Name: composite scaffolds of interfacial polyelectrolyte fibers for temporally controlled release of biomolecules
Uploaded: 2015-08-19T14:00:00.000+00:00
Duration: 11 min 13 s
Description: Watch this Scientific Journal Video about Composite Scaffolds of Interfacial Polyelectrolyte Fibers for Temporally Controlled Release of Biomolecules at JoVE.com

This work was supported by the Singapore National Research Foundation administered by one of its Research Centers of Excellence, the Mechanobiology Institute, Singapore. MFAC is supported by the Agency for Science, Technology and Research (Singapore) and National Agency for Research (France) joint program under project number 1122703037. BKKT is supported by the Mechanobiology Institute. We thank Mr. Daniel HC Wong for proof-reading the manuscript and Ms. Dawn JH Neo for assisting in the video production.

1. Preparation of Polyelectrolyte Solutions
<ol>
	<li>Purify chitosan, as detailed in Liao et al. Briefly, create a 1% (w/v) solution of chitosan in 2% (v/v) acetic acid and vacuum filter using grade 93 filter paper. Neutralize the filtrate using 5M NaOH until the pH stabilized to 7. Centrifuge the precipitated chitosan at 1,200 x g for 10 min. Decant the supernatant and add deionized water to wash the chitosan. Repeat the centrifugation and washing step two more times. Freeze the precipitated chitosan at -80 &#176;C and lyophilize O/N to obtain the purified form. Store purified chitosan in a dehumidified cabinet.</li>
	<li>Weigh out 1 g of purified chitosan into a sterile tissue culture dish. Place the chitosan in the tissue culture dish as close as possible to the UV lamp in the biological safety cabinet and expose to UV light for 15 min. Using sterile forceps, place the sterilized chitosan into a glass container. Dissolve chitosan using filtered 0.15M acetic acid to a final concentration between 0.5% and 1% (w/v).</li>
	<li>Weigh out 0.1 g of alginic acid sodium salt and dissolve in 10 ml distilled deionized (DDI) water to obtain a 1% (w/v) solution. Mix the alginic acid sodium salt for at least 2 hr on the vortex mixer to ensure complete dissolution. Filter the alginate solution through 0.2 &#181;m syringe filter. Store the alginate solution at 4 &#176;C.</li>
	<li>Reconstitute human recombinant growth factors such as vascular endothelial growth factor (VEGF) or beta - nerve growth factor (NGF), as recommended by manufacturer.</li>
</ol>
2. Drawing of IPC Fibers
<ol>
	<li>Mix proteins, growth factors or other biomolecules into 10-20 &#181;l aliquot of the polyelectrolyte solution that has a similar net charge. Biological molecules with net negative charge (eg bovine serum albumin [BSA]) should be mixed with alginate solution. Biological molecules with net positive charge (eg VEGF) should be mixed with chitosan solution.</li>
	<li>Place small aliquots (10-20 &#181;l) of chitosan and alginate on a stable flat surface that is covered with parafilm. The droplets of chitosan and alginate should be placed in close proximity but not in contact with each other.</li>
	<li>Lightly dip each tip on a pair of forceps into the chitosan and alginate droplets. Bring the droplets of polyelectrolytes together by pinching the forceps. When the droplets come into contact with each other, slowly pull the forceps vertically upward to draw the IPC fiber from the interface of the two droplets (Figure 1A).</li>
	<li>Carefully place the end of the drawn IPC fiber on the forceps on a collector, such as a flat polymeric scaffold affixed on a rotating mandrel (see section 3 and 4). Rotate the mandrel at a fixed speed of 10 mm/sec to allow formation of uniform and beadless IPC fibers. Increasing the speed of drawing the IPC fibers will form beads, which will cause a burst release of incorporated biochemical and premature fiber termination.10</li>
	<li>To determine incorporation efficiency, collect all the remaining liquids left on the parafilm by diluting with 500 &#181;l of 1X phosphate buffered saline (PBS). Measure the protein or growth factor content in the residue through BCA assay (for BSA), ELISA (for VEGF and NGF) or an appropriate assay to detect incorporated biomolecule.</li>
</ol>
3. Fabrication of&#160;Composite&#160;Hydrogel Scaffold of Pullulan-Dextran (PD) Polysaccharide and IPC Fibers
<ol>
	<li>Fabricating sacrificial pullulan frame for IPC fiber collection
	<ol>
		<li>Weigh out pullulan polysaccharide and mix with distilled deionized (DDI) water to create a 20% (w/v) aqueous solution. Mix the pullulan solution O/N to ensure homogeneity.</li>
		<li>Cast 15 g of pullulan solution into a 10&#160;cm diameter tissue culture polystyrene (TCPS) dish. Dry the pullulan solution O/N at 37&#160;&#176;C. Cut the pullulan films into 7 mm x 7 mm square frames.</li>
	</ol>
	</li>
	<li>Preparing pullulan-dextran polysaccharide solution
	<ol>
		<li>Create a 30% (w/v) solution of the polysaccharides pullulan and dextran (3:1 ratio) in DDI water. Mix O/N to ensure homogeneity of the polysaccharide solution. Slowly add in sodium bicarbonate to the polysaccharide solution to achieve a final concentration of 20% (w/v). Mix O/N to ensure homogeneity of the solution. Store the polysaccharide solution at 4 &#176;C.</li>
	</ol>
	</li>
	<li>Collecting IPC fibers on pullulan frame
	<ol>
