Synergistic coupling between 3D bioprinting and vascularization strategies

The essential components of the pre-bioprinting stage entail minimally-invasive tissue biopsy to obtain viable cells, efficient protocols to maintain these during culture, high-resolution imaging to acquire structural and anatomic data, and creation of 3D models of the target tissues or organs using various modeling approaches [18]. Prior to bioprinting, designing of injury-specific and biomimetic bioprinted constructs requires the understanding of the anatomy of targeted tissues or organs. In addition, accurate information on an integrated vascular network is highly essential to ensure adequate and proper nutrient supply. Thus, high-resolution medical imaging exhibit great significance in the process of bioprinting [19].

Currently, magnetic resonance imaging (MRI), computed tomography (CT), ultrasound imaging (USG) and various other imaging modalities (i.e., optical coherence tomography (OCT), angiography, etc [20]) are used to obtain 3D anatomies of tissues and organs (figure 1(a)). MRI is a highly preferred modality for imaging soft tissues, which utilizes pulsed radiofrequency electromagnetic waves to visualize the target [21]. CT provides high-resolution images compared to MRI images, with reduced time of scan. Anatomy can be conveniently obtained from CT images, which involves radiation. Further, the advancement from CT to micro-CT enables to characterize not only bulk anatomy but also the microstructural and mechanical properties of scaffolds, which is widely used for imaging bone density alteration and regeneration of the tissue. On the other hand, UGS is considered the safest imaging modality and has recently been suggested as a way to give bioprinted part sub-surface information. OCT is widely utilized in the disciplines of biological (cell dynamics and tissue growth in engineered tissues) and industrial testing because it is non-destructive, label-free, high resolution, and quick [22–27]. As opposed to fluorescent or ionizing radiation, OCT depends on the scattering characteristics of the sample [28].

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The field of medical imaging strategies has witnessed a transition from general tissue imaging to specialized vascular imaging. In the past, the focus primarily rested on capturing detailed anatomical structures and tissue composition. However, with the advent of advanced technologies, such as contrast-enhanced ultrasound, magnetic resonance angiography, and CT angiography, the emphasis has shifted towards non-invasive visualization of blood vessels and circulatory systems. This shift has allowed clinicians to study blood flow dynamics, and plan minimally invasive interventions with greater precision. For example, CT angiography combines a CT scan with an injection of a contrast agent and produces images of arteries in a body part [29]. Spiral CT scanners with 4–64 sections enables quick capture of isotropic data sets [30]. Utilizing the intrinsic nuclear spin characteristics of the nucleus to selectively detect and assess flow in a system of moving parts is a clinically beneficial application of MR angiography. Combining the flow experiment with surface coil techniques will result in high-resolution MR angiographic images [31, 32]. Typically, imaging vascularization is important for understanding healing processes of injured tissues and the relationship between blood vessels and host tissues, which can accelerate the development of treatment modalities. One of the advanced imaging techniques is the quantitative 3D imaging platform, which is a combinatory system of optical tissue clearing, light-sheet microscopy, and 3D image analysis [33]. While conventional techniques allow imaging of small regions (mm² areas), the quantitative 3D imaging platform enables the capturing of 3D maps of skeletal progenitors and vessel subtypes in the whole murine calvarium at single-cell resolution. Also, the 3D maps can be quantitatively analyzed including platelet endothelial cell adhesion molecule-1, endomucin, and osterix expression, which suggests a detailed framework for vascularized craniofacial bone remodeling. In another study, hierarchical imaging was performed to attain 3D brain vascular networks using scanning electron microscopy, synchrotron radiation and desktop micro-CT imaging, and computational network analysis [34]. Network features were quantified in various aspects, such as directionality, vessel diameter, length, and tortuosity, branch point density, and vascular volume fraction, from the early postnatal to adult brain. The hierarchical imaging technique can provide a detailed morphological development of the entire murine brain vasculature, which can also be applied to biologically-relevant questions including angiogenesis, pathophysiology, and drug effects. In addition to above techniques, a hybrid in-vivo imaging technique, photoacoustic microscopy, was introduced and customized for a high signal-to-noise ratio and high resolution for microvascular imaging [35]. While many imaging techniques require contrast agents, photoacoustic microscopy successfully captured microvessels by localizing photoacoustic signals from red blood cells without using contrast agents. Then, vascular network in ears, eyes, and the brain of a mouse were reconstructed as a 3D photoacoustic map image, and the flow of red blood cells in the mouse ear was represented. Overall, vascular imaging techniques are rapidly developing in terms of high resolution and computational analysis, and the attained anatomical atlas will be a useful tool for confirming the reconstruction of vascularized tissue and organ substitutes.

Completion of 3D models requires designing of the internal architecture of tissue substitutes after the acquisition of 3D anatomy. The internal architecture includes vascular networks for the supply of nutrients, internal channels, and pores for proper cell attachments, proliferation, and migration. The most common approaches utilized in designing of internal structure entail computer-aided design (CAD), freeform design, image-based design, internal surfaces, etc [36]. The generated 3D model is then converted into stereolithography (STL) format, which is the most readable and accepted file format for most commercial bioprinters (figure 1(a)). Once a blueprint of 3D anatomical structures is obtained, selection of biomaterials should be followed according to tissue-specific mechanical and biological characteristics.

