based printing being the most common approach. These studies demonstrate excellent viability and proliferation of printed ECs. Relatively few studies have tested the mechanical properties of printed structures which need to be robust under arterial pressure and will inform in vivo trials [103,107,108,109].

7. 3D bioprinting of heart valves

Valvular heart disease is a serious global health problem, affecting 3.1% of the adult population and projected to claim 23.6 million lives by 2030 [115,117]. Of the four valves in the heart, the aortic valve bears the greatest hemodynamic load, pumping 3-5 liters per minute and undergoing 30-40 million cycles per year [117]. Degenerative aortic valve disease is the most common cause of valvular heart disease representing at least 65.2% of cases (aortic stenosis 47.2%, aortic regurgitation 18.0%) [115,118]. The current clinical solution for a diseased valve is replacement with a prosthetic one which can be classified as either mechanical (made from metal or carbon) or bioprosthetic (made from chemically fixed animal cardiac tissue). These prosthetic valves have transformed the treatment of valvular heart disease, however, are not without their problems. Although mechanical valves are the most durable they require lifelong anticoagulation therapy [119,120]. Conversely, bioprosthetic valves do not require anticoagulation therapy but suffer from limited durability [122,123]. Furthermore, both types of prosthetic valves require replacement (mechanical valves ~20 years, bioprosthetic valves as short as 5 years in younger patients) which is of growing concern as the population ages, and as younger patients are being treated with the adoption of lower-risk procedures such as transcatheter valve implantation [116,124].

Tissue engineered heart valves (TEHVs) present a possible solution to this problem, with the potential to provide a valve that does not require anticoagulation therapy and could last a lifetime [125,126,127]. However, a successful TEHV has complex requirements including the ability to encourage cells to infiltrate, differentiate and proliferate in a highly specific manner, as well as exhibit adequate mechanical and haemodynamic functionality [128,129,130]. These, along with complex regulatory approvals and patient to patient variability means that TEHVs are yet to reach the clinic.

Bioprinting is one possible avenue to creating a successful TEHV, with the ability to construct patient-specific, hierarchical, cell-laden constructs that could also match the complex microstructure of a native valve. A handful of attempts have been made using extrusion-based, light-based and bioplotting bioprinting techniques and are described below.

7.1. Extrusion based TEHVs

Extrusion-based printing can rapidly fabricate anatomically correct heterogenous TEHVs using two types of cell-free poly-ethylene glycol-diacrylate (PEGDA) which are able to support porcine aortic valve interstitial cells (PAVICs) after 21 days [131]. Furhtermore, the same researchers incorporated SM cells and valve interstitial cells (VICs) into anatomically correct extrusion bioprinted valve roots and leaflets made from alginate and gelatin (Figure 9A) [132]. Both SMCs and VICs were viable (>80%) after 7 days and expressed elevated alpha-smooth muscle actin (aSMA) and vimentin respectively. Cell-laden constructs initially exhibited weaker ultimate tensile strength, failure strain and Young’s modulus than their cell-free counterparts, although over 7 days the cell-laden valves maintained their mechanical strength as opposed to cell-free constructs which weakened. Subsequently, human aortic valvular interstitial cells (HAVIC) were extrusion bioprinted into a simplified heart valve geometry (Figure 9B) [133]. Different ratios of methacrylated hyaluronic acid (Me-HA) and methacrylated gelatin (Me-Gel) impacted cell spreading, mechanical properties, and printability of constructs. After 3 days in a static culture tube, the encapsulated cells below the surface started to remodel the hydrogel by depositing collagen and glyosaminoglycans (GAGs).

These initial studies demonstrated the feasibility of extrusion bioprinting for the creation of TEHVs, however did not address concerns regarding the extreme mechanical requirements. Accordingly, there have been multiple investigations to reinforce the mechanical integrity of cell-laden hydrogels. These include the incorporation of nanoparticles such as nanocrystalline cellulose [134], variations in photocrosslinking methods [135], and reinforcement with electrospun polyacrylonitrile (PAN) fibers [136]. Despite demonstrating the ability to tailor mechanical properties, none of these studies have shown mechanical properties sufficient to match those of native heart valves.

7.2. Light-based TEHVs

Stereolithography printing was the first technique to successfully 3D print a TEHV, albeit cell-free [137]. However, since then there has only been one attempt at light-based bioprinting which used PEGDA:GelMA seeded with cardiac fibroblasts [138]. The valve showed 80% cell viability after 7 days although was not tested for haemodynamic or mechanical properties. Since then there have been no further attempts at light-based bioprinted TEHV, perhaps due to the lack of material heterogeneity achievable with the method, or more likely due to the relatively low mechanical strength.

