Bone tissue engineering has been widely studied using bioprinting as bioprinting has the ability to fabricate anatomically correct patient-specific tissue constructs. In a recent study (Gao et al., 2014), a thermal inkjet bioprinter was used to fabricate poly(ethylene glycol) dimethacrylate (PEGDMA) scaffolds. Bone marrow–derived human mesenchymal stem cells (hMSCs) were coprinted with nanoparticles of bioactive glass and hydroxyapatite (HA) under simultaneous polymerization. Bioprinting in that study enabled uniform distribution of hMSCs compared to manually pipetted hMSCs, which accumulated at the bottom of the scaffold due to gravity (see Fig. 9.2A). The results revealed that the bioprinted constructs encapsulating hMSCs and HA demonstrated the highest cell viability, collagen production, and alkaline phosphate activity with increased compressive modulus after 21-day culture in vitro. Bioprinting HA particles were also performed for in situ bioprinting purposes, where an LBB system was used to deposit HA nanoparticles (n-HAs) into mouse calvarial defects in a framework study (Keriquel et al., 2010), as detailed in Section 9.3.

Cardiac tissue engineering has been a growing interest as heart failure is a devastating disease (Hirt et al., 2014). While myocardium tissue has a limited regeneration capability as myocytes proliferation rapidly ceases after birth (Eulalio et al., 2012), tissue engineering of such a structurally and functionally complicated organ is essential. In the literature, limited attempt has been made in bioprinting of cardiac tissue models. Jakab et al. (2008) demonstrated EBB of tissue spheroids comprising human vascular endothelial cells (HUVECs) and cardiac cells isolated from myocardial tubes of chicken embryos. Tissue spheroids, which were adhesive and scaffold-free and which possessed rapid self-assembly capabilities, were bioprinted next-to-each other on collagen type-I biopaper in a single-layer grid pattern. Upon bioprinting, tissue spheroids were fused together in approximately 70 h and formed a single cardiac tissue patch that can synchronously beat. In addition to the scaffold-free approach undertaken in the abovementioned work, scaffold-based bioprinting has been investigated in a few studies. Xu et al. (2009a,b) bioprinted cardiac tissue constructs in a half-heart shape with connected ventricles using inkjet-based bioprinting as shown in Fig. 9.2B1–B2. In their study, primary feline adult and H1 cardiomyocytes were encapsulated in alginate/gelation composite hydrogels, and the cross-linker (calcium chloride solution) was selectively sprayed layer by layer. The resulting tissue construct with connected ventricles was electrically stimulated, and functional excitation–contraction coupling was successfully demonstrated. In addition to these studies, patterning of cells was also applied in cardiac tissue engineering. Gaebel et al. (2011) utilized a laser-induced forward transfer (LIFT) technique to pattern HUVECs and hMSCs in a geometrically defined pattern on polyester urethane urea (PEUU), and the fabricated samples were transplanted to the infarcted zone of rat hearts after LAD ligation. In 8 weeks posttransplantation, samples with LIFT-derived patterns facilitated increased vessel formation compared to randomly bioprinted cells as control groups, and the resulted myocardium patch provided significant functional improvement. Besides primary cells, human cardiac–derived cardiomyocyte progenitor cells (hCMPCs) were also bioprinted in mesh pattern made of alginate hydrogel (Gaetani et al., 2012). Bioprinted hCMPCs demonstrated phenotypic properties of cardiac lineage with enhanced expression of early cardiac transcription factors Nkx2.5, Gata-4, and Mef-2c.

Current tissue engineering techniques for cartilage regeneration cannot produce cartilage tissue that is indistinguishable from native tissue in terms of zonal properties and architectures (Makris et al., 2015). Thanks to its great potential for precise spatial and temporal deposition of cells and biomaterials with sophisticated patterns, bioprinting has recently gained increasing attention for engineering cartilage tissues that can closely mimic native tissues with zonally differentiated cells and extracellular matrix (ECM) composition (Kesti et al., 2015). Due to absence of blood vessels, cartilage tissue bioprinting has been extensively studied using various bioprinting modalities. LBB of stem cell–differentiated chondrocytes was attempted by Gruene et al. (2010) in which a computer-aided biofabrication technique was used with the assistance of LIFT. They successfully bioprinted porcine bone marrow–derived mesenchymal stem cells (MSCs) with high viability, where cells maintained their functionality and differentiation ability into osteogenic and chondrogenic lineage.

