This project was supported by the Graduate Student Study Abroad Program (NSC-101-2917-I-007-010); the Biomedical Engineering Program (NSC-101-2221-E-007-032-MY3); and the Nanotechnology National Program (NSC-101-2120-M-007-001-); and the Ministry of Science and Technology (MOST-104-2221-E-007-072-MY3), National Science Council of the R.O.C., Taiwan. The authors would like to thank Prof. Ali Khademhosseini, Gulden Camci-Unal, Arghya Paul, and Ronglih Liao at Harvard Medical School for sharing the hydrogel and cell encapsulation technology.

1. GelMA Synthesis
<ol>
	<li>Weigh 10 g of gelatin powder and add it to a glass flask with 100 mL ofDulbecco&#39;s phosphate-buffered saline (DPBS). Put a magnetic stir bar into the flask and place the flask on a stirring hot plate.</li>
	<li>Cover the flask with aluminum foil to avoid water evaporation. Set the hot plate temperature to 50-60 &#176;C and the stirrer at 100 rpm for 1 h to dissolve the gelatin powder well.</li>
	<li>After the gelatin has dissolved, add 8 mL of methacrylic anhydride very slowly (one drop per second) using a pipette. Let it react at 60 &#176;C for 3 h.</li>
	<li>Add pre-warmed DPBS (40 &#176;C) to the flask to a final volume of 500 mL and allow this mix well for 15 min to stop the reaction.</li>
	<li>Meanwhile, cut a dialysis membrane (14 kDa cut-off molecular weight) into several 25 cm-long tubes. Immerse them in deionized (DI) water for 15 min and make a knot to close one end of the dialysis tubes.</li>
	<li>Load the appropriate amount (30 - 60 mL) of the polymer solution into the dialysis tubes and close the other end. Place them in a 5-L plastic beaker with DI water for a week. Renew the DI water twice a day and maintain the solution at 40 - 50 &#176;C during the dialysis process.</li>
	<li>Collect the solution from the tube in a 500-mL glass bottle. Pour ~450 - 500 mL of solution in a 500-mL filter cup (pore size of 0.22 &#181;m) and apply a vacuum to the filter cup to force the solution to pass through the filter membrane for sterilization.</li>
	<li>Transfer the sterilized polymer into several 50-mL sterilized tubes and store them in a -80 &#176;C freezer for 3 - 5 days.</li>
	<li>Freeze-dry the -80 &#176;C polymer for 1 week using a freeze dryer to form GelMA. Store the GelMA in a -80 &#176;C freezer.</li>
</ol>
2. 3-(Trimethoxysilyl)propyl Methacrylate (TMSPMA) Modification
<ol>
	<li>&#160;Cut commercial glass slides into two small pieces (25 mm x 37.5 mm) and immerse them in 0.5 M NaOH solution for 4 h. Wash the slides with a large amount of DI water.</li>
	<li>Place the slide on a rack inside a glass container with 95% ethanol and clean using an ultrasonicator at 43 kHz for 15 min. Air dry the glass slide.</li>
	<li>Immerse the glass slides in 5% TMSPMA in 99.5% ethanol for 1 h.</li>
	<li>Wash the slides in 95% ethanol, air dry the slides, and anneal the TMSPMA coating in an oven at 80 &#176;C for 2 h.</li>
</ol>
3. Chip Fabrication
<ol>
	<li>Take one 2 mm- and one 0.3 mm-thick polymethylmethacrylate (PMMA) plate, apply double-sided tape to one side of the PMMA surface, and release the liner on one side. Leave two 2 mm-thick PMMA plates and one 1 mm-thick plate without double-sided tape.</li>
	<li>Laser-cut a 2 mm-thick PMMA plate without double-sided tape to 42 mm x 30 mm to make the bottom plate. Cut a 2 mm-thick PMMA plate with double-sided tape to make the boundary frame, with outer dimensions of 42 mm x 30 mm and inner dimensions of 37.5 mm x 25 mm.</li>
	<li>Laser-cut the 0.3 mm-thick PMMA plate with double-sided tape into a 12-mm center circle with two 2 mm-wide and 8 mm-long flow channels on opposite sides of the circle (Figure 1a).</li>
	<li>To prepare the PMMA mold for casting the PDMS cover, assemble the three pieces of PMMA components from steps 3.2-3.3 (Figure 1b) using double-sided tape.</li>
	<li>Laser-cut a 2 mm-thick PMMA plate with double-sided tape into a 5 cm x 5 cm piece with a 3 x 3 array of 8.5 mm x 8.5 mm hollow rectangles. Cut another 5 cm x 5 cm piece with a 3 x 3 array of 4-mm hollow circles. Laser-cut a 1 mm-thick PMMA plate without double-sided tape into a 5 cm x 5 cm PMMA bottom plate.</li>
	<li>Prepare the PMMA mold for casting the PDMS plug by assembling the three pieces of PMMA components from step 3.5 (Figure 1c) using double-sided tape.</li>
	<li>Prepare the PDMS cover and PDMS plug by properly mixing 30 g of PDMS elastomer and 3 g of PDMS curing agent; degas the mixture under a vacuum chamber for 1 h.</li>
	<li>Cast 1.8-2.0 g of the mixture into the PMMA mold for the PDMS cover and use the appropriate amount to fill each cavity of the PMMA mold for the PDMS plug. Cast 10 g of uncured PDMS mixture in a blank 10-cm plastic plate. Put these molds in a vacuum chamber to degas for 30 min.</li>
	<li>Cure the PDMS for 2 h at 80 &#176;C. After cooling down the cured PDMS on the mold, detach the PDMS covers and the PDMS plugs from the PMMA molds.</li>
	<li>Punch two holes with diameters of about 3 mm at the ends of flow channels of the PDMS covers.</li>
	<li>Cut the PDMS sheet molded from the 10-cm plastic plate into many 1 cm x 1 cm cubes and punch a 3-mm hole in each PDMS cube. Glue two uncured 1 cm x 1 cm PDMS cubes onto the openings of the PDMS cover (to serve as medium reservoirs and to aid in the curing process of the gradient circular hydrogel patterns) and cure for 1 h at 80 &#176;C.</li>
	<li>Bond the PDMS covers with the two PDMS reservoirs onto a TMSPMA-coated slide by pre-treating the bonding side of the PDMS cover and the TMSPMA-coated slide under an oxygen plasma machine (30 W RF power (high mode) and 600 mTorr compressed air) or a high-frequency electronic corona generator (115 V, 50/60 Hz, 0.35 A) for 90 s of O2 plasma treatment.</li>
