This work was supported by the National Natural Science Research Foundation of China Grants-in-Aid (grant nos. 11827803, 31971244, 31570947, 11772036, 61533016, U20A20390 and 32071311), National key research and development program of China (grant nos. 2016YFC1101101 and 2016YFC1102202), the 111 Project (B13003), and the Beijing Natural Science Foundation (4194079).

1. Wafer preparation
NOTE: This protocol is specific for the SU-8 2075 negative photoresist used during this research.
<ol>
	<li>Clean the silicon wafer 3 to 5 times with methanol and isopropanol on a spin coater as follows: first spin for 15 s at 500 rpm, and then spin for 60 s at 3,000 rpm.</li>
	<li>Transfer the silicon wafer to a hotplate, which is preheated to 180 &#176;C and bake the wafer for 10 min.</li>
	<li>Remove the silicon wafer from the hotplate and cool it to room temperature. Clean the wafer again with compressed air before proceeding with the spin coating. Apply 4 mL of the SU-8 2075 photoresist to the center of the wafer.</li>
	<li>Obtain a feature height of 70 &#181;m on the spin coater as follows: first spin for 12 s at 500 rpm, and then spin for 50 s at 2,100 rpm.</li>
	<li>Soft bake the wafer on a hotplate as follows: first bake for 8 min at 65 &#176;C, then bake for 20 min at 95 &#176;C. Next, remove the silicon wafer from the hotplate and cool it to room temperature before lithography.</li>
	<li>Place a photomask onto the photoresist film. Expose the wafer with a lithography equipment to achieve a total exposure of 200 mJ/cm2. 
	NOTE: The photomask contains nine sets of patterns of the chip, so it can fabricate nine microfluidic sprouting chips each time.</li>
	<li>Post expose the wafer on a hotplate as follows: first bake for 5 min at 65 &#176;C, and then bake for 20 min at 95 &#176;C. Next, remove the silicon wafer from the hotplate and cool it to room temperature before developing.</li>
	<li>Transfer the wafer to a glass Petri dish filled with SU-8 developer (PGMEA) to start developing. 
	CAUTION: The developer is irritating to the eyes and respiratory tract. Perform developing in a fume hood. Wear splash goggles, nitrile gloves, and airline mask during the operation.</li>
	<li>Shake the dish gently along the direction of the flow channel and change the developer after 10 min.</li>
	<li>Repeat step 1.9 3-5 times until the patterns can be clearly observed. 
	NOTE: Make sure there is no photoresist residue left on the wafer, otherwise rinse the wafer in the developer again.</li>
	<li>Transfer the wafer to a preheated hotplate set to 120 &#176;C and bake it for 30 min.</li>
	<li>Pipet 35 &#181;L of silane onto a coverslip,&#160;then put the coverslip along with the wafer into a desiccator and pull vacuum. Seal the desiccator and leave the wafer under vacuum for 4 h to silanize the wafer to prevent the adhesion of PDMS during the soft-lithography processes. 
	CAUTION: The silane is toxic. To prevent poisoning, perform silanization in the fume hood and wear nitrile gloves while handling.</li>
	<li>Release the vacuum from the desiccator. Remove the silanized wafer onto a preheated hotplate and bake it at 65 &#176;C for 2 h.</li>
	<li>Store the wafer in a clean Petri dish until required.</li>
</ol>
2. Microfluidic sprouting chip fabrication
<ol>
	<li>Combine 20 g of base agent and 2 g of curing agent (10:1 ratio) in a plastic beaker and mix them thoroughly with a mixing rod.</li>
	<li>Place the beaker into a desiccator and pull vacuum for 1 h to remove air bubbles in the PDMS mixture.</li>
	<li>Pour the PDMS mixture onto the wafer in the Petri dish and place the Petri dish back into the desiccator, degassing for another 30 min. 
	NOTE: It is helpful to use double-sided adhesive tape to glue the wafer onto the bottom of the Petri dish to ensure the wafer is kept horizontal during degassing and curing.</li>
	<li>Remove the Petri dish from the desiccator and place it into an 80 &#176;C dry oven for 3 h to cure.</li>
	<li>Carefully separate the PDMS layer from the wafer and cut the layer to nine chips with a scalpel according to the pattern. 
	NOTE: Keep the feature side up after separation.</li>
	<li>Punch two hydrogel injection ports and four media injection ports out of each chip using a 1 mm and a 3.5 mm biopsy punch, respectively.</li>
	<li>Clean the punched chips with residue-free tape to remove PDMS residue. Place the chips and nine glass coverslips into a plasma cleaner and treat them with oxygen plasma for 30 s to form covalent bonding on the surface.</li>
	<li>Take out the chips and coverslips. Attach the feature side of the chips onto the coverslips.</li>
	<li>Place the attached chips into an 80 &#176;C dry oven for 1 h to intensify the bonding.</li>
	<li>Autoclave the chips before use and keep them sterile for the rest of the procedure.</li>
</ol>
3. Surface modification and hydrogel injection
<ol>
	<li>For each chip, pipet 40 &#181;L of 1 mg/mL poly-D-lysine (PDL) and inject it into channels in the chip from hydrogel injection port. 
	NOTE: Make sure the PDL fills channels, especially in narrow channels near the ports.</li>
	<li>Incubate the chips at 37 &#176;C for 4 h to modify the surface of the PDMS .</li>
	<li>Pipet 200 &#181;L of sterile water and inject it into channels in the chip from hydrogel injection port to wash out PDL.</li>
	<li>Remove the chips into a 120 &#176;C dry oven for 3 h to restore hydrophobicity.</li>
	<li>Place on ice a sterile tube and calculate the volume of Type I collagen to be used as the following equation. 
	Final volume (50 &#181;L) x Final collagen concentration (3 mg/mL in this research) / Concentration in bottle = volume collagen to be added</li>
	<li>Calculate the volume of NaOH to be used as the following equation. 
	(volume collagen to be added) x 0.023 = volume 1 N NaOH</li>
	<li>Prepare 50 &#181;L of hydrogel in the tube as the following protocol: add 5 &#181;L of 10x PBS, 0.5 &#181;L of phenol red, calculated volume of 1 N NaOH, 4 &#181;L of 1 mg/mL Fibronectin, calculated volume of Type I collagen in turn. Add the proper amount of dH2O so that the total volume reaches 50 &#181;L. Mix all the contents in the tube thoroughly. 
	NOTE: Perform all the operations on ice. Use phenol red as an acid-based indicator to assist in visual determination of the pH of the hydrogel. The final hydrogel ends up orange when the pH is about 7.4 and the stiffness is about 15 kPa49.</li>