In addition to cardiac tissue engineering, engineering heart valves is also important as heart valves do not possess regeneration capability and dysfunctional heart valves, if the damage or disease is detrimental, need to be replaced by mechanical or biological prosthetic counterparts (Hockaday et al., 2014). Such replacement valves, however, are limited by thrombogenicity and calcification (Jana and Lerman, 2015). Despite its critical role in cardiovascular system, only a very limited number of work has been demonstrated in bioprinting of heart valves. Butcher's group at Cornell University demonstrated the bioprinting of a heart valve for the first time (Hockaday et al., 2012) using a dual-head bioprinter modified from a Fab@Home printer (Malone and Lipson, 2007). In that work, a dual crosslinking mechanism, consisting of ionic and physical crosslinking, was used to print polyethylene glycol diacrylate (PEGDA) mixed with sodium alginate. After printing, porcine aortic valve interstitial cells were seeded and cultured for up to 21 days. In their study, anatomically accurate axisymmetric aortic valve geometries, composed of a root wall and trileaflets, were demonstrated. Although the first-time presented work does not fall under bioprinting as cells were not involved during the printing process, the group later demonstrated bioprinting of aortic valves using different hydrogels and cell phenotypes. In another study by that group (Duan et al., 2013), dual-nozzle bioprinting of composite alginate/gelatin hydrogel was performed to fabricate thin hydrogel discs. Aortic root sinus smooth muscle cells (SMCs) and aortic valve interstitial cells (VICs) were spatially bioprinted, and samples were incubated for a week. The results revealed a cell viability of 81.4 ± 3.4% and 83.2 ± 4.0% SMCs and VICs, respectively. Acellular aortic valve constructs, on the other hand, exhibited reduced modulus, ultimate strength, and peak strain. In a recent study (Duan et al., 2014), the same group presented bioprinting of composite hydrogels using methacrylated hyaluronic acid (Me-HA) and methacrylated gelatin (Me-Gel) loaded with human aortic valvular interstitial cells (HAVICs). Fabricated samples (see Fig. 9.2D1) represented the designed trileaflet valve shape accurately, and the cells bioprinted with increased Me-Gel concentration exhibited better spreading. HAVICs encapsulated within the composite hydrogel expressed alpha smooth muscle actin (α-SMC) and vimentin (see Fig. 9.2D2) and remodeled ECM with deposition of collagen and GAGs.

Although liver has excellent regenerative and recuperative properties, liver tissue engineering has been a growing interest while liver failure, in association with failure of multiple organs, is a significant cause of morbidity and mortality (Yoon No et al., 2015). Engineering liver tissues stands as a promising direction for future organ transplantation needs in addition to the great potential of bioprinted liver tissue models in drug testing and high-throughput screening as liver tissue is highly sensitive to drug toxicity. Faulkner-Jones et al. (2015) demonstrated the bioprinting of human-induced pluripotent stem cells (hiPSCs), where bioprinted hiPSCs were stimulated and differentiated into hepatocytes for liver microorgan engineering. The presented work systematically analyzed the effect of the bioprinting process and parameters on stem cell fate and the influence of pressure and nozzle length on the viability of hiPSC and human embryonic stem cells, concluded that the utilized inkjet bioprinting process was gentle enough to maintain viability and pluripotency of cells, and directed their differentiation into hepatic lineage. A dual-head valve-based inkjet bioprinter was used to deposit sodium alginate and calcium chloride to fabricate multilayer tissue constructs (see Fig. 9.2E1), and the results demonstrated that bioprinted stem cells differentiated into hepatic lineages successfully after a 17-day differentiation period and expressed hepatocyte markers including HNF4α, albumin, and ZO-1, where albumin secretion peaked on day 21 (see Fig. 9.2E2).

