This study was supported by the Ministry of National Defense (MAB-105-070; MAB-106-077; MAB-107-032; MAB-107-065), the Ministry of Science and Technology (MOST 107-2320-B016-016), Tri-Service General Hospital, the National Defense Medical Center, Taiwan (TSGH-C106-046; TSGH-C106-115; TSGH-C107-041), and Cheng-Hsin General Hospital and National Defense Medical Center Cooperation (CH-NDMC-107-8).

The fat tissues were obtained from orthopedic surgeries as certified by the Institutional Review Board of Tri-Service General Hospital, Taipei, Taiwan, R.O.C. Procedures involving animal subjects have been approved by the Animal Care Committee at National Defense Medical Center, Taiwan (R.O.C).
1. Wet Spinning Process
<ol>
	<li>Solution preparation
	<ol>
		<li>Dissolve 5 g of gelatin powder in 100 mL of double-distilled water to obtain 5% (w/v) solution concentration.</li>
		<li>Stir the mixture slowly at 60 - 70 &#176;C overnight to achieve a completely homogeneous dispersion without any bubbles.</li>
	</ol>
	</li>
	<li>Wet spinning
	<ol>
		<li>Fiber formation 
		NOTE: A schematic of the method is shown in Figure 1A. The spinning setup is equipped with a peristaltic pump machine (see Table of Materials) that provides high-precision speed and controls the smooth delivery of the flow.
		<ol>
			<li>Plug a 26 G x 1/2 inch (0.45 mm x 12 mm) syringe as the solution injector into clear vinyl tubing.</li>
			<li>Prepare 100 mL of 99.5% acetone solution in a beaker glass to be used as the coagulation bath.</li>
			<li>Run the peristaltic pump machine at 21 rpm (3.25 mL/min) and let the gelatin solution stand for several seconds in acetone solution before rolling it up.</li>
			<li>Take out the gelatin fibers from the acetone solution and immerse them into 2.5% (w/v) of polycaprolactone/dichloromethane (PCL/DCM) solution with 1:20 (w/w).</li>
			<li>Let the gelatin fibers in the PCL/DCM solution dry overnight, under the hood at room temperature.</li>
		</ol>
		</li>
		<li>Tube formation 
		NOTE: A schematic of the method is shown in Figure 1B.
		<ol>
			<li>Use a peripheral venous catheter of 24 G x 3/4 inch (0.7 mm x 19 mm) as tube mold.</li>
			<li>Load the catheter into the gelatin solution and hold it there for 3 s.</li>
			<li>Load the catheter into the acetone solution and hold for 1 min.</li>
			<li>Load the catheter again into the gelatin solution and hold it there for 3 s.</li>
			<li>Repeat this alternate procedure 20x.</li>
			<li>Take out the catheter from the acetone solution and let the molded tube dry at room temperature for 5 min.</li>
			<li>Gently take out the gelatin tube from the catheter by using a hemostat and carefully transfer the gelatin tube into a glass Pasteur pipette with the immersion of 2.5% (w/v) of the PCL/DCM solution.</li>
			<li>Let the Pasteur pipette containing the gelatin tube and PCL/DCM solution dry overnight, under the hood at room temperature.</li>
		</ol>
		</li>
	</ol>
	</li>
</ol>
2. Morphology of the Gelatin Tube
<ol>
	<li>Mount the piece of dried gelatin tube on a carbon stub.</li>
	<li>Put the sample into the ion sputter coater machine (see Table of Materials) and coat the sample with gold for 60 s; then, observe its morphology by scanning electron microscopy(see Table of Materials).</li>
</ol>
3. Culture of Human Adipose Stem Cells
<ol>
	<li>Place the fat tissues (obtained from oral adipose tissue by clinical surgery) into a Petri dish containing 10 mL of transfer solution consisting of 0.1 M phosphate-buffered saline (PBS), 1% penicillin/streptomycin, and 0.1% glucose.</li>
	<li>Cut the tissues into small pieces (less than 1 mm2) with a scalpel blade.</li>
	<li>Transfer the tissues into a 15 mL plastic tube and centrifuge at 500 x g for 5 min.</li>
	<li>Remove the supernatant and incubate the sediment in a plastic tube containing 10 mL of Dulbecco&#39;s modified Eagle&#39;s medium (DMEM, consisting of 4 mM L-glutamine, 1 mM sodium pyruvate, and 15 mg/L phenol red) with 0.1% collagenase at 37 &#176;C, in a humidified atmosphere containing 95% air and 5% CO2, for 1 day.</li>
	<li>Centrifuge the plastic tube at 500 x g for 5 min, discard the supernatant carefully (do not touch the sediment), and then, suspend the sediment gently by adding 10 mL of DMEM containing 10% fetal bovine serum (FBS) and let it stand for 1 day at 37 &#176;C, in a humidified atmosphere containing 95% air and 5% CO2.</li>
	<li>Centrifuge the plastic tube at 500 x g for 5 min and discard the supernatant carefully; then, add 1 mL of stem cell medium (consisting of keratinocyte serum-free medium (K-SFM), 5% of FBS, N-acetyl-L-cysteine, ascorbic acid-2-phosphate, streptomycin, and amphotericin) to suspend the sediment, transfer it into a T25 culture flask containing 4 mL of stem cell medium, and let it stand for 3 days at 37 &#176;C, in a humidified atmosphere containing 95% air and 5% CO2.</li>