		<li>Affix the sacrificial pullulan frame (section 3.1) using an alligator clip. Stick the alligator clip and pullulan frame on the end of the rotating mandrel using plastic-coated adhesive tape. Rotate the mandrel with the affixed frame at a constant speed of 10 mm/sec.&#160;The pullulan frame can be affixed onto the rotating mandrel in desired orientations.</li>
	</ol>
	</li>
	<li>Draw the IPC fibers using a pair of forceps (section 1) and attach the drawn end of the IPC fibers onto the rotating pullulan frame. Draw the IPC fibers at a constant speed. Upon reaching the terminal end of the IPC fiber, dry the fibers-on-frame construct O/N at RT.</li>
	<li>Embedding IPC fibers into PD hydrogel scaffold</li>
	<li>To crosslink every gram of pullulan-dextran solution, add 100 &#181;l of 11% (w/v) sodium trimetaphosphate aqueous solution and 100 &#181;l 10M sodium hydroxide.7 Mix the solution at 60 rpm using a stirplate for 1 to 2 min. After the addition of sodium trimetaphosphate and sodium hydroxide, the polysaccharide solution will crosslink almost immediately. Pour the viscous polysaccharide solution onto the fibers-on-frame construct to fully embed the IPC fibers. Incubate the combined pullulan-dextran-IPC fibers (PD-IPC) at 60 &#176;C for 30 min&#160;to form a chemically crosslinked composite scaffolds (Figure 1B).</li>
	<li>CAUTION: Perform step 3.3.2 in the fume hood and use proper protective equipment as acetic acid is a corrosive and flammable.</li>
	<li>To induce pore formation in the PD-IPC scaffold, submerge the whole scaffold in 20% (w/v) acetic acid for 20 min.</li>
	<li>Remove unreacted reagents by washing PD-IPC scaffolds in 1X PBS for 5 min&#160;while shaking at 100 rpm. Repeat this step 2 times.</li>
	<li>Remove the excess PBS and immediately freeze the PD-IPC scaffolds at -80 &#176;C O/N. Lyophilize the scaffolds at least 24 hr before use in any controlled release or bioactivity assays.</li>
</ol>
4. Fabrication of Composite Scaffold of PCL and IPC Fibers
CAUTION: Dichloromethane is a hazardous material. Use the fume hood and personal protective equipment when handling dichloromethane.&#160;
<ol>
	<li>Creating pristine and patterned PDMS substrates
	<ol>
		<li>Create a pristine polydimethylsiloxane (PDMS) elastomeric substrate using a piece of TCPS of desired dimension using soft lithography process. Create patterned PDMS substrates by using standard soft lithographic methods on poly(methyl methacrylate) templates with the desired topography.12</li>
	</ol>
	</li>
	<li>Fabricating sacrificial PCL frame for IPC fiber collection
	<ol>
		<li>Weigh out PCL and dissolve in dichloromethane to create a 0.9% (w/v) solution. For every 1 cm2 area of the PDMS substrate, drop 500 &#181;l of 0.9% PCL solution. Allow all of the dichloromethane solvent to fully evaporate in the fume hood. Repeat the process of casting 0.9% PCL to thicken the film to the desired thickness. Remove the dried PCL film from the PDMS substrate. Create a hole in the PCL frame using a suitably-sized puncher.2</li>
	</ol>
	</li>
	<li>Collecting IPC fibers on the PCL frame
	<ol>
		<li>Affix the sacrificial PCL frame with hole (from 4.2.1) on an alligator clip. Stick the alligator clip onto the rotating mandrel by using plastic-coated adhesive tape. Attach the drawn end of the IPC fiber onto the PCL frame before starting the rotation at a constant speed of 10mm/sec (section 2). After the end of IPC fiber drawing, dry the fiber-on-frame construct O/N at 4 &#176;C.</li>
	</ol>
	</li>
	<li>Embedding fiber-on-frame construct into patterned PCL substrate
	<ol>
		<li>Drop 500 &#181;l of 0.9% PCL solution onto the PDMS substrate to create a pristine or patterned PCL base, as required. Cast multiple layers of 0.9% PCL solution to obtain a scaffold with the desired thickness. Allow all of the dichloromethane solvent to fully evaporate in the fume hood.</li>
		<li>Place the fiber-on-frame construct (section 4.3.1) on top of the PCL base. Add 0.9% PCL solution on the fiber-on-frame construct multiple times to get the desired thickness and fully embed the IPC fibers, fabricating a PCL-IPC composite scaffold (Figure 1C).</li>
	</ol>
	</li>
</ol>
5. Measurement of Release of Biological Agents from Composite IPC Scaffolds
<ol>
	<li>Place composite PD-IPC or PCL-IPC scaffolds and stand-alone IPC fibers separately in a 24-well plate.</li>
	<li>Immerse the scaffold and stand-alone IPC fibers with 500 &#181;l of 1X PBS. Incubate the samples at 37 &#176;C. Collect PBS at various time points (release media) and replace with 500 &#181;l of 1X PBS.</li>
	<li>Measure the amount of protein or growth factor in the release media using a BCA assay (BSA), ELISA (VEGF and NGF) or other appropriate assay to calculate the cumulative release profile for the incorporated biomolecule.</li>
</ol>
6. Seeding of Cells on Composite IPC Scaffolds to Test Bioactivity of Released Biological Agents
<ol>
	<li>Sterilize the lyophilized PD-IPC or PCL-IPC composite scaffolds using UV light in the biological safety cabinet for at least 20 min.</li>
	<li>Use standard cell culture techniques to obtain a cell suspension of 2 x 105 cells in 200 &#181;l growth media. Seed the concentrated cell suspension onto the composite scaffolds. After 20 min, top-up the volume of growth media to fully submerge the scaffolds.</li>
	<li>Measure cell activity through standard techniques such as Alamar blue metabolic activity assay, PC12 neurite outgrowth assay or immunofluorescence.</li>
</ol>

