The author thanks Dr. Jin Gao for PGS synthesis, Dr. Peter Crapo for insightful discussion for bioreactor setup, Drs. Mohamed Ezzelarab and Wei Wu for explanting baboon carotid arteries. This study was supported by a grant from the National Institutes of Health (R01 HL089658).

1. Tubular Scaffold Fabrication
<ol>
<li>Prepare glass tubing (length = 70 mm, outer diameter = 9.0 mm, wall thickness = 1.0 mm; Small Parts, Miramar, FL) as a mold for scaffold.</li>
<li>Prepare 1.0 wt /vol % hyaluronic acid (HA) (Sigma-Aldrich, St. Louis, MO) by dissolving 100 mg of HA into 10 ml of deionized water. Mix using a rotator overnight until there are no air bubbles in solution.</li>
<li>Pour 1.0 % HA solution into the top of glass tubing. Let it flow slowly down along the inner wall of the mold. Flip over the tubing when the solution reached the bottom of the mold using the rotator and repeat this step until the inner wall of the mold is evenly coated by the solution.</li>
<li>After coating, dry all glass tubings into the vacuum oven at 37&deg;C for 24 h.</li>
<li>Grind and sieve salt particles (size = 25-32 &mu;m) used as porogens.</li>
<li>Assemble a glass tubing (coated with 1.0 % HA solution inside), mandrel (stainless steel encased by PTFE tubing), heat-shrinkable (HS) sleeve, and PTFE ring as described previously8.</li>
<li>Add salt particles of the desired size to the glass mold using a spatula through the silicon rubber funnel. Tap the mold gently to ensure even distribution of salt particles. Scrape off the excess salt particles across the top of the mold using the side of spatula.</li>
<li>Turn on the hybridization incubator for 30 min before transferring the salt-packed mold. Turn off the hybridization incubator and Put the salt-packed mold into the hybridization incubator for salt fusion.</li> 
<li>Remove the molds from the hybridization incubator and dry them into the vacuum oven at 37&deg;C for 24 h.</li>
<li>Remove the salt-packed molds from the vacuum oven and allow them to cool at room temperature. Remove the stainless steel mandrel by pushing it down with holding the PTFE ring. If the mandrel is too tight to the mold, use the needle pliers to pull it out. Remove the PTFE ring from the bottom of the mold.</li>
<li>Put the molds into the oven at 120 &deg;C for 5 min to shrink HS sleeve. Check if the HS sleeve is shrunk. Remove the HS sleeve from the salt-packed molds and allow them to cool at room temperature. Keep the molds into the dessicator until they are used.</li>
<li>Prepare the PGS solution by dissolving 3 g of PGS into 15 ml of tetrahydrofuran (THF) (Sigma-Aldrich, St. Louis, MO) to form a 20 wt/vol % solution.</li>
<li>Add the PGS solution to the lumen of salt-packed molds by dropping it using a spoid in the hume hood. Lean the glass mold about 45&deg; and drop the PGS solution into inner lumen with rotating the mold slowly. Check if the PGS solution flows down along the wall of mold. If dry spot is observed, add more solution.</li> 
<li>After adding the PGS solution, place the molds in the hume hood for at least 30 min for evaporating THF.</li> 
<li>Put the molds into the vacuum oven for pre-determined curing time (ex. 24/ 36/ 48 h) at 100 mTorr and 150&deg;C for curing PGS.</li>
<li>After curing, take out the molds and place them at room temperature to cool down. Dip each mold into the deionized water with vertical position slowly. Don't dip them into the water fast because air bubbles might cause tear the scaffold.</li>
<li>Transfer the molds to the water bath carefully and place them at an angle for 1 h using silicon tube to dissolve HA easily. Check if HA is dissolved from the mold. If no HA is released from the mold, push the one end of the scaffold slowly using spatula to release it from the mold. Repeat this step at the other end of the scaffold and shake the mold back and forth slowly in the water bath. Check if the scaffold is moved inside the mold.</li>
<li>Transfer the scaffold to another water bath with stirrer carefully. Don&#x2019;t grab the scaffold firmly, which might crack the scaffold. Place the scaffold in the water bath for at least 3 days with changing water to leach out salt particles.</li>
<li>After leaching out salt particles, transfer each scaffold to the 15 ml centrifuge tube filled with the deionized water and place them in the dry-ice box for 1 h to freeze. Put centrifuge tubes into the lyophilizer (all tube cabs must be opened) and lyophilize them for 2 days.</li>
<li>Put all scaffolds into the dessicator before until they are used.</li>
</ol>
2. Scaffold Preparation for Cell Seeding
<ol>
<li>Take out scaffolds from the dessicator and cut them as 25-30mm of length.</li> 