	<li>Contact the plasma-treated surface of the PDMS covers and the TMSPMA slide and press them closely for permanent bonding through the formation of an Si-O-Si bond. 
	NOTE: Placing the chip in an oven at 80 &#176;C for 1 h can further enhance the bonding strength.</li>
	<li>After cooling, immerse the chips in 95% ethanol for 15 min and air dry. Then, sterilize the chips under UV irradiation for 1 h and store them in a box wrapped in aluminum foil in the laminar hood.</li>
</ol>
4. Static Gradient Strain on the Cell-laden Hydrogel
<ol>
	<li>Print and cut a piece of photomask, 25 mm x 37.5 mm in size, by printing the layout in Figure 2b on a transparent film. Adjust the printed size of Figure 2b to match the dimension in Figure 2a.</li>
	<li>Prepare 100 mL of DMEM medium with 10% FBS, 1% Pen-Strep, and 250 mL of DPBS in a 37 &#176;C water bath to use as the cell culture medium.</li>
	<li>Weigh 25 mg of freeze-dried GelMA into 0.3 mL of prewarmed (37 &#176;C) cell culture medium in a 1.5-mL black microcentrifuge tube. Put the microcentrifuge tube on a laboratory stirrer/hot plate until the GelMA dissolves in the medium.</li>
	<li>Weigh 50 mg of photoinitiator into 1 ml of DPBS in a microcentrifuge tube and place it in an 80 &#176;C oven for 15 min or until the photoinitiator has dissolved.</li>
	<li>Take 25 &#181;L of the 10% photoinitiator from step 4.4 and add it to the microcentrifuge tube from step 4.3. Pipette several times to mix well.</li>
	<li>Count 3 x 106 NIH 3T3 cells using an automated cell counter. Centrifuge the suspension at 200 x g for 5 min, discard the supernatant, and re-suspend the cells in 175 &#181;L of cell culture medium.</li>
	<li>Add the cell solution from step 4.6 to the microcentrifuge tube from step 4.5 to get a prepolymer cell solution of 5% GelMA, 0.5% photoinitiator, and ~6 x 106 3T3 cells/mL. After mixing, load 100 &#181;L of cell prepolymer in a 100-&#181;L micro syringe.</li>
	<li>Manually align (see Figure 3a) a piece of the photomask onto the bottom slide of the sterilized gradient strain chip and simply fix the position using a small drop of DI water in between. Connect the 100-&#181;L micro-syringe loaded with prepolymer cell solution to the inlet of the chip.</li>
	<li>Place (see Figure 3b) 50 &#181;L of prepolymer cell solution in the flow channel using the micro-syringe and then plug the outlet using a PDMS plug. Inject an extra 40 &#181;L of solution to create a convex bulge in the circular PDMS membrane.</li>
	<li>Move the chip with the photomask (bottom), micro-syringe (inlet), and PDMS plug (outlet) from step 4.9 under a UV lamp (365 nm, 9 mW/cm2) and expose it for 30 - 45 s to crosslink the concentric circular hydrogel in the fluidic chamber.</li>
	<li>Remove the PDMS plug and the micro-syringe to release the liquid pressure (see Figure 3c). Use a 1-mL syringe loaded with prewarmed DPBS to wash out uncrosslinked resins 3 times by flushing from the inlet to the outlet.</li>
	<li>Fill the flow channel with about 100 &#181;L of cell culture medium.</li>
	<li>Place the chip in a sterilized culture dish and culture in a 5% CO2 atmosphere at 37 &#176;C for a week. Refresh the medium every day.</li>
	<li>Take images of three chips on day 0 after 4 h of incubation, as a control group, and three chips on day 3, as the experimental set, using a microscope with a 20X objective. Measure the line width of each hydrogel from line 1 to line 12 using software (e.g., ImageJ) to calculate the compress strains ( Figure 4). 
	NOTE: The elongation percentage is calculated by dividing the value of the line width difference between 40 &#181;L and 0 &#181;L by the line width at 40 &#181;L.</li>
</ol>
5. Cell Staining for Alignment Analysis
<ol>
	<li>Use a syringe to inject 4% paraformaldehyde (PFA) in DPBS at RT into the flow chip for 15 min for the fixation of the cell-laden hydrogel. 
	NOTE: Caution. PFA is toxic and should be handled with care.</li>
	<li>Replace the solution with 0.5% cell membrane permeating solution in DPBS for 10 min to permeabilize the cell membrane at RT.</li>
	<li>PBS wash the samples 3 times with 5- to 10-min interval between washes (using a loaded syringe, as in step 5.1).</li>
	<li>Load 1% BSA solution into the fluidic channel for 45-60 min at RT (using a loaded syringe through the inlet port).</li>
	<li>Add 1.5 mL of methanol into the vial with Alexa Fluor 488 phalloidin to yield a final concentration of 6.6 &#181;M stock solution.</li>
	<li>Take 5 &#181;L of Alexa Fluor 488 phalloidin from step 5.5 and dilute it in 200 &#181;L of DPBS with 0.1% BSA to form a final concentration of 0.165 &#181;M Alexa Fluor 488 phalloidin.</li>
	<li>Add 200 &#181;L of the mixture solution to the fluidic channel using a micropipette and incubate the chip at 37 &#176;C for 45 - 60 min for actin staining. PBS wash (as above) the samples 3 times.</li>
	<li>Prepare 1 &#181;g/mL DAPI in PBS, flow it through the chip, and stain cell nuclei at 37 &#176;C for 5 min.</li>
	<li>Pipette DPBS into the fluidic channel to wash out the staining solution and refill the fluidic chamber with PBS solution to take images with a fluorescence microscope.</li>
	<li>Capture the fluorescent images of the 3T3 cells in the hydrogels using an inverted fluorescence microscope under 40X magnification with a CCD detector and filter sets of ex/em at 488/520 nm and 358/461 nm for Alexa Fluor 488 phalloidin (actin) and DAPI (nucleus), respectively.</li>
</ol>