2.1.1. Selection of biomaterials

2.1.2. Angiogenic biomaterials

2.1.3. Angiogenic factors in bioprinted vascular constructs

As an angiogenic growth factor, VEGF is a key contributor in formation of new blood vessels; and endothelial cells respond favorably to VEGF, which triggers their proliferation and sprouting to produce new, immature artery sprouts [116]. Even though VEGF can induce angiogenesis, other elements are needed to encourage vessel maturation. For example, PDGF promotes development of the nascent arteries by stimulating and attracting pericytes and smooth muscle cells that interact with endothelial sprouts, maintaining them and limiting regression [117]. Other growth factors that directly or indirectly regulate angiogenesis include the acidic and basic fibroblast growth factors (FGF-1and FGF -2 [118], transforming growth factor-beta (TGFβ) [119], angiogenin [120] and epidermal growth factor (EGF) [121]. For example, Meng et al developed 3D bioprinted metastatic model by replicating the structure of peripheral blood vessels and growth factors (such as EGF and VEGF) were released from stimuli-responsive capsules. Using 3D bioprinting, a molecular gradient of growth factors was created to replicate the biochemical properties of the tumor microenvironment, where lung cancer cells migrated towards arteries and invaded into the vascular system [122]. However, the model was relatively simple and a holistic, step-by-step integration of more constituents, such as immune cells, lymphatic vessels, etc., can be considered to mimic the heterogeneity of the tumor microenvironment. In another study, Park et al reported a method for pre-vascularizing bone tissue with dual growth factors. Human dental pulp stem cells that have both osteogenic and vasculogenic potential were bioprinted with bone morphogenetic protein-2 (BMP-2) in the peripheral zone of constructs and VEGF in the central zone, which was prone to hypoxia (figure 2(a)). Pre-vascularized constructs showed faster bone repair compared to non-vascularized tissues [123]. However, the constructs were tested in mouse dorsa only for 4 weeks and thus requiring their prolonged testing, if possible in larger animal models. Application of growth factors, in particular VEGF, has been the focus of efforts to stimulate blood vessel formation. However, it still raises concerns about the effect of released VEGF into the blood stream as there is a strong likelihood of undesired cell proliferation or tumorigenesis locally or distally.

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In other works, endothelial cell-specific microRNA-126 (miR-126) have shown to promotes angiogenesis in response to angiogenic growth factors, such as VEGF or FGF, by suppressing opposing signal transduction pathway regulators [125]. It has also been shown that some stem cell populations can be inducted with endotheliogenic differentiation through miR-210, commonly known as the master hypoximiR (hypoxia-inducible miR) [126]. Using miRNA co-differentiation for 3D heterotypic pre-vascularized bone formation, Celik et al fabricated doublet structures using spheroids of ADSCs transfected with miR-148b and miR-210, and showed their osteogenic and endothelial differentiation, mineralization, and bone formation potential (figure 2(b)) [124]. Concurrently, as gene and DNA-based therapies continue to advance, there is an increasing need for plasmid DNA (pDNA) production and application. The use of pDNA to produce growth factors in transfected cells provides a powerful alternative and is beneficial in long-term effects with low costs as compared to purified protein. Chen et al formed mature blood vessels using electrospun fibers with loaded multiple pDNA-calcium phosphate nanoparticles. Fibers with encapsulated nanoparticles containing plasmids encoding VEGF (pVEGF) and basic FGF (pbFGF) led to significantly higher density of mature blood vessels compared to those containing individual plasmid [127]. However, the use of pDNAs poses a safety concern and needs more testing, which limits their clinical translation.

To induce vascularization in implanted tissues, the application of pro-angiogenic growth factors to recruit a host vasculature is a widely used strategy. However, the use of growth factors in large scale is a costly strategy and less effective in vivo due to their relatively slow release, which may also be unfavorable to cell viability especially immediately after implantation. Despite bioprinting of angiogenic factors is attractive in inducing angiogenesis, generation of perfusable macro-scale vascular constructs requires the use of bioprinting technologies in order to build such from scratch.

In addition to the bioactivity, angiogenic performance, and angiogenic factors of biomaterials, bioprinting modalities differ and add more criteria on selection of biomaterials, such as viscosity, print fidelity, and mechanical strength. Hence, biomaterials and bioprinting modalities are often required to be contemplated simultaneously; in this sense, the bioprinting stage including bioprinting modalities and related strategies on vascularization is elaborated.

The bioprinting stage encompasses several different steps including the preparation of a bioink, bioprinter, and the bioprinting process. Bioprintability of the targeted tissue is influenced by several factors associated with the bioink and process. The bioink provide the process-ability for bioprinting to develop 3D constructs, where cells can proliferate and mature to form tissues or organs. The rheological properties of the bioink, such as viscosity, yield stress, and storage and loss modulus influence the print fidelity and the structural and mechanical strength of bioprinted constructs [18]. Although secondary cell lines have often been utilized, patient-specific primary or stem cells are ideal and suitable for the fabrication of transplantable tissue constructs to minimize rejection rates.