	<li>Pipet 2-3 &#181;L of hydrogel for each chip and slowly inject it into the central hydrogel channel from hydrogel injection port.</li>
	<li>Incubate the chips at 37 &#176;C for 30 min to allow gelation. 
	&#8203;NOTE: Seal the chips into a sealed box with 1 mL of sterile water if they are not used immediately. Sealed chips can be stored at 37 &#176;C for at most 24 h.</li>
</ol>
4. Cell seeding
<ol>
	<li>Pipet 20 &#181;L of 125 &#181;g/mL Fibronectin into one media injection port of the cell culture channel.</li>
	<li>Cut a pipette tip to fit the port of the cell culture channel with scissors.</li>
	<li>Insert the pipette tip into the other media injection port of the cell culture channel. Then, pipet out air from the cell culture channel to fill it with Fibronectin.</li>
	<li>Incubate the chips at 37 &#176;C for 1 h.</li>
	<li>Before cell seeding, pipet 20 &#181;L of ECM media into each media injection port and incubate the chips at 37 &#176;C for 30 min.</li>
	<li>Then, pipet out all the media in all media injection ports.</li>
	<li>Next, pipet 5 &#181;L of cell suspension into one media injection port of cell culture channel. Then, endothelial cells quickly spread over the entire channel under differential hydrostatic pressure. 
	NOTE: Prepare the cell suspension by trypsinizing human umbilical vein endothelial cells (HUVECs) from the culture flask and centrifuging them at 400 x g. Then, resuspend the cells to 107 cells/mL in a tube.</li>
	<li>Add about 4-6 &#181;L of ECM media to the other port to adjust the hydrostatic pressure and stop cell moving.</li>
	<li>Remove the chips to the cell incubator. Then, turn over the chips every 30 min until endothelial cells coat around the internal surface of the cell culture channel 2 h later (Figure 1). 
	NOTE: To make the chip upside down, pipet a little water on the back of the coverslip. Then the chip can attach to the cover of the Petri dish.</li>
	<li>Use the pipette tip to remove the attached cells in the injection ports very carefully.</li>
	<li>Then, insert four barbed female Luer adaptors into the media injection ports and fill with ECM media. 
	&#8203;NOTE: The adaptors can function as fluid reservoirs to provide nutrients for the cells in the channel.</li>
	<li>Remove the chips to cell incubator. Change ECM media in Luer adaptors every 12 h.</li>
</ol>
5. Measurement of FITC-dextran diffusional permeability
NOTE: To assess barrier function of the micro-vessel, diffusional permeability of the EC culture channel with or without cell lining is assessed.
<ol>
	<li>Take out a microfluidic sprouting chip with hydrogel injected.</li>
	<li>Repeat steps 4.1-4.6.</li>
	<li>Remove all the Luer adaptors and pipet out all the media in four media injection ports.</li>
	<li>Place the chip onto the confocal laser scanning microscope.</li>
	<li>Pipet 5 &#181;L of culture media containing 500 &#181;g/mL 40 kDa FITC-dextran to one port of cell culture channel. 
	NOTE: 40 kDa FITC-dextran has similar molecular size to VEGF-165 (39-45 kDa).</li>
	<li>Capture images every 3 s for 30 s. Thus, the diffusional permeability without cell lining is measured.</li>
	<li>Take out the microfluidic sprouting chip after HUVECs are confluent in the cell culture channel.</li>
	<li>Repeat steps 5.3-5.5.</li>
	<li>Capture images every 3 s for 30 s. Thus, the diffusional permeability with cell lining is measured.</li>
	<li>Calculate the diffusional permeability by quantifying changes of fluorescent intensity over time using the following modified equation50. 
	Pd= (I2- I1) / ((I1- Ib)&#183;&#916;t) &#183;S/d 
	where, Pd is the diffusional permeability coefficient, I1 is the average intensity at an initial time point, I2 is the average intensity after delta time (&#916;t), Ib is the background intensity, S is the area of the channel in fluorescence images, and d is the total interval of micro-posts in fluorescence images. In the present work, &#916;t is set as 9 s. 
	&#8203;NOTE: The present equation is slightly different from the original50, due to the diffusion direction of the fluorescence in the chip is only from the EC channel to hydrogel channel, which is not like the circular tube diffusing in all the radial direction.</li>
</ol>
6. Microfluidic control system setup
NOTE: The microfluidic control system in the present study is consisted of a micro-syringe pump, an electromagnetic pinch valve, a bubble trap chip, a microfluidic chip, a micro-peristaltic pump, and a reservoir. Each part of the system can be replaced by alternatives able to perform the same function.
<ol>
	<li>Bubble trap chip fabrication 
	NOTE: The bubble trap chip is used to remove air bubbles in circulation. The chip consists of three PDMS layers. The top layer is constructed using soft lithography to form grid structure as the liquid chamber. Each channel of the grid structure is 100 &#181;m wide. The bottom layer is a PDMS chunk with a hole. Between the two layers, a 100 &#181;m thin PDMS film is laid.
	<ol>
		<li>Repeat step 1 to prepare the wafer for bubble trap chip.</li>
		<li>Repeat steps 2.1-2.5 to fabricate the top layer and bottom layer of the bubble trap chip. 
		NOTE: The photomask of the top layer contains three sets of patterns, so cut the PDMS layer to three chips according to the pattern.</li>
		<li>Punch six holes on the end of inlet and outlet channels of the top layer using a 3.5 mm biopsy punch.</li>
		<li>Punch two holes on the corresponding position of bottom layer using a 6 mm biopsy punch according to the pattern on the top layer.</li>
		<li>Repeat steps 2.1-2.2 to prepare 10 g of PDMS mixture and pour it on the center of a cleaned silicon wafer.</li>
		<li>Fabricate a 100 &#181;m PDMS film by applying the following spin protocol: spin for 15 s at 500 rpm, increase the spin speed to 1,300 rpm and hold here for 45 s.</li>
		<li>Transfer the wafer to a preheated hotplate set to 180 &#176;C and bake it for 30 min.</li>
		<li>Clean the punched layers with residue-free tape to remove PDMS residue. Place the top layers and wafer into a plasma cleaner and treat them with oxygen plasma for 30 s to form covalent bonding on the surface.</li>
		<li>Take out top layers and wafer. Attach the feature side of top layers onto the PDMS film.</li>
		<li>Cut the film carefully along the edge of the top layers with a needle.</li>
		<li>Slowly separate the film from the wafer and turn over the chips to make the film side up after separation.</li>
