Although there are a wide variety of bioink materials, including hydrogels, dECM, microcarriers, and cell aggregates, not all of the currently available bioprinting modalities are compatible with every bioink material. EBB has the greatest flexibility among existing bioprinting modalities, due to the extrusion mechanism as well as the larger nozzle diameters. EBB can facilitate bioprinting of hydrogels, cell aggregates (in spheroid, strand, and cell pellet form), dECM components, and microcarriers in bulk hydrogels. Other bioprinting modalities including DBB and LBB support bioprinting of hydrogels only. Since the nozzle orifice diameter in DBB is very small [<120 μm (Ozbolat and Hospodiuk, 2016)], bioprinting of cell aggregates and microcarriers is challenging as the nozzle tip can easily clog and does not support formation of droplets as previously discussed (Gudapati et al., 2016). Although not yet attempted, DBB of dECM is technically feasible assuming the dECM viscosity is low and does not generate a fibrous gel. In LBB, bioprinting of cell aggregates, dECM, and microcarriers has not been attempted; however, due to the large mass of microcarriers and cell aggregates, bioprinting with LBB will represent a technical challenge.

The bioprintability of hydrogels is superior to that of other bioink types. As long as the crosslinking mechanism is incorporated into the layer-by-layer fabrication scheme, bioprinting of hydrogels is relatively straightforward as other bioink types require additional processes and customized systems for the extrusion setup. For example, cell aggregates need to be loaded in a custom-made extrusion head. Tissue spheroids need to be loaded into a pipette unit and delivered in a medium (Tan et al., 2014). The act of loading cell aggregates into the pipette unit can be problematic as tissue spheroids can start to aggregate; this can result in nozzle clogging. In addition, it is difficult to attain the necessary close contact between spheroids postbioprinting, resulting in poor tissue formation. Extrusion of cell pellets, on the other hand, is relatively straightforward; however, an additional support structure (such as agarose mold) should be printed in tandem to support aggregation of cells. Tissue strands, on the other hand, can be bioprinted in solid form without the need for a molding structure or a delivery medium. However, it is crucial that the strands are engineered to be both malleable and mechanically strong to facilitate proper bioprinting (Yu et al., 2016). Bioprinting dECM requires 3D printing of a frame to structurally support the dECM until full gelation is complete. Recent work demonstrated 3D printing of dECM with PCL (Pati et al., 2014), but degradation of the PCL structure is problematic when the printed construct is implanted in vivo. Bioprinting microcarriers is relatively straightforward, but loading them at high densities can cause nozzle clogging issues.

Cell encapsulation in hydrogels allows patterning of cells; however, subsequent ECM formation, digestion and degradation of the hydrogel matrix, interactions and proliferation of encapsulated cells are all issues that require careful attention. There are intrinsic limitations associated with cell encapsulation in hydrogels such as restricted cell interactions, proliferation, and colonization of immobilized cells within the hydrogel matrix, and the inability of cells to spread, stretch, and migrate to successfully generate the new tissue, particularly at high hydrogel concentrations (Ozbolat, 2015a). Cell aggregate–based bioink materials, however, exhibit better biomimetic characteristics facilitating both homo- and heterocellular interactions due to high cell densities and lack of exogenous matrix immobilization. These tissue constructs closely resemble the native tissue and preserve cell phenotype and functionality for extended periods of times. Microcarriers in hydrogels also allow better interaction between cells due to inherent high cell density. Decellularized matrix components provide a highly biomimetic environment for cell proliferation and differentiation as they are comprised of matrix proteins essential for cellular activities.

Resolution of bioprinting processes depends on the bioprinting modality as well as the bioink. When bioprinting cells in hydrogels using LBB, 5.6 ± 2.5 μm resolution can be achieved (Ozbolat and Yu, 2013). In DBB and EBB, the range can increase to approximately 50 and 100 μm, respectively (Ozbolat and Yu, 2013). Incorporating cell aggregates and microcarriers into EBB can further deteriorate the resolution. Particularly, cell aggregates in spheroid and strand forms have a diameter ranging from 250 to 450 μm (Mehta et al., 2012; Yu et al., 2016). Although bioprinting of cell pellet can result in higher resolution compared to bioprinting of tissue spheroids and strands, a reduction in nozzle diameter can have detrimental effect on the cell pellet as cells are directly exposed to the shear stress rather than cushioned within a hydrogel matrix. Although original design structure can be closely fabricated using these bioink materials, most do not retain shape and will spread, shrink, or swell depending on the bioink and culture conditions.

Most bioprintable hydrogels are affordable although some naturally derived hydrogels such as Matrigel™, fibrin, and collagen can be expensive. Microcarriers are also quite affordable as they are made of synthetic polymers (Malda and Frondoza, 2006). Cell aggregates, on the other hand, are expensive to produce and can require highly sophisticated approaches to their construction, as discussed previously. Hundreds of millions of cells are needed to prepare them in sufficient quantity, depending on the cell size and the rate of ECM deposition. Expansion of cells in these numbers is labor intensive, costly, and time-consuming. dECM-based bioink can also be expensive as a large amount of native tissue is needed to produce a small volume of bioink.

In general, bioprintable hydrogels are scalable; the majority can be obtained in abundant quantities compared to cell aggregates and dECM. Although tissue spheroids are not easy to scale up, cell pellets or tissue strands provide relatively better scalability (Yu et al., 2016). dECM, on the other hand, currently lacks scalability as the original tissue yields a very small amount of dECM. Therefore, dECM can be considered scalable for bioprinting when reinforced with an affordable hydrogel. Microcarriers are smaller in size; however, they are scalable and can be loaded in hydrogels in higher densities as long as they do not clog the nozzle unit during bioprinting.