2.2.1. Bioprinting modalities

2.2.2. Bioprinting approaches for vascularization strategies

2.2.2.1. Direct bioprinting for vascular tissue fabrication

The direct bioprinting approach utilizes cell-encapsulated bioinks to actively bioprint hollow vascular constructs. Thus, researchers preferred blending different materials to optimize a suitable balance between mechanical strength and bioactivity for bioprinting. For example, Collins and Birkinshaw developed a bioink comprising of GelMA, sodium alginate, and 4-arm poly(ethylene glycol)-tetra-acrylate (PEGTA) to 3D bioprint perfusable vascular structures with a multilayered coaxial extrusion system [85]. However, the compressive moduli of the constructs significantly decreased after 21 d of culture, mainly due to the degradation of the GelMA component. Thus, it was challenging to test some of essential mechanical parameters, such as suture retention and burst pressure, and to keep the perfusability of constructs beyond 21 d. The bioink supported the growth and proliferation of encapsulated endothelial and stem cells, resulting in the development of viable vasculature [85]. Hinton et al reported a customized EBB technique called, freeform reversible embedding of suspended hydrogels (FRESH), for precise deposition of bioinks into a secondary support bath to maintain their structural integrity. The study demonstrated the ability to fabricate a part of the right coronary artery vascular tree with a hollow lumen with a wall thickness of <1 mm [155]. FRESH bioprinting showed capabilities to engineer soft hydrogels into complex structures; however, the direct bioprinting of functional tissues requires further research and development to become fully realized. In another study, Cui and Boland used DBB (inkjet bioprinting) to precisely manufacture micron-sized fibrin tubes. Human microvascular endothelial cells and fibrin were bioprinted together where the cells eventually arranged themselves and proliferated inside the channels (figure 3(a)) [156]. Using LBB, Wu et al printed simple branch/stem structures of human umbilical vein endothelial cells (HUVECs) and human umbilical vein smooth muscle cells (HUVSMCs) onto a thin hydrogel substrate. The structure mimicked the vascular networks in native tissue and allowed cells to develop new, finer structures away from the stem and branches. When HUVSMCs and HUVECs were bioprinted near one another, they exhibited a symbiotic interaction and form cell–cell junctions around lumen-like structures [156]. However, these were preliminary results in the creation of a self-assembling, artificially-guided vascular network, which requires comprehending the connection between HUVSMCs and HUVECs.

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The use of coaxial nozzles in EBB revolutionized the fabrication of vascular constructs and resulted in the ability to bioprint lumen-incorporated strands. It involves a configuration of concentric nozzles wherein a bioink within the core is encapsulated by a shell crosslinker and on reversing this configuration allows for the deposition of hollow fibers. The diameter of the nozzle often requires optimization wherein the resolution could be improved upon decreasing diameter. The second factor is the flow rates of the shell and the core, which could be tuned to produce fibers of different configuration and shapes. In this regard, Zhang et al used coaxial printhead for direct hollow fiber fabrication, eliminating the need for any post-processing steps that are vital in other methods. The bioprinted hollow fibers were composed of either cell-laden chitosan or alginate hydrogels with a lumen diameter of less than 200 μm [159]. In another study, Zhang et al encapsulated HUVSMCs in sodium alginate and bioprinted them in the form of vascular conduits using a coaxial nozzle, where cells exhibited growth with deposition of collagen and smooth muscle matrix on and around the lumenal surface (figure 3(b)) [158]. Hong et al reported coaxial bioprinting of vascularized tissues using a human dermal fibroblast (HDF)-laden gelatin-PEG-tyramine (GPT) prepolymer-based bioink in the shell and gelatin-loaded with HUVECs in the core. Tyramine and gelatin were separated by PEG in the GPT process, resulting in a quick gelation time of about 4.5 s. The perfusable vascular constructs were maintained for up to 8 days in vitro [160]. Nevertheless, the study requires further understanding on controlling the spatial distribution of bioprinted cells during long-term tissue maturation. In another study, Shao et al employed coaxial bioprinting with HUVECs-laden GelMA bioink in the core surrounded by alginate in the shell resulting in formation of vessel-like structures [161]. Yet, these endothelialized microfibers were not hollow during culture, which requires further investigation. In another study, Gao et al created a hybrid bioink composed of VdECM and alginate mixed with EPCs and proangiogenic drugs/PLGA microspheres, which promoted EPC differentiation, proliferation, and neovascularization allowing direct fabrication of blood-vessel structures [102]. However, low mechanical strength of vessels was a weakness of the study hampering its surgical anastomosis with the host blood vessel in the implantation process. Recently, Yu et al adopted a modified coaxial nozzle with two core needles to extrude double-channel gel fibers to simultaneously build perfusable multi-material constructs [162]. The filament in one channel can play a core/shell role, while the filament in the other channel can play a perfusion role. These parallel channels within filaments were separated by a ∼50 μm wall of alginate, which is advantageous for nutrient supplementation via perfusion. However, the customized dual-core coaxial nozzle has not been tested for vascular applications. In another approach, small-diameter blood vessel grafts containing both functional endothelial and muscular cell layers were fabricated. For this, direct reservoir-assisted triple-coaxial cell printing and ECM bioinks obtained from the vascular tissue were used. The prematurely developed vessel was implanted in rat abdominal aorta for 3 weeks, which demonstrated patency, re-endothelialization, modified smooth muscle, and engraftment with the host tissue [163]. Although this approach accurately recapitulates small blood vessels (i.e., arteriole), it may not be suitable to study larger vessels (i.e. arteries, aorta). In another study, Gold et al developed a nanoengineered ECM bioink, composed of GelMA, poly(ethylene glycol) diacrylate (PEGDA), and 2D nanosilicates. The bioink was bioprinted into 3D cylindrical constructs comprising co-culture of vascular smooth muscle cells and endothelial cells. The 3D bioprinted constructs recapitulated thromboinflammatory reactions observed in advanced preclinical models or in vivo upon cytokine stimulation [164]. Of late, strategies leveraging the advantages of both microfluidics and coaxial printheads have also been reported [165]. Using this approach, it is possible to bioprint with low viscosity bioinks and still get a higher level of histoarchitectural complexity by mounting the chip upstream of a coaxial dispenser. Additionally, this can simultaneously deliver bioinks and crosslinking agents as separate flow streams through the coaxial nozzle, which allows single-step generation of stand-alone, hollow vascular conduits [166]. For example, Pi et al developed a multichannel coaxial extrusion system for microfluidic bioprinting of circumferentially multilayered tubular tissues in a single step, using customized bioinks consisting of GelMA, alginate, and eight-arm poly(ethylene glycol) acrylate with a tripentaerythritol core [167]. Recently, Ching et al developed a microfluidics-enabled molding approach combined with coaxial bioprinting to fabricate cell-laden vascular models. Freestanding and perfusable vascular structures with complicated shapes were fabricated using 3D porous molds of poly(ethylene glycol) diacrylate as casting templates that gradually release calcium ions as a crosslinking agent. The fabricated vasculatures could recapitulate physiologically-relevant conditions related to vascular diseases [168].