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The authors have nothing to disclose.

IPC fibers are formed by the interaction of two oppositely charged polyelectrolytes. The process utilizes the extraction of the complex from the interface of the polyelectrolytes, facilitating a self-assembly process for stable fiber formation. The mechanism of IPC fiber formation ensures that any biomolecule added into a similarly charged polyelectrolyte can be incorporated during the complexation process.10,11 Conversely, addition of a biomolecule into the oppositely charged polyelectrolyte will result in instantaneous precipitation. The simple fabrication methodology for IPC fibers lends versatility in incorporating various biological materials such as cells, growth factors and small molecules. This critical feature of IPC fibers was also observed in both the PD-IPC and PCL-IPC composite scaffolds, where growth factors with different physical and biomolecular properties (charge and molecular weight) were incorporated with high efficiency. In contrast, encapsulation efficiency using microsphere methodologies for VEGF and NGF can be as low as 16%6 and 8%13, respectively.
Both PD-IPC and PCL-IPC scaffolds also demonstrated temporal control of biomolecule release. VEGF from PD-IPC scaffolds showed linear release for at least 7 days. In contrast, reported VEGF release profiles from polymeric microspheres showed an initial burst release within 24 hr, releasing at least 60% of the total VEGF content.5,14,15 Similarly, NGF was shown to have sustained growth factor delivery from PCL-IPC scaffolds, while other polymeric-based growth factor delivery systems show a plateau of release from 20 days of release.13,16 Presumably, the different components of the composite scaffolds both contribute to growth factor release kinetics. Polymer relaxation mechanisms, porosity and tortuosity can greatly affect the release of hydrophilic biomolecules.17 In addition, the chemical characteristics of the polymeric scaffold may have electrostatic attraction and repulsion towards the growth factors that affects biomolecule release. Thus, the characteristics of the polymeric scaffold are critical in determining release profiles and biomolecules with temporally-controlled release. For instance, while the PD scaffold showed near-linear BSA release, a similar composite scaffold using poly(vinyl alcohol) hydrogel lacked permeability to BSA.18 Polymeric scaffolds that can be used in combination with IPC fibers may require preliminary testing to determine its permeability range.
In addition to controlled biochemical release, the capacity to retain the bioactivity of growth factors is an important feature for any biomolecule delivery system. The harsh alkaline environment of the PD solution and the presence of organic DCM solvent in PCL have the potential to degrade any sort of biomolecules. For example, microsphere delivery system that used dichloromethane showed a consistent trend of diminishing bioactivity with increasing timepoint due to the degradation of NGF.16 Nonetheless, we observed that the bioactivity of VEGF and NGF is maintained, further reiterating that a key advantage of IPC fibers are assimilated within the composite scaffolds.
The use of a sacrificial, biocompatible polysaccharide or polymeric frame that can be easily incorporated into the bulk scaffold gives the versatility to create a composite scaffold with multiple IPC fibers aligned in different configurations.7 Thus, it is necessary that the polymeric scaffolds to be applied with IPC fibers have the capacity for further assimilation into a bulk scaffold. The assimilation process is important in fully embedding IPC fibers into the polymeric scaffold. We observed this phenomenon in the sacrificial pullulan and PCL frames, which were both easily dissolved in the solvent of the bulk PD or PCL scaffold. Furthermore, the single-step method for IPC fabrication and biomolecule incorporation gives flexibility to the number of biomolecules that can be delivered by a single composite scaffold. The simplicity of the methodology also allows fine-tuning of the incorporation efficiency and release kinetics by changing polyelectrolyte identity or concentration. The natural polyelectrolytes chitosan and alginate were used due to its high charge density and similarity to various ECM carbohydrates found in animal tissues. Yet methylated collagen and terpolymer8,19 or chitin and alginate20 may also be used for IPC fiber formation to varying effects. Multiple IPC fibers release different biochemical cues with different kinetics can also be achieved to create a multi-functional composite scaffold. For example, extracellular matrix proteins such as fibronectin loaded into IPC fibers may provide a platform for spatially controlled cell adhesion.7 Sustained release of antibiotics and growth factors are also desirable for tissue regeneration applications. This may be possible by using PCL-IPC, which can incorporate hydrophobic, small molecule drugs and hydrophilic, protein-based growth factors in different compartments of the composite scaffold.2
Our presented methodology also permitted the fabrication of a scaffold with both topographical and biochemical cues. We observed that the NP-PCL-IPC scaffold had the highest enhancement of hMSC differentiation into the neuronal lineage, implying that mimicking multiple aspects of the biophysical and biochemical cellular niche is beneficial in directing cell behavior. The ease of the presented methodology permits its application to other patternable polymeric systems such as polyacrylamide21 or polyethylene glycol22, provided that the crosslinking process will not significantly affect IPC fiber integrity. For example, PD scaffolds were chosen in this study for its unique crosslinking mechanism that occurs in ambient and aqueous conditions.24 This can potentially provide more physiologically relevant substrates for in vitro and in vivo studies. In addition, embedding IPC fibers in a composite scaffold can overcome the lack of mechanical strength of IPC fibers23 by providing tensile strength. Indeed, PCL25 was chosen for its high mechanical strength.
In summary, a simple method for creating composite scaffolds for biomolecule delivery was described. We demonstrated how IPC fibers can be used for sustained delivery of biomolecules without compromising its bioactivity and the incorporation efficiency. We showed this by creating composite scaffolds with two variations: PD-IPC and PCL-IPC scaffolds. Applicability of IPC fibers is not limited to incorporation in PD- and PCL-based scaffolds, but can be potentially extended to other polymeric systems and to deliver other biomolecules.

The extracellular matrix has inherent biochemical and biophysical cues that direct cell behaviors. Mimicking this physiological three-dimensional (3D) microenvironment is a widely explored strategy for regenerative medicine and tissue engineering applications. For example, both naturally-derived and synthetic substrates have been modified with topographical cues as a means to mimic the biophysical cellular environment.1 For example, polycaprolactone (PCL) scaffolds can be easily patterned by casting on patterned PDMS substrates.2 However, most synthetic scaffolds inadequately recapitulate the controlled biochemical environment in vivo. Bulk or surface modification of synthetic materials only present biochemical cues for cell attachment but still lack temporal regulation of biochemical delivery.3 Thus, there is a need for optimal scaffolds that can mimic the temporally regulated biochemical delivery system of the extracellular matrix. 
Biochemical delivery systems such as microspheres are plagued by problems of loss of bioactivity and low incorporation efficiency due to the severity and complexity of multi-step synthesis process.4-6 Alternative methods that use a one-step fabrication and incorporation method were proven to have excellent potential to create a favorable biochemical microenvironment without the accompanying inefficiency in incorporation and loss of bioactivity. One viable solution is the use of interfacial polyelectrolyte complexation (IPC) fibers to deliver and protect biological agents. When two oppositely charged polyelectrolyte aqueous solutions are brought together, IPC fibers can be drawn out from the interface. Virtually any type of hydrophlic biomolecule in aqueous solution can be added into either the negatively- or positively-charged polyelectrolyte solution, thus facilitating the incorporation of useful biomolecules into the IPC fiber during the complexation process. Furthermore, this process only requires aqueous and ambient conditions, thereby decreasing the risk of loss of bioactivity. Using this method, active growth factors2,7 even cells8,9 have been successfully delivered. In addition, the simple method of forming IPC fibers allows molding into any shape or orientation. The stability of such fibers has been advantageous in its incorporation into both hydrophobic2 and hydrophilic polymers7 to create composite scaffolds. These composite scaffolds with IPC fibers are beneficial for creating a physiologically relevant biochemical environment while providing physical anchorage for cells. 
In this study, we show a method to incorporate IPC fibers into a hydrophilic and a hydrophobic scaffold with topography for controlled release of active biomolecules. As a proof-of-concept, we incorporate IPC fibers made from chitosan and alginate into the biocompatible, non-immunogenic and non-antigenic pullulan-dextran hydrophilic hydrogel or the biocompatible polycaprolactone hydrophobic scaffold.