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The authors have nothing to disclose.

In this paper, we report on a simple approach to compare cell alignment behavior after hydrogel shape guidance and tensile stretch. A flexible PDMS membrane creates a dome-shaped curvature for generating different heights of concentric circular hydrogels. After releasing the pressure, the PDMS membrane automatically applies force to the micro-patterned hydrogels to form gradient strain/elongation, with a maximum at the center and a minimum at the outer boundary. As the formation of the gradient strain is designed by the flexible PDMS membrane and the handling of the fluidic chip, there are several important parameters that should be attended to: (i) Precise control of the thickness of the PDMS membrane is crucial to adjust the gradient strain value. If the membrane is too thick, even the maximum injection volume of cell prepolymer will not be able to generate a proper convex curve in the PDMS membrane for crosslinking a gradient height of cell-laden hydrogels. In contrast, a too-thin PDMS membrane cannot apply sufficient force to the hydrogels. Please check that the weight of uncured PDMS in the PDMS cover mold is around 1.6-2.0 g per chip. (ii) Contamination prevention is very important during the crosslink handling of the cell-laden hydrogel in the fluidic chip. Before cell culturing in the incubator, thoroughly washing with sterilized PBS in the fluidic channel and using 75% ethanol to wipe the surface of the chip can help to avoid the contamination issue. (iii) The concentration of photoinitiator and the dosage of UV exposure should be careful controlled and in the range of ~0.1% - 2% (0.5% is recommended). Over-crosslinking the hydrogel and overdosing the UV irradiation will result in low cell viability. (iv) The line-width of the patterned hydrogel should not be too large. Otherwise, the nutrient replacement rate in the thick hydrogels will not be able to support cell proliferation. Usually, less than 300 &#181;m is recommended. The spacing between two hydrogel circles can be varied, and a 50% duty cycle is recommended. (v) While washing or refilling the flow channel with solution, the formation of bubbles should be avoided. Gently pipetting the solution in the chip can help to remove bubbles.
The concept of gradient strain generated by PDMS-deformed curvature can be further upgraded to applying dynamic gradient strains and can be integrated with biochemical stimulation, which can benefit many studies on functional tissue regeneration. The simple fluidic injection module with a PDMS plug can be replaced by any advanced fluidic system for extended experimental control. The PMMA mold can also be replaced by a microfabricated SU-8 mold or a bulk-etched silicon mold.
This gradient strain chip with a circular cell-laden hydrogel can generate static compressive force on the 3D hydrogel without external mechanical or electrical machines. Therefore, it provides a fast screening platform for investigating cell behaviors in a series of strain conditions, without the risk of contamination issues caused by the operation of external machines. However, time-controlled strain stimulation is not achievable because the PDMS membrane generates strains until the degradation of the hydrogel.

Serving as a block material that mimics a native microenvironment, a hydrogel containing extracellular matrix (ECM) can re-build biomimetic tissue scaffolds to support cell growth. To possess the functions of a tissue, organized cell alignment is an essential requirement. Various 2D (i.e., cells cultured on a surface) and 3D (i.e., cells encapsulated in a hydrogel) cell alignments have been achieved by culturing or encapsulating cells in or on flexible substrates with micro-or nano-patterns1. 3D cell alignment in microarchitecture is more attractive, as the microenvironment is closer to the native tissue construct2,3,4. One common approach for 3D cell alignment is the geometric cue of hydrogel shape2,3. Because of the restricted space for cell proliferation in the short-axis direction, cells aim to align along the long-axis direction in a micro-patterned hydrogel. Another approach is to apply tensile stretch to the hydrogels to achieve cell alignment parallel to the stretch direction4,5.
Biophysical stimulation on ECM hydrogels, such as compressive strain or an electrical field, can regulate cell functions for proper tissue integration, proliferation, and differentiation1,2,3. Much research has been done to investigate cellular behavior by applying one strain condition at a time using multiple mechanical control units4,6,7,8,9. For example, the use of mechanical step motors squeezed or stretched on a 3D cell-encapsulated collagen hydrogel has been a common approach7,10. However, such controlling equipment requires extra space and faces the issue of contamination in the incubator7,9,11,12. In addition, the large instrument cannot give a precise control environment to provide high reproducibility13.
Considering that cell-laden hydrogels are usually employed on the micro-scale for biomedical applications, it is advantageous to combine MEMS techniques to generate a range of strain/stretch stimulation to simultaneously investigate cell behaviors in 3D biomimetic constructs in vitro2,14,15,16,17,18. For example, using gas pressure to deform the PDMS membrane in microfluidic chips can give rise to various strains, driving cell differentiation to different lineages9,16. However, there are many technical challenges, such as complicated chip fabrication processes in a clean room and the software control integration of motors, pumps, valves, and compressed gases.
In this work, we demonstrate a simple approach to obtain a self-sustaining gradient static-strain microfluidic chip by employing a concentric circular hydrogel pattern and a flexible PDMS membrane. Unlike most of the existing approaches, our platform is a portable and disposable miniature device that can be fabricated outside a yellow room and that possesses self-generating gradient strains on concentric cell-encapsulated hydrogels, without external mechanical equipment during the incubation. 3T3 fibroblast cell behaviors influenced by a combination of hydrogel shape and a variety of tensile stretch guidance cues were demonstrated during the observation of cell alignment within 3D ECM-mimetic environments in the gradient strain chip for 3 days.

