Hydrogels are the most favorable material type for cell support; however, there are several weaknesses related to bioprinting. First, hydrogels do not contain specific ECM proteins for particular cell types. Therefore, they are unable to provide a native environment. Second, hydrogels encapsulate and confine cells, limiting their interactions. Also, it is difficult to achieve the same high cell density as in native tissues. The concentration of hydrogels also plays an important role as increased concentration improves mechanical properties, but limits biological activities. Also, more rigid hydrogels require higher pressure levels for extrusion or generation of droplets. A higher concentration of hydrogel lowers the mobility of cells resulting in limited cell proliferation and deposition of ECM proteins. The cellular microenvironment provided in hydrogels may differ significantly from native tissues. Hydrogels degrade much slower in vitro, due to the lack of enzymes, immune response, and blood flow. However, hydrogels must be designed to degrade in vivo in a synchronized manner with respect to the new tissue formation. Another important shortcoming is the toxicity of degradation products that can be harmful to cells or the newly formed tissue. Instability is another important limitation, which may occur at various stages of the bioprinting process. Most hydrogels are extremely soft and can easily spread and lose their structural integrity during the bioprinting process.

Microcarrier technology is an efficient method for cell expansion in a small volume; however, this technology requires bioreactors to provide constant movement of microcarriers and appropriate environmental parameters for cell growth. One of the limitations is the need for a more complex detachment system to harvest cells, depending on the scaffold material. Another limitation is the possibility of nozzle clogging due to inherent adhesive properties of microcarriers, gravitational effects, or high density in hydrogels. In addition, the bioprinting process requires a balance between the media dilution necessary for deposition on the bioprinting stage and close contact between microcarriers in the bioprinted construct. Moreover, degradation products of microcarriers can be toxic to cells if hard polymeric ones are used. Despite its limitations, microcarrier technology has been utilized in other application areas (Malda and Frondoza, 2006) and may be an important and powerful tool for tissue engineering in the near future.

The cell pellet is a dense mass of cells in a suspension with a minimal amount of culture media. Cells quickly form aggregates in the shape that they are cast, i.e., tissue strands or spheroids. It is the most practical scaffold-free bioprinting approach, but the major hurdle is the limited size of constructs to ensure the availability of oxygen in the center of constructs. Also, the need for a temporary molding structure limits the size of tissues to very thin tubular shapes [i.e., blood vessels (Norotte et al., 2009) and nerve conduits (Owens et al., 2013)]. The construct properties ultimately depend on the cell type, the number of cells, and the final construct dimensions.

Despite the variety of tissue spheroid fabrication techniques, there are still limitations with the use of tissue spheroids within the bioprinting domain. Tissue spheroids may experience necrosis in their core, but this can be overcome by vascularization or fabricating lumenized spheroids (Fleming et al., 2010). However, if the culture relies on endothelial cells to create vascularization, the angiogenic process can take more than 48 h. This method is especially important for scale-up fabrication of tissues, where the primary issue is vascularization within the construct (Ozbolat, 2015b). Several difficulties can be encountered during the bioprinting of tissue spheroids. First of all, loading tissue spheroids into a nozzle (i.e., glass pipette) is highly challenging. Spheroids are not a uniform size and can be easily deformed or broken, depending on the cell type and maturity of spheroid. The nozzle must be designed to be large enough to hold the spheroids inside and allow extrusion without clogging. Also, the medium surrounding the tissue spheroids should also have appropriate properties. A characteristic feature of cell aggregates is rapid fusion, which is a desirable attribute postbioprinting, but not during the bioprinting process. On the other hand, if spheroids are too sparse in hydrogels, they will not contact one another upon bioprinting.

Tissue strands are an alternative approach to spheroids, differing in shape and preparation methods. The major drawback is the lack of full automation of the process as loading of tissue strands is performed manually. Also, bioprinting parameters need to be optimized for each tissue-specific strand. Another disadvantage is the requirement for high numbers of cells, depending on the desired length of the strand; additionally, tissue strands shrink substantially during maturation due to contraction. For constructs that require larger diameters for construction of scalable tissues, vascularization of the construct needs to be incorporated.

The most important limitation of dECM-based bioink materials is the need for a live organ source for acquisition of dECM. It can be a complex procedure requiring surgical protocols, donor sources, and involvement of other specialists. Even if a successful extraction is executed, the risk of an immunoreaction by the recipient remains. Moreover, the volume of final product is extremely limited compared to the size of the original organ. Decellularized matrix also loses its mechanical and structural integrity as well as some biochemical properties when homogenized. In addition, toxic residues from the decellularization process can remain. The high cost of the bioink, the need for rigorous protocols, and low yields of product are limitations of dECM-based bioink materials. The major weakness of dECM relative to the bioprinting process is its weak mechanical properties; therefore, 3D printing of a support frame is essential. The cost of dECM depends upon the quantity of ECM in the tissue. If the ECM content is low, the cost of harvesting dECM is high. With some tissue types, the preparation process may result in loss of a large portion of ECM. Other procedures such as chemical treatment can contribute to dECM contamination and toxicity. Finally, cells seeded in dECM can rapidly degrade the bioink due to the production of cellular matrix metalloproteinases (MMPs). Thus, further research is needed to optimize the use of dECM in bioprinting.

While each bioink type has its own advantages over others, hybrid tissue constructs can be envisioned as the combination of two or more complementary bioink types. In hybrid constructs, two types of bioink types can be integrated; in bioprinting a pancreas-on-a-chip model, engineered islets (in tissue spheroid form) can be loaded in hydrogels to create an in vitro vascularized pancreatic tissue model (Peng et al., 2016).