Direct bioprinting of vascular constructs has resulted in some successful in-vivo outcomes [169] but factors including immunological response, adverse host response, appropriate degradation rate, fibrosis, and necrosis due to limited nutrient diffusion and mechanical compatibility with the native tissue limit their wide applications [170]. This has motivated researchers to develop scaffold-free strategies. The first seminal study towards this was accomplished by Norotte et al, who used vascular cells including smooth muscle cells and fibroblasts, which were assembled into distinct units that were either multicellular spheroids or cylinders with regulated diameters (300–500 µm). These were bioprinted layer-by-layer alongside with agarose rods, used here as a molding template. The individual components were combined after bioprinting to create single- and double-layered small-diameter vascular tubes (outer diameter ranging from 0.9 to 2.5 mm) with distinct shapes and branched structures [171]. However, limitations of the study include cell apoptosis due to thick vascular wall, limited resolution, and the inability to form complex geometries. In another study, Tan et al used 3D printed alginate hydrogel molds to facilitate spheroid fusion processes. Ring-shaped alginate molds were fabricated to obtain toroid-shaped tissue units where tissue spheroids, consisting of endothelial and smooth muscle cells, were then robotically deposited into the molds, which underwent fusion to form vascular constructs. However, additional calibrations and system enhancements will be required for the creation of non-open-structured molds to generate vasculature at smaller diameters [172]. Despite these intense efforts to bioprint vascular constructs, factors such as sufficient mechanical strength for structural integrity, requiring greater time to mature before enough ECM can be deposited, and scaling up structures to produce larger tissues still pose significant challenges. As an alternative to direct bioprinting, indirect bioprinting has a higher degree of freedom to build intricate structures and capability to encapsulate biomaterials within perfusable channels. Hence, detailed principles and applications of indirect bioprinting are followed.

2.2.2.2. Indirect bioprinting for vascularized tissue fabrication

Indirect bioprinting utilizes a sacrificial ink, which is first deposited in a hydrogel matrix and then removed to create structures that resemble hollow vessels. These vessels can then be populated with endothelial as well as smooth muscle cells. Towards an effective technique for vascularization of tissue constructs, Bertassoni et al reported a method using 3D printed agarose template fibers, which were later removed for creating perfusable networks inside GelMA (figure 4(a)) [173]. Further, using a casting approach, Miller et al 3D printed rigid filament networks of carbohydrate glass comprising of a mixture of glucose, sucrose, and dextran as a fugitive phase. Several hydrogels, including PEGDA, fibrin-fibrinogen-thrombin, and alginate were investigated to form different vascular architectures with viable endothelial cells inside perfusable channels (figure 4(b)) [174]. For example, using DBB, Lee et al fabricated perfusable functional vascular channels using bioprinting of collagen blocks and HUVEC-gelatin mixture, where gelatin was eventually liquified to obtain tubular channels. Confluent endothelium was formed inside the bioprinted constructs, and it was demonstrated that under physiological conditions, cells up to 5 mm away from the channel could still survive [175]. In another study, Kolesky et al layer-by-layer co-bioprinted human neonatal dermal fibroblast-loaded GelMA strands, fugitive ink, and fibroblast-laden GelMA. Fabricated vascular network was enclosed in GelMA, and the fugitive ink (Pluronic F-127) was eliminated by liquifying it at 4 °C. More than 95% of HUVECs were found to be viable with a confluent EC layer formed 48 h after seeding [176].