<table><tbody><tr><td>1.5 mL black microcentrifuge tube</td><td>Argos Technologies&amp;nbsp;</td><td>03-391-161</td><td>This one can be replaced with a neutral color of 1.5 mL tube covered with aluminun foil</td></tr><tr><td>10x DPBS</td><td>Sigma-Aldrich</td><td>56064C</td><td></td></tr><tr><td>Alexa Fluor 488 phalloidin&amp;nbsp;</td><td>Invitrogen</td><td>A12379&amp;nbsp;</td><td></td></tr><tr><td>BSA</td><td>Sigma</td><td>A1595</td><td></td></tr><tr><td>Calcein</td><td>Molecular Probe</td><td>C1430</td><td>For labeling viable cells</td></tr><tr><td>CCD</td><td>PCO. Imaging</td><td>Pixelfly qe</td><td></td></tr><tr><td>Cell membrane permeating solution</td><td>Sigma-Aldrich</td><td>X100</td><td>0.5% Triton X-100 for permeating cell membrane</td></tr><tr><td>DAPI</td><td>Sigma-Aldrich</td><td>D8417</td><td>Cell nucleus staining</td></tr><tr><td>Dialysis membrane</td><td>Sigma-Aldrich</td><td>D9527</td><td>Molecular weight cut-off = 14,000</td></tr><tr><td>DMEM</td><td>Gibco</td><td>11995-065</td><td></td></tr><tr><td>Double-side tape</td><td>3M</td><td>8003</td><td></td></tr><tr><td>FBS</td><td>Hyclone</td><td>SH30071.03</td><td></td></tr><tr><td>Gelatin</td><td>Sigma-Aldrich</td><td>G2500</td><td>gel strength 300, type A, from porcine skin</td></tr><tr><td>High frequency electronic corona generator</td><td>Electro-technic products</td><td>MODEL BD-20</td><td></td></tr><tr><td>Methacrylic Anhydride</td><td>Sigma-Aldrich</td><td>276685</td><td></td></tr><tr><td>Micro syringe</td><td>Hamilton</td><td>80501</td><td>50 &amp;mu;L&amp;nbsp;</td></tr><tr><td>Microscope</td><td>Olympus</td><td>IX71</td><td>Include two filter sets: LF405/LP-B-000 and LF488/LP-C-000 from Semrock</td></tr><tr><td>Oxygen plasma machine</td><td>Harrick plasma</td><td>PDC-001</td><td></td></tr><tr><td>Paraformaldehyde</td><td>Sigma-Aldrich</td><td>P6148</td><td>For fixing cell</td></tr><tr><td>PDMS</td><td>DOW CORNING</td><td>Sylgard 184</td><td>Mixture for PDMS chip cast-molding fabrication</td></tr><tr><td>Pen-Strep</td><td>Gibco</td><td>10378-016</td><td>penicillin/streptomycin</td></tr><tr><td>Photoinitiator</td><td>CIBA</td><td>Irgacure 2959</td><td></td></tr><tr><td>Propidium iodide</td><td>Sigma-Aldrich</td><td>P4170</td><td>For labeling dead cells</td></tr><tr><td>Sterile Filtration cup</td><td>Millipore</td><td>SCGPT05RE</td><td></td></tr><tr><td>TMSPMA</td><td>Sigma-Aldrich</td><td>440159</td><td>For hydrogel immobilization</td></tr><tr><td>Ultrasonicator</td><td>Delta</td><td>D150H</td><td>150W, 43kHz</td></tr><tr><td>UV light</td><td>DAIHAN</td><td>WUV-L10</td><td></td></tr><tr><td>Freeze Dryer</td><td>FIRSTEK</td><td>150311025</td><td></td></tr><tr><td>NIH3T3(fibroblast)</td><td>Food Industry Research and Development Institute(FIRDI)</td><td>08C0011</td><td></td></tr><tr><td>MOXI Z Mini Automated Cell Counter</td><td>ORFLO</td><td>MXZ001</td><td></td></tr></tbody></table>

gradient strain chip for stimulating cellular behaviors in cell-laden hydrogel

Artificial guidance for cellular alignment is a hot topic in the field of tissue engineering. Most of the previous research has investigated single strain-induced cellular alignment on a cell-laden hydrogel by using complex experimental processes and mass controlling systems, which are usually associated with contamination issues. Thus, in this article, we propose a simple approach to building a gradient static strain using a fluidic chip with a plastic PDMS cover and a UV transparent glass substrate for the stimulation of cellular behavior in a 3D hydrogel. Overloading photo-patternable cell prepolymer in the fluidic chamber can generate a convex curved PDMS membrane on the cover. After UV crosslinking, through a concentric circular micropattern under the curved PDMS membrane, and buffer washing, a microenvironment for investigating cell behaviors under a variety of gradient strains is self-established in a single fluidic chip, without external instruments. NIH3T3 cells were demonstrated after observing the change in the cellular alignment trend under geometry guidance, in cooperation with strain stimulation, which varied from 15 - 65% on hydrogels. After a 3-day incubation, the hydrogel geometry dominated the cell alignment under low compressive strain, where cells aligned along the hydrogel elongation direction under high compressive strain. Between these, the cells showed random alignment due to the dissipation of the radical guidance of hydrogel elongation and the geometry guidance of the patterned hydrogel.