Bioprinting for lung tissue engineering is highly new as there is only a recent work attempted to fabricate a lung tissue model using bioprinting. Horváth et al. (2015) demonstrated bioprinting of an in vitro air–blood barrier model using BioFactory^® by regenHU Ltd. In this regard, they bioprinted a zonally stratified tissue construct layer by layer. First, a thin layer of Matrigel™ was bioprinted as a basement membrane followed by bioprinting of a single layer of EA.hy926 endothelial cells to facilitate attachment of cells on the Matrigel™ layer. Later, on day 2, a new layer of Matrigel™ was bioprinted on the top of the previously built construct followed by bioprinting of a single layer of A549 epithelial cells. Manually deposited layers were also constructed as control samples. On day 5, the samples were fixed for characterization, and cell viability of >95% and ≥86% were achieved for epithelial and endothelial cells, respectively. As can be seen in Fig. 9.2F1 and F3, epithelial and endothelial cells were uniformly distributed with epithelial cells on the top and endothelial cells at the bottom. Sagittal histological sections also confirmed formation of a highly thin, packed, and uniform tissue layer (see Fig. 9.2F2) compared to manually pipetted control samples. The barrier quality, such as tightness of the constructs, was investigated through measuring the translocation of blue Dextran molecules from the apical to the basolateral compartment of the samples after 3 days of culture, and results revealed that the tightness of the bioprinted samples was better than that of the manually pipetted samples.

Engineering tissues for nervous system offers tremendous promise to replace diseased, aged, or injured components of nervous system; however, there is limited work done in the context of bioprinting nerve grafts. Lee et al. (2010) studied the effect of vascular endothelial growth factor (VEGF) release on proliferation and migration of murine neural stem cells (C17.2). In their study, C17.2 cells were bioprinted on a collagen layer next to a fibrin disk loaded with VEGF. The study showed that cells migrated toward VEGF-releasing fibrin gel and proliferated successfully in contrast to the cells that could not proliferate in collagen matrix. Recently, Hsieh et al. (2015) demonstrated bioprinting of a thermo-responsive polyurethane (PU) hydrogel with tunable stiffness and gelation ability at 37°C without need for a crosslinker. They showed the effectiveness of the bioink by loading it with neural stem cells and injecting it into a zebrafish embryo neural injury model. The results revealed that the injected gel rescued the function of impaired nervous system in 6 days.

As primary pancreatic β-cells do not easily survive in vitro and only a very few attempts were taken place in differentiation of β-cells from human stem cells, regeneration of pancreas tissues is primarily embodied to the extent that only β-cells from mouse lines or insulinoma cells have been used to fabricate pancreatic islets (Pagliuca and Melton, 2013; Pagliuca et al., 2015; Raikwar et al., 2015). A very few work has been done in the context of bioprinting for pancreatic tissues. Recently, Marchioli et al. (2015) encapsulated human and mouse islets as well as rat insulinoma INS1E β-cells within alginate or alginate/gelatin hydrogels and bioprinted them in dual-layer scaffolds. The scaffolds were later implanted in diabetic mice and explanted 7 days thereafter. Although the viability and morphology of islets were not impaired by encapsulation and bioprinting processes in both alginate and alginate/gelatin hydrogels (see Fig. 9.2H), bioprinted islets and INS1E β-cells lost their functionality in 7 day as they were not responsive to the change in glucose level. This can be attributed to the high level of presence of calcium (Ca²⁺) ions within crosslinked alginate because transmembrane calcium ion gradient, mediated by voltage-gated calcium channels, stimulated higher insulin secretion at low glucose level (Proks and Ashcroft, 1995). In addition, a recent work demonstrated the microfabrication of scaffold-free tissue strands (with strong expression of insulin) for EBB (Akkouch et al., 2015), where tissue strands were made of rat fibroblasts and mouse insulinoma TC-3 β-cells in the core and shell, respectively. The authors envisioned to use the demonstrated tissue strands for scale-up tissue bioprinting purposes.