<table><tbody><tr><td>Pullulan&amp;nbsp;</td><td>Hayashibara Inc Okayama Japan</td><td></td><td>Molecular weight (MW) 200 kDa. This material is pharmaceutical grade pullulan used to make pullulan frames and PD-IPC scaffolds.</td></tr><tr><td>Dextran</td><td>Sigma Aldrich</td><td>D1037</td><td>MW 500 kDa. This material is no longer being produced by Sigma Aldrich. Alternative suggested is catalog number 31392 (Sigma Aldrich). This material is used to make PD-IPC scaffolds.</td></tr><tr><td>Sodium Bicarbonate&amp;nbsp;</td><td>Sigma Aldrich</td><td>S5761</td><td>Sodium bicarbonate must be slowly added to the pullulan-dextran polysaccharide solution. Rapid addition of sodium bicarbonate will result in precipitation.&amp;nbsp;</td></tr><tr><td>Sodium Trimetaphosphate</td><td>Sigma Aldrich</td><td>T5508</td><td>This chemical is hygroscopic and must be stored in the dehumidifying cabinet. Aqueous solution of sodium trimetaphosphate must always be made fresh.</td></tr><tr><td>Sodium Hydroxide</td><td>Sigma Aldrich</td><td>S5881</td><td>This material is hazardous and must be handled with proper protective equipment such as nitrile gloves.</td></tr><tr><td>Chitosan</td><td>Sigma Aldrich</td><td>448877</td><td>MW 190-310 kDa. Acetylation degree of 75% to 85%. Purification of chitosan is required to create stable IPC fibers.</td></tr><tr><td>Acetic Acid</td><td>Merck</td><td></td><td>This can be replaced by another brand type. This material is corrosive and flammable. Protective equipment such as face shield, nitrile gloves, lab coat and shoe cover must be worn when handling this chemical in the fume hood.&amp;nbsp;</td></tr><tr><td>Alginic acid sodium salt from brown algae, low viscosity</td><td>Sigma Aldrich</td><td>A2158</td><td>Dissolve in water overnight. Filter through sterile 0.2&amp;nbsp;&amp;micro;m syringe filter before use. Store at 4 &amp;deg;C.</td></tr><tr><td>Bovine Serum Albumin</td><td>Sinopharm Chemical Reagent</td><td></td><td>Dissolve in sterile PBS and filter using 0.2 &amp;micro;m syringe filter before use.&amp;nbsp;</td></tr><tr><td>BCA assay kit</td><td>Pierce</td><td>23225</td><td>This kit was used to measure BSA release from PD-IPC scaffolds.&amp;nbsp;</td></tr><tr><td>Human Recombinant Vascular Endothelial Growth Factor</td><td>R&amp;amp;D systems</td><td>293-VE</td><td>Dissolve growth factor in 0.2% heparin solution to a final concentration of 5&amp;nbsp;mg/ml.</td></tr><tr><td>Heparin Sodium Salt From Porcine</td><td>Sigma Aldrich</td><td>H3393</td><td>This can be replaced by another brand type. Dissolve heparin salt in sterile water at a final concentration of 1% and filter through 0.2 &amp;micro;m syringe filter before use.&amp;nbsp;</td></tr><tr><td>Human Umbilical Vein Endothelial Cells (HUVEC)</td><td>Lonza</td><td>C2517A</td><td>This primary cell type was used in the assay to determine VEGF bioactivity after release from PD-IPC scaffolds.&amp;nbsp;</td></tr><tr><td>Alamar blue</td><td>Life Technologies</td><td>DAL1025</td><td>This is used to measure cell metabolic activity. Incubate Alamar blue with cells and maintain in standard cell culture conditions for 2 to 4 hours. Measure absorbance at 570 nm to determine Alamar blue percent reduction, which is correlated to the cell activity.&amp;nbsp;</td></tr><tr><td>ScanVac Coolsafe Lyophilizer</td><td>Labogene</td><td>7.001.200.060</td><td>This is a non-programmable freeze dryer that operates at -105 to -110&amp;nbsp;&amp;deg;C. This can be replaced by other standard lab lyophilizers.</td></tr><tr><td>Polycaprolactone (PCL)</td><td>Sigma Aldrich</td><td>181609</td><td>MW 65 kDa. This is no longer being manufactured by Sigma Aldrich. This can be replaced by Sigma Aldrich catalog number 704105.</td></tr><tr><td>Dichloromethane</td><td>Sigma Aldrich</td><td>V800151</td><td>This can be replaced by another brand type. This material is hazardous and must be handled in the fume hood. Protective equipment must be worn at all times when handling this chemical.</td></tr><tr><td>Polydimethylsiloxane (PDMS; 184 Silicone Elastomer Kit)</td><td>Dow Corning</td><td>(240)4019862</td><td>The elastomer kit comes with polymer base and crosslinker. Mixing the polymer base and crosslinker in different ratios will result in different stiffness of the PDMS.</td></tr><tr><td>Human Recombinant Beta-Nerve Growth Factor (NGF)</td><td>R&amp;amp;D systems</td><td>256-GF</td><td>Reconstituted in sterile DI water to a final concentration of 100&amp;nbsp;&amp;micro;g&amp;#8288;/&amp;#8288;ml. Aliquot and store in -20 &amp;deg;C until use.</td></tr><tr><td>Human Mesenchymal Stem Cells (hMSC)</td><td>Cambrex</td><td></td><td>This cell type was used in the assay to determine synergistic effect of NGF and nanotopography.</td></tr><tr><td>Rat PC12 Pheochromocytoma Cells&amp;nbsp;</td><td>ATCC</td><td></td><td>This cell type was used in the neurite outgrowth assay to determine bioactivity of NGF. After exposure to release media with NGF, measure number of cells with neurite extensions and normalize to total number of cells.</td></tr><tr><td>Grade 93 filter paper</td><td>Whatman</td><td>Z699675</td><td>This is used for the purification of chitosan after its precipitation with sodium hydroxide at pH&amp;nbsp;7.</td></tr><tr><td>Swing bucket centrifuge</td><td>Eppendorf</td><td>5810R</td><td>To be used during the purification of chitosan using 1,200 x g speed.</td></tr><tr><td>Motor with mandrel rotating at constant speed</td><td>Rhymebus</td><td>RM5E</td><td>The motor is used for the fabrication of IPC fibers on pullulan or PCL frame.</td></tr><tr><td>Phosphate buffered saline</td><td>FirstBase</td><td></td><td>Sterilize through filtration (0.2 &amp;micro;m filter) and autoclave.&amp;nbsp;</td></tr><tr><td>10-mm diameter Tissue Culture Polystyrene Dish (TCPS)</td><td>Greiner</td><td></td><td>The TCPS dish is used for casting of pullulan frame.&amp;nbsp;</td></tr><tr><td>Human VEGF ELISA kit</td><td>R&amp;amp;D systems</td><td>DVE00</td><td>The ELISA kit is used for detection of VEGF in the release medium.</td></tr><tr><td>Human NGF ELISA kit</td><td>R&amp;amp;D systems</td><td>DY256</td><td>The ELISA kit is used for detection of NGF in the release medium.</td></tr><tr><td>Plastic Coated Adhesive Tape</td><td>Bel-Art</td><td>9040336</td><td>The adhesive tape is used to securely stick the alligator clip to the rotating mandrel</td></tr></tbody></table>

composite scaffolds of interfacial polyelectrolyte fibers for temporally controlled release of biomolecules

