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Perspective Article
to inter-donor variation and ethical concerns. Synthetic arterial models, produced by seeding cells on a biodegrad- able tubular scaffold, can overcome these issues.29 To date, this model has not been perfused with blood to investigate blood-endothelial interactions.
Microvascular models involve the formation of <200 mm channels in polydimethylsiloxane, subendothelium or hydrogels.9,30,31 These channels are available in different branching geometries and can be lined by endothelial cells.31,32 Acellular models can accommodate smaller ves- sel diameters (5-70 mm) and higher hematocrits (up to 0.4) without blockage.30 In contrast, endothelialized microvas- cular models are limited by vessel diameters of 50-200 mm with 10-20% (v/v) packed RBC perfusates.9,31 Vessels less than 50 mm in diameter are difficult to produce in hydrogels, while vessels greater than 200 μm present issues during endothelial seeding and culture which pre- clude the formation of a confluent monolayer. Manufacturing limitations associated with vessel diame- ter can be abrogated by using in vivo microvascular mod- els such as muscle or dorsal skin flaps. However, these models lack the channel diameter and geometrical consis- tency resulting from in vitro-manufacturing techniques. The presence of endothelial cells in microvascular models enhances thrombosis, thereby hindering the hematocrits that can be tested. Physiological hematocrits are impor- tant for viscosity, shear stress and cell-to-cell interactions. For example, an increased hematocrit leads to more RBC- endothelial interactions which enhance shear stress at low flow rates. Furthermore, an increased hematocrit has been associated with altered RBC arrangement in the flow stream and margination of blood components.33
Designing vascular models
Perfusable in vitro macrovascular and microvascular models are composed of an endothelialized channel in a hydrogel scaffold. The unit is subsequently connected to a circuit and perfused with the desired “test” fluid. There are five key decision points in vascular model design (Figure 1): (i) channel geometry, (ii) scaffold moulding, (iii) endothelial cell seeding, (iv) circuit construction, and (v) perfusate selection.
The choice of channel size and shape determines the scaffold manufacturing method. Straight channels with diameters down to 120 mm can be injection-molded using rods, needles or wires stabilized on a mount. This method is cost-effective and technically easy to construct but has limited fidelity for endogenous vascular geome- try. Branched channel geometries with diameters of 50- 200 mm can be formed using a mixture of photolithogra- phy and soft lithography methods. Photolithography is used to create a hard negative-profile wafer. Complex channel geometries can be etched onto a photomask which is used to develop the desired pattern on a pho- toresist-coated crystalline silicon wafer.34 Notably, photo- lithography requires microfabrication facilities (e.g. clean room, photo pattern generator, spin coater) and special- ized consumables (e.g. crystalline silicon wafer, photore- sist, developer). The negative-profile wafer becomes the mold for the soft positive-profile stamp used to shape the hydrogel scaffold by soft lithography.31 The soft positive- profile stamp can be formed from polydimethylsiloxane
in any standard laboratory using a vacuum degasser and oven. Polydimethylsiloxane can be printed directly in three dimensions: however, this method has inferior res- olution compared to the photo/soft-lithography combi- nation (~760 mm versus 50 mm, resolution limited by soft lithography).35
The hydrogel scaffold can be constructed from colla- gen, alginates, agarose, poly(ethylene glycol) dimethacry- late or methacrylated gelatin.5 These materials are chosen for their transparency, absence of toxicity, fidelity for micropatterning and transport properties. Perivascular and/or tissue cells can be incorporated into the hydrogel to enhance biological approximation. The hydrogel is subsequently injection-molded onto a stamp and allowed to solidify at specified temperatures. Straight channel scaffolds can be formed as one piece. Scaffolds involving the use of a stamp are formed as two pieces (stamped pieced, flat slab), which are subsequently combined to form closed channels. The key obstacles at this stage are bubble formation during injection molding and channel damage during scaffold assembly.
The next decision involves the type of endothelial cells to use to seed the channels. Commercially available multi-donor human umbilical vein endothelial cells are the most commonly used cell type due to their availabili- ty, robustness and proliferation consistency. However, multi-donor cells can express multiple antigenic profiles, complicating donor-recipient cross-matching in transfu- sion simulations. Furthermore, human umbilical vein endothelial cells may behave differently from other pri- mary endothelial cells. Vessel-specific primary endothe- lial cells such as human lung microvascular endothelial cells and human aortic endothelial cells may improve physiological relevance but these cells are often more dif- ficult to culture and susceptible to variability. Endothelial derivatives from human pluripotent stem cells can also be used.36 The desired endothelial cells can be seeded on the channel surface under static or low flow conditions and grown to confluence under laminar flow. Notably, high seeding flow rates will prevent endothelial cells from attaching to the hydrogel scaffold. We recommend seed- ing endothelial cells under static conditions and allowing up to 3 hours for the cells to attach, prior to incubating the cells overnight under laminar flow conditions. Endothelial cell growth under laminar flow is important to cultivate the barrier function and monolayer present in endogenous blood vessels.
Variables in circuit construction include circulation vol- ume, tubing material and diameter, gas trap and pump. The circulation volume can be adjusted to accommodate the desired outcome measures (e.g. flow cytometry, enzyme-linked immunosorbent assay, mass spectrome- try, biochemical analysis). Tubing length and diameter can be varied to reach the desired circulation volume. Notably, the more tubing is used to maximize circulation volume, the greater the ratio between non-endothelial- ized and endothelized surface area in the circuit. Medical grade tubing made out of Tygon, fluoropolymers or polyvinylchloride are used as they are inert, biocompati- ble and gas permeable. In our set-up, the gas trap is con- nected just before the inlet port of the vascular model to minimize bubbles passing through the endothelialized
haematologica | 2019; 104(3)