Multiscale bioprinting of vascularized models
Introduction
The engineered tissue constructs consist of self-assembled cell aggregates, synthetic materials, biological scaffolds, or porous hydrogels and fibrous meshes with macroscale characteristics that promote cell adhesion, migration, and proliferation [1]. The balance of migration and proliferation of encapsulated cells over their life cycles requires consistent delivery of nutrients and oxygen [2], in particular when constructs implanted for ischemic heart disease [3] and diabetic ulcers [4]. It is thus important to create functional blood vessel networks ranging from a few micrometers to millimeters within complex structures, such as liver, kidney, and heart [5,6] (Fig. 1). Vascular systems in vivo include a mixture of different cell types that undergo continuous remodeling by stimuli from endothelial, nervous, immune, and endocrine cells [7]. They also play a key role in graft perfusion and integration of implants into the host vascular system [8,9]. Vascular growth is further associated with developmental processes of angiogenesis, wound healing, and the progression of various pathologies such as cancer [9,10]. Despite significant improvements of current technologies to create three-dimensional (3D) blood vessels, the formation of a functional engineered vascular system with multiscale vessel networks from capillaries to large vessels has remained challenging in this field [11]. Inability in fabrication of 3D vascular networks has limited tissue engineering in the growth of thick tissue or organ-level constructs.
Many approaches have been proposed to induce the growth of a vascular system within 3D engineered tissue constructs [12]. Vascularization has been induced within tissue constructs by incorporation of growth factors into the construct [12] (which yields vasculogenesis), seeding endothelial cells in the construct to stimulate angiogenesis [13], or engineering multiscale microfluidic channels inside biocompatible materials using various microfabrication technologies [14]. Either naturally- or synthetically-derived scaffolds when mixed with proangiogenic factors yield de novo formation of capillaries [15]. In another strategy, co-cultures of endothelial cells and other cell types of the desired organ were used within tissue scaffolds prior to their implantation [16]. This strategy depends on the biological properties of endothelial cells to promote the formation of neo-capillaries in vivo. Such process can be described by remodeling phenomenon driven by the endothelial cells that are activated by the proangiogenic agents [17]. For example, in the case of human umbilical vein endothelial cell (HUVEC)-loaded collagen hydrogel implant, a tree-like structure of branched capillaries was observed over two months of implantation [17]. Several other studies have shown the ability of endothelial cells along with mesenchymal stem cells (MSCs) to form a stable vascular system under both in vitro and in vivo conditions [18,19]. In addition to these strategies, microfabrication techniques have been developed to generate perfusable channels in hydrogels [[20], [21], [22]]. Similarly, pre-implant fabrication of capillary networks into tissues constructs has shown accelerated implant integration into the host tissue as well as the formation of capillary-like structures coupled with the host vasculature [23,24]. However, these approaches lack the ability to directly generate multiscale vessel networks along with anastomosis in the host tissue, which is one of the major challenges in tissue integration.
Conventional engineering of vascular tubes includes sheet rolling and tubular molding [25]. In sheet rolling, a sheet of cellular or acellular biomaterials is rolled over a predefined mandrel and it remodels to a stable and homogeneous structure [26]. The tubular molding can also be used to shape injected biomaterials and crosslink or solidify the tube inside the annular mold [27]. These fabrication methods lack proper control on the structure and are limited to mesoscale blood vessel size (Fig. 1). Recent advancements in materials engineering and microfabrication techniques have further led to the development of new strategies [11] that offer precise control over various aspects of the cell-laden tissue implants including cell patterning [28], cell phenotype [29], cellular alignment [30], as well as the microenvironmental cues including mechanical properties [6], chemical properties (e.g. ligand density), and topographic features (e.g. cell adhesiveness) [9]. Such manipulations along with the use of 3D bioprinting techniques have been successfully applied to overcome the technical challenges of fabricating a functional microvasculature (as shown in Fig. 1 and summarized in Table 1) [6]. Bioprinting has become an essential tool for fabricating vascularized tissue constructs, in which the design of printable, tunable and biocompatible bioinks is essential [9,31]. Common examples of bioinks with proper biophysical properties and endogenous cellular cues are alginate [31,32] and gelatin methacryloyl (GelMA) hydrogels [[33], [34], [35], [36]]. In addition, the integration of time into bioprinting process known as four-dimensional (4D) bioprinting yields variations in configurational and/or biophysical characteristics of tissue constructs in response to environmental stimuli, which lead to a more realistic formation of capillaries [37]. In this review, we summarize recent 3D bioprinting techniques including sacrificial, core/shell bioprinting techniques and innovations in bioprinting approaches such as 4D bioprinting for making vascularized tissue constructs [9,11]. We then discuss how nanotechnology can be integrated into the fabrication process to advance delivery of proangiogenic factors and therapeutic compounds. Temporal controls over distributions of neo-capillaries within pre-vascularized implants may solve existing challenges in cell encapsulation of thick tissue constructs, while the designed vasculature can help anastomosis after implantation. Such vascularized models can be designed to deliver therapeutic molecules into the body circulation in a controlled fashion.
Section snippets
Sacrificial bioprinting
Classical bioprinting techniques include active printing of cellular and extracellular components in a predefined shape. Such approaches require rapid gelation or crosslinking process to provide a stable hollow tube construct [9,38]. To overcome challenges associated with printability of hydrogels and the lack of structural support in fabricating hollow tube constructs, researchers print sacrificial channels that are supported by surrounding walls and then dissolve the channels, known as
Alginate-based bioprinting
Among various bioprinting strategies, core/shell geometry seems efficient and promising in creating vascular networks (Fig. 4 A, B) [67]. The main benefit of the core/shell structure is the capacity of fabricating hierarchical, multi-layer tissue constructs with desirable biological and mechanical properties [67]. As a low-cost material, numerous bioengineers employ alginate hydrogel as a component in the design and fabrication of bioinks. Alginate is a naturally occurring, nontoxic,
Biomaterials in 4D bioprinting
The dynamic environment of hydrogel scaffolds can be designed with chemical or physical functionalities to control time dimension. 4D bioprinting has been introduced in tissue engineering due to its intrinsic features such as (i) biofabrication of 3D scaffolds with programmable architectures using stimuli-responsive materials capable of changing their shape and function in response to external stimuli, and (ii) post-maturation of printed cells within scaffolds, generating complex tissue
Concluding remark and future directions
The classical strategy to create multiscale capillaries is to fabricate a large endothelialized vascular channel within the construct and tune the surrounding microstructure to induce capillary formation from the channel. These models lack directionality of vessel formation and cannot resemble the native vessels. We have reviewed current bioprinting techniques for fabricating vascularized models (see also Table 1). The use of sacrificial strategy with extrusion bioprinting has been employed by
Acknowledgements
The authors acknowledge funding from the National Institutes of Health (EB021857, AR066193, AR057837, CA214411, HL137193, EB024403, EB023052, EB022403), and Air Force Office of Sponsored Research under award # FA9550-15-1-0273. A.K.M. acknowledges the fellowships from the Fonds de Recherche du Québec–Santé (FRQS) and Canadian Institutes of Health Research (CIHR). S.M. acknowledges the American Fund for Alternatives to Animal Research for AFAAR Postdoctoral Fellowship. S.R.S. would like to thank
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