<li>Prepare two silicone rubber stoppers (diameter = 20 mm, wall thickness = 6.35 mm, middle hole diameter = 3 mm) and connect two PTFE tubings (outer diameter = 3.97 mm, inner diameter = 2.38 mm, wall thickness = 0.79 mm, length = 60 mm and 40 mm) through the middle holes of each stopper.</li>
<li>Cut a HS ring as a 1.5 mm of length and put it on top of one end of the scaffold. Push one PTFE tubing (connected silicone rubber stopper) into the one end of scaffold as small overlapping portion as possible.</li>
<li>Put the scaffold and PTFE tubing into the oven at 120 &deg;C for 10 min to shrink HS ring. Check if HS ring is shrunk and scaffolds are connected to PTFE tubing firmly.</li>
<li>Take out the scaffold and PTFE tubing and place them at room temperature to cool down. Put the scaffold-PTFE tubing assembly into the polycarbonate tube (outer diameter = 15.5 mm, inner diameter = 9 mm, length = 50 mm) as a bioreactor chamber.</li> 
<li>Connect another end of PTFE tubing to the scaffold as steps 2.3) and 2.4).</li> 
<li>Adjust two silicone rubber stoppers to attach their inner surfaces to both ends of the polycarbonate tube. Attach both outer surfaces of silicone rubber stoppers to two aluminum alloy plates (top and bottom; width = 20 mm, length = 70 mm, thickness = 6.35 mm). Put two threaded rods into the side hole of plate and fix the polycarbonate tube (included the scaffold inside) to plates by tightening thumb nut screw on the plate.</li>
<li>Measure the seedable length of each scaffold (a length between two HS rings) and calculate inner surface area (&pi; x inner diameter x seedable length) of each scaffold for cell seeding.</li> 
<li>Wrap one bioreactor chamber unit with aluminum foil and sterilize it by autoclaving at 120 &deg;C for 30 min. Sterilize each part (medium reservoir, pump tubing, gas exchanger, manifolds, and needle valve) of the bioreactor by autoclaving at 120 &deg;C for 30 min and assemble them inside the cell culture hood.</li>
<li>Pre-treat and rinse scaffolds with a series of perfusions (70, 50, and 25 % ethanol for 1 h, phosphate-buffered saline (PBS; Mediatech, Herndon, VA) for 2 h, and SMC culture medium for 24 h by using a peristaltic pump at 1.0 ml/min in the flow circuit.</li>
</ol>
3. Cell Seeding and Culture 
<ol>
<li>Isolate primary arterial smooth muscle cells (SMCs) from the carotid arteries of juvenile male baboons (Papio Anubis) and characterize them in two-dimensional (2D) culture before passaging and seeding as described previously9.</li>
<li>Expand SMCs using the MCDB 131 medium (Mediatech, Herndon, VA) with 10% fetal bovine serum (Lonza, Walkersville, MD), 1 % L-glutamine (Mediatech), 50 &mu;g/ml ascorbic acid (Fisher Scientific, Fair Lawn, NJ), and an antibiotic-antimycotic solution (10,000 IU/ml penicillin, 10,000 &mu;g/ml streptomycin, and 25 &mu;g/ml amphotericin B; Mediatech).</li>
<li>Detach each bioreactor chamber from the flow circuit in the bioreactor. Seed SMCs (passage 4-6) to the scaffold with a density of 2x106 cells/cm2 by injecting suspension with 5 ml syringe into the lumen of scaffolds after cutting off both abluminal (inlet) and luminal (outlet) flows.</li> 
<li>Put all bioreactor chambers into the hybridization incubator at 37 &deg;C and rotate them at 2 rpm for 4 h to allow the cells to uniformly spread into the scaffold.</li>
<li>Take out bioreactor chambers from the hybridization incubator reattach them to flow circuit of the bioreactor in the hood. Transfer all bioreactor system into cell culture incubator.</li>
<li>Connect the pump tubing (inner diameter = 1.6 mm) to the cartilage of peristaltic pump and load fresh culture medium in the reservoir. Start the through-wall perfusion (luminal to abluminal) at 1.0 ml/min for 15 min followed by luminal flow only.</li> 
<li>Increase the flow rate and pressure gradually from 1.0 ml/min (luminal shear stress: 1.1 dynes/cm2) and 20 mmHg on day 1 to 14.2 ml/min (15.3 dynes/cm2) and 120 mmHg on day 21, respectively. Change the pump tubing (inner diameter = 3.2 mm) at day 4 and the medium once every week.</li>
</ol>
4. Tissue Harvesting and Sample Preparation for Analysis
<ol>
<li>After 21-day culture, stop the pump and clamp tubings (inlet and outlet) of the medium reservoir. Detach the pump tubing from the cartilage of peristaltic pump and transfer bioreactor system into the hood.</li>
<li>Clamp inlet and outlet tubings of each bioreactor chamber and detach it from the flow loop. Open the inlet tubing and prepare a petri-dish at the outlet tubing to collect culture medium inside the polycarbonate tube. Open the outlet tubing slowly and remove the medium from the chamber.</li>