		<li>Place the bottom layers and attached chips into a plasma cleaner and treat them with oxygen plasma for 30 s again.</li>
		<li>Take out bottom layers and attached chips. Attach the bottom layers onto the PDMS film and the bubble trap chips are done. 
		NOTE: Align the holes on the bottom layer with the patterns on the top layer when attaching.</li>
		<li>Place the chips into an 80 &#176;C dry oven for 1 h to intensify the bonding.</li>
	</ol>
	</li>
	<li>Assemble the microfluidic control system 
	NOTE: All the parts of the system, such as bubble trap chip, reservoir, tubes, and connectors are used after autoclave sterilization, except electronic equipment. Assemble the microfluidic control system on a clean bench.
	<ol>
		<li>To assemble the microfluidic control system, prepare two polytetrafluoroethylene tubes, two short silicone tubes, three long silicone tubes, one barbed female Luer adaptor, one Y type connector, and three L type connectors. 
		NOTE: The advantage of polytetrafluoroethylene tube is low elasticity, therefore, less media is needed in pipeline. Silicone tube, in contrast, has high elasticity, therefore, it is suitable for the pinch valve and peristaltic pump.</li>
		<li>Fill the syringe with 10 mL of preheated (37 &#176;C) ECM medium. 
		NOTE: Preheating helps the medium release dissolved gas.</li>
		<li>Connect a polytetrafluoroethylene tube to the syringe by a barbed female Luer adaptor. Then, connect the other end of polytetrafluoroethylene tube to a Y type connector.</li>
		<li>Next, connect two long silicone tubes to the other two ends of Y type connector with one tube connecting to the reservoir and the other tube connecting to the bubble trap chip.</li>
		<li>Connect another long silicone tube to the reservoir.</li>
		<li>Next, use two short silicone tubes to connect all inlet and outlet holes on the top layer of the bubble trap chip. Connect a polytetrafluoroethylene tube to the backend of the chip.</li>
		<li>Next, fix the syringe onto the micro-syringe pump.</li>
		<li>Clip two long silicone tubes into the electromagnetic pinch valve.</li>
		<li>Next, switch the electromagnetic pinch valve to open the pipeline between the syringe and the reservoir. Inject the media to the reservoir using a micro-syringe pump to exhaust air in the tube.</li>
		<li>Then, switch the valve again to open the pipeline between the syringe and the bubble trap chip. Inject the media to fill the liquid chamber and the backend tube of the bubble trap chip.</li>
	</ol>
	</li>
</ol>
7. Endothelial sprouting assay
NOTE: A stage top incubator assembled with phase contrast microscope is used in the present study to observe the process of sprouting in real time. The stage top incubator can maintain the temperature, humidity, and CO2 control on microscope stages, being good for live cell imaging. But the equipment is not necessary for the assay. The protocols provided here can also be worked in a basic cell incubator.
<ol>
	<li>Take out endothelial sprouting chips from the cell incubator.</li>
	<li>Then, remove Luer adaptors on the cell culture side. Insert two pipe plugs into the hydrogel injection ports of the microfluidic sprouting chip. 
	NOTE: The plugs can transform from needles and their main function is to prevent media from spilling.</li>
	<li>Connect the backend tube of bubble trap chip to one port of cell culture channel. 
	NOTE: Make sure there are no bubbles in the tube before connection.</li>
	<li>Insert a T type connector to the other port and connect it to long silicone tube connected with the reservoir.</li>
	<li>Clip the long silicone tube into the micro-peristaltic pump.</li>
	<li>Then, insert an air filter to the reservoir.</li>
	<li>Next, assemble the microfluidic sprouting chip to the stage top incubator.</li>
	<li>Next, connect the vacuum pump to the holes in the bottom layer of bubble trap chip with a TPU tube.</li>
	<li>Set up the circulation volume and flow rate in the custom program, which controls the micro-syringe pump and electromagnetic pinch valve simultaneously (Figure 2). 
	NOTE: The flow rate across the endothelial cell culture channel is calculated according to the classic equation. 
	&#964;&#160;= 6&#181;Q / Wh2 
	where, &#964; is shear stress (dyn/cm2), &#181; is viscosity of the medium (8.8 x 10-4 Pa&#8226;s), Q is the flow rate across the endothelial cell culture channel (ml/s), h is channel height (70 &#181;m), and W is channel width (1,000 &#181;m). The viscosity of the medium is measured using a coaxial cylinder type rotational viscometer. The height and width of the channel are predetermined and manually confirm using a phase contrast microscope at 4x magnification. The circulation volume is 5 mL and flow rate is 85 &#181;L/min (average 0.2 m/s in cell culture channel) for 15 dyn/cm2 shear stress according to calculation.</li>
	<li>Next, set up the flow rate of micro-peristaltic pump. 
	NOTE: The flow rate of micro-peristaltic pump is slightly higher than micro-syringe pump, in order to prevent the media from spilling.</li>
	<li>Turn on the micro-syringe pump. Then, the circulation control system is established.</li>
</ol>
8. Data analysis
NOTE: To quantify the sprouts, the normalized area of sprouting, average sprout length, and longest sprout length were calculated. Results represent mean &#177; SEM obtained from three independent studies. Statistical significance (P &#60; 0.05) is assessed by Student&#39;s t-test.
<ol>
	<li>Fix cells and sprouts in the microfluidic sprouting chip after experiments with 4% paraformaldehyde in PBS for 15 min. Stain nuclei with 4&#8242;,6-diamidino-2-phenylindole dihydrochloride (DAPI; 1: 1000; Sigma-Aldrich) for 10 min and then stain cytoskeleton with TRITC phalloidin (P5285, 1: 100; Sigma-Aldrich) for 1 h. Wash cells with PBS three times at 5 min intervals between each step.</li>
	<li>Take confocal images of the chips in a tiling mode and stitch them using image editing software.</li>
	<li>Count the number of fluorescent pixels of Z-projection images using a custom code in programming software to quantify normalized area of sprouting (See Supplemental File).</li>
	<li>Manually identify and label each tip of sprouts in Z-projection images and calculate distances between sprouts tips to ECs basement membrane using another custom code to quantify average sprout length and longest sprout length (See Supplemental File).</li>
</ol>