Hydrogel-based bioink materials can be considered the most practical bioink type for bioprinting living tissues; the process can be as simple as loading cells into a hydrogel precursor solution. Different hydrogel types possess varying levels of ease in bioprinting, depending on their crosslinking mechanisms. Other bioink types can be challenging in terms of preparation and processing as well as bioprinting. For example, preparation of cell aggregates is labor intensive and time-consuming. The addition of hydrogels to the process may be necessary for delivery in an extrusion process, support aggregation of cells, or fusion of cell aggregates making the entire process quite complicated. For dECM preparation, well-established protocols are needed to generate the bioink with appropriate biological and biochemical characteristics. Moreover, the bioink needs to be loaded in a compatible hydrogel solution adding complexity to the process. Microcarriers, on the other hand, require the fabrication of porous biodegradable delivery particles for initial seeding of the cells. Upon satisfactory attachment and spreading of cells, microcarriers are then loaded in a hydrogel solution for the bioprinting mission. With the exception of hydrogels (excluding blend hydrogels requiring multiple crosslinking mechanisms), all other bioink types are a multistep approach and, therefore, are not considered practical.

Hydrogels and dECM components provide superior structural support for cells in 3D until they are able to secrete their own ECM. Mechanical and structural properties of hydrogels depend on their type and concentration; a lower concentration is more supportive of cell proliferation and growth, but demonstrates inferior mechanical properties. Microcarriers, on the other hand, provide a mechanically strong 3D scaffold as an environment for cells. Cell aggregates have no structural and mechanical support at the beginning of neotissue formation; later, cells support each other through cadherin-mediated cell–cell adhesion followed by ECM deposition (Takeichi, 1988), which lends further integrity to the 3D structure. Depending on the cell type used, mechanical properties alter over time. Parenchymal cells of metabolically highly active organs such as liver, pancreas, and heart do not secrete many ECM components, and the resulting cell aggregates are very weak in their mechanical and structural integrity. Therefore, supporting stromal cells should be cocultured to provide enough strength for bioprinting purposes. Although mechanical properties are important, they can affect the diffusion and permeability of the bioink adversely. In general, permeability of cell aggregates is lower than that of hydrogels, and the diffusion of media and oxygen is limited. Cell aggregates over 400 μm can experience hypoxia at their core, curtailing cell survival (Mehta et al., 2012).

The bioprinting time depends on the size and porosity of the construct as well as the resolution of the bioink. Generally, the bioprinting time of hydrogels and microcarriers loaded in hydrogels is shorter (i.e., few minutes). Bioprinting of dECM takes longer than hydrogels as bioprinting and solidification of the structural PCL support adds extra time. Bioprinting of cell aggregates such as tissue spheroids may take a significant amount of time using a pick-and-place type of robotic dispenser (i.e., minutes to an hour depending on the construct size and complexity) (Itoh et al., 2015). In addition, printing the supporting mold structure also contributes to prolonged fabrication times. Although the bioprinting step takes longer using cell aggregates, postprinting maturation and tissue formation time is significantly shorter than that of hydrogels and microcarrier-loaded hydrogels. Complete tissue fusion and maturation can be obtained successfully in a week, whereas this process is much longer with hydrogels (Yu and Ozbolat, 2014).

Degradability of bioink materials depends on the selected hydrogels or their blend, concentrations, temperature, conditions (in vitro or in vivo), and presence of external additives. Hydrogels that are highly sensitive to external conditions are thermosensitive hydrogels, which are easy to dissolve, and thereby lose their shape if environmental conditions are not constant. Also, degradation can be varied by hydrolytic or enzymatic labile components within a hydrogel (Nicodemus and Bryant, 2008). Additionally, the presence of cells within hydrogels limits the choice of chemical formulations while still maintaining mechanical strength. Because of their hard polymer composition, microcarriers degrade slower than hydrogels and may release substance that is toxic to cells. In summary, the degradation rate of 3D constructs must be designed according to the application.

Both natural and synthetic hydrogels and microcarriers made of polymers, including PLA, glass, alginate, dextran, and acrylamide in different forms, are commercially available (Jakob et al., 2016; Malda and Frondoza, 2006). As tissue engineering and biofabrication technologies have advanced by leaps and bounds over the past decade, 3D culturing of cells has gained the attention of commercial enterprises and a wide array of spheroid models (such as tumor or other tissue models for drug screening) are now available (Alhasan et al., 2016). Despite substantial progress, there is still a great need for representative spheroid models for several tissue types in the research community. Tissue strands and dECM-based bioink materials are relatively new and are not currently available commercially.

The introduction of any biological or nonbiological material has the potential to induce an immune response when implanted into a host. In a recent study, hydrogels obtained from other species such as collagen type I obtained from rat tail generated a low-level of immune response as demonstrated by an in vitro lymphocyte proliferation assay (Murphy et al., 2013). Therefore, the bioink, such as hydrogels and dECM components, obtained from other species can induce an immune response. However, cell aggregates fabricated from autologous cells are not expected to elicit immunogenic issues.

Bioprinted tissue and organ constructs have been used in various fields including tissue engineering and regenerative medicine, drug screening, cancer research, and high-throughput assays (Ozbolat et al., 2016); however, not all bioink types have been utilized in these application areas yet. Hydrogels have been extensively used in bioprinting for tissue engineering and regenerative medicine as well as drug testing, cancer research, and high-throughput assays. Recently, cell aggregate–based bioink materials have been used in fabrication of tissue models for drug screening (Peng et al., 2016). Although the present bioink types detailed in this chapter have been used in tissue engineering and regenerative medicine research extensively, a clinical trial has yet to be performed.