Research reported in this publication was supported by the National Institute of Arthritis and Musculoskeletal and Skin Diseases of the National Institute of Health under award numbers R03AR068087 and R21AR071585 and by the Boettcher Foundation (#11219) to MDK. CBE was supported by NIH/NCATS Colorado CTSA Grant Number TL1 TR001081.

All animal procedures were approved by the University of Colorado Denver Institutional Animal Care and Use Committee. 6-week old male Sprague-Dawley rats were used for the present study. The rat growth plate injury model was created following a previously published report30.
1. Preparation of the chitosan polymer
<ol>
	<li>Obtain purified and lyophilized low molecular weight (LMW) chitosan from commercially available sources (see Table of Materials).</li>
	<li>Add 495 mL of double-distilled water (ddH2O) and a stir bar to a 1 L beaker. Add 5 g of chitosan (Step 1.1.) and mix well. 
	NOTE: Chitosan is only sparingly soluble in an aqueous solution at physiological pH, so the chitosan will not dissolve easily at this step.</li>
	<li>Add 5 mL of glacial acetic acid to the above-prepared chitosan solution.</li>
	<li>Stir covered at 300 rpm for 18 h with the beaker set in a water bath held at 50 &#176;C.</li>
	<li>Using a B&#252;chner flask and funnel, filter the chitosan solution through decreasing sizes of filter paper: 22 &#181;m, 8 &#181;m, and 2.7 &#181;m (see Table of Materials).</li>
	<li>Add the filtered chitosan solution to cellulose dialysis tubing (see Table of Materials) and allow to dialyze in ddH2O at room temperature for 4 days, changing the ddH2O every day. 
	NOTE: Use ultrapure ddH2O water for the last change.</li>
	<li>Transfer the dialyzed chitosan solution to a beaker and adjust the pH to 8.0 using 1 M NaOH.</li>
	<li>Aliquot the chitosan into centrifuge tubes and centrifuge at 4000 x g for 5 min at room temperature.</li>
	<li>Decant the supernatant to a waste stream and resuspend the chitosan in ddH2O, repeating 2x.</li>
	<li>Freeze and then lyophilize the chitosan pellet.
	<ol>
		<li>Each day, remove the lyophilized product and record the mass. 
		&#8203;NOTE: When the mass of the lyophilized product is no longer changing, the product is fully dried and can be stored at -20 &#176;C until ready for use.</li>
	</ol>
	</li>
</ol>
2. Fabrication of chitosan hydrogel
<ol>
	<li>Add 2 mL of 6% acetic acid and 120 mg of purified chitosan (Step 1) to a 10 mL Luer-lock syringe to form a 6% w/v chitosan solution.</li>
	<li>Connect the Luer-lock syringe to another identical syringe using a female-female Luer-lock connector and mix the solution back and forth for 30 s, or until the chitosan has fully dissolved in the acetic acid.</li>
	<li>Before crosslinking, add any therapeutic or bioactive agent to the chitosan solution (if needed). For the present study, 200 ng of SDF-1a and TGF-b3 (see Table of Materials) were added to the microgels. 
	NOTE: SDF-1a and TGF-b3 are bioactive agents relevant to growth plate tissue regeneration. SDF-1a promotes migration of mesenchymal stem cells to the defect site, and TGF-b3 serves as a chondrogenic factor to induce differentiation of these stem cells down the chondrogenic lineage31.&#160; 
	NOTE: Mix the chitosan again between the syringes to fully incorporate the therapeutic.</li>
	<li>Prepare 100 mM of stock crosslinker solution of genipin (see Table of Materials) in 100% ethanol.</li>
	<li>Add 100 &#181;L of the prepared genipin solution (Step 2.4.) to the chitosan-containing syringe, and again mix back and forth between syringes for 30 s.</li>
	<li>Extrude the mixture from the syringe onto a 35 mm Petri dish, cover it with paraffin film, and incubate it at 37 &#176;C overnight in a humidified atmosphere. 
	NOTE: The solution will turn dark blue, indicating that the crosslinking reaction between chitosan and genipin has occurred, leading to the formation of chitosan microgel.</li>
	<li>Filter the prepared chitosan microgel following the steps below.
	<ol>
		<li>Gently break the hydrogel into smaller pieces using a spatula. 
		NOTE: The pieces should be small enough to be transferred to the back of a 10 mL syringe, ~1-2 cm in diameter.</li>
		<li>Place a filter of desired mesh size into the back of a clean 10 mL syringe. 
		NOTE: The typical size range for microgels is between 50-200 &#181;m.</li>
		<li>Transfer the broken gel pieces into the syringe fitted with the filter and add 6 mL of ddH2O. 
		&#8203;NOTE: The chitosan gel will swell significantly in the aqueous medium, so a large change in the gel volume is expected.</li>
		<li>Connect the syringe via a Luer-lock connector to another clean 10 mL syringe.</li>
		<li>Force the gel + water mixture through the syringe with the filter to create microgels with a specified maximum diameter.
		<ol>
			<li>After the first filtration, open the back of the syringe containing the filter and extrude the mixture back into this syringe.</li>
			<li>Replace the back of the syringe and force the mixture through the filter again.</li>
			<li>Repeat the filtration 5-6x or until there is little resistance through the filter.</li>
		</ol>
		</li>
	</ol>
	</li>
	<li>Rinse and purify the filtered microgels.
	<ol>
		<li>Transfer the filtered gel mixture to a 50 mL conical tube, bring the total volume up to 20 mL with ddH2O, and then vortex the mixture to ensure homogeneous dispersion.</li>
		<li>Centrifuge the microgels at 100 x g for 5 min at room temperature and decant the upper aqueous phase.</li>
		<li>Resuspend the microgels in 10 mL of 70% ethanol, vortex, and place under UV light for 1 h to sterilize.</li>
		<li>Centrifuge the microgels at 1,000 x g for 5 min at room temperature, discard the ethanol, and rinse 3x with ddH2O.</li>
	</ol>
	</li>
</ol>
3. Preparation of microgels for in vitro or in vivo applications
NOTE: For the present study, cartilage regeneration in growth plate injuries was studied in a rat model. For details, see reference31.
<ol>
	<li>Resuspend the microgel pellets 1:1 in ddH2O.&#160;Microgels can be stored for up to 1 month suspended in ddH2O at 4 &#176;C. If a bioactive agent is used, the microgels must be used immediately.</li>
	<li>Create the injury site in the animal following a previously published report30.</li>
	<li>Flush the injury site with saline and either keep the animal untreated (for control study) or inject the chitosan microgels only or the microgels loaded with the bioactive agents (Step 3.2.).</li>
	<li>Close the wound in the animal and administer post-surgical analgesics30.</li>
	<li>At days 7 or 28 post-surgery, euthanize the rat by CO2 overdose, excise the limbs and perform histology to assess the tissue repair at the injury site31.</li>
</ol>