Various scaffolds used in tissue engineering require a controlled biochemical environment to mimic the physiological cell niche. Interfacial polyelectrolyte complexation (IPC) fibers can be used for controlled delivery of various biological agents such as small molecule drugs, cells, proteins and growth factors. The simplicity of the methodology in making IPC fibers gives flexibility in its application for controlled biomolecule delivery. Here, we describe a method of incorporating IPC fibers into two different polymeric scaffolds, hydrophilic polysaccharide and hydrophobic polycaprolactone, to create a multi-component composite scaffold. We showed that IPC fibers can be easily embedded into these polymeric structures, enhancing the capability for sustained release and improved preservation of biomolecules. We also created a composite polymeric scaffold with topographical cues and sustained biochemical release that can have synergistic effects on cell behavior. Composite polymeric scaffolds with IPC fibers represent a novel and simple method of recreating the cellular niche.

In this article, we sought to create composite scaffolds with IPC fibers for the sustained release of various biomolecules. Characteristics of the biomolecules used in this study are found in Table 1. IPC fibers were first embedded into a hydrophilic PD hydrogel to create a PD-IPC composite scaffold (Figure 1B). Model molecule BSA was first tested to determine the feasibility of using a composite scaffold for controlled biomolecule release. BSA was incorporated into PD-IPC scaffolds with an efficiency of 45 &#177; 0.97%. BSA released from the PD-IPD showed near-linear kinetics with an initial attenuated burst release followed by a concomitant steady state (Figure 2). After 2 months, BSA achieved a total release of 97%. In contrast, standalone IPC fibers exhibited a rapid release of 80% of BSA within 4 hr.
Next, we checked the release profile and bioactivity of VEGF using PD-IPC scaffolds. VEGF was incorporated with an efficiency of 75.5 &#177; 2.7%, and showed a sustained release for at least 1 week (Figure 3A). Human umbilical vein endothelial cells (HUVECs) were seeded on the PD-IPC scaffolds to determine the bioactivity of VEGF. HUVECs on the PD-IPC scaffolds showed a significant increase in Alamar blue reduction and metabolic activity compared with plain PD scaffolds at day 1, indicating good preservation of VEGF function after being released from PD-IPC scaffolds (Figure 3B). Alamar blue reduction at days 3, 6 and 7 decreased to achieve comparable levels with the plain PD scaffold (Figure 3B).
The PCL-IPC composite scaffolds were also tested for controlled release and compared with standalone IPC fibers. We incorporated NGF as a representative molecule into the PCL-IPC composite scaffolds with an incorporation efficiency of 66.38 &#177; 2.71%. PCL-IPC composite scaffolds showed linear sustained release and approximately 80% cumulative release after 18 days (Figure 4A). On the other hand, IPC stand-alone fiber showed a 70% burst release within 24 hr followed by a stagnant release rate. Using a PC12 neurite outgrowth assay, we examined the bioactivity of the released NGF (Figure 4B). The neurite outgrowth of PC12 cells grown on PCL-IPC composite scaffold release media showed similar levels with PC12 cells cultured in 30 ng/ml NGF supplemented media. This indicates that the released NGF remained bioactive for at least 7 days.
Combination of topography and sustained growth factor delivery may mimic the cellular microenvironment better. The versatile methodology of PCL-IPC fabrication allowed the fabrication of a biochemically- and topographically-controlled composite scaffold. We fabricated a PCL-IPC composite scaffold with nano-sized gratings structure (NP-PCL-IPC scaffold). We observed the synergistic effect of topography and sustained NGF release by assessing neuronal differentiation of human mesenchymal stem cells (hMSCs) (Figure 5). hMSCs cultured on the NP-PCL-IPC composite scaffolds showed higher expression of the Microtubule associated protein 2 (MAP2), indicative of neuronal differentiation. On the other hand, MAP2 protein expression was substantially lower in hMSCs cultured on PCL-IPC with only NGF release or patterned PCL (NP-PCL).
<img alt="Figure 1" src="/files/ftp_upload/53079/53079fig1.jpg" /> 
Figure 1. Incorporation of IPC fibers into hydrophilic and hydrophobic scaffolds. (A) Drawing of IPC fibers at the interface of positively (chitosan) and negatively (alginate) charged polyelectrolyte solutions. (B) Schematic diagram showing incorporation of IPC fibers in hydrophilic PD solution to create PD-IPC composite scaffold. (C) Schematic diagram showing incorporation of IPC fibers in hydrophobic PCL scaffold to create PCL-IPC composite scaffold. This figure is adapted from Cutiongco et al., 2014.7 <a href="https://www.jove.com/files/ftp_upload/53079/53079fig1large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 2" src="/files/ftp_upload/53079/53079fig2.jpg" /> 
Figure 2. Controlled release of BSA from PD-IPC composite scaffold. BSA was incorporated into PD-IPC composite scaffolds and its release was measured at various time points using BCA assay. Cumulative BSA released is provided as a percentage of the total amount of BSA (in &#181;g) incorporated in the IPC fibers and are presented as mean percentage &#177; standard deviation. This figure is adapted from Cutiongco et al., 2014.7 <a href="https://www.jove.com/files/ftp_upload/53079/53079fig2large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 3" src="/files/ftp_upload/53079/53079fig3.jpg" /> 
Figure 3. Controlled release and bioactivity of VEGF from PD-IPC composite scaffold. (A) Cumulative release profile of VEGF from PD-IPC composite scaffolds. VEGF release was measured at various time points using ELISA specific for VEGF. (B) Cell viability of endothelial cells grown on PD-IPC composite scaffolds, as measured by Alamar blue metabolic assay. Statistical analysis was performed using one-way ANOVA with Tukey&#8217;s post-hoc test. *denotes p &#60; 0.05. This figure is adapted from Cutiongco et al., 2014.7 <a href="https://www.jove.com/files/ftp_upload/53079/53079fig3large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 4" src="/files/ftp_upload/53079/53079fig4.jpg" /> 
Figure 4. Controlled release and bioactivity of NGF from PCL-IPC composite scaffold. (A) Cumulative release profile of NGF from PCL-IPC composite scaffolds. NGF release was measured at various time points using ELISA specific for NGF. Insert shows cumulative release profile of NGF from PCL-IPC scaffold for the first 8 hr. (B) Bioactivity of NGF as measured by outgrowth of PC12 neural cells. PC12 outgrowth was measured through image analysis. This figure is adapted from Teo et al., 2014.2 <a href="https://www.jove.com/files/ftp_upload/53079/53079fig4large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 5" src="/files/ftp_upload/53079/53079fig5.jpg" /> 
Figure 5. Differentiation of hMSC on NP-PCL-IPC scaffold. Confocal scanning microscopy image of hMSC cultured on different composite scaffolds. (A) NP-PCL-IPC, (B) NP-PCL, (C) PCL-IPC, (D) Pristine PCL scaffolds. Green stain denotes MAP2, a neuronal lineage marker. Red stain denotes F-actin, showing the cellular cytoskeleton. Blue stain denotes the nucleus. This figure is adapted from Teo et al., 2014.2 <a href="https://www.jove.com/files/ftp_upload/53079/53079fig5large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
Table 1. Characteristics of biochemicals used for controlled release from composite IPC scaffolds.