The authors have nothing to disclose.

For a long time, real-time observation of neovascularization has been a problem. Several approaches have been developed recently to create perfused vessels lining with ECs and adjacent to extracellular matrix for sprouting22,32,40,46,54, but the mechanical microenvironment is still hard to maintain constantly. It remains a difficult subject to mimic the initia...

Neovascularization happens in many normal and pathological processes1,2,3,4, which include two major processes in adults, angiogenesis and arteriogenesis5. Besides the best-known growth factors, such as vascular endothelial growth factor (VEGF)6, mechanical stimulations, in particular the blood flow induced shear stress, is important in the regulation of neovascularization7. As we know, the magnitude and forms of shear stress vary dramatically and dynamically in different parts of the vasculature, resulting in important effects on vascular cells8,9,10,11,12. Previous studies have shown that shear stress may affect various aspects of ECs, including cell phenotypic changes, signal transduction, gene expression, and the communication with mural cells13,14,15,16,17,18,19,20; hence, regulate neovascularization21,22,23,24.
Therefore, to better understand neovascularization, it is important to reconstruct the process in natural cellular microenvironment in vitro. Recently, many models have been established to create micro-vessels and provide precise control of microenvironment25,26,27, taking advantage of advances in microfabrication and microfluidic technology. In these models, micro-vessels can be generated by hydrogel28,29, polydimethylsiloxane (PDMS) microfluidic chips30,31,32 or 3D bioprinting33,34. Some aspects of the microenvironment, such as luminal shear stress22,23,35,36, transendothelial flow37,38,39,40, biochemical gradient of angiogenic factors41,42, strain/stretch43,44,45, and co-cultured with other types of cells32,46 have been mimicked and controlled. Usually, a large reservoir or syringe pump was used to provide perfused medium. Transendothelial flow in these models was created by pressure drop between the reservoir and micro-tube22,23,38,40. However, the mechanical microenvironment was hard to maintain constantly in this way. Transendothelial flow would increase and then exceed the physiological level if a high flow rate with high shear stress was used for perfusion. Previous study showed that at the initial period of neovascularization, the velocity of transendothelial flow is very low due to the intact ECs and basement membrane, usually under 0.05 &#181;m/s8. Meanwhile, though luminal shear stress in vascular system varies greatly, it is relatively high with mean values of 5-20 dyn/cm2,11,47. For now, the velocity of transendothelial flow in previous works have been generally kept between 0.5-15 &#181;m/s22,38,39,40, and the luminal shear stress was usually under 10 dyn/cm2&#160;23. It remains a difficult subject to constantly expose ECs to high luminal shear stress and physiological level of transendothelial flow simultaneously.
In the present study, we describe an in vitro 3D model to mimic the initial event of neovascularization (MIEN). We developed a microfluidic chip and an automatic control, highly efficient circulation system to form perfusion micro-tubes and simulate the process of sprouting48. With the MIEN model, the microenvironment of ECs stimulated at the initial period of neovascularization are firstly recapitulated. ECs can be stimulated by high luminal shear stress, physiological level of transendothelial flow and various VEGF distribution simultaneously. We describe the steps of establishing the MIEN model in detail and the key points to be paid attention to, hoping to provide a reference for other researchers.