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Although the indirect bioprinting approach for vascularization of tissue constructs has advanced significantly due to the ease with building thick or slab gels as opposed to free-standing vascular constructs, such thick gel constructs may exhibit challenges including the inability to recapitulate multiple layers of blood vessels, inappropriate degradation of the bulk gel, and due to the intricacies involved with sacrificial inks, the size, morphology, and functionality of the obtained vessel-like structures may be constrained. Although indirect approaches are highly appealing, for fabrication of organ-on-chip platforms, their utilization in implantable tissues is limited as they usually lack suturable vascular pedicles to anastomose to the host vessel system.

After bioprinting tissue and organ substitutes, the stage of post-bioprinting plays a critical role to differentiate stem cells and mature bioprinted substitutes in suitable bioreactors, which is a highly time-dependent process and performed under tightly controlled conditions (figure 1(c)). Cell–cell interactions, cellular differentiation and proliferation, and degradation of bioinks are controlled and coordinated during the post-processing stage to mimic the natural biological environment. Various stimulating factors are regulated inside bioreactors to modulate and enhance the maturation and vascularization of bioprinted tissue and organ substitutes. During this stage, sacrificial materials (such as vascular templates) can be removed either immediately or in controlled programmed manner to generate vascular networks [177]. This stage also includes the removal of the support material, if only, incorporated at the bioprinting stage to stabilize the constructs [18]. After vascular networks are generated, several strategies can be considered for maturation of blood vessels using bioreactors such as perfusion, rotational, or pressure bioreactors. To elaborate, perfusion bioreactors are composed of media, sensors, pumps, grafts, and chambers [178]. They enable dynamic stimulation on vasculature with controlled pressure, flow, and shear stress, mimicking physiological forces. For instance, a study proposed perfusable tubular scaffolds (20 mm in diameter and 2 mm in thickness) grown in a customized perfusion bioreactor [179]. The luminal surface of scaffolds was coated with fibronectin for functionalization purposes. When controlled flow rate was applied, the scaffolds revealed fluid flow without leakage and accelerated endothelial cell proliferation. Although perfusion bioreactors have proven to effectively mimic physiological stimulation and induce various cellular activities, their complex setup may hinder broader accessibility and clinical applicability. Rotational bioreactors can be utilized with a relatively simpler setup, which can be employed for cell seeding and development of cell-loaded tubular sheet structures [180, 181]. Also, smooth-muscle cells have been shown to elongate toward the flow direction, so bilayer blood vessel-like constructs can be achieved with control of rotating directions and use of multiple cell types (i.e., endothelial, smooth muscle, and fibroblast cells). However, the developed construct should be eventually tested with fluid flow as carotid arteries experience dynamic conditions with 60–120 mmHg pressure, 15–43 cm s⁻¹ flow velocity, and 6–21 dyne cm⁻² wall shear stress [182, 183], in which hemodynamic conditions can be recapitulated with pressure bioreactors in vitro. For the proof of this concept, several customized or commercially-available pressure bioreactors have been reported with different levels of complexity [184–188]. As compared to static conditions, smooth muscle organization and endothelial coverage were preserved in the vessel wall under pulsating conditions created by a pressure bioreactor [189]. Although bioreactors often require re-intervention and have not yet met off-the-shelf simplicity, the development of bioreactors will suggest a promising way to culture viable vessels and a patient-relevant system.

Several allogeneic multilayered skin grafts have been tested for impaired non-healing cutaneous wounds but such skin grafts have failed over time permanently due to lack of complex skin microvasculature crucial for integration with the host tissue [191]. To meet such requirements, perfusable, vascularized skin substitutes were demonstrated to improve dermal and epidermal maturation [191]. The grafts were fabricated via EBB using two different types of type I collagen-based bioinks: human-derived (1) dermal fibroblasts (FBs), EC, and placental pericytes (PCs) for the dermis layer, and (2) keratinocytes (KCs) for the epidermis layer. After bioprinting, grafts were engrafted onto the skin wound site of C.B-17 SCID/bg mice and in-vivo perfusion was confirmed at Week 4 [191]. For a complete xeno-free implantation using immunodeficient (C.B-17 SCID/bg) mice, a bioink comprising of human collagen type I and fibronectin was developed and embedded with cells in the same manner (a dermal layer with EC, FBs, and PCs and an epidermal layer with KCs) [192]. The bioprinted bilayered substitutes revealed skin-specific features like rete ridge-like structures found a mature stratified epidermis and perfusable microvessels within 2 weeks. These studies suggest that the anatomically accurate layers and crosstalk among epithelial and other cells (i.e., KCs) are important, which results in beneficial paracrine effects attained from PCs [193, 194]and maturation on KCs. Another study demonstrated that Zn₂SiO₄ (ZS) nanoparticles supported vascularized, innervated skin regeneration (figure 6(a)) [195]. In this regard, 8 week-old BALB/c mice were used to establish a 2nd-degree burn model and transplanted with ZS nanoparticles-incorporated into PCL fibrous scaffolds. The local release of Zn and Si ions was favorable to innervation and re-vascularization, which was illustrated by accelerated wound healing process, endothelial-specific marker (CD31), and neuron-specific marker (PGP 9.5) expressions. Along with vascularization, pigmentation is another important feature of the skin tissue protecting epidermal layers from natural UV radiation and substantially influencing social and phychological well-being of patients [196, 197]. Thereby, reproduction of pigmented and vascularized skin substitutes have been attempted. Among various bioprinting approaches, a robotic platform was introduced, which is a hybrid process combined with EBB, plastic compression, and DBB [198]. For the basal layer, human-derived FBs and ECs were encapsulated in collagen type I-based bioink and bioprinted in a single layer via EBB. The plastic compression was firstly introduced and performed for even distribution of the collagen layer in 3D without affecting cell viability. Lastly, DBB was conducted using human-derived KCs and melanocytes to form a patterned array for the epidermis layer. After transplantation to immunodeficient rats, the bioprinted substitutes showed similar percentage of HMB45⁺ melanocytes to the control foreskin and successfully anastomosed to the vascular plexus of the host. In brief, while recent advances in skin substitutes allowed vascularized, innervated and/or pigmented skin wound healing using multiple cell types, in-vitro skin grafts have not yet shown similar maturity to native skin in terms of vascularized dermis or stratum corneum structure and should be improved for developing off-the-shelf clinical products [191, 199].

