Keywords

Introduction

Cardiovascular tissue bioprinting occupies a critical crossroads position between the fields of biomaterials engineering, cardiovascular biology, three-dimensional (3D) design and modeling, and biomanufacturing [1,2,3,4]. This complex area of research requires expertise from all these disciplines to provide a multidisciplinary approach that enables fabrication of functional and living tissues and organs, whether for basic science or translational research applications [5]. A major challenge that hampers this field is the lack of systematic characterization of the physical and chemical properties of hydrogel-based bioinks that are applicable to organ and tissue bioprinting [6,7,8]. Tailoring bioink properties to mimic the complex native tissue extracellular matrix (ECM) is of great importance and a slight divergence could result in pathological or loss of function manifests [9, 10].

Functional tissue bioprinting holds great promise to combine rationally designed biomaterials, functional cells, and macromolecules into 3D constructs that closely recapitulate the mechanical, structural, and functional microenvironment of native tissues [11]. With precise control over spatial arrangement of the cell-biomaterial architecture, 3D bioprinting can provide complex physiochemical and biological cues that are necessary for the maintenance and maturation of functional tissue analogues. To date, a range of bioprinting platforms have been used to create artificial tissue constructs, such as extrusion-based [12,13,14,15], droplet-based [16,17,18], and laser-based printing [19,20,21] (Fig. 4.1).

Fig. 4.1
figure 1

Major methods that are available to bioprint tissue analogues. (a) Extrusion-based bioprinter s use single- or multi-nozzle print heads, squeezing out the bioink using mechanical or pneumatic forces. (b) Droplet-based bioprinting uses thermal or piezoelectric forces to discharge droplets of bioinks. (c) Laser-based bioprinters use high-energy pulsed laser to eject bioink droplets from a donor layer onto the receiving substrate. (Adapted from [32])

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Recent advances in in vitro tissue development has made 3D bioprinting an attractive means for the next-generation regenerative medicine, specifically as a platform for tissue replacement to rescue failed organs in patient-specific therapies [6, 22,23,24]. Some of these applications include pancreatic tissue printing, to address loss of function in diabetes [25, 26], bioprinted kidney tissue analogues that can be used instead of, or in conjunction with, dialysis in kidney failure therapies [27, 28] and cardiac tissue constructs which can be bioprinted to repair the damaged heart tissue post injury (e.g., myocardial infarction), or in the case of congenital heart diseases [29,30,31].

This chapter explores the critical parameters of hydrogel-based bioinks that are necessary for their successful application in functional cardiovascular tissue engineering, as tissue analogues for disease modeling and drug screening in vitro, or as implantable tissue grafts to treat a range of congenital and acquired cardiovascular diseases. We will explore the biophysical, biochemical, and biological considerations for the candidate bioinks that enable 3D bioprinting of functional cardiovascular constructs.

Types of Bioinks

Bioink can be defined as a printable biomaterial, based on naturally occurring or synthetically derived polymers (hydrogels), that can recreate some aspects of native tissue ECM [6, 7, 14]. In addition to the hydrogel component, bioinks typically consist of cells and/or small molecules (e.g., growth or angiogenic factors) to enhance their bioactivity. Successful bioinks maintain (or promote) encapsulated cell survival, adhesion, proliferation, and function in vitro and ultimately in vivo. Some of the desirable features of bioinks include the ability to form stable filaments during printing and gentle crosslinking mechanisms, which allow for spatial control of hydrogel deposition while maintaining cell viability [7, 8, 15, 33]. Currently available bioink formulations are often based on existing hydrogel biomaterials, such as alginate [34, 35], fibrin [36, 37], hyaluronic acids [38, 39], and gelatin [13, 40,41,42]. These bioinks can be divided into several broad categories based on the specific crosslinking/solidification characteristics of their parent hydrogel (Table 4.1). Critical for a successful bioink, it should be able to incorporate and protect bioactive compounds within the formed hydrogel. This could enhance the bioink functional mimicry to native tissues and enhance the function of encapsulated cells [43].