Watch this Scientific Journal Video about Composite Scaffolds of Interfacial Polyelectrolyte Fibers for Temporally Controlled Release of Biomolecules at JoVE.com

Composite Scaffolds of Interfacial Polyelectrolyte Fibers for Temporally Controlled Release of Biomolecules

Scaffolds for tissue engineering need to recapitulate the complex biochemical and biophysical microenvironment of the cellular niche. Here, we show the use of interfacial polyelectrolyte complexation fibers as a platform to create composite, multi-component polymeric scaffolds with sustained biochemical release.

Various scaffolds used in tissue engineering require a controlled biochemical environment to mimic the physiological cell niche. Interfacial ...

composite-scaffolds-interfacial-polyelectrolyte-fibers-for-temporally

Nanyang Technological University

Research

JoVE Journal

Bioengineering

8.3K Views.  National University of Singapore. Scaffolds for tissue engineering need to recapitulate the complex biochemical and biophysical microenvironment of the cellular niche. Here, we show the use of interfacial polyelectrolyte complexation fibers as a platform to create composite, multi-component polymeric scaffolds with sustained biochemical release.

Video: Composite Scaffolds of Interfacial Polyelectrolyte Fibers for Temporally Controlled Release of Biomolecules

Density Gradient Multilayered Polymerization (DGMP): A Novel Technique for Creating Multi-compartment, Customizable Scaffolds for Tissue Engineering

Complex tissue culture matrices, in which types and concentrations of biological stimuli (e.g. growth factors, inhibitors, or small molecules) or matrix structure (e.g. composition, concentration, or stiffness of the matrix) vary over space, would enable a wide range of investigations concerning how these variables affect cell differentiation, migration, and other phenomena. The major challenge in creating layered matrices is maintaining the structural integrity of layer interfaces without diffusion of individual components from each layer1. Current methodologies to achieve this include photopatterning2-3, lithography4, sequential functionalization5, freeze drying6, microfluidics7, or centrifugation8, many of which require sophisticated instrumentation and technical skills. Others rely on sequential attachment of individual layers, which may lead to delamination of layers9. 
DGMP overcomes these issues by using an inert density modifier such as iodixanol to create layers of varying densities10. Since the density modifier can be mixed with any prepolymer or bioactive molecule, DGMP allows each scaffold layer to be customized. Simply varying the concentration of the density modifier prevents mixing of adjacent layers while they remain aqueous. Subsequent single step polymerization gives rise to a structurally continuous multilayered scaffold, in which each layer has distinct chemical and mechanical properties. The density modifier can be easily removed with sufficient rinsing without perturbation of the individual layers or their components. This technique is therefore well suited for creating hydrogels of various sizes, shapes, and materials.
A protocol for fabricating a 2D-polyethylene glycol (PEG) gel, in which alternating layers incorporate RGDS-350, is outlined below. We use PEG because it is biocompatible and inert. RGDS, a cell adhesion peptide11, is used to demonstrate spatial restriction of a biological cue, and the conjugation of a fluorophore (Alexa Fluor 350) enables us to visually distinguish various layers. This procedure can be adapted for other materials (e.g. collagen, hyaluronan, etc.) and can be extended to fabricate 3D gels with some modifications10.

Density Gradient Multilayered Polymerization DGMP: A Novel Technique for Creating Multi-compartment, Customizable Scaffolds for Tissue Engineering

Here we describe a unique strategy for creating biocompatible, layered matrices with continuous interfaces between distinct layers for tissue engineering. Such a scaffold could provide an ideal customizable environment to modulate cell behavior by various biological, chemical or mechanical cues

Complex tissue culture matrices, in which types and concentrations of biological stimuli (e.g. growth factors, inhibitors, or small molecules) or matrix ...

Expansion of Two-dimension Electrospun Nanofiber Mats into Three-dimension Scaffolds

Electrospinning has been the preferred technology in producing a synthetic, functional scaffold due to the biomimicry to extracellular matrix and the ease control of composition, structure, and diameter of fibers. However, despite these advantages, traditional electrospun nanofiber scaffolds come with limitations including disorganized nanofiber orientation, low porosity, small pore size, and mainly two-dimensional mats. As such, there is a great need for developing a new process for fabricating electrospun nanofiber scaffolds that can overcome the above limitations. Herein, a novel and simple method is outlined. A traditional 2D nanofiber mat is transformed into a 3D scaffold with desired thickness, gap distance, porosity, and nanotopographic cues to allow for cell seeding and proliferation through the depressurization of subcritical CO2 fluid. In addition to providing a scaffold for tissue regeneration to occur, this method also provides the opportunity to encapsulate bioactive molecules such as antimicrobial peptides for local drug delivery. The CO2 expanded nanofiber scaffolds hold great potential in tissue regeneration, wound healing, 3D tissue modeling, and topical drug delivery.

This article demonstrates the technique of expanding a traditional, two-dimension (2D) electrospun nanofiber mat into a three-dimension (3D) scaffold through the depressurization of subcritical CO2 fluid. These augmented scaffolds are 3D, closely mimic cellular nanotopographic cues, and preserve the functions of biologic molecules encapsulated within the nanofibers.