<li>Detach abluminal gauge needle and luminal silicone tubing from the chamber and detach aluminum plates by releasing thumb screws. Remove the polycarbonate tube and detach one end of the tissue construct from the PTFE tubing.</li>
<li>Prepare a petri-dish filled with 10 ml of PBS and detach other end of the construct from the PTFE tubing. Put construct into the PBS and rinse three times.</li>
<li>Cut the construct with a razor blade as 5-mm of length. Freeze it at -80 &deg;C freezer for biochemical assay or transfer it to 15 ml centrifuge tube filled with 5 ml of fresh medium for mechanical testing or snap freeze into the Tissue-Tek optimal cutting temperature compound (Sakura Finetek Inc., Torrance, CA) for histology/immunofluorescence staining.</li> 
</ol>
5. Representative Results:
The tubular PGS scaffolds were fabricated using biodegradable elastomer by salt fusion method (Fig. 1A). Each bioreactor chamber provided scaffolds with both luminal and abluminal flows and could be detached as a separate unit from a main flow loop (Fig. 1B). Bioreactor system was designed for culturing four scaffolds at a time by controlling and monitoring flow as well as pressure (Fig. 1C &amp; D). 
Schematic of bioreactor culture was shown in Fig. 2. After seeding SMCs, each bioreactor chamber was rotated at 37 &deg;C for 4 h to distribute cells uniformly in the lumen of scaffold. And then, pulsatile flow was applied to scaffolds until day 14 with gradually increasing flow rate (Fig. 2A) and pressure (Fig. 2B). After day 14, flow and pressure kept constant until the end of culture (day 21).
Surface morphology of the PGS scaffold was examined by scanning electron microscopy (SEM). Scanning electron micrographs showed that scaffold had consistent wall thicknesses (539 &plusmn; 18 &mu;m) (Fig. 3A) and randomly distributed macro- and micro-pores on the luminal surface (Fig. 3B). Morphometric parameters of the scaffold was measured from microcomputed tomography (micro-CT) and imaging analysis. Average pore size is 23.3 &plusmn; 3.9 &#x03BC;m and pore interconnectivity is 99.4 &plusmn; 0.62 %, which means that all pores are fully interconnected in scaffold. Porosity measured by ethanol displacement is 75.6 &plusmn; 2.7 %.
Cellular morphology of the PGS construct was examined by SEM (Fig. 4). Multilayered SMCs completely covered the luminal surface and they were perpendicularly oriented to flow direction. These results show that mechanical conditioning from pulsatile flow bioreactor supports SMC orientation in the scaffold. 
The presence of extracellular matrix (ECM) and elastic fibers were examined by H&amp;E staining and elastin autofluorescence (Fig. 5). H&amp;E staining demonstrated that cells and ECM proteins completely covered the lumen of the PGS construct. Elastin autofluorescence also showed circumferentially organized elastic fibers at the luminal surface of the construct. Production of ECM proteins in PGS constructs were measured from biochemical assays. The insoluble elastin and collagen contents were 20.2 &plusmn; 9.1 &mu;g/mg of tissue and 6.3 &plusmn; 1.9 &mu;g/mg of tissue, respectively.
<img src="/files/ftp_upload/2691/2691fig1.jpg" alt="Figure 1"/> 
Figure 1. Scaffold fabrication and bioreactor system. (A) Schematic of tubular scaffold fabrication. (B) Bioreactor chamber. Scaffolds were connected to PTFE tubing, placed into the polycarbonate tube, and fixed by silicone rubber stoppers and aluminum alloy plates. Each chamber has two flow paths: luminal (by silicone tubing) and abluminal (by gauge needle) flows. (C) Bioreactor system placed inside the incubator. It includes medium reservoir, peristaltic pump module, gas exchanger, pressure transducers, two manifolds (top and bottom), and needle valve. (D) Bioreactor system placed outside the incubator. It includes pressure monitor, flow control unit, data acquisition system, and computer.
<img src="/files/ftp_upload/2691/2691fig2.jpg" alt="Figure 2"/> 
Figure 2. Schematic of the bioreactor culture. (A) Culture protocol. (B) Applied pressure profiles at each time point (Day 1, 4, 7, and 14). 
<img src="/files/ftp_upload/2691/2691fig3.jpg" alt="Figure 3"/> 
Figure 3. Surface morphology of the PGS scaffold. (A) Cross-section. (B) Lumen. 
<img src="/files/ftp_upload/2691/2691fig4.jpg" alt="Figure 4"/> 
Figure 4. Cellular morphology of the PGS construct. (A) Lumen. (B) Cross-section of 45&deg; cut. Arrows in both figures represent the flow direction.
<img src="/files/ftp_upload/2691/2691fig5.jpg" alt="Figure 5"/> 
Figure 5. Histology and elastin autofluorescence of the PGS construct. (A) H&amp;E staining. (B) Corresponding elastin autofluorescence. L: lumen. Magnification: 40X. Scale bar: 50 &mu;m.