7.2. Bioplotted TEHVs

More recently, TEHVs have been fabricated using a novel bioplotting method called freeform reversible embedding of suspended hydrogels (FRESH). Notably, this method has been used to fabricate an entire neo-natal scale human heart exhibiting multiple anatomical structures such as the ventricular chambers and aortic valves (Figure 9D) [139]. In the same study, tri-leaflet valves were fabricated using a cell-free collagen based ink. While this valve strictly speaking, did not contain cells, the same technique was used to print cell-laden vessels with diameters of 8-50 µm. It must be noted however, that prior to mechanical testing the valves were decellularized to enhance mechanical properties. The valve was tested in a physiological flow loop setting as per ISO 5840 and demonstrated adequate total regurgitation fraction, although insufficiently high effective orifice area and transvalvular pressure gradient for the aortic setting.

The same technique has been used to evaluate the recellularization potential of simplified cell-laden TEHVs in a subcutaneous rat model (Figure 9C) [140]. Rat mesenchymal stem cells (rMSCs) were printed using a bioink of highly concentrated Type I collagen hydrogels (Lifeink 200) and implanted for 2, 4, 8, and 12 weeks to assess cell infiltration, inflammation, and uni-axial tensile mechanical properties. The scaffolds expressed aSMA and vimentin biomarkers, demonstrating an acceptable inflammatory response although there was limited cellular infiltration into the scaffold. Tensile mechanical properties varied throughout the duration of the 12 weeks potentially showing variations of scaffold resorption and remodelling, although statistically insignificant. Following this, a similar study was conducted accompanied by finite element computational analysis in an attempt to characterise the remodelling of mechanical properties [141]. The study concluded that subcutaneous culture was insufficient for replicating native tissue mechanical properties and a dynamic cellularization environment such as a bioreactor would be necessary.

Thus, despite the promise shown in printing of heterogenous cell-laden valves with patient-specific geometries, bioprinting of purely bioink- based TEHVs have insufficient mechanical properties. Novel methods of mechanical reinforcement such as incorporation of nanoparticles [142,143], new methods of physical or chemical crosslinking [144], and soft network composites may help [145,146].

8. 3D bioprinting of myocardium and heart

Injury to the heart in the form of myocardial infarction (MI) or heart failure, can be treated with medication, angioplasty and stenting, bypass surgery and lifestyle changes, however, the only remaining option for end stage disease is a heart transplant. With the well-known drawbacks that transplants entail, including donor shortage and procedure-associated risks, tissue engineering and regenerative medicine (TE&RM) approaches have been gaining ground as potential solutions.

Current views in TE&RM support the notion that the best way to achieve adequate engineered-tissue capabilities is by trying to replicate native composition, structure and properties as closely as possible [147,148]. Bioprinting technologies, with their capability to pattern cells and biomaterials with great spatial resolution, are standing out as an unparalleled tool for myocardial TE&RM and will be examined in the following paragraphs and also summarised in Table 1.

Many myocardial bioprinting studies focus on optimisation of printing conditions, bioink formulation and/or functionalisation to incorporate additional benefits. Hence, one research group formulated a bioink based on cardiac extracellular matrix (ECM) [149], while another incorporated gold nanorods (GNRs) under the hypothesis that it would beneficially contribute to the electrical conduction of the construct [150]. For their part, Erdem and co-workers attempted to create an oxygen-releasing bioink by adding calcium peroxide to their GelMA-based formulation, which improved survival and function of the construct under hypoxic conditions [151]. All reported good feasibility and cell viability, together with the desired effect albeit to a limited degree.

More recently, Ahrens et al focused on imparting cardiomyocyte (CM) alignment in their bioprinted cardiac constructs. The particularity of this approach relies on the fact that the cellular component of the bioink is not dissociated cells, but rather preassembled aggregates of tens of thousands of cardiac cells, denominated anisotropic organ building blocks (OOBs), and generated by seeding and culturing a mixture of CMs derived from iPSCs, human neonatal dermal fibroblasts and a collagen hydrogel in a micropillar array. This resulted in highly aligned, densely cellularized (+200·10⁶ cells/ml) cardiac struts and sheets.

The anisotropy of the bioprinted constructs came from boththe auxotonic mechanical conditioning of the OOBs by the micropillars, and the shear stress direction imparted by the extrusion through the nozzle. The cardiac constructs contracted synchronously and with greater force than spheroid-based controls and could sustain pacing up to 3 Hz [152].

Similarly, Ong et al also chose to print with cell aggregates (spheroids) rather than cell suspensions, but they did not add any supporting hydrogel under the assumption that it may hinder cell-cell interactions. Their constructs beat spontaneously and synchronously, and responded to electrical pacing. When varying the ratio between CMs, ECs and cardiac fibroblasts, results suggested that the latter can delay and even block electrical conduction if present in too high quantity [153]. When implanted in a rat MI model, the cardiac patch resulted in a smaller scar area and greater vascularization than the control (omental patch), as well as better functional output (ejection fraction and cardiac output), however these were not statistically significant [154].