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The authors have nothing to disclose.

Microgels have been widely researched in recent years due to their high level of applicability for various purposes, such as drug delivery or cell encapsulation9. The ease of manufacturing of micro-scale biomaterial constructs is of significant relevance in tissue engineering, as it allows researchers to develop hydrogel-based strategies at a specific size and time scale. However, most methods for fabricating chitosan microgels require expensive equipment and reagents, emulsions, or cytotoxic solv...

The growth plate, also known as the physis, is the cartilage structure located at the end of long bones that mediates growth in children. If the growth plate becomes injured, repair tissue known as a &#34;bony bar&#34; can form, which interrupts normal growth and can cause growth defects or angular deformities. Epidemiological data have shown that 15%-30% of all childhood skeletal injuries are related to the growth plate. Bony bar formation occurs in up to 30% of these injuries, making growth plate injuries and their associated treatment a significant clinical manifestation issue1,2,3,4. When bony bar formation occurs, the most common treatment avenue involves resectioning the bony bar and inserting an interpositional material, such as silicon or adipose tissue5. However, patients that undergo bony bar resection surgery often have a poor prognosis for full recovery, as there is currently no treatment that can fully repair an injured growth plate6,7,8. In light of these shortcomings, there is a critical need for effective strategies for treating growth plate injuries, both in preventing the formation of a bony bar and regenerating healthy physeal cartilage tissue.
Hydrogel microparticles, or microgels, have recently gained interest as injectable scaffolds that can provide sustained release of therapeutics9. Due to their high tunability and biocompatibility, microgels are also well-suited for bioactive factor or cell encapsulation. Microgels can be made of various materials, ranging from synthetic polymers, such as polyethylene glycol (PEG), to natural polymers like alginate or chitosan10,11,12.&#160;Chitosan has been shown to have several beneficial effects for tissue engineering, such as its ability to destabilize the outer membrane of gram-negative bacteria, thereby offering inherent antimicrobial activity13,14. Additionally, chitosan is cost-effective, cell-interactive, and easily modified using its amine-containing structure. Chitosan-based microgels promise a biomaterial strategy for drug delivery and material signaling that can promote tissue regeneration while preventing bacterial infection. However, chitosan microgels are often fabricated with a wide range of techniques that require special equipment, emulsion techniques, or cytotoxic solvent rinses. For example, some studies have fabricated chitosan microgels with emulsion-based methods, but these protocols require solvent rinses and cytotoxic crosslinkers, potentially negating their translation to clinical settings15,16. Other studies have used microfluidics or electrospray approaches to fabricate chitosan microgels, which require special equipment, preparation, and training17,18. Chitosan microgels are also commonly made with a dropwise process of crosslinker into chitosan solution; however, this method is highly dependent on solution viscosity, polymer concentration, and flow rate, making it difficult to control the size and dispersity of the microgels19,20. Conversely, the method for microgel fabrication described herein requires no specialist equipment or solvent rinses, making these microgels viable for fabrication in nearly any lab or setting. Therefore, these microgels represent highly relevant biomaterials for a quick, cost-effective, and easy-to-produce drug delivery vehicle for many applications.