Table 4.1 Classification of hydrogel bioinks based on crosslinking or solidification mechanism
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To achieve an optimal bioink formulation and successfully print functional tissues, the choice of specific bioprinting process and post-print tissue culture, maintenance, and maturation is critically important [6, 7, 14]. Further, hydrogel properties such as viscosity, crosslinking mechanism, stiffness, mass transfer properties (e.g., diffusion and permeability), and biodegradability must be taken into consideration, depending on the specific application [6, 8, 15, 19, 44]. To generate a biomimetic niche that can support tissue functionality and cell maturation, the chosen bioink would need to also allow for specific chemical modifications such as small molecule conjugation and ECM proteins immobilization within the 3D printed construct.

Cardiac Bioink Characteristics

Printability

Printing resolution is dependent on the volume of deposited layer. To maintain a fine balance between thin prints (high resolution) and cell viability, the bioink should generate relatively low shear stress levels under modest pressures [33, 45, 46]. Shape fidelity at high-resolution prints is critical for building up functional tissue analogues, particularly for organs that are highly vascularized and have complex tissue organization such as the heart [8, 15, 40, 41]. To maintain fidelity, bioinks should have low reflow rate during the printing process and facile crosslinking steps and culture conditions (Fig. 4.2). This requires the ability of printed construct to be self-supporting at the macroscale, ideally with little to no supporting materials. While some support bioinks might be required to maintain complex/hollow shapes during printing, they would have to be either incorporated as a functional component of the tissue analogue or allow for full removal post-printing (i.e., sacrificial inks such as pluronic). This is particularly important in bioprinting of cardiovascular tissues, considering the remarkably high blood vessel density in the tissue (about 160 capillaries per mm2 of myocardial tissue [47]). For cardiac tissue bioprinting, therefore, successful fabrication of self-standing and stable, hollow channels at diameters ranging from micrometers (capillaries) to centimeters (arteries) would be a major challenge. These perfusable vascular cardiac constructs can provide an invaluable platform for drug screening [27, 48, 49] and disease modeling studies [2, 27, 48, 50] in vitro, and a new generation of cardiac patch devices for in vivo regenerative therapies [38, 51, 52].

Fig. 4.2
figure 2

Schematic demonstration of the typical approach to assess bioink printability. 1: Initial screening of bioink formulations to (a) establish fiber versus droplet formation, and (b) successfully stack multiple layers without fusing between the layers. 2: Rheological evaluations are performed to determine (a) the flow initiation and yield stress properties, (b) degree of shear thinning, and (c) recovery from shear thinning after printing. (Adapted from [14])

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Post-print Processing

Post-print processes are required to achieve the adequate print fidelity, as bioinks often require crosslinking [7, 53, 54]. Introducing crosslinking agents via an additional liquid phase may be detrimental to the shape fidelity. Therefore, employing a crosslinking step via long-wave ultraviolet (UV) radiation [53, 55] or visible light [40, 56] exposure at appropriate durations and intensities would be beneficial. Other types of post-print processes have often been used. For instance, promising results have been shown for salt-based post-processing after the constructs were printed [57,58,59]. Regardless of the chosen process, the structural and chemical stability of the initial print are critical to maintain reproducibility of engineered tissues. One challenge with light-based crosslinking is, however, the possible cell damage due to UV light irradiation [60]. This can be mitigated by combining or replacing UV with other more cytocompatible post-processing methods. Some alternative processes include aerosolized salt solution spray [61], incubation at elevated temperatures (>30 °C) [62], and enzymatic reactions [63] (Table 4.1). Additionally, less reliance upon ionic-based crosslinking may mitigate the precipitation or salting-out of adjunct proteins [64].

A successful cardiac bioprint will rely on a fast-acting crosslinking reagent with negligible cytotoxicity, while capable of retaining high printing fidelities. UV-crosslinked hydrogels, such as methacrylate modified gelatin (gelMA), have been extensively used for cardiac tissue printing as these hydrogels can generate stable yet relatively soft matrices that mimic the physiomechanical and biochemical properties of the native heart tissue [5, 65]. GelMA crosslinking via UV light occurs at relatively short times and moderate intensities (~2 minutes and between 1 and 20 mW/cm2), which could help to avoid excessive cell damage and death [16, 65].

