As we move toward developing ever more complex in vitro systems, the ability to incorporate functional vasculature constitutes a critical step toward building physiologically relevant tissue and disease models. In line with this goal, recent decades have witnessed an explosive growth in the use of engineered microvascular networks (MVNs) for modeling vascularized tissues. Various techniques have been employed to generate perfusable microvasculature in vitro, including the traditional top-down approach of seeding endothelial cells (ECs) into scaffolds with pre-defined architecture and properties as well as the bottom-up strategy of allowing endothelial cells to self-organize into MVNs when exposed to the appropriate in vivo-like cues, as summarized by Song and colleagues.1 Together, these techniques have led to the design of MVNs for a wide spectrum of scales and purposes, such as recapitulating organ-specific vasculature, providing a foundation for macroscopic tissue assembly, and facilitating in vivo vascularization of bioartificial tissue upon graft implantation.2–9 Vascular models have also been applied to investigate mechanisms of physiological and pathological vessel remodeling, leukocyte–endothelial interactions in vascular diseases, vasculature-dependent tumor cell invasion and metastasis, and drug action, among other applications.10–17 

Despite these advances, there remains one major challenge that limits the scope and scale of microphysiological models that can be created: how to develop perfusable vasculature within organoids—self-organizing in vitro tissue constructs that recapitulate organ-specific structural and functional properties.18 Without a perfusable vasculature, organoids in 3D culture rely solely on passive diffusion for transport of nutrients, oxygen, and metabolic waste, which gradually leads to the development of a necrotic core. Although many have attempted to incorporate functional vasculature into organoid models (for a detailed review, see Ref. 19) to date, vasculature within organoids connecting with a perfusable external vascular network has only been demonstrated through transplantation into host animals, where native vasculature is able to penetrate into the ectopic implant or anastomose with pre-vascularized organoids. Even brain organoids, generated by promoting the development of ectoderm without any noticeable population of endothelial cells, can still be fully vascularized after xenotransplantation into mice.20 

These animal studies suggest that under appropriate in vivo-like conditions, organoids are capable of interacting with surrounding MVNs and can eventually become vascularized. Although it is impractical to replicate every aspect of the in vivo microenvironment in microfluidic systems, many groups have focused on identifying key biochemical factors that are crucial for the development of microvasculature. On the other hand, relatively less attention has been given to specific biomechanical factors, which may be as important as biochemical ones when determining the parameter space for engineering vasculature and organoids.

In vivo, a developing vasculature is subjected to various mechanical cues, ranging from viscoelastic properties of extracellular matrix (ECM), shear stresses induced by interstitial flow, to traction stresses generated by perivascular cells (Fig. 1). In addition, because endothelial cells (ECs) form the interface between circulating blood flow and surrounding tissue, they are also exposed to hemodynamic shear stress from pulsatile or unidirectional blood flow and transmural flow. Tremendous efforts have sought to understand how these mechanical factors affect the behavior of ECs,21 especially during vasculogenesis or angiogenesis, which are two major mechanisms for forming and developing vascularization. In this Editorial, we summarize the mechanical factors that are beneficial—or even indispensable—for the proper development and maintenance of functional vascular networks in vitro to provide insight into possible strategies of incorporating functional vascular perfusion of organoids.

FIG. 1.

Schematic representation of possible mechanisms and relevant mechanical factors to vascularize an organoid through co-culturing with an engineered vascular bed in a microfluidic chip.

FIG. 1.

Schematic representation of possible mechanisms and relevant mechanical factors to vascularize an organoid through co-culturing with an engineered vascular bed in a microfluidic chip.

Close modal

Based on the successful vascularization of organoids in vivo, an intuitive approach to incorporating functional vasculature within organoids in vitro is to embed organoids in an engineered vascular bed and induce angiogenesis from neighboring vessels (Fig. 1), recapitulating the natural process observed in vivo. However, with the exception of certain fibroblast spheroid and cancer models,22 efforts to vascularize organoids in a similar manner have failed. To overcome this limitation, it is important to consider the key mechanical cues present in the in vivo microenvironment that must be carried over into in vitro cultures. One mechanical factor of particular importance is fluid shear stress (FSS), which is widely established to play a critical role in regulating vascular physiology. ECs are able to detect changes in flow conditions through various mechanisms, including mechanosensory hubs located at cell–cell junctions (e.g., PECAM-1) and cell–matrix interfaces (e.g., integrin-mediated cell adhesion structures), as summarized in a review by Gordon and colleagues. These mechanosensing pathways are, thus, heavily involved in development and homeostasis as ECs experience different levels of FSS during blood vessel formation, maturation, maintenance, and remodeling.23 