By modulating a microgel&#39;s composition and material characteristics, researchers can gain precise control over the cellular microenvironment, thus directing cell behavior in a material-dependent manner. Microgels can be employed on their own or combined with bulk biomaterial systems to impart specific functionalities, such as the extended release of bioactive factors or precise special signaling for native or exogenous cells. Biomaterials and microgels have emerged as attractive treatment avenues for growth plate injuries. Significant effort has been dedicated to developing alginate and chitosan-based biomaterials to treat growth plate injuries21,22,23,24,25. Due to the dynamic temporal nature of growth plate ossification and bone elongation, the mechanism of bony bar formation is not fully understood.&#160;Therefore, several animal models have been developed to better elucidate the mechanisms of endochondral ossification and bony bar formation, such as in rats, rabbits, and sheep26,27,28. One such model is a rat growth plate injury model, which uses a drill-hole defect in the rat tibia to produce a bony bar in a predictable and reproducible manner and mimics human injuries across all three zones of the growth plate29,30. Several biomaterial-based strategies for treating growth plate injuries have been tested using this model. Additionally, two different methods for fabricating chitosan microgels have been developed, which can be used as an injectable biomaterial system that releases therapeutics in a sustained manner10,31. These microgels have been employed in a rat physeal injury model, and they showed improved cartilage regeneration31 when releasing SDF-1a and TGF-b3. The techniques provided in this protocol describe methods developed to fabricate these chitosan microgels, which can then be employed in a wide variety of tissue engineering applications. For example, recent studies have used thermo- or magento-responsive chitosan microgels for controlled oncological drug delivery applications32,33.

<table border="1" fo:keep-together.within-page="1" fo:keep-with-next.within-page="always">
	<tbody>
		<tr>
			<td>Acetic acid</td>
			<td>SigmaAldrich</td>
			<td>AX0073</td>
			<td></td>
		</tr>
		<tr>
			<td>BD Luer-Lock Syringe</td>
			<td>Fisher Scientific</td>
			<td>14-823-16E</td>
			<td></td>
		</tr>
		<tr>
			<td>B&#252;chner Funnel</td>
			<td>Fisher Scientific</td>
			<td>FB966F</td>
			<td>100 mm diameter</td>
		</tr>
		<tr>
			<td>Chitosan (low molecular weight)</td>
			<td>SigmaAldrich</td>
			<td>448869</td>
			<td>75-80% deacetylation</td>
		</tr>
		<tr>
			<td>Dialysis Membrane Tubing</td>
			<td>Fisher Scientific</td>
			<td>08-670-5C</td>
			<td>3500 MWCO</td>
		</tr>
		<tr>
			<td>Ethanol</td>
			<td>SigmaAldrich</td>
			<td>493538</td>
			<td></td>
		</tr>
		<tr>
			<td>Genipin</td>
			<td>SigmaAldrich</td>
			<td>G4796</td>
			<td></td>
		</tr>
		<tr>
			<td>Heracell 150i Incubator</td>
			<td>ThermoFisher</td>
			<td>50116047</td>
			<td></td>
		</tr>
		<tr>
			<td>Parafilm</td>
			<td>Fisher Scientific</td>
			<td>13-374-12</td>
			<td></td>
		</tr>
		<tr>
			<td>Recombinant human SDF-1a</td>
			<td>Peprotech</td>
			<td>300-28A</td>
			<td></td>
		</tr>
		<tr>
			<td>Recombinant human TGF-b3</td>
			<td>Peprotech</td>
			<td>100-36E</td>
			<td></td>
		</tr>
		<tr>
			<td>Whatman Filter Paper Grade 540</td>
			<td>SigmaAldrich</td>
			<td>Z241547</td>
			<td>8 mm pore size</td>
		</tr>
		<tr>
			<td>Whatman Filter Paper Grade 541</td>
			<td>SigmaAldrich</td>
			<td>WHA1541055</td>
			<td>22 mm pore size</td>
		</tr>
		<tr>
			<td>Whatman Filter paper Grade 542</td>
			<td>SigmaAldrich</td>
			<td>WHA1542185</td>
			<td>2.7 mm pore size</td>
		</tr>
		<tr>
			<td>Wire Mesh Sieve</td>
			<td>McMaster-Carr</td>
			<td>9317T86</td>
			<td>No. 100 Mesh</td>
		</tr>
	</tbody>
</table>