To compare the mechanical variations between each circular hydrogel in the completed gradient strain stimulation chip, we measured the line widths of each circular hydrogel in two of the same chips, with injection volumes of 0 &#181;L (Figure 4a) and 40 &#181;L (Figure 4b), respectively. The percent elongations at each circle were calculated by dividing the elongations in the 40 &#181;L-injected chip by the line widths of the corresponding hydrogels in the 0 &#181;L-injected chip (Figure 4c). The hydrogel is an incompressible material, so the vertical contracting strain will be equivalent to the lateral elongation strain. Therefore, the percent elongations in the chip with the 40-&#181;L injection volume show 15 - 65% elongation, which can be converted to the compressive strain (see Supplementary File 1).
Fluorescent staining of the cell nuclei and F-actin with DAPI and phalloidin, respectively, was done to analyze the cellular alignment. The DAPI staining provided data on nuclear orientation, and the phalloidin staining was applied to assess cell spreading. Figure 5a-c shows the cell alignment in the gradient strain chip. In line 1, the 3T3 cells aligned along the radial direction. In hydrogel 7, the cells aligned randomly, and cells aligned along the circular direction in line 12. In accordance with the fluorescent staining images, a 90&#176; shift from the angle of cell alignment of the 3T3 cells in line 1 (maximum hydrogel elongation in the radial direction) to the alignment angle of the cells with long-axis alignment in line 12 (lowest hydrogel elongation in the radial direction) was discovered.
In previous research2,9, cells in 200-&#181;m line-patterned hydrogels aim to align along the long-axis direction of the hydrogel. However, in this study, we observed that the elongated stretch on the 200-&#181;m hydrogels in the short-axis direction provided another factor to affect and dominate the cellular alignment by controlling the percentage of the strain on the hydrogel. For the 65% strain on line 1, the radical alignment proved that the elongation stretch of the hydrogel dominates the cell alignment. For the 15% strain on line 12, the circular alignment proved that the long-axis effect dominated the cell alignment. For the 40% strain on hydrogel 7, the cells aligned randomly due to the neutralization of the geometry guidance and strain effect.
<img alt="Figure 1" class="xfigimg" src="/files/ftp_upload/53715/53715fig1.jpg" /> 
Figure 1. PMMA Mother Mold for PDMS Sheet and Plug Fabrication. (a) The separated components of the PMMA mold, including a bottom plate, boundary frame, and flow channel. After assembly with double-sided tape, (b) the PMMA mother mold for the flow sheet is formed. (c) Another PMMA mold is assembled for the PDMS plug. The red number represents the depth. (unit: mm) <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/53715fig1large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 2" class="xfigimg" src="/files/ftp_upload/53715/53715fig2.jpg" /> 
Figure 2. Plastic Photomask Design for the Cell-laden Hydrogel Micropattern. There are two openings with triangle shapes (2 mm at the bottom line and 6.5 mm in height) connecting to the flow channels to supply fresh cell culture medium. (a) The plastic photomask is labeled with the dimension. The period of the concentric circle is 400 nm, and the duty cycle is 50%. The diameter of the center circle is 2 mm. (b) The photomask layout, without labels for laser printing on a plastic transparent film. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/53715fig2large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 3" class="xfigimg" src="/files/ftp_upload/53715/53715fig3.jpg" /> 
Figure 3.Fabrication Processes of the Gradient Strain Cell-laden Hydrogel Along the Radial Direction in a PDMS Fluidic Chip. (a) A plastic photomask is aligned and stuck under the chip with a TMSPMA-coated fluidic channel. A micro-syringe with prepolymer cell solution is inserted into the inlet of the chip and used to inject about 50 &#181;L to fill the flow channel. (b) The outlet of the flow channel is closed with the PDMS plug and an additional 40 &#181;L of prepolymer cell solution is injected. The glass bottom is UV-patterned for 30 s to fabricate the concentric circular hydrogel in the flow chip. (c) The liquid pressure is released in the flow channel by unplugging the outlet and the un-crosslinking mixture is washed out with DPBS. (d) A chip with static gradient strain applies concentric cell-laden hydrogels that are ready for cell culturing. During the UV crosslinking process, (e) a non-continuous gradient height of hydrogel along the radius is formed. (f) After unplugging the outlet, the PDMS membrane becomes flat and applies gradient stress on the cell-encapsulated hydrogels. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/53715fig3large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 4" class="xfigimg" src="/files/ftp_upload/53715/53715fig4.jpg" /> 
Figure 4.The Pure Hydrogel Elongation Gradient. The elongation line width and percentage of pure hydrogel without cell encapsulation in the gradient strain chip on day 3 with the (a) 0-&#181;L (control group) and (b) 40-&#181;L injection volumes. (c) The elongation percentage is calculated by dividing the value of the line width difference between 40 &#181;L and 0 &#181;L by the line width of 40 &#181;L. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/53715fig4large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 5" class="xfigimg" src="/files/ftp_upload/53715/53715fig5.jpg" /> 
Figure 5. Fluorescent Actin-nucleus Stain images of the 3T3 Cells Encapsulated in the Gradient Chips on Day 3. The cell alignment direction in (a-c) line 1, (d-f) line 7, and (g-i) line 12 reveal radial alignment, random alignment, and circular alignment, respectively. The green and blue colors show the actin and nucleus stains, respectively. The dotted white line represents the boundary of the hydrogel. Scale bar: 200 &#181;m. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/53715fig5large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
Supplementary Figure 1. Calculation of the Dome-shaped PDMS Curvature. H(x): the convex PDMS curve; H0: the maximum height difference between the PDMS dome before and after deformation; r: the radius of the dome; V: the over-injection volume of the blue region, which causes the PDMS deformation as a dome. See Supplementary File 1 for details. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/Supplementary_Figure_1.tif" target="_blank">Please click here to download this figure.</a>
Supplementary File 1. <a href="https://cloudflare-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/53715/Supplementary_File_1-SI.docx" target="_blank">Please click here to download this file.</a>