	<li>Discard the supernatant, add 1 mL of 0.25% trypsin-ethylenediaminetetraacetic acid (EDTA), and incubate it at 37 &#176;C, in a humidified atmosphere containing 95% air and 5% CO2 for 3.5 min, to detach the cells from the bottom of the culture flask.</li>
	<li>Add 1 mL of FBS, suspend the mixture, and transfer it to a microcentrifuge tube.</li>
	<li>Centrifuge the suspension at 500 x g for 3.5 min, suspend the sediment with 1 mL of stem cell medium, and transfer it to the culture flask containing 5 mL of stem cell medium; then, keep it at 37 &#176;C in a humidified atmosphere containing 95% air and 5% CO2.</li>
	<li>Maintain the subculture by changing the stem cell medium every 3 days.</li>
</ol>
4. Cultivation of Cells on the Gelatin Tube
<ol>
	<li>Sterilize the gelatin tube with UV light for 2 h; then, immerse it in 75% (v/v) ethanol and wash it 2x with stem cell medium to remove the residual ethanol.</li>
	<li>Collect the human adipose stem cells (hASCs) from the culture flask.
	<ol>
		<li>Remove the medium from the culture flask.</li>
		<li>Add 1 mL of 0.25% trypsin-EDTA and incubate at 37 &#176;C in a humidified atmosphere containing 95% air and 5% CO2 for 3.5 min, to detach hASCs from the bottom of the culture flask.</li>
		<li>Add 1 mL of FBS, suspend the mixture, and transfer it to a microcentrifuge tube.</li>
		<li>Centrifuge the microcentrifuge tube at 500 x g for 3.5 min; then, discard the supernatant carefully, and suspend the sediment with 1 mL of stem cell medium.</li>
	</ol>
	</li>
	<li>Place the gelatin tube in a 6-well plate containing 3 mL of stem cell medium and seed 4 x 104 of hASCs to the tube; incubate them for 2 weeks at 37 &#176;C, in a humidified atmosphere containing 95% air and 5% CO2.</li>
	<li>Change the stem cell medium every 3 days to provide enough nutrition for cell growth.</li>
</ol>
5. Immunocytochemistry
<ol>
	<li>Remove the stem cell medium and wash the well with PBS.</li>
	<li>Fix the tube with cells with 0.1% acetic acid glacial for 30 min at room temperature.</li>
	<li>Wash the well 2x with PBS.</li>
	<li>Add a surfactant (NP-40, 0.05%) and incubate for 10 min at room temperature.</li>
	<li>Wash the well 3x with PBS.</li>
	<li>Add 2% normal goat serum and incubate at room temperature for 30 min.</li>
	<li>Decant the solution, add primary antibody nestin (neural progenitor cell marker), and incubate overnight at 4 &#176;C.</li>
	<li>Wash the well 3x with PBS.</li>
	<li>Add secondary antibody donkey anti-mouse-fluorescein isothiocyanate (FITC) and keep the sample in the dark for 2 h.</li>
	<li>Wash the well 3x with PBS.</li>
	<li>Add 1 mL of Hoechst 33342 to stain the nuclei and incubate for 10 min at room temperature.</li>
	<li>Wash the well gently and observe it under a fluorescence microscope.</li>
</ol>
6. In Vivo Biocompatibility Test
NOTE: Rats with a weight between 201 - 225 g have been successfully tested using this protocol.
<ol>
	<li>Anesthesia
	<ol>
		<li>Induce and maintain anesthesia; preferably, anesthetize a rat (8-week-old Sprague Dawley rat, female) via an intramuscular injection of tiletamine and zolazepam (25 mg/kg + 25 mg/kg, respectively) and xylazine (5 mg/kg). Perform a pain stimulation test by pressing the rat&#39;s fingers to confirm the anesthetization.</li>
	</ol>
	</li>
	<li>Sterilization of the surgical site
	<ol>
		<li>Shave and clean the skin of rat&#8217;s trapezius area (over the back of the neck) and wipe it with povidone-iodine lotion to prepare the aseptic condition of skin (100 mg/mL).</li>
		<li>Cover the rat with sterile surgical drapes to reduce bacterial transfer and subsequent contamination of the surgical site.</li>
	</ol>
	</li>
	<li>Implantation of the gelatin tube
	<ol>
		<li>Gently cut the skin of rat&#8217;s trapezius area with a surgical scissor.</li>
		<li>Create a wound of 2 cm and place the gelatin tube directly on the layer between the fascia and the muscle.</li>
	</ol>
	</li>
	<li>Postimplantation surgery
	<ol>
		<li>Close the wound with nylon suture and wipe it with povidone-iodine solution (100 mg/mL).</li>
		<li>Induce the rat via an intramuscular injection of ketoprofen (2.5 mg/kg) and cefazolin (15 mg/kg).</li>
		<li>Keep the rat under aseptic conditions for 7 days.</li>
	</ol>
	</li>
	<li>Euthanasia
	<ol>
		<li>Rat is euthanized via CO2 asphyxiation in accordance with AVMA Guidelines for Euthanasia. Confirm the death of rat by toe and tail pinch, followed by chest touching to ensure the heartbeat.</li>
		<li>Cut the rat gently with a scalpel blade, remove the implanted tissue, and take photos for observation.</li>
	</ol>
	</li>
</ol>