fabrication of size-controlled and emulsion-free chitosan-genipin microgels for tissue engineering applications

Chitosan microgels are of significant interest in tissue engineering due to their wide range of applications, low cost, and immunogenicity. However, chitosan microgels are commonly fabricated using emulsion methods that require organic solvent rinses, which are toxic and harmful to the environment. The present protocol presents a rapid, non-cytotoxic, non-emulsion-based method for fabricating chitosan-genipin microgels without the need for organic solvent rinses. The microgels described herein can be fabricated with precise size control. They exhibit sustained release of biomolecules, making them highly relevant for tissue engineering, biomaterials, and regenerative medicine. Chitosan is crosslinked with genipin to form a hydrogel network, then passed through a syringe filter to produce the microgels. The microgels can be filtered to create a range of sizes, and they show pH-dependent swelling and degrade over time enzymatically. These microgels have been employed in a rat growth plate injury model and were demonstrated to promote increased cartilage tissue repair and to show complete degradation at 28 days in vivo. Due to their low cost, high convenience, and ease of fabrication with cytocompatible materials, these chitosan microgels present an exciting and unique technology in tissue engineering.

Successful fabrication of chitosan microgels relies on the crosslinking reaction between genipin and chitosan, specifically involving the amines on the chitosan polymer chains. In contrast to other microgel fabrication techniques, this method does not require emulsions or solvent rinses and can be quickly and easily conducted with inexpensive equipment. A hallmark indicator for successful microgel fabrication is the distinct color change from off-white to dark blue after the chitosan and genipin have been mixed. The cros...

Watch this Scientific Journal Video about Fabrication of Size-Controlled and Emulsion-Free Chitosan-Genipin Microgels for Tissue Engineering Applications at JoVE.com

Fabrication of Size-Controlled and Emulsion-Free Chitosan-Genipin Microgels for Tissue Engineering Applications

The present protocol describes a non-emulsion-based method for the fabrication of chitosan-genipin microgels. The size of these microgels can be precisely controlled, and they can display pH-dependent swelling, degrade in vivo,&#160;and be loaded with therapeutic molecules that release over time in a sustained manner, making them highly relevant for tissue engineering applications.

Chitosan microgels are of significant interest in tissue engineering due to their wide range of applications, low cost, and immunogenicity. However, ...

fabrication-size-controlled-emulsion-free-chitosan-genipin-microgels

Research

JoVE Journal

Bioengineering

3.3K Views.  Colorado School of Mines. The present protocol describes a non-emulsion-based method for the fabrication of chitosan-genipin microgels. The size of these microgels can be precisely controlled, and they can display pH-dependent swelling, degrade in vivo, and be loaded with therapeutic molecules that release over time in a sustained manner, making them highly relevant for tissue engineering applications.