Electrospinning has been the preferred technology in producing a synthetic, functional scaffold due to the biomimicry to extracellular matrix and the ease ...

Genetics

Interview: Bioreactors and Surfaced-Modified 3D-Scaffolds for Stem Cell Research

A Nature Editorial in 2003 asked the question "Good-bye, flat biology?" What does this question imply? In the past, many in vitro culture systems, mainly monolayer cultures, often suffered from the disadvantage that differentiated primary cells had a relatively short life-span and de-differentiated during culture. As a consequence, most of their organ-specific functions were lost rapidly. Thus, in order to reproduce better conditions for these cells in vitro, modifications and adaptations have been made to conventional monolayer cultures.
The last generation of CellChips -- micro-thermoformed containers -- a specific technology was developed, which offers the additional possibility to modify the whole surface of the 3D formed containers. This allows a surface-patterning on a submicron scale with distinct signalling molecules. Sensors and signal electrodes may be incorporated. Applications range from basic research in cell biology to toxicology and pharmacology. Using biodegradable polymers, clinical applications become a possibility. Furthermore, the last generation of micro-thermoformed chips has been optimized to allow for cheap mass production.

In the past many in vitro culture systems -- mainly monolayer cultures -- often suffered from the disadvantage that differentiated primary cells had a relatively short life-span and de-differentiated during culture. As a consequence, most of their organ-specific functions were lost rapidly. Thus, in order to reproduce better conditions for these cells in vitro, modifications and adaptations have been made to conventional monolayer cultures.

A Nature Editorial in 2003 asked the question  "Good-bye, flat biology?"  What does this question imply? In the past, many in vitro culture systems,  ...

Biology

Self-reporting Scaffolds for 3-Dimensional Cell Culture

Culturing cells in 3D on appropriate scaffolds is thought to better mimic the in vivo microenvironment and increase cell-cell interactions. The resulting 3D cellular construct can often be more relevant to studying the molecular events and cell-cell interactions than similar experiments studied in 2D. To create effective 3D cultures with high cell viability throughout the scaffold the culture conditions such as oxygen and pH need to be carefully controlled as gradients in analyte concentration can exist throughout the 3D construct. Here we describe the methods of preparing biocompatible pH responsive sol-gel nanosensors and their incorporation into poly(lactic-co-glycolic acid) (PLGA) electrospun scaffolds along with their subsequent preparation for the culture of mammalian cells. The pH responsive scaffolds can be used as tools to determine microenvironmental pH within a 3D cellular construct. Furthermore, we detail the delivery of pH responsive nanosensors to the intracellular environment of mammalian cells whose growth was supported by electrospun PLGA scaffolds. The cytoplasmic location of the pH responsive nanosensors can be utilized to monitor intracellular pH (pHi) during ongoing experimentation.

Biocompatible pH responsive sol-gel nanosensors can be incorporated into poly(lactic-co-glycolic acid) (PLGA) electrospun scaffolds. The produced self-reporting scaffolds can be used for in&#160;situ monitoring of microenvironmental conditions whilst culturing cells upon the scaffold. This is beneficial as the 3D cellular construct can be monitored in real-time without disturbing the experiment.

Culturing cells in 3D on appropriate scaffolds is thought to better mimic the in vivo microenvironment and increase cell-cell interactions. The resulting ...

Fabrication of a Bioactive, PCL-based "Self-fitting" Shape Memory Polymer Scaffold

Tissue engineering has been explored as an alternative strategy for the treatment of critical-sized cranio-maxillofacial (CMF) bone defects. Essential to the success of this approach is a scaffold that is able to conformally fit within an irregular defect while also having the requisite biodegradability, pore interconnectivity and bioactivity. By nature of their shape recovery and fixity properties, shape memory polymer (SMP) scaffolds could achieve defect &#8220;self-fitting.&#8221; In this way, following exposure to warm saline (~60 &#186;C), the SMP scaffold would become malleable, permitting it to be hand-pressed into an irregular defect. Subsequent cooling (~37 &#186;C) would return the scaffold to its relatively rigid state within the defect. To meet these requirements, this protocol describes the preparation of SMP scaffolds prepared via the photochemical cure of biodegradable polycaprolactone diacrylate (PCL-DA) using a solvent-casting particulate-leaching (SCPL) method. A fused salt template is utilized to achieve pore interconnectivity. To realize bioactivity, a polydopamine coating is applied to the surface of the scaffold pore walls. Characterization of self-fitting and shape memory behaviors, pore interconnectivity and in vitro bioactivity are also described.

Scaffolds capable of fitting within cranio-maxillofacial (CMF) bone defects while exhibiting osteoconductivity and bioactivity are of interest. This protocol describes the preparation of a shape memory scaffold based on polycaprolactone diacrylate (PCL-DA) using a solvent-casting particulate-leaching (SCPL) method employing a fused salt template and application of a bioactive polydopamine coating.

Tissue engineering has been explored as an alternative strategy for the treatment of critical-sized cranio-maxillofacial (CMF) bone defects. Essential to ...

Three-dimensional Biomimetic Technology: Novel Biorubber Creates Defined Micro- and Macro-scale Architectures in Collagen Hydrogels

Tissue scaffolds play a crucial role in the tissue regeneration process. The ideal scaffold must fulfill several requirements such as having proper composition, targeted modulus, and well-defined architectural features. Biomaterials that recapitulate the intrinsic architecture of in vivo tissue are vital for studying diseases as well as to facilitate the regeneration of lost and malformed soft tissue. A novel biofabrication technique was developed which combines state of the art imaging, three-dimensional (3D) printing, and selective enzymatic activity to create a new generation of biomaterials for research and clinical application. The developed material, Bovine Serum Albumin rubber, is reaction injected into a mold that upholds specific geometrical features. This sacrificial material allows the adequate transfer of architectural features to a natural scaffold material. The prototype consists of a 3D collagen scaffold with 4 and 3 mm channels that represent a branched architecture. This paper emphasizes the use of this biofabrication technique for the generation of natural constructs. This protocol utilizes a computer-aided software (CAD) to manufacture a solid mold which will be reaction injected with BSA rubber followed by the enzymatic digestion of the rubber, leaving its architectural features within the scaffold material.

An innovative biofabrication technique was developed to engineer three-dimensional constructs that resemble the architectural features, components, and mechanical properties of in vivo tissue. This technique features a newly developed sacrificial material, BSA rubber, which transfers detailed spatial features, reproducing the in vivo architectures of a wide variety of tissues.