Watch this Scientific Journal Video about Gradient Strain Chip for Stimulating Cellular Behaviors in Cell-laden Hydrogel at JoVE.com

Gradient Strain Chip for Stimulating Cellular Behaviors in Cell-laden Hydrogel

This article introduces a simple approach to providing non-continuous gradient static strains on a concentric cell-laden hydrogel to regulate cell alignment for tissue engineering.

Artificial guidance for cellular alignment is a hot topic in the field of tissue engineering. Most of the previous research has investigated single ...

gradient-strain-chip-for-stimulating-cellular-behaviors-cell-laden

Nanyang Technological University

Research

JoVE Journal

Bioengineering

7.9K Views.  National Tsing Hua University. Dit artikel introduceert een eenvoudige benadering aan het verstrekken van niet-continu verlopende statische spanningen op een concentrische cel-beladen hydrogel te reguleren celuitlijning voor weefselengineering.

Video: Gradient Strain Chip for Stimulating Cellular Behaviors in Cell-laden Hydrogel

Microfabricated Platforms for Mechanically Dynamic Cell Culture

The ability to systematically probe in vitro cellular response to combinations of mechanobiological stimuli for tissue engineering, drug discovery or fundamental cell biology studies is limited by current bioreactor technologies, which cannot simultaneously apply a variety of mechanical stimuli to cultured cells. In order to address this issue, we have developed a series of microfabricated platforms designed to screen for the effects of mechanical stimuli in a high-throughput format. In this protocol, we demonstrate the fabrication of a microactuator array of vertically displaced posts on which the technology is based, and further demonstrate how this base technology can be modified to conduct high-throughput mechanically dynamic cell culture in both two-dimensional and three-dimensional culture paradigms.

In this protocol, we demonstrate the fabrication of a microactuator array of vertically displaced posts on which the technology is based, and how this base technology can be modified to conduct high-throughput mechanically dynamic cell culture in both two-dimensional and three-dimensional culture paradigms.

The ability to systematically probe in vitro cellular response to combinations of mechanobiological stimuli for tissue engineering, drug discovery or ...

A Gradient-generating Microfluidic Device for Cell Biology

 The fabrication and operation of a gradient-generating microfluidic device for studying cellular behavior is described.  A microfluidic platform is an enabling experimental tool, because it can precisely manipulate fluid flows, enable high-throughput experiments, and generate stable soluble concentration gradients.  Compared to conventional gradient generators, poly(dimethylsiloxane) (PDMS)-based  microfluidic devices can generate stable concentration gradients of growth factors with well-defined profiles.  Here, we developed simple gradient-generating microfluidic devices with three separate inlets.  Three microchannels combined into one microchannel to generate concentration gradients.  The stability and shape of growth factor gradients were confirmed by fluorescein isothyiocyanate (FITC)-dextran with a molecular weight similar to epidermal growth factor (EGF).  Using this microfluidic device, we demonstrated that fibroblasts exposed to concentration gradients of EGF migrated toward higher concentrations.  The directional orientation of cell migration and motility of migrating cells were quantitatively assessed by cell tracking analysis.  Thus, this gradient-generating microfluidic device might be useful for studying and analyzing the behavior of migrating cells.

We describe a protocol for the microfabrication of the gradient-generating microfluidic device that can generate spatial and temporal gradients in well-defined microenvironment. In this approach, the gradient-generating microfluidic device can be used to study directed cell migration, embryogenesis, wound healing, and cancer metastasis.

 The fabrication and operation of a gradient-generating microfluidic device for studying cellular behavior is described.  A microfluidic platform is an ...

Biology

Designing Microfluidic Devices for Studying Cellular Responses Under Single or Coexisting Chemical/Electrical/Shear Stress Stimuli

Microfluidic devices are capable of creating a precise and controllable cellular micro-environment of pH, temperature, salt concentration, and other physical or chemical stimuli. They have been commonly used for in vitro cell studies by providing in vivo like surroundings. Especially, how cells response to chemical gradients, electrical fields, and shear stresses has drawn many interests since these phenomena are important in understanding cellular properties and functions. These microfluidic chips can be made of glass substrates, silicon wafers, polydimethylsiloxane (PDMS) polymers, polymethylmethacrylate (PMMA) substrates, or polyethyleneterephthalate (PET) substrates. Out of these materials, PMMA substrates are cheap and can be easily processed using laser ablation and writing. Although a few microfluidic devices have been designed and fabricated for generating multiple, coexisting chemical and electrical stimuli, none of them was considered efficient enough in reducing experimental repeats, particular for screening purposes. In this report, we describe our design and fabrication of two PMMA-based microfluidic chips for investigating cellular responses, in the production of reactive oxygen species and the migration, under single or coexisting chemical/electrical/shear stress stimuli. The first chip generates five relative concentrations of 0, 1/8, 1/2, 7/8, and 1 in the culture regions, together with a shear stress gradient produced inside each of these areas. The second chip generates the same relative concentrations, but with five different electric field strengths created within each culture area. These devices not only provide cells with a precise, controllable micro-environment but also greatly increase the experimental throughput.