Video: Fabrication of Size-Controlled and Emulsion-Free Chitosan-Genipin Microgels for Tissue Engineering Applications

Fabrication of Biologically Derived Injectable Materials for Myocardial Tissue Engineering

This protocol provides methods for the preparation of an injectable extracellular matrix (ECM) gel for myocardial tissue engineering applications. Briefly, decellularized tissue is lyophilized, milled, enzymatically digested, and then brought to physiological pH. The lyophilization removes all water content from the tissue, resulting in dry ECM that can be ground into a fine powder with a small mill. After milling, the ECM powder is digested with pepsin to form an injectable matrix. After adjustment to pH 7.4, the liquid matrix material can be injected into the myocardium. Results of previous characterization assays have shown that matrix gels produced from decellularized pericardial and myocardial tissue retain native ECM components, including diverse proteins, peptides and glycosaminoglycans. Given the use of this material for tissue engineering, in vivo characterization is especially useful; here, a method for performing an intramural injection into the left ventricular (LV) free wall is presented as a means of analyzing the host response to the matrix gel in a small animal model. Access to the chest cavity is gained through the diaphragm and the injection is made slightly above the apex in the LV free wall. The biologically derived scaffold can be visualized by biotin-labeling before injection and then staining tissue sections with a horse radish peroxidase-conjugated neutravidin and visualizing via diaminobenzidine (DAB) staining. Analysis of the injection region can also be done with histological and immunohistochemical staining. In this way, the previously examined pericardial and myocardial matrix gels were shown to form fibrous, porous networks and promote vessel formation within the injection region.

Methods for preparing an injectable matrix gel from decellularized tissue and injecting it into rat myocardium in vivo are described.

This protocol provides methods for the preparation of an injectable extracellular matrix (ECM) gel for myocardial tissue engineering applications.  ...

Postproduction Processing of Electrospun Fibres for Tissue Engineering

Electrospinning is a commonly used and versatile method to produce scaffolds (often biodegradable) for 3D tissue engineering.1, 2, 3 Many tissues in vivo undergo biaxial distension to varying extents such as skin, bladder, pelvic floor and even the hard palate as children grow. In producing scaffolds for these purposes there is a need to develop scaffolds of appropriate biomechanical properties (whether achieved without or with cells) and which are sterile for clinical use. The focus of this paper is not how to establish basic electrospinning parameters (as there is extensive literature on electrospinning) but on how to modify spun scaffolds post production to make them fit for tissue engineering purposes - here thickness, mechanical properties and sterilisation (required for clinical use) are considered and we also describe how cells can be cultured on scaffolds and subjected to biaxial strain to condition them for specific applications.
Electrospinning tends to produce thin sheets; as the electrospinning collector becomes coated with insulating fibres it becomes a poor conductor such that fibres no longer deposit on it. Hence we describe approaches to produce thicker structures by heat or vapour annealing increasing the strength of scaffolds but not necessarily the elasticity. Sequential spinning of scaffolds of different polymers to achieve complex scaffolds is also described. Sterilisation methodologies can adversely affect strength and elasticity of scaffolds. We compare three methods for their effects on the biomechanical properties on electrospun scaffolds of poly lactic-co-glycolic acid (PLGA).
Imaging of cells on scaffolds and assessment of production of extracellular matrix (ECM) proteins by cells on scaffolds is described. Culturing cells on scaffolds in vitro can improve scaffold strength and elasticity but the tissue engineering literature shows that cells often fail to produce appropriate ECM when cultured under static conditions. There are few commercial systems available that allow one to culture cells on scaffolds under dynamic conditioning regimes - one example is the Bose Electroforce 3100 which can be used to exert a conditioning programme on cells in scaffolds held using mechanical grips within a media filled chamber.4 An approach to a budget cell culture bioreactor for controlled distortion in 2 dimensions is described. We show that cells can be induced to produce elastin under these conditions. Finally assessment of the biomechanical properties of processed scaffolds cultured with or without cells is described.

Electrospun scaffolds can be processed post production for tissue engineering applications. Here we describe methods for spinning complex scaffolds (by consecutive spinning), for making thicker scaffolds (by multi-layering using heat or vapour annealing), for achieving sterility (aseptic production or sterilisation post production) and for achieving appropriate biomechanical properties.

Electrospinning is a commonly used and versatile method to produce  scaffolds (often biodegradable) for 3D tissue engineering.1, 2, 3 Many tissues in vivo ...