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Skeletal muscle is a dynamic, heterogeneous, and innervated tissue composed of highly differentiated bundles of muscle fibers [204]. Although muscle fibers have potential to regenerate and self-repair from small injuries, large volume of muscle defects and genetic disorders pose a constant clinical challenge. The considerable progress in the domain of bioprinting and the advent of muscle-specific bioinks have led to successful fabrication of bioprinted vascularized muscle constructs mimicking the mechanical, structural and biochemical features of human muscles, which have been tested in vivo [10]. This has been evidenced by 3D implantable muscle constructs which were fabricated using a human primary muscle progenitor cell-laden hydrogel bioink, a sacrificing acellular gelatin, and a supporting PCL frame (figure 6(b)) [200]. In-vitro evaluation of the muscle constructs was conducted using myosin heavy chain, which confirmed the maturation. Following in-vitro analysis, the bioprinted muscle graft was implanted in rodents with tibialis anterior muscle defects and showed densely packed myofiber-like structures with 82% recovery, von Willebrand factor (vWF⁺) vessels, and neurofilament (NF⁺) nerves. Ngan et al developed a bioprinting approach to generate muscle constructs composed of myoblast encapsulated GelMA bioink via photo-crosslinking [205]. Bioprinted GelMA constructs were implanted in nude rats, and spontaneous formation of myofibers was observed. Immunohistochemistry analysis revealed characteristic tissue maturity with sarcomeric striations, where enhanced maturation and well-organized interconnections between ingrowth vessels and nerve axons were observed. In addition, Lee et al utilized EBB using skeletal muscle-specific bioink comprising of dECM-methacrylate (MA), polyvinyl alcohol (PVA), and human primary muscle progenitor cells, which allows self-alignment and release of bioactive factors [206]. Owing to synergistic effects of topographical and biochemical cues, the bioprinted substitutes revealed neuromuscular junction characteristics by vWF⁺ vessels, NF⁺ axons, and myosin heavy chain (MHC⁺) myofibers when implanted into rats with a tibialis anterior muscle injury model. Also, a coaxial nozzle was utilized to bioprint human skeletal muscle cells (hSKMs) in the skeletal muscle-based dECM via an inner core nozzle and HUVECs in VdECM via an outer shell nozzle [101]. The core–shell filaments were bioprinted in a parallel manner and transplanted into Sprague–Dawley rats with VML injuries. Interestingly, as compared to hSKM-HUVEC mixed filaments, coaxially-bioprinted group revealed higher muscle weight and number of blood vessels and CD31⁺ cells at Week 4. The compartmentalization of the structure (allocation of different dECM) and cell types induced each cell to fully appreciate its environment and promote myogenesis and angiogenesis efficiently. Furthermore, innvervation was observed in the coaxially bioprinted group with acetylcholine receptors (AChRs⁺) and anti-beta III tubulin (TUJ1⁺) cells. Overall, vascular and neural components, such as endothelial and neural cells and growth factors, can be considered to support muscle cell survival, differentiation, and muscle function.