In the context of linking organoids to functional vascular beds using the aforementioned strategy, angiogenesis is critical in that it enables sprouting of blood vessels from the perfused vascular bed into an avascular organoid and/or anastomosis with a pre-vascularized organoid. Flow dynamics have been shown to intricately regulate angiogenesis, with Ghaffari et al. observing that shear stress profiles and pressure gradients can affect the direction of sprouting and the rate of sprout elongation during development.24 Galie and colleagues further demonstrated that, above a certain shear stress threshold, both luminal flow and transmural flow can induce angiogenic sprouting, and continuous flow is needed to sustain sprouting and prevent retraction.25 Importantly, given that the successful vascularization of organoids requires extensive remodeling of the vascular bed, long-term, continuous perfusion of luminal flow through MVNs may be an essential factor.

The effect of FSS on developing organoids has also become increasingly recognized. To date, perhaps the most compelling demonstration that FSS can impact organoid vascularization comes from Homan and colleagues, who showed that subjecting human pluripotent stem cell-derived kidney organoids to external FSS in microfluidic chips leads to enhanced vascularization, as demonstrated by an increase in vessel area coverage, vessel length, and junctional density and including sprouting of the internal network outside of the organoid.26 While there is no strong evidence to support the formation of consistently and fully perfusable vessels in this system, this study reveals that FSS should be considered as a significant factor in generating culture systems with fully vascularized organoids.

Taken together, these findings suggest that successfully connecting organoids to a functional vascular bed may in large part rely on our ability to both tune and sustain defined flows over long time scales in culture systems. Toward this goal, one area of active exploration is the development and integration of new tools for inducing and controlling FSS in engineered MVNs. For example, the design of mechanical pumping systems can support continuous circulation of cell culture media to produce luminal or transmural flow at defined physiological levels.27,28 As the tissue engineering field increasingly turns to microfluidic devices to model more complex biological systems in high-throughput formats, innovation in microphysiological systems will play a key role in optimizing fluid flow for promoting functional vascularization of organoids seeded onto a vascular bed.

Interstitial flow is the movement of fluid through the extracellular matrix of tissues. It provides a necessary transport mechanism through the interstitium and constitutes an important component of microcirculation. In particular, interstitial flow plays a critical role during vasculogenesis—before perfusable vasculature is established—by supplying nutrients and oxygen, removing metabolic waste, and providing additional mechanical cues to cells. Various microfluidic platforms have been engineered to investigate the effects of interstitial flow on ECs and vessel formation with most studies reporting a beneficial role of physiological interstitial flow on angiogenesis, vasculogenesis, and 3D capillary morphogenesis in vitro.29,30

As such, another possible strategy for vascularization is to seed organoids and endothelial cells together within a hydrogel scaffold for co-development under the action of interstitial flow. In order to self-organize into functional MVNs in such a system, the endothelial cells must be able to intensively remodel the surrounding ECM and interact with the embedded organoid. In this case, interstitial flow could benefit the formation of self-organized MVNs, possibly through upregulating matrix metalloproteinase (MMP) production.25,30 Indeed, using both generic MVNs formed with human umbilical vein endothelial cells and human lung fibroblasts as well as brain-specific MVNs that include brain endothelial cells, pericytes, and astrocytes, as model systems, colleagues have demonstrated that interstitial flow enhances vessel formation via significant remodeling of vessels, leading to functional MVNs with increased vessel area, diameter, connectivity, and longevity.31,32 With the careful characterization of matrix permeability of the hydrogels employed to co-culture organoids and endothelial cells, desirable interstitial flow profiles can be established by tuning parameters of aforementioned pumping systems, or even with gravity-driven flow.