Micro-fabricated devices integrated with fluidic components provide an in vitro platform for cell studies mimicking the in vivo micro-environment. We developed polymethylmethacrylate-based microfluidic chips for studying cellular responses under single or coexisting chemical/electrical/shear stress stimuli.

Microfluidic devices are capable of creating a precise and controllable cellular micro-environment of pH, temperature, salt concentration, and other ...

A Microfluidic Platform for Stimulating Chondrocytes with Dynamic Compression

Mechanical stimuli are known to modulate biological functions of cells and tissues. Recent studies have suggested that compressive stress alters growth plate cartilage architecture and results in growth modulation of long bones of children. To determine the role of compressive stress in bone growth, we created a microfluidic device actuated by pneumatic pressure, to dynamically (or statically) compress growth plate chondrocytes embedded in alginate hydrogel cylinders. In this article, we describe detailed methods for fabricating and characterizing this device. The advantages of our protocol are: 1) Five different magnitudes of compressive stress can be generated on five technical replicates in a single platform, 2) It is easy to visualize cell morphology via a conventional light microscope, 3) Cells can be rapidly isolated from the device after compression to facilitate downstream assays, and 4) The platform can be applied to study mechanobiology of any cell type that can grow in hydrogels.

This article provides detailed methods for fabricating and characterizing a pneumatically actuating microfluidic device for chondrocyte compression.

Mechanical stimuli are known to modulate biological functions of cells and tissues. Recent studies have suggested that compressive stress alters growth ...

Traction Microscopy Integrated with Microfluidics for Chemotactic Collective Migration

Cells change migration patterns in response to chemical stimuli, including the gradients of the stimuli. Cellular migration in the direction of a chemical gradient, known as chemotaxis, plays an important role in development, the immune response, wound healing, and cancer metastasis. While chemotaxis modulates the migration of single cells as well as collections of cells in vivo, in vitro research focuses on single-cell chemotaxis, partly due to the lack of the proper experimental tools. To fill that gap, described here is a unique experimental system that combines microfluidics and micropatterning to demonstrate the effects of chemical gradients on collective cell migration. Furthermore, traction microscopy and monolayer stress microscopy are incorporated into the system to characterize changes in cellular force on the substrate as well as between neighboring cells. As proof-of-concept, the migration of micropatterned circular islands of Madin-Darby canine kidney (MDCK) cells is tested under a gradient of hepatocyte growth factor (HGF), a known scattering factor. It is found that cells located near the higher concentration of HGF migrate faster than those on the opposite side within a cell island. Within the same island, cellular traction is similar on both sides, but intercellular stress is much lower on the side of higher HGF concentration. This novel experimental system can provide new opportunities to studying the mechanics of chemotactic migration by cellular collectives.

Collective cell migration in development, wound healing, and cancer metastasis is often guided by the gradients of growth factors or signaling molecules. Described here is an experimental system combining traction microscopy with a microfluidic system and a demonstration of how to quantify the mechanics of collective migration under biochemical gradient.

Cells change migration patterns in response to chemical stimuli, including the gradients of the stimuli. Cellular migration in the direction of a chemical ...

Controlled Strain of 3D Hydrogels under Live Microscopy Imaging

External forces are an important factor in tissue formation, development, and maintenance. The effects of these forces are often studied using specialized in vitro stretching methods. Various available systems use 2D substrate-based stretchers, while the accessibility of 3D techniques to strain soft hydrogels, is more restricted. Here, we describe a method that allows external stretching of soft hydrogels from their circumference, using an elastic silicone strip as the sample carrier. The stretching system utilized in this protocol is constructed from 3D-printed parts and low-cost electronics, making it simple and easy to replicate in other labs. The experimental process begins with polymerizing thick (&#62;100 &#956;m) soft fibrin hydrogels (Elastic Modulus of ~100 Pa) in a cut-out at the center of a silicone strip. Silicone-gel constructs are then attached to the printed-stretching device and placed on the confocal microscope stage. Under live microscopy the stretching device is activated, and the gels are imaged at various stretch magnitudes. Image processing is then used to quantify the resulting gel deformations, demonstrating relatively homogenous strains and fiber alignment throughout the gel&#8217;s 3D thickness (Z-axis). Advantages of this method include the ability to strain extremely soft hydrogels in 3D while executing in situ microscopy, and the freedom to manipulate the geometry and size of the sample according to the user&#8217;s needs. Additionally, with proper adaptation, this method can be used to stretch other types of hydrogels (e.g., collagen, polyacrylamide or polyethylene glycol) and can allow for analysis of cells and tissue response to external forces under more biomimetic 3D conditions.

The presented method involves uniaxial stretching of 3D soft hydrogels embedded in silicone rubber while allowing live confocal microscopy. Characterization of the external and internal hydrogel strains as well as fiber alignment are demonstrated. The device and protocol developed can assess the response of cells to various strain regimes.

External forces are an important factor in tissue formation, development, and maintenance. The effects of these forces are often studied using specialized ...

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.

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 ...