Over the past few decades, there has been an exceedingly huge demand for patient-specific bone grafts to repair bone defects [21]. Recently, a group of researchers has developed cuboid shaped, prevascularized bone tissue constructs with human adipose-derived mesenchymal stem cells (ASCs) and HUVECs (figure 6(c)) [201]. EBB and DBB were used to bioprint the constructs. For bioprinting of HUVECs, fibrinogen with various growth factors, such as VEGF, bFGF, and aprotinin, were prepared. Osteo-hydrogel, containing fibrinogen, gelatin, hyaluronic acid, glycerol, hydroxyapatite, VEGF, bFGF, aprotinin, and osteogenically differentiated ASCs, was prepared. Subcutaneous implantation of these constructs into immunodeficient mice revealed the formation of blood vessels and microvessels of various calibers from the bioprinted HUVECs. A calcified bone matrix was produced by bioprinted ASCs indicating ectopic bone formation. However, the implanted substitutes did not maintain their original shapes because of deformation, and a physical support made of hard polymers or bioceramics can be considered to prevent deformation. Regarding this, a bioceramic bioink consisting of collagen, β-tricalcium phosphate (β-TCP), human-derived ADSCs, and HUVECs was utilized [207]. Bone substitutes were bioprinted via EBB using the bioink formulation (80% β-TCP; ADSC:HUVEC at 7.5:2.5 ratio) opted from in-vitro studies, in which a high ratio of β-TCP had been proved to activate in-vitro osteogenesis of ADSCs [208]. After inserting into a mouse model of posterolateral lumbar spinal fusion, all the index of bone regeneration (i.e. fusion rate, bone volume, bone mineral density, and trabecular thickness and number) significantly improved in β-TCP bioprinted substitutes. Also, a significant increment in number of blood vessels and CD31⁺ area was observed. The crosstalk between ADSCs and HUVECs secreted growth factors, chemokines, and cytokines, which enhanced angiogenesis and osteogenesis via tumor necrosis factor alpha and Notch signaling pathways [209, 210]. On the other hand, a chemically-modified nanocomposite bioink was proposed using amine-functionalized copper-doped glass nanoparticles, oxidized alginate, and gelatin to improve biological performance and printability via EBB [211]. This composite bioink was bioprinted with osteosarcoma cells and bone marrow stem cells (BMSCs), which demonstrated high cell viability, rapid spreading, and differentiation owing to the presence of cell-adhesive ligands in the nanoparticle-laden bioink. Without growth factors in an ex-vivo environment, copper, known to induce secretion of VEGF from BMSCs [212, 213], promoted ion stimulation leading to both osteogenesis and angiogenesis. Further, a GelMA-based bioink added with BMP-4 and BMSCs was employed to reconstitute osteon-mimetic vascularized bone substitutes via EBB [214]. This osteon-mimetic substitute was designed in an heterogeneous structure consisting of the alternation of PCL and BMSC-laden filament, which was seeded with EPCs. Also, their architectures were altered in terms of canal structures (central medullary, peripheral Haversian, and transverse Volkmann canals), and different canal compositions were implanted into BALB/c nude mice subcutaneously and rabbit lateral femoral condyle. Despite low mechanical properties, the hierarchical structure including all canals revealed the highest percent vascularity and bone regeneration. Also, EPCs and BMSCs mutually promoted osteogenesis and angiogenesis by secreting exosomes and related factors [215–217]. Overall, bone substitutes should have adequate vascularization to support bone regeneration, and more bone microenvironment-related cells, such as osteoclasts and endothelial, neural, and immune cells, should be considered.

Currently, myocardial infarction and congenital heart disorders are the leading cause of death worldwide [218]. Cardiac tissues are conglomeration of different involuntary muscle types. Due to the inherent limitation of host immune rejection to artificial materials, scaffold-free cardiac tissues were established utilizing a commercially-available spheroid bioprinting platform [8]. Spheroids of cardiomyocytes derived from iPSCs, endothelial cells, and fibroblasts were bioprinted on a needle array using the Kenzan method and assembled into scaffold-free cardiac tissues. Cardiac tissues were then transferred to a chamber to electrically stimulate them, where an increased beating rate was observed in all samples. However, due to the absence of a perfusable vasculature, this approach did not demonstrate scalable tissues. In this regard, Skylar et al bioprinted cardiac tissue construct utilizing iPSCs-derived organoids (figure 6(d)) [202]. Perfusable channels were embedded utilizing the sacrificial writing into functional tissue (SWIFT) approach demonstrating its scalability. In another study, 3D bioprinted collagen-based constructs were fabricated utilizing the FRESH approach to rebuild the components of heart [219]. The developed collagen scaffold allowed rapid vascularization of the tissue, while maintaining its mechanical strength. Although fully-functional organs remain a challenge to achieve, FRESH allows to recapitulate the anatomical properties of native tissues. In a recent study, in-vivo analysis of a bioprinted cardiac patch, consisting of GelMA-cardiac ECM (cECM) and human c-kit+ progenitor cells (hCPC) has shown improved vascularization and reparative functionality in the right ventricle of a rat heart [220]. Specifically, hCPC-GelMA-cECM revealed improved therapeutic effects in tissue remodeling and cardiac functional outcomes, which was shown by increased right ventricular vessel density, improved cell density, and reduced fibrosis area. However, further tissue-level analysis, such as cell types, cell density, and autologous or allogenic cell sourcing, should be optimized to improve therapeutic capability against right ventricle dysfunction. In another study, a highly elastic bioink with the capacity to induce angiogenesis is formulated using GelMA, human tropoelastin (MeTro), and gelatin for cardiac tissue replacement [221]. For bioprinting cardiovascular tissues via EBB, the bioink was encapsulated with cardiomyocytes (CMs), cardiac fibroblasts (CFs), and HUVECs. More specifically, a heterogeneous lattice construct was bioprinted using a CM/CF-laden bioink and a HUVEC-laden bioink, and its biocompatibility was proven with 85% cell viability during 7 d of culture. In addition, the expression of sarcomeric α-actinin and CD31 confirmed a mature stage of muscle differentiation and vascularization, which was also demonstrated by the endothelium barrier function. Notably, beating of cardiac cells were identified at Day 5 post bioprinting, and its coordination improved from 10 to 15 d of culture. Despite achievement of beating cardiac tissues, therapeutic effects required to be assessed using a relevant animal model. To address this, cardiac substitutes were developed using a non-mulberry silk-polyethylene glycol di-methacrylate (PEGDMA)-GelMA conductive bioink containing carbon nanotubes (CNTs) via EBB [222]. For the in-vitro study, HUVECs in the conductive bioink were bioprinted and matured in a perfusion bioreactor for 14 d, and their proliferation and maturation were displayed with increased DNA content and CD31 gene expression. Further, when neonatal rat primary cardiomyocytes were seeded, beating was observed on Day 2. For the in-vivo study, Sprague–Dawley rats with myocardial infarction sites were implanted with the bioprinted cardiac substitutes and revealed minimal immune response, wound healing, and vascular recruitment. Taken together, investigations in cardiac disease model and healing process should be conducted simultaneously for pre-clinical studies.