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No conflicts of interest declared.

The fabrication technique using a biodegradable elastomer described here has several features. (1) We used hyaluronic acid (HA) as a mold release. Since HA is water soluble, scaffold was easily released from the glass mold after soaking it into water. In this report, we used 1.0 wt/vol % of HA solution because low concentration (&lt; 0.5 wt/vol %) of solution is not viscous and flows down so fast when we pouring it on top of glass tubing. To coat HA solution uniformly, we flipped over the glass tubing when the solution flew down the bottom of the tubing and repeated this step. This HA coating is a critical to our fabrication procedure for releasing final scaffolds. (2) We used heat-shrinkable (HS) sleeve for retaining salts in the glass tubing. Since salts were densely packed in the space between the inner wall of glass tubing and the HS sleeve, HS sleeve retained salts after removing mandrel and PTFE ring in the bottom of the tubing. We could remove HS sleeve easily by putting the mold into an oven at 120 &deg;C for 5 min, and then get tubular salt templates. (3) We used the salt fusion method. It is well known that salt fusion method can improve pore interconnectivity and mechanical properties by varying fusion time10. Furthermore, since we used PGS, macro-pores were produced by the salt particles during leaching process, while the micro-pores were likely generated by glycerol vapor formed during PGS curing as we described previously11. Thus, this method has a potential to fabricate porous tubular scaffolds with different macro- and micro-structures by varying salt particles as well as PGS curing condition.
The mechanical conditioning from the bioreactor has provided pulsatile flow perfusion (maximum mean flow = 14 ml/min, maximum shear stress = 15.3 dyne/cm2, frequency = 0.5 - 1.7 Hz) and physiologically relevant pressure with the PGS scaffold, which led to SMC growth and orientation (Fig. 4). These results are consistent with previous studies reporting that cyclic stretch at this frequency and shear stress increases SMC proliferation12, and ECM protein production13,14. In addition to SMC growth and orientation, PGS construct supported ECM protein production, especially circumferentially organized elastic fibers (Fig. 5) within 3-week culture in the bioreactor. Some studies using an elastomeric scaffold as a small-diameter arterial construct have demonstrated mechanical strength and burst pressure comparable to native arteries15, and rapid SMC integration in compliant scaffolds using spinner flask16,17, while no elastic fibers were found in these constructs. Our results suggest that cyclic radial distension from the bioreactor improved mechanical signal transduction more effectively to SMCs in PGS scaffold, which likely contributed to elastin synthesis and organization.
Since vascular SMCs were the only cells that produced ECM proteins in our approach, quiescent endothelium and improve mechanical strength are necessary to develop a clinically successful small-diameter arterial constructs. We have reported that endothelial cells co-cultured with SMCs generated a confluent monolayer and supported phenotype protein expression under our culture conditions and mechanical conditioning9. Therefore, based on our approach described here, modification of co-culture experiment conditions would be a next step to improve the functions of resultant constructs and generate nonthrombogenic, robust, and compliant arterial construct similar to native arteries.