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The authors have nothing to disclose.

We presented the development of gelatin-based biomaterials by using a simple wet spinning technique that can be applied in the study of natural polymers for tissue regeneration. This work demonstrated the possibility of gelatin fabrication as a great protein source without the addition of other sources, with the aim to optimize the properties of gelatin itself. The development of gelatin-based biomaterials was entirely carried out in room temperature (22 - 26 &#176;C). A gentle solution preparation is a critical step within the protocol, as is maintaining an exact solution concentration and stirring the solution very smoothly to avoid any bubbles. The appearance of bubbles in the solution, especially when loaded into the syringe, will affect the form of gelatin in the coagulation medium.
Gelatin-based biomaterials were obtained through the phase transition of the gelatin solution in a suitable coagulation medium that causes the desolvation of the polymer molecules and precipitates the gelatin solution into a solid polymer phase11. By employing this basic concept of the wet spinning method, we utilized the shape of the gelatin into tube form to increase the functionality of its surface and to mimic the features of human tissues through the molding process. In this case, the adequate times to load, hold, and repeat the steps are necessary to build up an exact tube form. In this protocol, the homogeneity of the tube shape during the process cannot be controlled, considering the asynchronously of the transition phase from the bottom to the up part of the tube. Additionally, the preserved and sterilization methods for this material are limited because of its protein content.
Examined by in vitro and in vivo tests, the current results demonstrated that the gelatin tubes created with this protocol exhibit an excellent biocompatibility and have great potential applicability in tissue engineering. The inner diameter of the gelatin tubes obtained in this study was more than 200 &#181;m (Figure 3B), which allows the cells to pass through and easily adhere to the tubes&#39; inner surface16. In addition, the thickness and mildly asperous surface of the tube provides a proper template for cell attachment (Figure 3A,C). Consistently, the immunocytochemistry results showed that the cells penetrated and adhered to the gelatin tube-particularly the neural stem cells, as observed through the positive staining of the neural cells with the marker nestin (Figure 4B). As the nestin marker was found in the area of the growth cones during the stages of axon extension, the positive staining of nestin is indicative of the cone growth17. Ultimately, this positive gelatin-hASCs combination can be implemented in the treatment of nervous system regeneration. The spinal cord is made up of groups of neuron axons and bundles of nerve fibers along the spinal vertebrae. Thus, the gelatin tube may not only provide appropriate filling gap material but will also guide nerve fiber growth and regeneration through the tube. Moreover, the great biocompatibility observed by the implantation of a gelatin tube into the fascia layer did not demonstrate any harmful and inflammation effect on the local tissues, and the tube was surrounded by well-developed connective tissues (Figure 5A,B).
In conclusion, we successfully developed a simple and useful protocol for the construction of gelatin-based biomaterials with excellent biocompatibility and cell environment that can be used in prospect tissue engineering applications, such as in nervous system regeneration, blood vessel replacement, urinary tract reconstruction, and in the rebuilding of organ parts. This protocol can be implemented by tissue engineers who are currently developing correlative natural polymers without any synthetic polymer addition. The protocol can be easily expanded upon and can be used to customize the design of specific biomaterials. In the future, this protocol can be adapted to produce protein-based biomaterials on a mass scale, thus helping to realize tissue regeneration.

The latest development in tissue regeneration involves the application of tissue engineering, which represents a challenge for the improvement of new therapeutic strategies in medical treatments. For example, the limited potential of nervous system regeneration, following injury or disease, poses a significant health problem worldwide. Due to the complexity of pathophysiological processes associated with the nervous system, the use of traditional autograft or the implementation of stabilization surgery has been shown to offer benefits in functional outcomes, but there is no strong evidence for the effects of spinal fixation surgery1,2. The tissue at the damaged area is lost and replaced with hypertrophically induced astrocytes3, eventually forming a dense glial scar4,5. This matrix acts as a barrier that blocks the recovery of nerve function6,7 and is, thus, greatly hinders regeneration. Therefore, a suitable filling gap material is expected to prevent the loss of tissue and reduce the formation of scar-associated connective tissue by maintaining the integrity of the damaged area, as well as by providing the direct replacement of neural cells and circuitry to promote axon regeneration.
Polymeric biomaterials have been preferred as scaffolds for tissue regeneration therapy, based on the regulation of cell or axon behavior and tissue progression through natural extracellular matrix (ECM) support. The fiber format is commonly considered as a building block for various materials, owing to its one-dimensional structure8. The fibers can generally be obtained by melt extrusion or wet spinning method; however, the large size and cost of the equipment and the difficulty to perform these methods are challenging. In addition, the majority of the work related to polymer fibers has been focused on synthetic or composite materials. Natural polymers as a source of biomaterial offer better biocompatibility properties for the human body. Nonetheless, to obtain the alignment of natural polymer fibers is relatively more difficult than of synthetic polymer sources9. Hence, the conversion of a natural polymer as a rich source of protein into biomaterial fibers is an important strategy &#8212; not only can the biomaterial fibers be directly isolated from the raw material, thus avoiding an unnecessary transformation to monomers, but the protein fibers also have a good appearance and favorable characteristics10.
In this regard, we describe an inexpensive processing method for the manufacturing of natural polymer fibers through the basic concept of wet spinning, that can be implemented on the laboratory scale for tissue engineering. Wet spinning is performed by the extrusion and coagulation of a polymer solution into a suitable polymer nonsolvent. An appropriate, viscous solution doped into coagulation medium causes the polymer molecules to dissolve. Through the phase transition, the filaments then lose their solubility and are precipitated in the form of a solid polymer phase11. Referring to this concept, we then expanded the development of gelatin into the tube form by a molding process, which is considered proper for tissue regeneration application. In addition, intrinsically, we can also develop any shape of material from gelatin fibers (e.g., gelatin conduit rolled up from several gelatin fibers), for other desired applications.
Gelatin, a biodegradable natural polymer, is formed from denatured and hydrolyzed collagen, including any semicrystalline, amorphous, or triple helical state of collagen12. It is well known that collagen is the essential structural protein in all connective tissues of vertebrates and invertebrates13,14, which is similar to the protein structure of the main ECM that induces nerve growth and, simultaneously, replaces a large amount of glycosaminoglycan secreted during spinal cord injuries. Therefore, the use of gelatin as a source would be a great choice for any medical vehicle. Besides being an inexpensive source, gelatin is also biodegradable and cytocompatible and clinically proven to be a temporary defect filler15. Developed into a tube form, in vitro and in vivo tests described here demonstrate that gelatin has an excellent biocompatibility and suitability for future tissue engineering applications. Cultured with human adipose stem cells, gelatin tubes improve cell differentiation into neural progenitor cells by using positive nestin staining as a neural cell marker. Furthermore, gelatin as filling gap material, as produced by the method established in this study, is expected to be manageable and safe and to greatly benefit tissue engineers who are currently developing correlative natural polymers for the enhancement of tissue regeneration strategies.