<table border="1" fo:keep-together.within-page="1" fo:keep-with-next.within-page="always">
	<tbody>
		<tr>
			<td>0.25% Trypsin-EDTA</td>
			<td>Genview</td>
			<td>GP3108</td>
			<td></td>
		</tr>
		<tr>
			<td>Collagen I, rat tail</td>
			<td>Corning</td>
			<td>354236</td>
			<td></td>
		</tr>
		<tr>
			<td>DAPI</td>
			<td>Sigma-Aldrich</td>
			<td>D9542</td>
			<td></td>
		</tr>
		<tr>
			<td>Electromagnetic pinch valve</td>
			<td>Wokun Technology</td>
			<td>WK02-308-1/3</td>
			<td></td>
		</tr>
		<tr>
			<td>Endothelial cell medium (ECM)</td>
			<td>Sciencell</td>
			<td>1001</td>
			<td></td>
		</tr>
		<tr>
			<td>Fetal bovine serum (FBS)</td>
			<td>Every Green</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Fibronectin</td>
			<td>Corning</td>
			<td>354008</td>
			<td></td>
		</tr>
		<tr>
			<td>FITC-dextran</td>
			<td>Miragen</td>
			<td>60842-46-8</td>
			<td></td>
		</tr>
		<tr>
			<td>Graphical programming environment</td>
			<td>Lab VIEW</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Image editing software</td>
			<td>PhotoShop</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Image processing program</td>
			<td>ImageJ</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Isopropanol</td>
			<td>Sigma-Aldrich</td>
			<td>91237</td>
			<td></td>
		</tr>
		<tr>
			<td>Lithography equipment</td>
			<td>Institute of optics and electronics, Chinese academy of sciences</td>
			<td>URE-2000/35</td>
			<td></td>
		</tr>
		<tr>
			<td>Methanol</td>
			<td>Sigma-Aldrich</td>
			<td>82762</td>
			<td></td>
		</tr>
		<tr>
			<td>Micro-peristaltic pump</td>
			<td>Lead Fluid</td>
			<td>BT101L</td>
			<td></td>
		</tr>
		<tr>
			<td>Micro-syringe pump</td>
			<td>Lead Fluid</td>
			<td>TYD01</td>
			<td></td>
		</tr>
		<tr>
			<td>Oxygen plasma</td>
			<td>MING HENG</td>
			<td>PDC-MG</td>
			<td></td>
		</tr>
		<tr>
			<td>Paraformaldehyde</td>
			<td>Sigma-Aldrich</td>
			<td>P6148</td>
			<td></td>
		</tr>
		<tr>
			<td>PBS (10x)</td>
			<td>Beyotime</td>
			<td>ST448</td>
			<td></td>
		</tr>
		<tr>
			<td>Permanent epoxy negative photoresist</td>
			<td>Microchem</td>
			<td>SU-8 2075</td>
			<td></td>
		</tr>
		<tr>
			<td>Phenol Red sodium salt</td>
			<td>Sigma-Aldrich</td>
			<td>P5530</td>
			<td></td>
		</tr>
		<tr>
			<td>Polydimethylsiloxane (PDMS)</td>
			<td>Dow Corning</td>
			<td>Sylgard 184</td>
			<td></td>
		</tr>
		<tr>
			<td>Poly-D-lysine hydrobromide (PDL)</td>
			<td>Sigma-Aldrich</td>
			<td>P7886</td>
			<td></td>
		</tr>
		<tr>
			<td>Polytetrafluoroethylene</td>
			<td>Teflon</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Program software</td>
			<td>MATLAB</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Recombinant Human VEGF-165</td>
			<td>StemImmune LLC</td>
			<td>HVG-VF5</td>
			<td></td>
		</tr>
		<tr>
			<td>Sodium hydroxide (NaOH)</td>
			<td>Sigma-Aldrich</td>
			<td>1.06498</td>
			<td></td>
		</tr>
		<tr>
			<td>Stage top incubator</td>
			<td>Tokai Hit</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>SU-8 developer</td>
			<td>Microchem</td>
			<td>NA</td>
			<td></td>
		</tr>
		<tr>
			<td>Trichloro(1H,1H,2H,2H-perfluorooctyl)silane</td>
			<td>Sigma-Aldrich</td>
			<td>448931</td>
			<td></td>
		</tr>
		<tr>
			<td>TRITC Phalloidin</td>
			<td>Sigma-Aldrich</td>
			<td>P5285</td>
			<td></td>
		</tr>
	</tbody>
</table>

microfluidic model to mimic initial event of neovascularization

Neovascularization is usually initialized from an existing normal vasculature and the biomechanical microenvironment of endothelial cells (ECs) in the initial stage varies dramatically from the following process of neovascularization. Although there are plenty of models to simulate different stages of neovascularization, an in vitro 3D model that capitulates the initial process of neovascularization under the corresponding stimulations of normal vasculature microenvironments is still lacking. Here, we reconstructed an in vitro 3D model that mimics the initial event of neovascularization (MIEN). The MIEN model contains a microfluidic sprouting chip and an automatic control, highly efficient circulation system. A functional, perfusable microchannel coated with endothelium was formed and the process of sprouting was simulated in the microfluidic sprouting chip. The initially physiological microenvironment of neovascularization was recapitulated with the microfluidic control system, by which ECs would be exposed to high luminal shear stress, physiological transendothelial flow, and various vascular endothelial growth factor (VEGF) distributions simultaneously. The MIEN model can be readily applied to the study of neovascularization mechanism and holds a potential promise as a low-cost platform for drug screening and toxicology applications.

The in vitro 3D model to mimic the initial event of neovascularization (MIEN) presented here consisted of a microfluidic sprouting chip and a microfluidic control system. The microfluidic sprouting chip was optimized from previous publications22,23,37,40,51,52,53. Briefly, it contai...

Watch this Scientific Journal Video about Microfluidic Model to Mimic Initial Event of Neovascularization at JoVE.com

Microfluidic Model to Mimic Initial Event of Neovascularization

Here, we provide a microfluidic chip and an automatically controlled, highly efficient circulation microfluidic system that recapitulates the initial microenvironment of neovascularization, allowing endothelial cells (ECs) to be stimulated by high luminal shear stress, physiological level of transendothelial flow, and various vascular endothelial growth factor (VEGF) distribution simultaneously.