Capillary Force Lithography for Cardiac Tissue Engineering

Cardiovascular disease remains the leading cause of death worldwide1. Cardiac tissue engineering holds much promise to deliver groundbreaking medical discoveries with the aims of developing functional tissues for cardiac regeneration as well as in vitro screening assays. However, the ability to create high-fidelity models of heart tissue has proven difficult. The heart&#8217;s extracellular matrix (ECM) is a complex structure consisting of both biochemical and biomechanical signals ranging from the micro- to the nanometer scale2. Local mechanical loading conditions and cell-ECM interactions have recently been recognized as vital components in cardiac tissue engineering3-5.
A large portion of the cardiac ECM is composed of aligned collagen fibers with nano-scale diameters that significantly influences tissue architecture and electromechanical coupling2. Unfortunately, few methods have been able to mimic the organization of ECM fibers down to the nanometer scale. Recent advancements in nanofabrication techniques, however, have enabled the design and fabrication of scalable scaffolds that mimic the in vivo structural and substrate stiffness cues of the ECM in the heart6-9.
Here we present the development of two reproducible, cost-effective, and scalable nanopatterning processes for the functional alignment of cardiac cells using the biocompatible polymer poly(lactide-co-glycolide) (PLGA)8 and a polyurethane (PU) based polymer. These anisotropically nanofabricated substrata (ANFS) mimic the underlying ECM of well-organized, aligned tissues and can be used to investigate the role of nanotopography on cell morphology and function10-14.
Using a nanopatterned (NP) silicon master as a template, a polyurethane acrylate (PUA) mold is fabricated. This PUA mold is then used to pattern the PU or PLGA hydrogel via UV-assisted or solvent-mediated capillary force lithography (CFL), respectively15,16. Briefly, PU or PLGA pre-polymer is drop dispensed onto a glass coverslip and the PUA mold is placed on top. For UV-assisted CFL, the PU is then exposed to UV radiation (&#955; = 250-400 nm) for curing. For solvent-mediated CFL, the PLGA is embossed using heat (120 &#176;C) and pressure (100 kPa). After curing, the PUA mold is peeled off, leaving behind an ANFS for cell culture. Primary cells, such as neonatal rat ventricular myocytes, as well as human pluripotent stem cell-derived cardiomyocytes, can be maintained on the ANFS2.

In this protocol, we demonstrate the fabrication of biomimetic cardiac cell culture substrata made from two distinct polymeric materials using capillary force lithography. The described methods provide a scalable, cost-effective technique to engineer the structure and function of macroscopic cardiac tissues for in vitro and in vivo applications.

Cardiovascular disease remains the leading cause of death worldwide1. Cardiac tissue engineering holds much promise to deliver groundbreaking medical ...

Human Cartilage Tissue Fabrication Using Three-dimensional Inkjet Printing Technology

Bioprinting, which is based on thermal inkjet printing, is one of the most attractive enabling technologies in the field of tissue engineering and regenerative medicine. With digital control&#160;cells, scaffolds, and growth factors can be precisely deposited to the desired two-dimensional (2D) and three-dimensional (3D) locations rapidly. Therefore, this technology is an ideal approach to fabricate tissues mimicking their native anatomic structures. In order to engineer cartilage with native zonal organization, extracellular matrix composition (ECM), and mechanical properties, we developed a bioprinting platform using a commercial inkjet printer with simultaneous photopolymerization capable for 3D cartilage tissue engineering. Human chondrocytes suspended in poly(ethylene glycol) diacrylate (PEGDA) were printed for 3D neocartilage construction via layer-by-layer assembly. The printed cells were fixed at their original deposited positions, supported by the surrounding scaffold in simultaneous photopolymerization. The mechanical properties of the printed tissue were similar to the native cartilage. Compared to conventional tissue fabrication, which requires longer UV exposure, the viability of the printed cells with simultaneous photopolymerization was significantly higher. Printed neocartilage demonstrated excellent glycosaminoglycan (GAG) and collagen type II production, which was consistent with gene expression. Therefore, this platform is ideal for accurate cell distribution and arrangement for anatomic tissue engineering.

The methods described in this paper show how to convert a commercial inkjet printer into a bioprinter with simultaneous UV polymerization. The printer is capable of constructing 3D tissue structure with cells and biomaterials. The study demonstrated here constructed a 3D neocartilage.

Bioprinting, which is based on thermal inkjet printing, is one of the most attractive enabling technologies in the field of tissue engineering and ...

Surgical Retrieval, Isolation and In vitro Expansion of Human Anterior Cruciate Ligament-derived Cells for Tissue Engineering Applications