Tissue scaffolds play a crucial role in the tissue regeneration process. The ideal scaffold must fulfill several requirements such as having proper ...

A Facile and Eco-friendly Route to Fabricate Poly(Lactic Acid) Scaffolds with Graded Pore Size

Over the recent years, functionally graded scaffolds (FGS) gaineda crucial role for manufacturing of devices for tissue engineering. The importance of this new field of biomaterials research is due to the necessity to develop implants capable of mimicking the complex functionality of the various tissues, including a continuous change from one structure or composition to another. In this latter context, one topic of main interest concerns the design of appropriate scaffolds for bone-cartilage interface tissue. In this study, three-layered scaffolds with graded pore size were achieved by melt mixing poly(lactic acid) (PLA), sodium chloride (NaCl) and polyethylene glycol (PEG). Pore size distributions were controlled by NaCl granulometry and PEG solvation. Scaffolds were characterized from a morphological and mechanical point of view. A correlation between the preparation method, the pore architecture and compressive mechanical behavior was found. The interface adhesion strength was quantitatively evaluated by using a custom-designed interfacial strength test. Furthermore, in order to imitate the human physiology, mechanical tests were also performed in phosphate buffered saline (PBS) solution at 37 &#176;C. The method herein presented provides a high control of porosity, pore size distribution and mechanical performance, thus offering the possibility to fabricate three-layered scaffolds with tailored properties by following a simple and eco-friendly route.

A Facile and Eco-friendly Route to Fabricate PolyLactic Acid Scaffolds with Graded Pore Size

In this work, poly(lactic acid)/polyethylene glycol (PLA/PEG) scaffolds were prepared by using a combination of melt mixing and selective leaching. The method herein discussed permitted to develop three-layer scaffolds by highly controlling both porosity and pore size. The mechanical properties were also evaluated in a physiological environment.

Over the recent years, functionally graded scaffolds (FGS) gaineda crucial role for manufacturing of devices for tissue engineering. The importance of ...

Fabrication of Extracellular Matrix-derived Foams and Microcarriers as Tissue-specific Cell Culture and Delivery Platforms

Cell function is mediated by interactions with the extracellular matrix (ECM), which has complex tissue-specific composition and architecture. The focus of this article is on the methods for fabricating ECM-derived porous foams and microcarriers for use as biologically-relevant substrates in advanced 3D in vitro cell culture models or as pro-regenerative scaffolds and cell delivery systems for tissue engineering and regenerative medicine. Using decellularized tissues or purified insoluble collagen as a starting material, the techniques can be applied to synthesize a broad array of tissue-specific bioscaffolds with customizable geometries. The approach involves mechanical processing and mild enzymatic digestion to yield an ECM suspension that is used to fabricate the three-dimensional foams or microcarriers through controlled freezing and lyophilization procedures. These pure ECM-derived scaffolds are highly porous, yet stable without the need for chemical crosslinking agents or other additives that may negatively impact cell function. The scaffold properties can be tuned to some extent by varying factors such as the ECM suspension concentration, mechanical processing methods, or synthesis conditions. In general, the scaffolds are robust and easy to handle, and can be processed as tissues for most standard biological assays, providing a versatile and user-friendly 3D cell culture platform that mimics the native ECM composition. Overall, these straightforward methods for fabricating customized ECM-derived foams and microcarriers may be of interest to both biologists and biomedical engineers as tissue-specific cell-instructive platforms for in vitro and in vivo applications.

The tissue-specific extracellular matrix (ECM) is a key mediator of cell function. This article describes methods for synthesizing pure ECM-derived foams and microcarriers that are stable in culture without the need for chemical crosslinking for applications in advanced 3D in vitro cell culture models or as pro-regenerative bioscaffolds.

Cell function is mediated by interactions with the extracellular matrix (ECM), which has complex tissue-specific composition and architecture. The focus ...

Core/shell Printing Scaffolds For Tissue Engineering Of Tubular Structures

Three-dimensional (3D) printing of core/shell filaments allows direct fabrication of channel structures with a stable shell that is cross-linked at the interface with a liquid core. The latter is removed post-printing, leaving behind a hollow tube. Integrating an additive manufacturing technique (like the one described here with tailor-made [bio]inks, which structurally and biochemically mimic the native extracellular matrix [ECM]) is an important step towards advanced tissue engineering. However, precise fabrication of well-defined structures requires tailored fabrication strategies optimized for the material in use. Therefore, it is sensible to begin with a set-up that is customizable, simple-to-use, and compatible with a broad spectrum of materials and applications. This work presents an easy-to-manufacture core/shell nozzle with luer-compatibility to explore core/shell printing of woodpile structures, tested with a well-defined, alginate-based scaffold material formulation.

Presented here is a simple-to-use, core/shell, three-dimensional bioprinting set-up for one-step fabrication of hollow scaffolds, suitable for tissue engineering of vascular and other tubular structures.

Three-dimensional (3D) printing of core/shell filaments allows direct fabrication of channel structures with a stable shell that is cross-linked at the ...

Interlinked Macroporous 3D Scaffolds from Microgel Rods

A two-component system of functionalized microgels from microfluidics allows for fast interlinking into 3D macroporous constructs in aqueous solutions without further additives. Continuous photoinitiated on-chip gelation enables variation of the microgel aspect ratio, which determines the building block properties for the obtained constructs. Glycidyl methacrylate (GMA) or 2-aminoethyl methacrylate (AMA) monomers are copolymerized into the microgel network based on polyethylene glycol (PEG) star-polymers to achieve either epoxy or amine functionality. A focusing oil flow is introduced into the microfluidic outlet structure to ensure continuous collection of the functionalized microgel rods. Based on a recent publication, microgel rod-based constructs result in larger pores of several hundred micrometers and, at the same time, lead to overall higher scaffold stability in comparison to a spherical-based model. In this way, it is possible to produce higher-volume constructs with more free volume while reducing the amount of material required. The interlinked macroporous scaffolds can be picked up and transported without damage or disintegration. Amine and epoxy groups not involved in interlinking remain active and can be used independently for post-modification. This protocol describes an optimized method for the fabrication of microgel rods to form macroporous interlinked scaffolds that can be utilized for subsequent cell experiments.

Microgel rods with complementary reactive groups are produced via microfluidics with the ability to interlink in aqueous solution. The anisometric microgels jam and interlink into stable constructs with larger pores compared to spherical-based systems. Microgels modified with GRGDS-PC form macroporous 3D constructs that can be used for cell culture.

A two-component system of functionalized microgels from microfluidics allows for fast interlinking into 3D macroporous constructs in aqueous solutions ...