Printed Tissue Stability and Controlled Degradation

Swelling and contraction of hydrogels upon crosslinking and during tissue culture could be detrimental to the printed construct fidelity and cellular interactions/functions [66]. This can also alter the final mechanical properties of the bioink and skew its 3D arrangement [67, 68]. Controlled degradation and remodeling of bioprinted construct, along with secretion of ECM proteins by encapsulated cells, are critical for engineering a biomimetic tissue microenvironment [69]. However, printed tissue breakdown and remodeling, in an uncontrolled manner, can severely limit its translational applications, as they could cause implant detachment and failure in vivo [70, 71]. To alleviate these challenges, extensive research has focused on the development of new bioink formulations with tunable degradation profile, to enable cell-mediated tissue remodeling, while maintaining the construct integrity [72,73,74,75,76,77,78,79]. Particularly in the case of cardiac tissue constructs, a significant degree of matrix remodeling by bioprinted cells is necessary to achieve the intercellular connectivity and the remarkably high cell packing density of the native heart tissue [50, 80,81,82,83]. To that end, gelMA-based bioinks are specifically favorable as they are both biodegradable and chemically defined.

Mass Transfer Properties of Cardiac Bioinks

Incorporating functional vascularization is a critical aspect of tissue bioprinting to generate and maintain large (clinically relevant) tissue constructs [84,85,86,87,88,89,90,91,92,93]. This is particularly important for bioengineered cardiac constructs, considering the remarkably high vascular density in the native myocardial tissue (approximately one capillary per cardiomyocyte) [94, 95]. With passive perfusion alone, the maximum thickness of a viable 3D tissue construct is usually around 100–250 μm [96, 97] before vascularization is required. This distance can potentially be extended to about 600 μm in bioprinted constructs that can be prevascularized prior to cell seeding. If the construct’s pore size is large enough to allow for more effective passive diffusion, a similar effect can occur [98,99,100].

Tissue-Specific and Chemically Modifiable Bioinks

The ability to incorporate tissue- and cell type-specific materials into the bioink is critically important for achieving in vivo-like functionality. Modification chemistry can provide the small molecules and ECM factors that are specific to the tissue/organ microenvironment [73, 78, 101]. For this purpose, covalent conjugation, or similar levels of immobilization, would be an effective approach to recapitulate the specific tissue cues for bioprinted cells and initiate self-directed environmental remodeling. Keeping the modification chemistry and bioink preparation steps simple would also be significantly beneficial by cutting down on preparation time and equipment and material expenses and by enhancing batch-to-batch consistency [6]. Furthermore, the hydrogel bioink should be xeno-free or consist of entirely chemically defined components, to facilitate translational use in regenerative medicine and to enhance the reliability for use in in vitro assays, such as drug screening and disease modeling [3].

To generate a bioink that is supportive to cardiac cells and recapitulates the organ/tissue-specific niche, high-throughput analysis techniques, such as transcriptome analysis (RNA-Seq) and proteomics can be used to characterize the native cardiac tissue ECM. For instance, bone morphogenic proteins (BMP2/BMP4) and Wnt inhibitors (IWP2) are known to play key roles in generation of early cardiomyocytes in vitro [102,103,104,105,106,107,108,109,110]. Incorporating certain concentrations of these factors in the tissue generation pipeline (e.g., in cardiac bioink) may promote the regenerative capacity of printed constructs. Further, functionalizing the bioink with ECM proteins, such as cadherins, connexins, and collagen, can be used to promote cell attachment, migration, and remodeling [86, 111,112,113,114,115,116]. ECM proteins coupled with secreted small molecules such as tumor necrosis factor alpha (TNFα), interleukin (IL)-1, IL-6, transforming growth factor beta (TGFβ), angiotensin II, and endothelin 1 can also help promote tissue maturation and vascularization in cardiac constructs (Fig. 4.3) [117, 118].