<table><tbody><tr><td>Hyaluronic acid sodium salt</td><td>Sigma-Aldrich</td><td>H7630</td><td></td></tr><tr><td>Tetrahydrofuran</td><td>Sigma-Aldrich</td><td>401757</td><td></td></tr><tr><td>MCDB 131</td><td>Mediatech, Inc.</td><td>15-100-CV</td><td></td></tr><tr><td>Fetal bovine serum</td><td>Lonza Inc.</td><td>BW14-502F</td><td></td></tr><tr><td>L-glutamine</td><td>Mediatech, Inc.</td><td>25-005-CV</td><td></td></tr><tr><td>Ascorbic acid</td><td>Fisher Scientific</td><td>A62-500</td><td></td></tr><tr><td>Antibiotic-antimycotic solution</td><td>Mediatech, Inc.</td><td>30-004-CI</td><td></td></tr><tr><td>Phosphate Buffered Saline (PBS)</td><td>Mediatech, Inc.</td><td>21-031-CV</td><td></td></tr><tr><td>Tissue-Tek optimal cutting temperature compound, 4583</td><td>Sakura Finetek</td><td>25608-930</td><td></td></tr></tbody></table>

elastomeric pgs scaffolds in arterial tissue engineering

Cardiovascular disease is one of the leading cause of mortality in the US and especially, coronary artery disease increases with an aging population and increasing obesity1. Currently, bypass surgery using autologous vessels, allografts, and synthetic grafts are known as a commonly used for arterial substitutes2. However, these grafts have limited applications when an inner diameter of arteries is less than 6 mm due to low availability, thrombotic complications, compliance mismatch, and late intimal hyperplasia3,4. To overcome these limitations, tissue engineering has been successfully applied as a promising alternative to develop small-diameter arterial constructs that are nonthrombogenic, robust, and compliant. Several previous studies have developed small-diameter arterial constructs with tri-lamellar structure, excellent mechanical properties and burst pressure comparable to native arteries5,6. While high tensile strength and burst pressure by increasing collagen production from a rigid material or cell sheet scaffold, these constructs still had low elastin production and compliance, which is a major problem to cause graft failure after implantation. Considering these issues, we hypothesized that an elastometric biomaterial combined with mechanical conditioning would provide elasticity and conduct mechanical signals more efficiently to vascular cells, which increase extracellular matrix production and support cellular orientation.
The objective of this report is to introduce a fabrication technique of porous tubular scaffolds and a dynamic mechanical conditioning for applying them to arterial tissue engineering. We used a biodegradable elastomer, poly (glycerol sebacate) (PGS)7 for fabricating porous tubular scaffolds from the salt fusion method. Adult primary baboon smooth muscle cells (SMCs) were seeded on the lumen of scaffolds, which cultured in our designed pulsatile flow bioreactor for 3 weeks. PGS scaffolds had consistent thickness and randomly distributed macro- and micro-pores. Mechanical conditioning from pulsatile flow bioreactor supported SMC orientation and enhanced ECM production in scaffolds. These results suggest that elastomeric scaffolds and mechanical conditioning of bioreactor culture may be a promising method for arterial tissue engineering.

Watch this Scientific Journal Video about Elastomeric PGS Scaffolds in Arterial Tissue Engineering at JoVE.com

Elastomeric PGS Scaffolds in Arterial Tissue Engineering

Elastomeric PGS scaffolds with vascular smooth muscle cells cultured in a pulsatile flow bioreactor may lead to promising small-diameter arterial constructs with native ECM production in a relatively short culture period.

Cardiovascular disease is one of the leading cause of mortality in the US and especially, coronary artery disease increases with an aging population and ...

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15.5K Views.  University of Pittsburgh. Elastomeric PGS scaffolds with vascular smooth muscle cells cultured in a pulsatile flow bioreactor may lead to promising small-diameter arterial constructs with native ECM production in a relatively short culture period.