<table><tbody><tr><td>&lt;strong&gt;Solution preparation:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Gelatin type B (porcine)</td><td>Ferak</td><td>Art. -Nr. 10733</td><td>500 g vial</td></tr><tr><td>&lt;strong&gt;Wet spinning process:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Peristaltic pump</td><td>Gilson</td><td>Model M312</td><td>Minipuls*3</td></tr><tr><td>Plastic tube connector</td><td>World Precision Instruments</td><td>14011</td><td>1 box</td></tr><tr><td>Syringe</td><td>Sterican</td><td>5A06258541</td><td>26Gx1/2&quot;(0.45 x 12mm)</td></tr><tr><td>Acetone</td><td>Ferak</td><td>Art. -Nr. 00010</td><td>2.5 L vial</td></tr><tr><td>Polycaprolactone CAPA 6500</td><td>Perstorp</td><td>24980-41-4</td><td>-</td></tr><tr><td>Dichloromethane&amp;nbsp;</td><td>Scharlau</td><td>CL03421000</td><td>1 L vial</td></tr><tr><td>Glass Pasteur pipette</td><td>Fisher Scientific</td><td>13-678-20A</td><td>-</td></tr><tr><td>Hemostat</td><td>Shinetec instruments</td><td>ST-B021</td><td>-</td></tr><tr><td>Peripheral venous catheter (Introcan Certo)</td><td>B. Braun</td><td>1B03258241</td><td>24Gx3/4&quot;(0.7 x 19mm)</td></tr><tr><td>&lt;strong&gt;Morphology of the gelatin tube:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Ion sputter coater machine&amp;nbsp;</td><td>Hitachi</td><td>e1010</td><td>-</td></tr><tr><td>Scanning electron microscopy</td><td>Hitachi</td><td>S-3000N</td><td>-</td></tr><tr><td>&lt;strong&gt;Cultivation of cells on the gelatin tube:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Trypsin-EDTA</td><td>Gibco</td><td>488625</td><td>100 mL vial</td></tr><tr><td>Fetal bovine serum</td><td>Gibco</td><td>923119</td><td>500 mL vial</td></tr><tr><td>Dulbecco&#039;s modified Eagle&#039;s medium&amp;nbsp;</td><td>Gibco</td><td>31600-034</td><td>Powder</td></tr><tr><td>Keratinocyte-SFM medium</td><td>Gibco</td><td>10744-019</td><td>500 mL vial</td></tr><tr><td>T25 culture flask</td><td>TPP</td><td>90025</td><td>VENT type</td></tr><tr><td>6-well plate</td><td>Falcon</td><td>1209938</td><td>-</td></tr><tr><td>&lt;strong&gt;Immunocytochemistry:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Phospate-buffered saline</td><td>Gibco</td><td>654471</td><td>500 mL vial</td></tr><tr><td>Acetic acid glacial</td><td>Ferak</td><td>Art. -Nr. 00697</td><td>500 mL vial</td></tr><tr><td>NP-40 surfactant (Tergitol solution)</td><td>Sigma</td><td>056K0151</td><td>500 mL vial</td></tr><tr><td>Normal goat serum</td><td>Vector Laboratories</td><td>S-1000-20</td><td>20 mL vial, concentrate</td></tr><tr><td>Nestin (primary antibody)</td><td>Santa Cruz Biotechnology</td><td>SC-23927</td><td>-</td></tr><tr><td>Donkey anti-mouse-fluorescein isothiocyanate (secondary antibody)</td><td>Santa Cruz Biotechnology</td><td>SC-2099</td><td>-</td></tr><tr><td>Hoechst 33342</td><td>Anaspec</td><td>AS-83218</td><td>5 mL vial</td></tr><tr><td>&lt;strong&gt;&lt;em&gt;In vivo&lt;/em&gt; biocompatibility test:&lt;/strong&gt;</td><td></td><td></td><td></td></tr><tr><td>Tiletamine+zolazepam&amp;nbsp;</td><td>Virbac</td><td>BC91</td><td>5 mL vial</td></tr><tr><td>Xylazine</td><td>Bayer korea</td><td>KR03227</td><td>10 mL vial</td></tr><tr><td>Ketoprofen</td><td>Astar</td><td>1406232</td><td>2 mL vial</td></tr><tr><td>Povidone-iodine solution</td><td>Everstar</td><td>HA161202</td><td>4 L barrel</td></tr><tr><td>Cefazolin</td><td>China Chemical &amp;amp; Pharmaceutical</td><td>18P909</td><td>1 g vial</td></tr><tr><td>Scalpel blade</td><td>Shinetec instruments</td><td>ST-B021</td><td>-</td></tr><tr><td>Surgical scissor</td><td>Shinetec instruments</td><td>ST-B021</td><td>-</td></tr></tbody></table>

wet-spinning-based molding process of gelatin for tissue regeneration

This article presents an inexpensive method to fabricate gelatin, as a natural polymer, into monofilament fibers or other appropriate forms. Through the wet spinning method, gelatin fibers are produced by smooth extrusion in a suitable coagulation medium. To increase the functional surface of these gelatin fibers and their ability to mimic the features of tissues, gelatin can be molded into a tube form by referring to this concept. Examined by in vitro and in vivo tests, the gelatin tubes demonstrate a great potential for application in tissue engineering. Acting as a suitable filling gap material, gelatin tubes can be used to substitute the tissue in the damaged area (e.g., in the nervous or cardiovascular system), as well as to promote regeneration by providing a direct replacement of stem cells and neural circuitry. This protocol provides a detailed procedure for creating a biomaterial based on a natural polymer, and its implementation is expected to greatly benefit the development of correlative natural polymers, which help to realize tissue regeneration strategies.