Injury to the ACL is a commonly encountered problem in active individuals. Even partial tears of this intra-articular knee ligament lead to biomechanical deficiencies that impair function and stability. Current options for the treatment of partial ACL tears range from nonoperative, conservative management to multiple surgical options, such as: thermal modification, single-bundle repair, complete reconstruction, and reconstruction of the damaged portion of the native ligament. Few studies, if any, have demonstrated any single method for management to be consistently superior, and in many cases patients continue to demonstrate persistent instability and other comorbidities.
The goal of this study is to identify a potential cell source for utilization in the development of a tissue engineered patch that could be implemented in the repair of a partially torn ACL. A novel protocol was developed for the expansion of cells derived from patients undergoing ACL reconstruction. To isolate the cells, minced hACL tissue obtained during ACL reconstruction was digested in a Collagenase solution. Expansion was performed using DMEM/F12 medium supplemented with 10% fetal bovine serum (FBS) and 1% penicillin/streptomycin (P/S). The cells were then stored at -80 &#186;C or in liquid nitrogen in a freezing medium consisting of DMSO, FBS and the expansion medium. After thawing, the hACL derived cells were then seeded onto a tissue engineered scaffold, PLAGA (Poly lactic-co-glycolic acid) and control Tissue culture polystyrene (TCPS). After 7 days, SEM was performed to compare cellular adhesion to the PLAGA versus the control TCPS. Cellular morphology was evaluated using immunofluorescence staining. SEM (Scanning Electron Microscope) micrographs demonstrated that cells grew and adhered on both PLAGA and TCPS surfaces and were confluent over the entire surfaces by day 7. Immunofluorescence staining showed normal, non-stressed morphological patterns on both surfaces. This technique is promising for applications in ACL regeneration and reconstruction.

Surgical Retrieval, Isolation and In vitro Expansion of Human Anterior Cruciate Ligament-derived Cells for Tissue Engineering Applications

For future applications as a patch to repair partial tears of the Anterior Cruciate Ligament (ACL), human ACL derived cells were isolated from tissue obtained during reconstructive procedures, expanded in vitro and grown on tissue engineered scaffolds. Cellular adhesion and morphology was then performed to confirm biocompatibility on scaffold surface.

Injury to the ACL is a commonly encountered problem in active individuals. Even partial tears of this intra-articular knee ligament lead to biomechanical ...

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. ...

Fabrication of Mechanically Tunable and Bioactive Metal Scaffolds for Biomedical Applications

Biometal systems have been widely used for biomedical applications, in particular, as load-bearing materials. However, major challenges are high stiffness and low bioactivity of metals. In this study, we have developed a new method towards fabricating a new type of bioactive and mechanically reliable porous metal scaffolds-densified porous Ti scaffolds. The method consists of two fabrication processes, 1) the fabrication of porous Ti scaffolds by dynamic freeze casting, and 2) coating and densification of the porous scaffolds. The dynamic freeze casting method to fabricate porous Ti scaffolds allowed the densification of porous scaffolds by minimizing the chemical contamination and structural defects. The densification process is distinctive for three reasons. First, the densification process is simple, because it requires a control of only one parameter (degree of densification). Second, it is effective, as it achieves mechanical enhancement and sustainable release of biomolecules from porous scaffolds. Third, it has broad applications, as it is also applicable to the fabrication of functionally graded porous scaffolds by spatially varied strain during densification.

Bioactive and mechanically reliable metal scaffolds have been fabricated through a method which consists of two processes, dynamic freeze casting for the fabrication of porous Ti, and coating and densification of the Ti scaffolds. The densification process is simple, effective and applicable to the fabrication of functionally graded scaffolds.

Biometal systems have been widely used for biomedical applications, in particular, as load-bearing materials. However, major challenges are high stiffness ...

Preparation of Thermoresponsive Nanostructured Surfaces for Tissue Engineering

Thermoresponsive poly(N-isopropylacrylamide) (PIPAAm)-immobilized surfaces for controlling cell adhesion and detachment were fabricated by the Langmuir-Schaefer method. Amphiphilic block copolymers composed of polystyrene and PIPAAm (St-IPAAms) were synthesized by reversible addition-fragmentation chain transfer (RAFT) radical polymerization. A chloroform solution of St-IPAAm molecules was gently dropped into a Langmuir-trough apparatus, and both barriers of the apparatus were moved horizontally to compress the film to regulate its density. Then, the St-IPAAm Langmuir film was horizontally transferred onto a hydrophobically modified glass substrate by a surface-fixed device. Atomic force microscopy images clearly revealed nanoscale sea-island structures on the surface. The strength, rate, and quality of cell adhesion and detachment on the prepared surface were modulated by changes in temperature across the lower critical solution temperature range of PIPAAm molecules. In addition, a two-dimensional cell structure (cell sheet) was successfully recovered on the optimized surfaces. These unique PIPAAm surfaces may be useful for controlling the strength of cell adhesion and detachment.