Several tissue engineering approaches have been applied in skin tissue fabrication and tissue substitutes including autologous split-thickness skin graft (gold standard) (Coruh and Yontar, 2012), allografts (Leon-Villapalos et al., 2010), acellular dermal substitutes, and cellularized graft-like commercial products (Leon-Villapalos et al., 2010), i.e., Dermagraft and Apligraf (Organogenesis Inc.) (Metcalfe and Ferguson, 2007). Recently, bioprinting technology has been adopted for skin tissue fabrication as well. Lee et al. (2013) presented bioprinting of skin tissue using an 8-channel valve-based bioprinter, where a 13-layer tissue construct was bioprinted layer by layer using collagen hydrogel. Keratinocytes were bioprinted on the top of alternating layers of human foreskin fibroblasts and acellular collagen layers, and the resulting constructs demonstrated densely packed cells in epidermis layers in oppose to the dermis with low density of cells and less ECM deposition. In addition to DBB, LBB has also been used for biofabrication of skin tissue substitutes (Koch et al., 2012). Cells from human immortalized keratinocyte cell line and NIH 3T3 fibroblasts were bioprinted in collagen matrix in alternating layers on a sheet of Matriderm™. Histological results demonstrated high density of keratinocytes and fibroblasts with expression of laminin protein, which is a major component of basement membrane in skin. The same group extended their work (Michael et al., 2013) and demonstrated implantation of the tissue constructs on dorsa of mice as shown in Fig. 9.2I1. Results revealed that the bioprinted tissues were engrafted with the hosts in 11 days (see Fig. 9.2I2) with stratified epidermis with early sign of differentiation and formation of stratum corneum as well as some blood vessels. Control samples at the air–liquid interface in vitro culture, on the other hand, demonstrated proliferation of cells with limited differentiation. In a recent study (Yanez et al., 2014), Boland's group demonstrated the effect of bioprinting endothelial cells within skin substitutes on formation of macrovasculature during new tissue remodeling. In this regard, they encapsulated neonatal human dermal fibroblast and neonatal human epidermal keratinocytes (NHEKs) in collagen and laid down the dermis layer followed by patterning human dermal microvascular endothelial cells (HMVECs) on dermis layer of the construct by selectively bioprinting thrombin-laden HMVECS on manually deposited fibrinogen layer. The process was completed by covering the fibrin layer with collagen-laden NHEK cells. Then, the fabricated skin substitutes were implanted on dorsa of mice and compared with implanted commercially available skin substitutes (control). The results revealed that bioprinted HMVECs formed microvessels and the bioprinted constructs barely generated contraction compared to the control groups. In abovementioned studies, tissue constructs were bioprinted in vitro and implanted to a host; however, Skardal et al. (2012) demonstrated in situ bioprinting of stratified skin substitutes by alternating layer of fibrinogen/collagen and thrombin loaded with amniotic fluid–derived stem cells due to their lack of immunogenicity. With in situ bioprinting, skin substitutes were 3D bioprinted directly onto full thickness wounds on pigs and recapitulated the native skin more closely than control groups including the bioink loaded with MSCs and acellular hydrogels. Despite the efforts in skin tissue bioprinting, biofabrication of skin substitutes that virtually mimic native skin is still a challenge as integrating sweat glands is remained elusive.