In this study, we successfully developed the gelatin into fibers (Figure 2A) and tubes (Figure 2B,C) through the user-friendly wet spinning concept. These gelatin-based materials can be utilized as any medical tool, depending on their shapes. Considering that the functional surface and frame of such materials are more suitable for tissue regeneration, we examined the biocompatibility of gelatin tube by performing in vitro and in vivo tests.
To obtain an overview of the gelatin tube, first, we conducted morphological observations by using scanning electron microscopy (Figure 3). The results showed that the surface of the gelatin tube was not very smooth (Figure 3A), its inner diameter was more than 200 &#181;m (Figure 3B), and its thickness was approximately 20 &#181;m (Figure 3C).
Then, immunocytochemistry was performed to examine the biocompatibility of the gelatin tube in vitro (Figure 4). Initially, the gelatin tube was incubated with human adipose stem cells for 2 weeks and was stained with nestin as a neural cell marker (Figure 4A). Under the fluorescence microscope, the staining showed the positive nestin staining marker (green color) along with a bunch of detected nuclei (blue color), which suggested that the cells could penetrate and adhere well to the gelatin tube and differentiate into neural progenitor cells (Figure 4B). Further, the in vivo biocompatibility test of the gelatin tube showed analogous results. The gelatin tube was inserted into the fascia layer in a rat&#39;s trapezius area for 7 days. The result demonstrates that the implantation of the gelatin tube into the fascia layer indicated its safety and great biocompatibility, showed no indication of redness, swelling, or other inflammation symptoms of the local tissue, and was surrounded by well-developed connective tissues (Figure 5A,B).
<img alt="Figure 1" class="xfigimg" src="/files/ftp_upload/58932/58932fig1.jpg" /> 
Figure 1: Scheme of the wet spinning setup. (A) Gelatin fibers. (B) Gelatin tube. <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/58932/58932fig1large.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/58932/58932fig2.jpg" /> 
Figure 2: Bright view images of gelatin-based material obtained through the wet spinning method. (A) Gelatin fibers. (B) Gelatin tube. (C) Gelatin tube in PBS solution. <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/58932/58932fig2large.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/58932/58932fig3.jpg" /> 
Figure 3: Morphology of the gelatin tube, observed by scanning electron microscopy. (A) A superficial view of the surface of the gelatin tube (Scale bar = 50 &#181;m). (B) Inclined plane view (Scale bar = 200 &#181;m). (C) Vertical view (Scale bar = 25 &#181;m). <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/58932/58932fig3large.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/58932/58932fig4.jpg" /> 
Figure 4: Observation of human adipose stem cell (hASC) growth on the gelatin tube. (A) Bright view image of gelatin tube-hASCs. (B) Under the fluorescence microscope, the staining showed the positive nestin staining marker along with a bunch of nuclei stained with Hoechst 33342.The gelatin tube provided an optimal environment for hASCs to adhere and grow inside the tube (Scale bar = 200 &#181;m; blue color = nuclei; green color = nestin). <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/58932/58932fig4large.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/58932/58932fig5.jpg" /> 
Figure 5: In vivo biocompatibility test of the gelatin tube. (A) The implantation of the gelatin tube into the fascia layer. (B) Observation of the local tissue environment implanted with the gelatin tube (red rectangle area) after 7 days. No redness, swelling, or other inflammation symptoms were observed, and the tube is surrounded by well-developed connective tissues 7 days after implantation. This is indicative of an excellent biocompatibility of the gelatin tube with the local tissue environment. <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/58932/58932fig5large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>

Watch this Scientific Journal Video about Wet-spinning-based Molding Process of Gelatin for Tissue Regeneration at JoVE.com

Wet-spinning-based Molding Process of Gelatin for Tissue Regeneration

We developed and describe a protocol based on the wet spinning concept, for the construction of gelatin-based biomaterials used for the application of tissue engineering.

This article presents an inexpensive method to fabricate gelatin, as a natural polymer, into monofilament fibers or other appropriate forms. Through the ...

wet-spinning-based-molding-process-of-gelatin-for-tissue-regeneration

Nanyang Technological University

Research

JoVE Journal

Bioengineering

8.7K Views.  National Taiwan University of Science and Technology. Dit protocol verbetert een basisconcept van nat-spinnen om een materiaal van één enkele vezelstrip in diverse stroken te ontwikkelen die voor om het even welke aangewezen medische toepassing kunnen worden gebruikt. Deze techniek is eenvoudig, goedkoop, en vereist niet de toevoeging van synthetische polymeren voor aanpassing van de natuurlijke polymeren in een bron van eiwitten voor materialen van belang. Deze techniek kan worden uitgebreid tot biomateriaal ontwikkeling voor het doden van ko...

Video: Wet-spinning-based Molding Process of Gelatin for Tissue Regeneration

Microfluidic Dry-spinning and Characterization of Regenerated Silk Fibroin Fibers

The protocol demonstrates a method for mimicking the spinning process of silkworm. In the native spinning process, the contracting spinning duct enables the silk proteins to be compact and ordered by shearing and elongation forces. Here, a biomimetic microfluidic channel was designed to mimic the specific geometry of the spinning duct of the silkworm. Regenerated silk fibroin (RSF) spinning doped with high concentration, was extruded through the microchannel to dry-spin fibers at ambient temperature and pressure. In the post-treated process, the as-spun fibers were drawn and stored in ethanol aqueous solution. Synchrotron radiation wide-angle X-ray diffraction (SR-WAXD) technology was used to investigate the microstructure of single RSF fibers, which were fixed to a sample holder with the RSF fiber axis normal to the microbeam of the X-ray. The crystallinity, crystallite size, and crystalline orientation of the fiber were calculated from the WAXD data. The diffraction arcs near the equator of the two-dimensional WAXD pattern indicate that the post-treated RSF fiber has a high orientation degree.