The liver is the largest gland in the human body and plays a crucial role in metabolism, detoxification, bile production and regulation of electrolytes and water among other functions. It contains several resident cell types, which are tightly arranged in a specific hexagonal block called the hepatic lobule, which is the fundamental building block of the liver [223]. Unfortunately, liver failure remains one of the major causes of mortality in the world including hepatitis (inflammation of the liver tissue), which is among top ten global diseases [224]. Due to its physiological complexity and involved multiple cell types, 3D bioprinting is one of the most feasible options for its regeneration. Further, to induce vascularization in bioprinted constructs, ECs can be incorporated. For example, Kang et al used ECs along with hepatic cells and created a hepatic lobule (∼1 mm) array. The developed hepatic lobule contained hepatocytes surrounded and compartmentalized by ECs with a lumen structure mimicking the vascular network of a native structure with higher albumin secretion, urea production, and albumin, and CD31 protein levels, along with greater cytochrome P450 enzyme activity [225]. Pimentel et al fabricated thick (∼1 cm) and densely populated (10 million cell/ml) tissue constructs with a 3D four-arm branch network and soft tissue-like (1–10 kPa) stiffness that can be directly perfused on a fluidic platform for a lengthy (>14 d) period of time. The encapsulated HepG2 cells demonstrated proliferation, and the microenvironment's control caused cells to aggregate in spheroids in particular locations [226]. For drug screening applications, Janani et al recapitulated the native liver architecture using (1) a human ADSC-derived hepatocyte-like cell (HLC)-laden ECM-based bioink and (2) a HUVEC/human hepatic stellate cell (HHSC)-laden ECM-based bioink [227]. The liver model comprising HLC/HUVEC/HHSC mimicked native alternate cords of hepatocytes and showed enhanced urea synthesis, albumin production, and cytochrome P450 (CPR) activity over 2 weeks. Further, drug toxicity was assessed in different conditions (i.e., non-hepatotoxicants aspirin and dexamethasone), and a follow-up metabolic evaluation demonstrated the response of DNA concentration, CPR activity, and lactate dehydrogenase activity. These results represented the possibility of clinically-relevant vascularized liver model by a dose-dependent hepatotoxic responses. Moreover, Liu et al suggested a vascularized liver platform fabricated via a hybrid system including EBB, polydimethylsiloxane (PDMS) encapsulation, perfusion, and transplantation (figure 6(e)) [203]. To begin, a bioink comprising gelatin/GelMA/fibrin embedded with MSCs and HUVECs was used as a supporting bath after crosslinking. Then, a sacrificial ink comprising gelatin/GelMA/fibrin without cells was used to bioprint perfusable channels and got removed. After seeding HUVECs into channels, PDMS was coated to protect the bioprinted structure and connect PU tubes. Perfusion was performed for 7 d, and the integrated pattern and CD31⁺ HUVECs were observed. To develop a liver model, HepG2/HUVEC/human foreskin fibroblast (HFF) aggregates were used to replace HUVECs in the supporting bath, which expressed albumin after 7 d of perfusion in vitro. At last, the perfusable liver platform was connected to the arterial vessel of Sprague–Dawley rats, in artery-to-vein mode. Although the implantation could be maintained for only a week due to thrombus formation, the perfusable liver platform suggested an alternative concept to anastomose with the host vascular network and support liver function. In short, using 3D bioprinting, structural and functional characteristics of liver lobules have been developed; however, simulating hepatocyte functions in accordance with zonal variation of the tissue along with the intricate network of blood vessels is still a challenge. Generation of perfusable vascular networks that carry oxygenated blood to hepatocytes at anatomically appropriate sites would produce biomimetic liver constructs.

Overall, advancements in the realm of bioprinting have enabled the fabrication of musculoskeletal and metabolically highly active functional tissues and organs. It has also expanded its arms in a promising concept of bioprinting directly at the site of injury to increase its effectiveness, which is discussed in the following section.