Mechanical Stimulation of Chondrocyte-agarose Hydrogels

Articular cartilage suffers from a limited repair capacity when damaged by mechanical insult or degraded by disease, such as osteoarthritis. To remedy this deficiency, several medical interventions have been developed. One such method is to resurface the damaged area with tissue-engineered cartilage; however, the engineered tissue typically lacks the biochemical properties and durability of native cartilage, questioning its long-term survivability. This limits the application of cartilage tissue engineering to the repair of small focal defects, relying on the surrounding tissue to protect the implanted material. To improve the properties of the developed tissue, mechanical stimulation is a popular method utilized to enhance the synthesis of cartilaginous extracellular matrix as well as the resultant mechanical properties of the engineered tissue. Mechanical stimulation applies forces to the tissue constructs analogous to those experienced in vivo. This is based on the premise that the mechanical environment, in part, regulates the development and maintenance of native tissue1,2. The most commonly applied form of mechanical stimulation in cartilage tissue engineering is dynamic compression at physiologic strains of approximately 5-20% at a frequency of 1 Hz1,3. Several studies have investigated the effects of dynamic compression and have shown it to have a positive effect on chondrocyte metabolism and biosynthesis, ultimately affecting the functional properties of the developed tissue4-8. In this paper, we illustrate the method to mechanically stimulate chondrocyte-agarose hydrogel constructs under dynamic compression and analyze changes in biosynthesis through biochemical and radioisotope assays. This method can also be readily modified to assess any potentially induced changes in cellular response as a result of mechanical stimuli.

The biosynthesis of cartilaginous extracellular matrix by chondrocytes can be affected by application of mechanical stimuli. This method describes the technique of applying dynamic compressive strains to chondrocytes encapsulated in 3D constructs and the evaluation of induced changes in chondrocyte metabolism.

Articular cartilage suffers from a limited repair capacity when damaged by mechanical insult or degraded by disease, such as osteoarthritis. To remedy ...

Pattern Generation for Micropattern Traction Microscopy

Micropattern traction microscopy allows control of the shape of single cells and cell clusters. Furthermore, the ability to pattern at the micrometer length scale allows the use of these patterned contact zones for the measurement of traction forces, as each micropatterned dot allows for the formation of a single focal adhesion that then deforms the soft, underlying hydrogel. This approach has been used for a wide range of cell types, including endothelial cells, smooth muscle cells, fibroblasts, platelets, and epithelial cells. 
This review describes the evolution of techniques that allow the printing of extracellular matrix proteins onto polyacrylamide hydrogels in a regular array of dots of prespecified size and spacing. As micrometer-scale patterns are difficult to directly print onto soft substrates, patterns are first generated on rigid glass coverslips that are then used to transfer the pattern to the hydrogel during gelation. First, the original microcontact printing approach to generate arrays of small dots on the coverslip is described. A second step that removes most of the pattern to leave islands of small dots is required to control the shapes of cells and cell clusters on such arrays of patterned dots. 
Next, an evolution of this approach that allows for the generation of islands of dots using a single subtractive patterning step is described. This approach is greatly simplified for the user but has the disadvantage of a decreased lifetime for the master mold needed to make the patterns. Finally, the computational approaches that have been developed for the analysis of images of displaced dots and subsequent cell-generated traction fields are described, and updated versions of these analysis packages are provided.

We describe improvements to a standard method for measuring cellular traction forces, based on microcontact printing with a single subtractive patterning step of dot arrays of extracellular matrix proteins on soft hydrogels. This method allows for simpler and more consistent fabrication of island patterns, essential for controlling cell cluster shape.

Micropattern traction microscopy allows control of the shape of single cells and cell clusters. Furthermore, the ability to pattern at the micrometer ...

Studying Tissue Rigidity and Shear Stress Effects on Endothelial Cells

In Vitro Model Integrating Substrate Stiffness and Flow to Study Endothelial Cell Responses

We present an innovative in vitro model aimed at investigating the combined effects of tissue rigidity and shear stress on endothelial cell (EC) function, which are crucial for understanding vascular health and the onset of diseases such as atherosclerosis. Traditionally, studies have explored the impacts of shear stress and substrate stiffness on ECs, independently. However, this integrated system combines these factors to provide a more precise simulation of the mechanical environment of the vasculature. The objective is to examine EC mechanotransduction across various tissue stiffness levels and flow conditions using human ECs. We detail the protocol for synthesizing gelatin methacrylate (GelMA) hydrogels with tunable stiffness and seeding them with ECs to achieve confluency. Additionally, we describe the design and assembly of a cost-effective flow chamber, supplemented by computational fluid dynamics simulations, to generate physiological flow conditions characterized by laminar flow and appropriate shear stress levels. The protocol also incorporates fluorescence labeling for confocal microscopy, enabling the assessment of EC responses to both tissue compliance and flow conditions. By subjecting cultured ECs to multiple integrated mechanical stimuli, this model enables comprehensive investigations into how factors such as hypertension and aging may affect EC function and EC-mediated vascular diseases. The insights gained from these investigations will be instrumental in elucidating the mechanisms underlying vascular diseases and in developing effective treatment strategies.

We synthesized and characterized a tunable gelatin-based substrate for culturing vascular endothelial cells (ECs) under relevant vascular flow conditions. This biomimetic surface replicates both physiological and pathological conditions, enabling the study of mechanical forces on EC behavior and advancing our understanding of vascular health and disease mechanisms.

We present an innovative in vitro model aimed at investigating the combined effects of tissue rigidity and shear stress on endothelial cell (EC) function, ...

Gradient Strain Chip for Stimulating Cellular Behaviors in Cell-laden Hydrogel

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