A protocol for the microfluidic spinning and microstructure characterization of regenerated silk fibroin monofilament is presented.

The protocol demonstrates a method for mimicking the spinning process of silkworm. In the native spinning process, the contracting spinning duct enables ...

Chemistry

Printing Thermoresponsive Reverse Molds for the Creation of Patterned Two-component Hydrogels for 3D Cell Culture

Bioprinting is an emerging technology that has its origins in the rapid prototyping industry. The different printing processes can be divided into contact bioprinting1-4 (extrusion, dip pen and soft lithography), contactless bioprinting5-7 (laser forward transfer, ink-jet deposition) and laser based techniques such as two photon photopolymerization8. It can be used for many applications such as tissue engineering9-13, biosensor microfabrication14-16 and as a tool to answer basic biological questions such as influences of co-culturing of different cell types17. Unlike common photolithographic or soft-lithographic methods, extrusion bioprinting has the advantage that it does not require a separate mask or stamp. Using CAD software, the design of the structure can quickly be changed and adjusted according to the requirements of the operator. This makes bioprinting more flexible than lithography-based approaches.
Here we demonstrate the printing of a sacrificial mold to create a multi-material 3D structure using an array of pillars within a hydrogel as an example. These pillars could represent hollow structures for a vascular network or the tubes within a nerve guide conduit. The material chosen for the sacrificial mold was poloxamer 407, a thermoresponsive polymer with excellent printing properties which is liquid at 4 &deg;C and a solid above its gelation temperature ~20 &deg;C for 24.5% w/v solutions18. This property allows the poloxamer-based sacrificial mold to be eluted on demand and has advantages over the slow dissolution of a solid material especially for narrow geometries. Poloxamer was printed on microscope glass slides to create the sacrificial mold. Agarose was pipetted into the mold and cooled until gelation. After elution of the poloxamer in ice cold water, the voids in the agarose mold were filled with alginate methacrylate spiked with FITC labeled fibrinogen. The filled voids were then cross-linked with UV and the construct was imaged with an epi-fluorescence microscope.

A bioprinter was used to create patterned hydrogels based on a sacrificial mold. The poloxamer mold was backfilled with a second hydrogel and then eluted, leaving voids which were filled with a third hydrogel. This method uses fast elution and good printability of poloxamer to generate complex architectures from biopolymers.

Bioprinting is an emerging technology that has its origins in the rapid prototyping industry. The different printing processes can be divided into contact ...

Immunology and Infection

Engineering 3D Cellularized Collagen Gels for Vascular Tissue Regeneration

Synthetic materials are known to initiate clinical complications such as inflammation, stenosis, and infections when implanted as vascular substitutes. Collagen has been extensively used for a wide range of biomedical applications and is considered a valid alternative to synthetic materials due to its inherent biocompatibility (i.e., low antigenicity, inflammation, and cytotoxic responses). However, the limited mechanical properties and the related low hand-ability of collagen gels have hampered their use as scaffold materials for vascular tissue engineering. Therefore, the rationale behind this work was first to engineer cellularized collagen gels into a tubular-shaped geometry and second to enhance smooth muscle cells driven reorganization of collagen matrix to obtain tissues stiff enough to be handled.
The strategy described here is based on the direct assembling of collagen and smooth muscle cells (construct) in a 3D cylindrical geometry with the use of a molding technique. This process requires a maturation period, during which the constructs are cultured in a bioreactor under static conditions (without applied external dynamic mechanical constraints) for 1 or 2 weeks. The &#8220;static bioreactor&#8221; provides a monitored and controlled sterile environment (pH, temperature, gas exchange, nutrient supply and waste removal) to the constructs. During culture period, thickness measurements were performed to evaluate the cells-driven remodeling of the collagen matrix, and glucose consumption and lactate production rates were measured to monitor the cells metabolic activity. Finally, mechanical and viscoelastic properties were assessed for the resulting tubular constructs. To this end, specific protocols and a focused know-how (manipulation, gripping, working in hydrated environment, and so on) were developed to characterize the engineered tissues.

In this work, we present a technique for the rapid fabrication of living vascular tissues by direct culturing of collagen, smooth muscle cells and endothelial cells. In addition, a new protocol for the mechanical characterization of engineered vascular tissues is described.

Synthetic materials are known to initiate clinical complications such as inflammation, stenosis, and infections when implanted as vascular substitutes. ...

Fabricating Degradable Thermoresponsive Hydrogels on Multiple Length Scales via Reactive Extrusion, Microfluidics, Self-assembly, and Electrospinning

While various smart materials have been explored for a variety of biomedical applications (e.g., drug delivery, tissue engineering, bioimaging, etc.), their ultimate clinical use has been hampered by the lack of biologically-relevant degradation observed for most smart materials. This is particularly true for temperature-responsive hydrogels, which are almost uniformly based on polymers that are functionally non-degradable (e.g., poly(N-isopropylacrylamide) (PNIPAM) or poly(oligoethylene glycol methacrylate) (POEGMA)). As such, to effectively translate the potential of thermoresponsive hydrogels to the challenges of remote-controlled or metabolism-regulated drug delivery, cell scaffolds with tunable cell-material interactions, theranostic materials with the potential for both imaging and drug delivery, and other such applications, a method is required to render the hydrogels (if not fully degradable) at least capable of renal clearance following the required lifetime of the material. To that end, this protocol describes the preparation of hydrolytically-degradable hydrazone-crosslinked hydrogels on multiple length scales based on the reaction between hydrazide and aldehyde-functionalized PNIPAM or POEGMA oligomers with molecular weights below the renal filtration limit. Specifically, methods to fabricate degradable thermoresponsive bulk hydrogels (using a double barrel syringe technique), hydrogel particles (on both the microscale through the use of a microfluidics platform facilitating simultaneous mixing and emulsification of the precursor polymers and the nanoscale through the use of a thermally-driven self-assembly and cross-linking method), and hydrogel nanofibers (using a reactive electrospinning strategy) are described. In each case, hydrogels with temperature-responsive properties similar to those achieved via conventional free radical cross-linking processes can be achieved, but the hydrazone cross-linked network can be degraded over time to re-form the oligomeric precursor polymers and enable clearance. As such, we anticipate these methods (which may be generically applied to any synthetic water-soluble polymer, not just smart materials) will enable easier translation of synthetic smart materials to clinical applications.