Similar to their ability to activate mechanosensing pathways in response to FSS, ECs are also able to sense their local mechanical microenvironment through focal adhesions. Mechanical stimuli propagate through the actin cytoskeleton, resulting in cytoskeletal remodeling and subsequent adaptation of the endothelial cells to their microenvironment.33 One important factor in this process is matrix stiffness, which can in addition alter how ECs respond to soluble, angiogenic factors released by stromal cells such as vascular endothelial growth factor. Most studies performed on 2D deformable substrates have reported that decreasing substrate stiffness promotes network formation. However, results differ when endothelial cells are cultured within 3D hydrogels, where an intermediate hydrogel stiffness in the range around 500 to a few thousands Pa has been shown to lead to increased spreading and sprouting from embedded ECs.34,35

In most hydrogel systems, stiffness is modulated by changing polymer density, which simultaneously alters other properties such as pore size and ligand density. In recent years, several groups have successfully decoupled hydrogel stiffness from matrix density via various methods and confirmed the beneficial role of hydrogel with intermediate stiffness of a few hundred Pa for angiogenesis and vessel formation.36,37 However, the effect of matrix stiffness on development of internal vasculature within organoid remains a relatively unexplored area. Considering the fact that matrix stiffness is known to affect downstream integrin signaling pathways and plays a role in determining stem cell fate and growth of organoid,38,39 it follows that the mechanical properties of ECM will likely influence the formation of internal vasculature within organoid as well.

Since almost all biological tissues are viscoelastic, the time-dependent mechanical response of the ECM also needs to be considered in order to faithfully recapitulate the in vivo like microenvironment. To date, however, the impact of viscoelastic properties of ECM on embedded ECs and the resultant formation of functional vessels have not been thoroughly investigated. Currently, natural materials are most commonly employed for culturing organoids and vascular networks, because they intrinsically possess desirable viscoelastic properties, in addition their excellent cell affinity, biocompatibility, and ease of use. It must be noted though historically, organoids and MVNs have usually required different types of natural ECMs. Many organoid protocols involve embedding in Matrigel at a certain stage for proper development of organoid but forming functional vasculature within Matrigel remains challenging with inconsistent results.40 Instead, in vitro MVNs are generally grown within collagen or fibrin, where ECs undergo local matrix remodeling to form patent lumens.41,42 For all of these natural matrices however, the range of stiffness that can be achieved is relatively narrow in addition to other challenges such as batch-to-batch variation and low modularity. In contrast, synthetic materials can be utilized to generate hydrogels with a much wider range of matrix stiffness and other properties. In addition, small molecules and soluble factors can be incorporated into synthetic hydrogels via covalent interactions in a controllable manner to independently modulate various mechanical and biochemical factors in the microenvironment. While synthetic hydrogels often exhibit a nearly linear elastic behavior, some novel cross-linking strategies have been used to introduce a viscoelastic response.43–45 

Degradability of hydrogels also significantly affects the behaviors of embedded ECs as they self-organized into functional MVN. Within natural hydrogels, ECs actively remodel the surrounding ECM to connect with each other and form patent lumens with the initial matrix gradually degraded and replaced by cell-generated ECM. As for synthetic hydrogels, novel cross-linking methods enable the replication of this process in a tunable manner. For instance, through the incorporating engineered MMP-degradable peptide crosslinkers in the hydrogel precursor solution, it is possible for ECs to undergo vasculogenesis and angiogenesis, thus remodeling their ECM toward formation of functional MVN.46–48 Although synthetic hydrogels require further innovation and optimization, the defined compositions of synthetic polymers enable the realization of precise and tunable mechanical conditions to regulate behaviors of embedded cells in vitro. As such, improved synthetic hydrogels could potentially offer optimized in vivo-like microenvironments to promote vascular bed formation and boost the development of internal vasculature within embedded organoids.

Mechanical factors, such as fluid shear stress, interstitial flow, and extracellular matrix properties (Fig. 1), have been increasingly appreciated as key players in regulating endothelial cell behavior and vessel formation. Incorporating all these factors into microphysiological models promises to enhance the physiological relevance of engineered in vitro microenvironments. By culturing organoids and MVNs in such mechanically defined conditions, combined with improving our understanding of the precise impacts of those factors, we anticipate that in vitro organoid models with physiologically realistic intravascular and interstitial perfusion will be achieved in the near future.

R.D.K. is a co-founder and a board member of AIM Biotech. He also has current research support from Boehringer-Ingelheim, Roche, Amgen, GSK, and Novartis.

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