Nanoscaled sea-island surfaces composed of thermoresponsive block copolymers were fabricated by the Langmuir-Schaefer method for controlling spontaneous cell adhesion and detachment. Both the preparation of the surface and the adhesion and detachment of cells on the surface were visualized.

Thermoresponsive poly(N-isopropylacrylamide) (PIPAAm)-immobilized surfaces for controlling cell adhesion and detachment were fabricated by the ...

Tissue Engineering by Intrinsic Vascularization in an In Vivo Tissue Engineering Chamber

In reconstructive surgery, there is a clinical need for an alternative to the current methods of autologous reconstruction which are complex, costly and trade one defect for another. Tissue engineering holds the promise to address this increasing demand. However, most tissue engineering strategies fail to generate stable and functional tissue substitutes because of poor vascularization. This paper focuses on an in vivo tissue engineering chamber model of intrinsic vascularization where a perfused artery and a vein either as an arteriovenous loop or a flow-through pedicle configuration is directed inside a protected hollow chamber. In this chamber-based system angiogenic sprouting occurs from the arteriovenous vessels and this system attracts ischemic and inflammatory driven endogenous cell migration which gradually fills the chamber space with fibro-vascular tissue. Exogenous cell/matrix implantation at the time of chamber construction enhances cell survival and determines specificity of the engineered tissues which develop. Our studies have shown that this chamber model can successfully generate different tissues such as fat, cardiac muscle, liver and others. However, modifications and refinements are required to ensure target tissue formation is consistent and reproducible. This article describes a standardized protocol for the fabrication of two different vascularized tissue engineering chamber models in vivo.

Tissue Engineering by Intrinsic Vascularization in an In Vivo Tissue Engineering Chamber

This is a guideline for constructing in vivo vascularized tissue using a microsurgical arteriovenous loop or a flow-through pedicle configuration inside a tissue engineering chamber. The vascularized tissues generated can be employed for organ regeneration and replacement of tissue defects, as well as for drug testing and disease modeling.

In reconstructive surgery, there is a clinical need for an alternative to the current methods of autologous reconstruction which are complex, costly and ...

Fabrication of Inverted Colloidal Crystal Poly(ethylene glycol) Scaffold: A Three-dimensional Cell Culture Platform for Liver Tissue Engineering

The ability to maintain hepatocyte function in vitro, for the purpose of testing xenobiotics&#39; cytotoxicity, studying virus infection and developing drugs targeted at the liver, requires a platform in which cells receive proper biochemical and mechanical cues. Recent liver tissue engineering systems have employed three-dimensional (3D) scaffolds composed of synthetic or natural hydrogels, given their high water retention and their ability to provide the mechanical stimuli needed by the cells. There has been growing interest in the inverted colloidal crystal (ICC) scaffold, a recent development, which allows high spatial organization, homotypic and heterotypic cell interaction, as well as cell-extracellular matrix (ECM) interaction. Herein, we describe a protocol to fabricate the ICC scaffold using poly (ethylene glycol) diacrylate (PEGDA) and the particle leaching method. Briefly, a lattice is made from microsphere particles, after which a pre-polymer solution is added, properly polymerized, and the particles are then removed, or leached, using an organic solvent (e.g., tetrahydrofuran). The dissolution of the lattice results in a highly porous scaffold with controlled pore sizes and interconnectivities that allow media to reach cells more easily. This unique structure allows high surface area for the cells to adhere to as well as easy communication between pores, and the ability to coat the PEGDA ICC scaffold with proteins also shows a marked effect on cell performance. We analyze the morphology of the scaffold as well as the hepatocarcinoma cell (Huh-7.5) behavior in terms of viability and function to explore the effect of ICC structure and ECM coatings. Overall, this paper provides a detailed protocol of an emerging scaffold that has wide applications in tissue engineering, especially liver tissue engineering.

Fabrication of Inverted Colloidal Crystal Polyethylene glycol Scaffold: A Three-dimensional Cell Culture Platform for Liver Tissue Engineering

This manuscript presents a detailed protocol for the fabrication of an emerging three-dimensional hepatocyte culture platform, the inverted colloidal crystal scaffold, and the concomitant techniques to assess hepatocyte behavior. The size-controllable pores, interconnectivity and ability to conjugate extracellular matrix proteins to the poly(ethylene glycol) (PEG) scaffold enhance Huh-7.5 cell performance.

The ability to maintain hepatocyte function in vitro, for the purpose of testing xenobiotics&#39; cytotoxicity, studying virus infection and developing ...