Protocols are described for the fabrication of degradable thermoresponsive hydrogels based on hydrazone cross-linking of polymeric oligomers on the bulk scale, microscale, and nanoscale, the latter for preparation of both gel nanoparticles and nanofibers.

While various smart materials have been explored for a variety of biomedical applications (e.g., drug delivery, tissue engineering, bioimaging, etc.), ...

Microfluidic Bioprinting for Engineering Vascularized Tissues and Organoids

Engineering vascularized tissue constructs and organoids has been historically challenging. Here we describe a novel method based on microfluidic bioprinting to generate a scaffold with multilayer interlacing hydrogel microfibers. To achieve smooth bioprinting, a core-sheath microfluidic printhead containing a composite bioink formulation extruded from the core flow and the crosslinking solution carried by the sheath flow, was designed and fitted onto the bioprinter. By blending gelatin methacryloyl (GelMA) with alginate, a polysaccharide that undergoes instantaneous ionic crosslinking in the presence of select divalent ions, followed by a secondary photocrosslinking of the GelMA component to achieve permanent stabilization, a microfibrous scaffold could be obtained using this bioprinting strategy. Importantly, the endothelial cells encapsulated inside the bioprinted microfibers can form the lumen-like structures resembling the vasculature over the course of culture for 16 days. The endothelialized microfibrous scaffold may be further used as a vascular bed to construct a vascularized tissue through subsequent seeding of the secondary cell type into the interstitial space of the microfibers. Microfluidic bioprinting provides a generalized strategy in convenient engineering of vascularized tissues at high fidelity.

We provide a generalized protocol based on a microfluidic bioprinting strategy for engineering a microfibrous vascular bed, where a secondary cell type could be further seeded into the interstitial space of this microfibrous structure to generate vascularized tissues and organoids.

Engineering vascularized tissue constructs and organoids has been historically challenging. Here we describe a novel method based on microfluidic ...

3D Microtissues for Injectable Regenerative Therapy and High-throughput Drug Screening

To upgrade traditional 2D cell culture to 3D cell culture, we have integrated microfabrication with cryogelation technology to produce macroporous microscale cryogels (microcryogels), which can be loaded with a variety of cell types to form 3D microtissues. Herein, we present the protocol to fabricate versatile 3D microtissues and their applications in regenerative therapy and drug screening. Size and shape-controllable microcryogels can be fabricated on an array chip, which can be harvested off-chip as individual cell-loaded carriers for injectable regenerative therapy or be further assembled on-chip into 3D microtissue arrays for high-throughput drug screening. Due to the high elastic nature of these microscale cryogels, the 3D microtissues exhibit great injectability for minimally invasive cell therapy by protecting cells from mechanical shear force during injection. This ensures enhanced cell survival and therapeutic effect in the mouse limb ischemia model. Meanwhile, assembly of 3D microtissue arrays in a standard 384-multi-well format facilitates the use of common laboratory facilities and equipment, enabling high-throughput drug screening on this versatile 3D cell culture platform.

This protocol describes the fabrication of elastic 3D macroporous microcryogels by integrating microfabrication with cryogelation technology. Upon loading with cells, 3D microtissues are generated, which can be readily injected in vivo to facilitate regenerative therapy or assembled into arrays for in vitro high-throughput drug screening.

To upgrade traditional 2D cell culture to 3D cell culture, we have integrated microfabrication with cryogelation technology to produce macroporous ...

Advancing 3D Bioprinting with &#954;-Carrageenan Sub-Microgel Medium

Embedded Bioprinting of Tissue-like Structures Using &#954;-Carrageenan Sub-Microgel Medium

Embedded three-dimensional (3D) bioprinting utilizing a granular hydrogel supporting bath has emerged as a critical technique for creating biomimetic scaffolds. However, engineering a suitable gel suspension medium that balances precise bioink deposition with cell viability and function presents multiple challenges, particularly in achieving the desired viscoelastic properties. Here, a novel &#954;-carrageenan gel supporting bath is fabricated through an easy-to-operate mechanical grinding process, producing homogeneous sub-microscale particles. These sub-microgels exhibit typical Bingham flow behavior with small yield stress and rapid shear-thinning properties, which facilitate the smooth deposition of bioinks. Moreover, the reversible gel-sol transition and self-healing capabilities of the &#954;-carrageenan microgel network ensure the structural integrity of printed constructs, enabling the creation of complex, multi-layered tissue structures with defined architectural features. Post-printing, the &#954;-carrageenan sub-microgels can be easily removed by a simple phosphate-buffered saline wash. Further bioprinting with cell-laden bioinks demonstrates that cells within the biomimetic constructs have a high viability of 92% and quickly extend pseudopodia, as well as maintain robust proliferation, indicating the potential of this bioprinting strategy for tissue and organ fabrication. In summary, this novel &#954;-carrageenan sub-microgel medium emerges as a promising avenue for embedded bioprinting of exceptional quality, bearing profound implications for the in vitro development of engineered tissues and organs.

Author Spotlight: Development of Homogeneous &#954;-Carrageenan Sub-Microgel Baths for High-Resolution 3D Bioprinting

This study introduces a novel &#954;-carrageenan sub-microgel suspension bath, displaying remarkable reversible jamming-unjamming transition properties. These attributes contribute to the construction of biomimetic tissues and organs in embedded 3D bioprinting. The successful printing of heart/esophageal-like tissues with high resolution and cell growth demonstrates high-quality bioprinting and tissue engineering applications.