Interestingly, bioprinting of cells can be complemented by the simultaneous deposition of other materials that will perform a specific function. For instance, one study incorporated a polycaprolactone (PCL) frame around the bioprinted construct to subject it to auxotonic stress during culture, as well as micro-springs that enabled measurement of the patches’ contractile force. After 4 weeks in culture, samples showed the presence of aligned and dense sarcomeres, as well as connexin 43. Moreover, the cardiac constructs displayed physiologic response when subjected to drugs known to alter contraction force and frequency, namely epinephrine and carbachol [44].

In another example, Asulin and co-workers combined the bioprinting of three distinct components to create a cardiac construct with integrated electronics allowing for both sensing of the patch’s activity and its pacing. One bioink consisted of a mixture of CMs and ECM, while the other two material inks were made of polydimethylsiloxane (PDMS) and either graphite flakes for conduction, or surfactant for passivation (passivation is essential for achieving accurate spatial sensing and stimulation while reducing noise) (Figure 10) [155].

In a major step towards implant personalization, Tal Dvir and team used patient omental biopsies to create patient-specific bioinks [156]. These consisted of a mixture of the patient’s own cells, reprogrammed into iPSCs and later differentiated into the desired phenotypes (endothelial and CMs), and the patient’s extracellular matrix, turned into a thermoresponsive collagen hydrogel. Using these they bioprinted thick (2 mm) vascularized cardiac patches, made of CMs and perfusable channels 300 µm in diameter, lined with endothelial cells. Furthermore, they sought to print constructs in a more architecturally relevant, larger size, for which they developed a support medium allowing bioink curing and construct extraction without damage to the cells. This allowed them to bioprint an anatomically accurate, miniaturized heart (20x18 mm) (Figure 11) [139]. The heart displayed two hollow chambers as ventricles, and was surrounded by perfusable channels modelled to imitate coronary arteries.

Similarly, Lee et al also utilized a support bath to print human CMs derived from embryonic stem cells and collagen into an ellipsoidal shell, which contracted in unison and could be paced up to 2 Hz (Figure 12) [157]. Finally, Kupfer and co-workers focused on bioprinting complex structures, this time adapting an MRI scan of the heart to contain an inlet and outlet channel through which unidirectional flow could be propagated. A key aspect of this work is that they chose to bioprint hiPSCs, in situ differentiating them to cardiac lineages after construct formation. This way the rupture of CM-to-CM connections for bioink generation is avoided, which the authors argue results in enhanced cell density and tissue connectivity [158].

Very recently, to tackle the issues of construct size and cell survival during and after bioprinting, a 6 degrees-of-freedom robotic arm was modified to serve as a bioprinter. This surpasses the capabilities of any cartesian printer, which only permits bottom-up manufacturing, and allows for printing in any direction. In this approach, tubular scaffolds in the manner of vascular conduits, through which culture medium is constantly circulated, act as a backbone onto which layers of endothelial cells and CMs are deposited. The process is iterative, consisting of rounds of bioprinting and culturing, to give time for cell attachment between the new layer and the previously deposited ones, and vascular network growth. The printing is performed in a chamber, filled either with an oil bath to ensure adhesion of the bioink to complex-shaped surfaces, or culture media. Although the authors did not demonstrate the printing of a full heart, this work provides proof of concept and shows clear potential to generate complex tissues and entire organs (Figure 13) [159].

9. Summary and concluding remarks

Recent developments in 3D printing technology have revolutionized medical practice by using personalised 3D printed heart and vascular models in enhancing diagnosis and treatment of cardiovascular disease. Technological advancement in 3D bioprinting has great potential and is a powerfil tool utilising tissue engineering principles to bioprint cardiac cells, vascular and tissue constructs. New advancements in 3D bioprinting have made it possible to construct organ-mimicking myocardial or cellular structures with high precision and flexibility. A variety of biomaterials in combination with various bioprinting methods have been shown to demonstrate great potential in 3D bioprinting cardiac constructs and grafts, myocardium and heart valves. Despite promising results and rapid progress further research and investigation is required to better understand the cross over between a native organ and amalgamation of bioprinted heterogeneous tissue constructs. Although mechanical properties are in some cases being reproduced, novel methods of mechanical reinforcement are essential to enable translation in vivo. Whilst the area of valve replacement is advancing rapidly, we see less exploration of bioprinting of vascular grafts. Endovascular minimally invasive treatments have rendered bypass surgery redundant in many patients, in both the coronary and peripheral circulations, however, bespoke patient derived bioprinted arteries may in fact sway the pendulum back from minimally invasive treatments by providing a more durable solution to occlusive arterial disease by virtue of biocomplatibilty, bespoke sizing and the reproduction of arterial mechanical properties.