Fig. 4.3
figure 3

The demand for effective vascularization in vitro increases with tissue construct size and complexity

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Tunable Mechanical Properties

Altering biomechanical characteristics of the bioinks can be achieved via initial or secondary crosslinking processes. Such modifications can provide the specific mechanical cues to encapsulated cells and promote desired cellular functionalities [14, 119]. The ability to independently tune chemical and mechanical properties of these hydrogels is critically important. For example, crosslinking of hydrogel matrices can tune their stiffness [120,121,122,123], while conjugation of various ECM proteins and small molecules can independently provide biochemical cues to the cells [7, 124]. Mechanical properties play a major role in successful application of bioprinted cardiac constructs, as these tissues require strictly regulated stiffness values to exhibit proper functionality both in vitro and in vivo (Fig. 4.4) [32, 82, 123, 125, 126]. The post-print crosslinking mechanisms that are used for most hydrogel bioinks allow for generating tissues with a relatively broad range of stiffness. Therefore, bioprinting technology holds a great potential for manufacturing a wide variety of functional tissues.

Fig. 4.4
figure 4

Range of ECM stiffness required for proper organ development and functionality of various organs and tissues. (Adapted from [127])

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It has been shown by different groups that cardiomyocytes exhibit maximal contractile function on matrix stiffnesses ranging from 1 to 16 kPa [51, 82, 115, 123, 126]. Thus, an optimal cardiac-specific bioink may be expected to show elastic modulus within this range (Fig. 4.5) [126]. Incorporating large numbers of nondividing cardiomyocytes in bioinks can compromise their mechanical properties and printing fidelity (resolution and stability). To address this issue, cardiomyocytes can be combined with proliferative cardiac precursor/stem cells, cardiac, and vascular cells (endothelial and smooth muscle cells) and be encapsulated in the bioink at remarkably lower cell densities [29, 128,129,130]. Multiplication, differentiation, and maturation of these multilineage cardiac cell populations in bioprinted constructs, in association with controlled matrix degradation and remodeling, can lead to generation of myocardial mimetic tissue constructs at appropriate cell density and configuration. The use of stem cell sources for cardiac tissue printing may require additional examination and characterization to avoid incomplete or undesirable cell differentiation (e.g., tumor formation or reduced functionality of the engrafted tissue) that could impact the clinical application of the printed constructs [131,132,133].

Fig. 4.5
figure 5

Demonstration of varying range of stiffness for various cardiac tissues and tissue engineering hydrogels. Optimal range of stiffness, resulting in maximum contractile work of cardiac myocytes, is highlighted in blue. (Adapted from [126])

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Outlook and Conclusions

To maintain cell viability and functionality, hydrogel-based bioinks must fulfill several key biophysical and biochemical requirements, before, during, and post-printing processes. These parameters, together with the need for a functional vascular network in the construct, are critical for generating high fidelity cardiac tissue analogues. Initial and/or secondary crosslinking processes would allow for better control over the chemical and mechanical cues within the 3D constructs and therefore, enable reconstructing diverse tissue microenvironments. Preparing commercially available cardiac bioink kits that can be optimized for a specific tissue bioprinting would be a significant advancement in the field, especially if different chemical and mechanical properties of the hydrogels could be decoupled and independently tuned. A balance must be obtained between cardiac bioink crosslinking degree, stiffness, and biodegradation to allow for bioprinted cells to remodel their microenvironment. This is a critical step toward achieving enhanced cardiac cell connectivity, maturation, and function. Additionally, keeping bioink synthesis and modification chemistry robust and simple would be highly beneficial for wider appeal to researchers in the field.

In summary, cardiac bioprinting aims to generate clinically applicable, cardiac tissue analogues that can replace damaged/diseased tissue in vivo or be used as biomimetic platforms in vitro to model various diseases. Recent advances in bioprinting technologies have enabled fabrication of complex, patient-specific, tissue architectures at an organ-relevant spatial resolution, while supporting viability and function of multiple cell types. However, there remain some challenges for the clinical application of bioprinted cardiac constructs. Development of new cardiac-specific bioinks, using tailored biomaterials and precisely tuned selection of macromolecules, could be a great step forward toward clinical bioprinting. New methods are also needed to incorporate functional, multiscale vascular networks within printed constructs that can be perfused to maintain functionality of large-scale tissue constructs. Further, enhanced temporal and spatial resolutions in the new generation of bioprinters can help engineering more advanced cardiac tissue substitutes for regenerative medicine.