Neovascularization is usually initialized from an existing normal vasculature and the biomechanical microenvironment of endothelial cells (ECs) in the ...

microfluidic-model-to-mimic-initial-event-of-neovascularization

Research

JoVE Journal

Bioengineering

4.5K Views.  Beihang University. Here, we provide a microfluidic chip and an automatically controlled, highly efficient circulation microfluidic system that recapitulates the initial microenvironment of neovascularization, allowing endothelial cells (ECs) to be stimulated by high luminal shear stress, physiological level of transendothelial flow, and various vascular endothelial growth factor (VEGF) distribution simultaneously.

Video: Microfluidic Model to Mimic Initial Event of Neovascularization

Tri-layered Electrospinning to Mimic Native Arterial Architecture using Polycaprolactone, Elastin, and Collagen: A Preliminary Study

Throughout native artery, collagen and elastin play an important role, providing a mechanical backbone, preventing vessel rupture, and promoting recovery under pulsatile deformations. The goal of this study was to mimic the structure of native artery by fabricating a multi-layered electrospun conduit composed of poly(caprolactone) (PCL) with the addition of elastin and collagen with blends of 45-45-10, 55-35-10, and 65-25-10 PCL-ELAS-COL to demonstrate mechanical properties indicative of native arterial tissue, while remaining conducive to tissue regeneration. Whole grafts and individual layers were analyzed using uniaxial tensile testing, dynamic compliance, suture retention, and burst strength. Compliance results revealed that changes to the middle/medial layer changed overall graft behavior with whole graft compliance values ranging from 0.8 - 2.8 % / 100 mmHg, while uniaxial results demonstrated an average modulus range of 2.0 - 11.8 MPa. Both modulus and compliance data displayed values within the range of native artery. Mathematical modeling was implemented to show how changes in layer stiffness affect the overall circumferential wall stress, and as a design aid to achieve the best mechanical combination of materials. Overall, the results indicated that a graft can be designed to mimic a tri-layered structure by altering layer properties.

The aim of this study was to mimic the native three layered architecture of the arterial wall. To accomplish this, electrospinning was employed with the use of a 3-1 (input-output) nozzle and blends of polycaprolactone, elastin, and collagen.

Throughout native artery, collagen and elastin play an important role, providing a mechanical backbone, preventing vessel rupture, and promoting recovery ...

Design of a Cyclic Pressure Bioreactor for the Ex Vivo Study of Aortic Heart Valves

The aortic valve, located between the left ventricle and the aorta, allows for unidirectional blood flow, preventing backflow into the ventricle. Aortic valve leaflets are composed of interstitial cells suspended within an extracellular matrix (ECM) and are lined with an endothelial cell monolayer. The valve withstands a harsh, dynamic environment and is constantly exposed to shear, flexion, tension, and compression. Research has shown calcific lesions in diseased valves occur in areas of high mechanical stress as a result of endothelial disruption or interstitial matrix damage1-3. Hence, it is not surprising that epidemiological studies have shown high blood pressure to be a leading risk factor in the onset of aortic valve disease4. 
The only treatment option currently available for valve disease is surgical replacement of the diseased valve with a bioprosthetic or mechanical valve5. Improved understanding of valve biology in response to physical stresses would help elucidate the mechanisms of valve pathogenesis. In turn, this could help in the development of non-invasive therapies such as pharmaceutical intervention or prevention. Several bioreactors have been previously developed to study the mechanobiology of native or engineered heart valves6-9. Pulsatile bioreactors have also been developed to study a range of tissues including cartilage10, bone11 and bladder12. The aim of this work was to develop a cyclic pressure system that could be used to elucidate the biological response of aortic valve leaflets to increased pressure loads. 
The system consisted of an acrylic chamber in which to place samples and produce cyclic pressure, viton diaphragm solenoid valves to control the timing of the pressure cycle, and a computer to control electrical devices. The pressure was monitored using a pressure transducer, and the signal was conditioned using a load cell conditioner. A LabVIEW program regulated the pressure using an analog device to pump compressed air into the system at the appropriate rate. The system mimicked the dynamic transvalvular pressure levels associated with the aortic valve; a saw tooth wave produced a gradual increase in pressure, typical of the transvalvular pressure gradient that is present across the valve during diastole, followed by a sharp pressure drop depicting valve opening in systole. The LabVIEW program allowed users to control the magnitude and frequency of cyclic pressure. The system was able to subject tissue samples to physiological and pathological pressure conditions. This device can be used to increase our understanding of how heart valves respond to changes in the local mechanical environment.

Design of a Cyclic Pressure Bioreactor for the Ex Vivo Study of Aortic Heart Valves

A cyclic pressure bioreactor capable of subjecting heart valve tissue to physiological and pathological pressure conditions has been designed. A LabVIEW program allows users to control pressure magnitude, amplitude and frequency. This device can be used to study the mechanobiology of heart valve tissue or isolated cells.

The aortic valve, located between the left ventricle and the aorta, allows for unidirectional blood flow, preventing backflow into the ventricle. Aortic ...

Microfluidic Chip Fabrication and Method to Detect Influenza

Fast and effective diagnostics play an important role in controlling infectious disease by enabling effective patient management and treatment. Here, we present an integrated microfluidic thermoplastic chip with the ability to amplify influenza A virus in patient nasopharyngeal (NP) swabs and aspirates. Upon loading the patient sample, the microfluidic device sequentially carries out on-chip cell lysis, RNA purification and concentration steps within the solid phase extraction (SPE), reverse transcription (RT) and polymerase chain reaction (PCR) in RT-PCR chambers, respectively. End-point detection is performed using an off-chip Bioanalyzer (Agilent Technologies, Santa Clara, CA). For peripherals, we used a single syringe pump to drive reagent and samples, while two thin film heaters were used as the heat sources for the RT and PCR chambers. The chip is designed to be single layer and suitable for high throughput manufacturing to reduce the fabrication time and cost. The microfluidic chip provides a platform to analyze a wide variety of virus and bacteria, limited only by changes in reagent design needed to detect new pathogens of interest.