Embedded three-dimensional (3D) bioprinting utilizing a granular hydrogel supporting bath has emerged as a critical technique for creating biomimetic ...

Engineering

Bioprintable Alginate/Gelatin Hydrogel 3D In Vitro Model Systems Induce Cell Spheroid Formation

The cellular, biochemical, and biophysical heterogeneity of the native tumor microenvironment is not recapitulated by growing immortalized cancer cell lines using conventional two-dimensional (2D) cell culture. These challenges can be overcome by using bioprinting techniques to build heterogeneous three-dimensional (3D) tumor models whereby different types of cells are embedded. Alginate and gelatin are two of the most common biomaterials employed in bioprinting due to their biocompatibility, biomimicry, and mechanical properties. By combining the two polymers, we achieved a bioprintable composite hydrogel with similarities to the microscopic architecture of a native tumor stroma. We studied the printability of the composite hydrogel via rheology and obtained the optimal printing window. Breast cancer cells and fibroblasts were embedded in the hydrogels and printed to form a 3D model mimicking the in vivo microenvironment. The bioprinted heterogeneous model achieves a high viability for long-term cell culture (&#62; 30 days) and promotes the self-assembly of breast cancer cells into multicellular tumor spheroids (MCTS). We observed the migration and interaction of the cancer-associated fibroblast cells (CAFs) with the MCTS in this model. By using bioprinted cell culture platforms as co-culture systems, it offers a unique tool to study the dependence of tumorigenesis on the stroma composition. This technique features a high-throughput, low cost, and high reproducibility, and it can also provide an alternative model to conventional cell monolayer cultures and animal tumor models to study cancer biology.

Bioprintable Alginate/Gelatin Hydrogel 3D In Vitro Model Systems Induce Cell Spheroid Formation

We developed a heterogeneous breast cancer model consisting of immortalized tumor and fibroblast cells embedded in a bioprintable alginate/gelatin bioink. The model recapitulates the in vivo tumor microenvironment and facilitates the formation of multicellular tumor spheroids, yielding insight into the mechanisms driving tumorigenesis.

The cellular, biochemical, and biophysical heterogeneity of the native tumor microenvironment is not recapitulated by growing immortalized cancer cell ...

Protocols of 3D Bioprinting of Gelatin Methacryloyl Hydrogel Based Bioinks

Gelatin methacryloyl (GelMA) has become a popular biomaterial in the field of bioprinting. The derivation of this material is gelatin, which is hydrolyzed from mammal collagen. Thus, the arginine-glycine-aspartic acid (RGD) sequences and target motifs of matrix metalloproteinase (MMP) remain on the molecular chains, which help achieve cell attachment and degradation. Furthermore, formation properties of GelMA are versatile. The methacrylamide groups allow a material to become rapidly crosslinked under light irradiation in the presence of a photoinitiator. Therefore, it makes great sense to establish suitable methods for synthesizing three-dimensional (3D) structures with this promising material. However, its low viscosity restricts GelMA's printability. Presented here are methods to carry out 3D bioprinting of GelMA hydrogels, namely the fabrication of GelMA microspheres, GelMA fibers, GelMA complex structures, and GelMA-based microfluidic chips. The resulting structures and biocompatibility of the materials as well as the printing methods are discussed. It is believed that this protocol may serve as a bridge between previously applied biomaterials and GelMA as well as contribute to the establishment of GelMA-based 3D architectures for biomedical applications.

Presented here is a method for the 3D bioprinting of gelatin methacryloyl.

Gelatin methacryloyl (GelMA) has become a popular biomaterial in the field of bioprinting. The derivation of this material is gelatin, which is hydrolyzed ...

Gelatin Methacryloyl Granular Hydrogel Scaffolds: High-throughput Microgel Fabrication, Lyophilization, Chemical Assembly, and 3D Bioprinting

The emergence of granular hydrogel scaffolds (GHS), fabricated via assembling hydrogel microparticles (HMPs), has enabled microporous scaffold formation in situ. Unlike conventional bulk hydrogels, interconnected microscale pores in GHS facilitate degradation-independent cell infiltration as well as oxygen, nutrient, and cellular byproduct transfer. Methacryloyl-modified gelatin (GelMA), a (photo)chemically crosslinkable, protein-based biopolymer containing cell adhesive and biodegradable moieties, has widely been used as a cell-responsive/instructive biomaterial. Converting bulk GelMA to GHS&#160;may open a plethora of opportunities for tissue engineering and regeneration. In this article, we demonstrate the procedures of high-throughput GelMA microgel fabrication, conversion to resuspendable dry microgels (micro-aerogels), GHS formation via the chemical assembly of microgels, and granular bioink fabrication for extrusion bioprinting. We show how a sequential physicochemical treatment via cooling and photocrosslinking enables the formation of mechanically robust GHS.&#160;When light is inaccessible (e.g., during deep tissue injection), individually crosslinked GelMA HMPs may be bioorthogonally assembled via enzymatic crosslinking using transglutaminases. Finally, three-dimensional (3D) bioprinting of microporous GHS&#160;at low HMP packing density is demonstrated via the interfacial self-assembly of heterogeneously charged nanoparticles.

This article describes protocols for high-throughput gelatin methacryloyl microgel fabrication using microfluidic devices, converting microgels to resuspendable powder (micro-aerogels), the chemical assembly of microgels to form granular hydrogel scaffolds, and developing granular hydrogel bioinks with preserved microporosity for 3D bioprinting.

The emergence of granular hydrogel scaffolds (GHS), fabricated via assembling hydrogel microparticles (HMPs), has enabled microporous scaffold formation ...

Wet-spinning-based Molding Process of Gelatin for Tissue Regeneration

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