Bioprinting of scale-up tissues and organs vitally depends on vascularization as integration of vascular network will essentially provide oxygen and media supply to cells for their survival and function (Ozbolat, 2015b). Vascular tissue fabrication has been performed by various bioprinting modalities including EBB (Zhang et al., 2013a,b; Yu et al., 2013; Norotte et al., 2009), DBB (Xu et al., 2012; Christensen et al., 2015; Blaeser et al., 2013), and LBB (Xiong et al., 2015b). In EBB, a wide variety of extrusion techniques have been utilized. The author's group used coaxial-nozzle extrusion, where hydrogels including sodium alginate and chitosan were bioprinted directly in tubular form with encapsulated cells (Zhang et al., 2013a,b). During coaxial (core–shell) flow, ejected crosslinker (flowing through the core) contacted the precursor hydrogel solution (flowing through the shell) and facilitated rapid gelation and formation of tubular constructs (Fig. 9.2J1). Six-week-cultured HUVSMC-laden samples demonstrated deposition of smooth muscle matrix (Fig. 9.2J2). That approach enabled direct bioprinting of vascular constructs in a practical manner. In addition, there are other direct vascular tissue bioprinting approaches, such as bioprinting droplets of cell-laden hydrogels layer by layer using inkjet-based bioprinting performed by Nakamura et al. (Nishiyama et al., 2008), Huang's group (Xu et al., 2012; Christensen et al., 2015), and Blaeser et al. (2013). With the ability of bottom-up construction, inkjet-based bioprinting enabled branched tubes built in both horizontal and vertical directions. A similar approach was also performed using LBB demonstrated by Huang's group (Xiong et al., 2015a,b). In abovementioned studies, scaffold-based approaches were utilized; however, Forgacs and his coworkers followed a scaffold-free approach in bioprinting vascular tissues, where tissue spheroids were bioprinted one by one and self-assembled into larger tissue units (Norotte et al., 2009). As agarose is inert to cell adhesion, the printed agarose mold facilitated rapid fusion of tissue spheroids and maturation of the tissue.

In addition to single tissue types, efforts have been geared toward bioprinting composite tissues to recapitulate the complex biology, anatomy, and functionality of organ-level structures. Merceron et al. (2015) recently demonstrated fabrication of muscle–tendon units by bioprinting hybrid constructs using a multihead nozzle assembly. In this regard, PCL and PU was 3D printed to construct a frame to support cellular constructs, where half of the unit was printed using PCL and the other half was printed using PU (Fig. 9.3A1–A4). A composite hydrogel-based bioink, comprising 3 mg/ml hyaluronic acid, 35 mg/ml gelatin, and 25 mg/ml fibrinogen in calcium-free high-glucose Dulbecco's Modified Eagle Medium (DMEM), was used to 3D bioprint 3T3 fibroblasts and myoblasts into the PCL and PU frames to construct tendon and muscle units, respectively. The results revealed cell viability of >80% with differentiated cells at the end of a 7-day culture. The final muscle–tendon units were elastic on the PU-C1C12 muscle section with an elastic modulus of 0.39 ± 0.05 MPa and stiff on the PCL-3T3 tendon side with an elastic modulus of 46.47 ± 2.67 MPa.

In addition to the presented tissue types, there is a few other work at the early fundamental study level for bioprinting of retinal and brain tissues. Lorber et al. (2014) presented piezoelectric inkjet bioprinting of retinal ganglion cells (RGCs) and glia and investigated the effect of bioprinting parameters on the viability of cells and their growth-promoting properties. They concluded that inkjet bioprinting did not adversely affect the cell viability and RGC neurite outgrowth, rather RGCs demonstrated further neurite growth when bioprinted on a glial substrate. Recently, Lozano et al. (2015) presented manual deposition of primary cortical neuron-laden gellan gum-RGD for brainlike tissue fabrication. Three-layer constructed tissue models, with cortical neurons encapsulated in the top and bottom layers, demonstrated axon growth and penetration toward the cell-free middle layer in 5 days. Although no computer-control motion system was applied, the presented work unveiled the first-time demonstration of layer-by-layer fabrication for brain tissue engineering.

Bioprinting for bone tissue fabrication has been extensively studied as bone is one of the most widely engineered tissue types in regenerative medicine. The majority of bone tissue bioprinting attempts were mainly for the investigation of the science basis of stem cell differentiation toward osteogenic lineage within bioprinted tissue constructs (Kim et al., 2010). In vivo implantation has been performed for calvarium reconstruction or bone formation under subcutaneous tissue to assess in vivo performance and engraftment into the host (Bose et al., 2013); however, bioprinting of scale-up bone tissues for significant defects or substantial bone loss is still a challenge due to the lack of integration of blood vessels. In addition, most bioprinting approaches utilize soft materials such as hydrogels for bone tissue fabrication, which is not easy to implant into load-bearing sites of human body. Therefore, hydrogels should be reinforced with mechanically strong materials or integrated with supporting frames for large-size regeneration of bone tissue in vivo (Temple et al., 2014).