An integrated microfluidic thermoplastic chip has been developed for use as a molecular diagnostic. The chip performs nucleic acid extraction, reverse transcriptase, and PCR. Methods for fabricating and running the chip are described.

Fast  and effective diagnostics play an important role in controlling infectious  disease by enabling effective patient management and treatment. Here, we ...

A Microfluidic Technique to Probe Cell Deformability

Here we detail the design, fabrication, and use of a microfluidic device to evaluate the deformability of a large number of individual cells in an efficient manner. Typically, data for ~102 cells can be acquired within a 1 hr experiment. An automated image analysis program enables efficient post-experiment analysis of image data, enabling processing to be complete within a few hours. Our device geometry is unique in that cells must deform through a series of micron-scale constrictions, thereby enabling the initial deformation and time-dependent relaxation of individual cells to be assayed. The applicability of this method to human promyelocytic leukemia (HL-60) cells is demonstrated. Driving cells to deform through micron-scale constrictions using pressure-driven flow, we observe that human promyelocytic (HL-60) cells momentarily occlude the first constriction for a median time of 9.3 msec before passaging more quickly through the subsequent constrictions with a median transit time of 4.0 msec per constriction. By contrast, all-trans retinoic acid-treated (neutrophil-type) HL-60 cells occlude the first constriction for only 4.3 msec before passaging through the subsequent constrictions with a median transit time of 3.3 msec. This method can provide insight into the viscoelastic nature of cells, and ultimately reveal the molecular origins of this behavior.

We demonstrate a microfluidics-based assay to measure the timescale for cells to transit through a sequence of micron-scale constrictions.

Here we detail the design, fabrication, and use of a microfluidic device to evaluate the deformability of a large number of individual cells in an ...

A Microfluidic Model of Biomimetically Breathing Pulmonary Acinar Airways

Quantifying respiratory flow characteristics in the pulmonary acinar depths and how they influence inhaled aerosol transport is critical towards optimizing drug inhalation techniques as well as predicting deposition patterns of potentially toxic airborne particles in the pulmonary alveoli. Here, soft-lithography techniques are used to fabricate complex acinar-like airway structures at the truthful anatomical length-scales that reproduce physiological acinar flow phenomena in an optically accessible system. The microfluidic device features 5 generations of bifurcating alveolated ducts with periodically expanding and contracting walls. Wall actuation is achieved by altering the pressure inside water-filled chambers surrounding the thin PDMS acinar channel walls both from the sides and the top of the device. In contrast to common multilayer microfluidic devices, where the stacking of several PDMS molds is required, a simple method is presented to fabricate the top chamber by embedding the barrel section of a syringe into the PDMS mold. This novel microfluidic setup delivers physiological breathing motions which in turn give rise to characteristic acinar air-flows. In the current study, micro particle image velocimetry (&#181;PIV) with liquid suspended particles was used to quantify such air flows based on hydrodynamic similarity matching. The good agreement between &#181;PIV results and expected acinar flow phenomena suggest that the microfluidic platform may serve in the near future as an attractive in vitro tool to investigate directly airborne representative particle transport and deposition in the acinar regions of the lungs.

Soft-lithography was utilized to produce a representative true-scale model of pulmonary alveolated airways that expand and contract periodically, mimicking physiological breathing motion. This platform recreates respiratory acinar flows on a chip, and is anticipated to facilitate experimental investigation of inhaled aerosol dynamics and deposition in the pulmonary acinus.

Quantifying respiratory flow characteristics in the pulmonary acinar depths and how they influence inhaled aerosol transport is critical towards ...

Using Adhesive Patterning to Construct 3D Paper Microfluidic Devices

We demonstrate the use of patterned aerosol adhesives to construct both planar and nonplanar 3D paper microfluidic devices. By spraying an aerosol adhesive through a metal stencil, the overall amount of adhesive used in assembling paper microfluidic devices can be significantly reduced. We show on a simple 4-layer planar paper microfluidic device that the optimal adhesive application technique and device construction style depends heavily on desired performance characteristics. By moderately increasing the overall area of a device, it is possible to dramatically decrease the wicking time and increase device success rates while also reducing the amount of adhesive required to keep the device together. Such adhesive application also causes the adhesive to form semi-permanent bonds instead of permanent bonds between paper layers, enabling single-use devices to be non-destructively disassembled after use. Nonplanar 3D origami devices also benefit from the semi-permanent bonds during folding, as it reduces the likelihood that unrelated faces may accidently stick together. Like planar devices, nonplanar structures see reduced wicking times with patterned adhesive application vs uniformly applied adhesive.

We demonstrate the use of patterned aerosol adhesives to construct 3D paper microfluidic devices. This method of adhesive application forms semi-permanent bonds between layers, enabling single-use devices to be non-destructively disassembled after use and to ease folding complex nonplanar structures.

We demonstrate the use of patterned aerosol adhesives to construct both planar and nonplanar 3D paper microfluidic devices. By spraying an aerosol ...

Microfluidic Flow Chambers Using Reconstituted Blood to Model Hemostasis and Platelet Transfusion In Vitro

Blood platelets prepared for transfusion gradually lose hemostatic function during storage. Platelet function can be investigated using a variety of (indirect) in vitro experiments, but none of these is as comprehensive as microfluidic flow chambers. In this protocol, the reconstitution of thrombocytopenic fresh blood with stored blood bank platelets is used to simulate platelet transfusion. Next, the reconstituted sample is perfused in microfluidic flow chambers which mimic hemostasis on exposed subendothelial matrix proteins. Effects of blood donation, transport, component separation, storage and pathogen inactivation can be measured in paired experimental designs. This allows reliable comparison of the impact every manipulation in blood component preparation has on hemostasis. Our results demonstrate the impact of temperature cycling, shear rates, platelet concentration and storage duration on platelet function. In conclusion, this protocol analyzes the function of blood bank platelets and this ultimately aids in optimization of the processing chain including phlebotomy, transport, component preparation, storage and transfusion.

Microfluidic Flow Chambers Using Reconstituted Blood to Model Hemostasis and Platelet Transfusion In Vitro

Platelet transfusion and hemostasis was modeled using blood reconstitution and microfluidic flow chambers to investigate the function of blood banking platelets. The data demonstrate the consequences of platelet storage lesion on hemostasis, in vitro.

Blood platelets prepared for transfusion gradually lose hemostatic function during storage. Platelet function can be investigated using a variety of ...

Multi-step Variable Height Photolithography for Valved Multilayer Microfluidic Devices

Microfluidic systems have enabled powerful new approaches to high-throughput biochemical and biological analysis. However, there remains a barrier to entry for non-specialists who would benefit greatly from the ability to develop their own microfluidic devices to address research questions. Particularly lacking has been the open dissemination of protocols related to photolithography, a key step in the development of a replica mold for the manufacture of polydimethylsiloxane (PDMS) devices. While the fabrication of single height silicon masters has been explored extensively in literature, fabrication steps for more complicated photolithography features necessary for many interesting device functionalities (such as feature rounding to make valve structures, multi-height single-mold patterning, or high aspect ratio definition) are often not explicitly outlined. 
Here, we provide a complete protocol for making multilayer microfluidic devices with valves and complex multi-height geometries, tunable for any application. These fabrication procedures are presented in the context of a microfluidic hydrogel bead synthesizer and demonstrate the production of droplets containing polyethylene glycol (PEG diacrylate) and a photoinitiator that can be polymerized into solid beads. This protocol and accompanying discussion provide a foundation of design principles and fabrication methods that enables development of a wide variety of microfluidic devices. The details included here should allow non-specialists to design and fabricate novel devices, thereby bringing a host of recently developed technologies to their most exciting applications in biological laboratories.

Multilayer microfluidic devices often involve the fabrication of master molds with complex geometries for functionality. This article presents a complete protocol for multi-step photolithography with valves and variable height features tunable to any application. As a demonstration, we fabricate a microfluidic droplet generator capable of producing hydrogel beads.

Microfluidic systems have enabled powerful new approaches to high-throughput biochemical and biological analysis. However, there remains a barrier to ...

Initial Evaluation of Antibody-conjugates Modified with Viral-derived Peptides for Increasing Cellular Accumulation and Improving Tumor Targeting

Antibody-conjugates (ACs) modified with virus-derived peptides are a potentially powerful class of tumor cell delivery agents for molecular payloads used in cancer treatment and imaging due to increased cellular accumulation over current ACs. During early AC in vitro development, fluorescence techniques and radioimmunoassays are sufficient for determining intracellular localization, accumulation efficiency, and target cell specificity. Currently, there is no consensus on standardized methods for preparing cells for evaluating AC intracellular accumulation and localization. The initial testing of ACs modified with virus-derived peptides is critical especially if several candidates have been constructed. Determining intracellular accumulation by fluorescence can be affected by background signal from ACs at the cell surface and complicate the interpretation of accumulation. For radioimmunoassays, typically treated cells are fractionated and the radioactivity in different cell compartments measured. However, cell lysis varies from cell to cell and often nuclear and cytoplasmic compartments are not adequately isolated. This can produce misleading data on payload delivery properties. The intravenous injection of radiolabeled virus-derived peptide-modified ACs in tumor bearing mice followed by radionuclide imaging is a powerful method for determining tumor targeting and payload delivery properties at the in vivo phase of development. However, this is a relatively recent advancement and few groups have evaluated virus-derived peptide-modified ACs in this manner. We describe the processing of treated cells to more accurately evaluate virus-derived peptide-modified AC accumulation when using confocal microscopy and radioimmunoassays. Specifically, a method for trypsinizing cells to remove cell surface bound ACs. We also provide a method for improving cellular fractionation. Lastly, this protocol provides an in vivo method using positron emission tomography (PET) for evaluating initial tumor targeting properties in tumor-bearing mice. We use the radioisotope 64Cu (t1/2 = 12.7 h) as an example payload in this protocol.

Viral-derived peptides coupled to antibody-conjugates (ACs) is an approach gaining momentum due to the potential of delivering molecular payloads with increased tumor cell accumulation. Utilizing common methods to evaluate peptide conjugation, AC and payload intracellular accumulation, and tumor targeting, this protocol helps researchers during the key initial development phases.

Antibody-conjugates (ACs) modified with virus-derived peptides are a potentially powerful class of tumor cell delivery agents for molecular payloads used ...

Perfusable Vascular Network with a Tissue Model in a Microfluidic Device

A spheroid (a multicellular aggregate) is regarded as a good model of living tissues in the human body. Despite the significant advancement in the spheroid cultures, a perfusable vascular network in the spheroids remains a critical challenge for long-term culture required to maintain and develop their functions, such as protein expressions and morphogenesis. The protocol presents a novel method to integrate a perfusable vascular network within the spheroid in a microfluidic device. To induce a perfusable vascular network in the spheroid, angiogenic sprouts connected to microchannels were guided to the spheroid by utilizing angiogenic factors from human lung fibroblasts cultured in the spheroid. The angiogenic sprouts reached the spheroid, merged with the endothelial cells co-cultured in the spheroid, and formed a continuous vascular network. The vascular network could perfuse the interior of the spheroid without any leakage. The constructed vascular network may be further used as a route for supply of nutrients and removal of waste products, mimicking blood circulation in vivo. The method provides a new platform in spheroid culture toward better recapitulation of living tissues.

The protocol describes how to engineer a perfusable vascular network in a spheroid. The spheroid's surrounding microenvironment is devised to induce angiogenesis and connect the spheroid to the microchannels in a microfluidic device. The method allows the perfusion of the spheroid, which is a long-awaited technique in three-dimensional cultures.