Bioprinting for transplantation and clinics is a trending area of application, and a number of bioprinted tissue types have been tested in vivo; however, translation of bioprinting technology into clinics for human has yet to be succeeded as there are several impediments down the road. First of all, due to the lack of hierarchical vascular network, bioprinting of scale-up tissues is still a challenge (Ozbolat and Yu, 2013). Tissues at the clinically relevant sizes, except cartilage and skin, have not been reported yet (Ozbolat, 2015a). In the literature, bioprinted tissues have been implanted into vascularized sites on animal models such as subcutaneous tissue (Marchioli et al., 2015), but they need to be implanted into less immune-responsive sites and connected to blood stream as an intravascular device when bioprinted in larger sizes. Although a myriad of bioprinted tissue constructs have been demonstrated, one of the major issues is the cell source considering immunogenicity issues. Therefore, stem cell use in bioprinted tissues is a key factor in enabling translation of bioprinted tissues into clinical practices. There is no bioprinted product undergone FDA approval yet, and FDA has not taken any consideration on the classification of bioprinters and bioprinted tissues but as the technology evolves further through clinical trials, then the first product going through the FDA regulations will be exemplary for preceding ones.

Bioprinting for drug screening and high-throughput assays is a highly powerful technology as 3D tissue models can be precisely formed in a high-throughput manner. Both extrusion- and inkjet-based bioprinting technologies can be used to create heterocellular tissue environment to predict drug toxicity (Feng et al., 2011). Currently employed techniques in liver fabrication employ HepG2 liver carcinoma cell line, which does not represent the human liver physiology closely (Palakkan et al., 2013). Although complex patterned structures have been successfully integrated within a perfusion system to assess drug toxicity in a control manner, the cell type used as well as the lack of other associated cell types reduces the reliability of the tissue model in predicting the performance of drugs (Chang et al., 2010). In this regard, Organovo's liver tissue model stands as a promising tool to represent a biomimetically developed liver tissue model for drug testing (Roskos et al., 2015). The employed scaffold-free technique comprising stem cell–derived hepatocytes as well as hepatic satellite cells and endothelial cells facilitates physiologically relevant liver tissue, and further progress is needed to create a standardized bioprinted liver tissue model for pharmaceutical industry. In addition to liver tissue models, cells can be bioprinted in a high-throughput manner to study their behavior as well as bacterial cells can be bioprinted along with drugs to understand the interactions between applied drugs and bioprinted different bacterial strains (Rodríguez-Dévora et al., 2012). Different dosages can be bioprinted in high-throughput manner to effectively analyze the influence of different antibiotics on bacterial cells; however, a translational effort is yet to be demonstrated for the pharmaceutical industry.

Bioprinting for cancer research is relatively new, and there are only a few efforts in fabrication of tumor models (i.e., ovarian, cervical, and breast tissue models) (Knowlton et al., 2015). To study cancer pathogenesis, inkjet-based bioprinting stands as a promising tool in generating 3D tumor models with multiple cell types in controlled composition; however, there is no well-established model in abovementioned tumor types that can be used in pharmaceutics industry yet. Formation of larger tissues as an in vitro cancer tissue model for exploration of cancer metastasis has not been attempted using bioprinting yet; however, there are various microfluidics-based platforms with closely recapitulated cancer microenvironment for cancer metastasis research (Jeon et al., 2015; Zervantonakis et al., 2012). These platforms lack precise placement of multiple cell types in a localized manner. Bioprinting has the capability of patterning the cellular microenvironment, but further research is required in angiogenesis work in bioprinting to establish standard cancer tissue models for cancer metastasis research.