Video: Elastomeric PGS Scaffolds in Arterial Tissue Engineering

Engineering Biological-Based Vascular Grafts Using a Pulsatile Bioreactor

Much effort has been devoted to develop and advance the methodology to regenerate functional small-diameter arterial bypasses. In the physiological environment, both mechanical and chemical stimulation are required to maintain the proper development and functionality of arterial vessels1,2.
Bioreactor culture systems developed by our group are designed to support vessel regeneration within a precisely controlled chemo-mechanical environment mimicking that of native vessels. Our bioreactor assembly and maintenance procedures are fairly simple and highly repeatable3,4. Smooth muscle cells (SMCs) are seeded onto a tubular polyglycolic acid (PGA) mesh that is threaded over compliant silicone tubing and cultured in the bioreactor with or without pulsatile stimulation for up to 12 weeks. There are four main attributes that distinguish our bioreactor from some predecessors. 1) Unlike other culture systems that simulate only the biochemical surrounding of native blood vessels, our bioreactor also creates a physiological pulsatile environment by applying cyclic radial strain to the vessels in culture. 2) Multiple engineered vessels can be cultured simultaneously under different mechanical conditions within a controlled chemical environment. 3) The bioreactor allows a mono layer of endothelial cells (EC) to be easily coated onto the luminal side of engineered vessels for animal implantation models. 4) Our bioreactor can also culture engineered vessels with different diameter size ranged from 1 mm to 3 mm, saving the effort to tailor each individual bioreactor to fit a specific diameter size. 
The engineered vessels cultured in our bioreactor resemble native blood vessels histologically to some degree. Cells in the vessel walls express mature SMC contractile markers such as smooth muscle myosin heavy chain (SMMHC)3. A substantial amount of collagen is deposited within the extracellular matrix, which is responsible for ultimate mechanical strength of the engineered vessels5. Biochemical analysis also indicates that collagen content of engineered vessels is comparable to that of native arteries6. Importantly, the pulsatile bioreactor has consistently regenerated vessels that exhibit mechanical properties that permit successful implantation experiments in animal models3,7. Additionally, this bioreactor can be further modified to allow real-time assessment and tracking of collagen remodeling over time, non-invasively, using a non-linear optical microscopy (NLOM)8. To conclude, this bioreactor should serve as an excellent platform to study the fundamental mechanisms that regulate the regeneration of functional small-diameter vascular grafts.

Our group has developed a bioreactor culture system that mimics the physiological pulsatile stresses of the cardiovascular system to regenerate implantable small-diameter vascular grafts.

Much effort has been devoted to develop and advance the methodology to regenerate functional small-diameter arterial bypasses. In the physiological ...

Engineering a Bilayered Hydrogel to Control ASC Differentiation

Natural polymers over the years have gained more importance because of their host biocompatibility and ability to interact with cells in vitro and in vivo. An area of research that holds promise in regenerative medicine is the combinatorial use of novel biomaterials and stem cells. A fundamental strategy in the field of tissue engineering is the use of three-dimensional scaffold (e.g., decellularized extracellular matrix, hydrogels, micro/nano particles) for directing cell function. This technology has evolved from the discovery that cells need a substrate upon which they can adhere, proliferate, and express their differentiated cellular phenotype and function 2-3. More recently, it has also been determined that cells not only use these substrates for adherence, but also interact and take cues from the matrix substrate (e.g., extracellular matrix, ECM)4. Therefore, the cells and scaffolds have a reciprocal connection that serves to control tissue development, organization, and ultimate function. Adipose-derived stem cells (ASCs) are mesenchymal, non-hematopoetic stem cells present in adipose tissue that can exhibit multi-lineage differentiation and serve as a readily available source of cells (i.e. pre-vascular endothelia and pericytes). Our hypothesis is that adipose-derived stem cells can be directed toward differing phenotypes simultaneously by simply co-culturing them in bilayered matrices1. Our laboratory is focused on dermal wound healing. To this end, we created a single composite matrix from the natural biomaterials, fibrin, collagen, and chitosan that can mimic the characteristics and functions of a dermal-specific wound healing ECM environment.

This protocol focuses on utilizing the inherent ability of stem cells to take cue from their surrounding extracellular matrix and be induced to differentiate into multiple phenotypes. This methods manuscript extends our description and characterization of a model utilizing a bilayered hydrogel, composed of PEG-fibrin and collagen, to simultaneously co-differentiate adipose-derived stem cells1.

Natural polymers over the years have gained more importance because of their host biocompatibility and ability to interact with cells in vitro and in ...

Electrospun Fibrous Scaffolds of Poly(glycerol-dodecanedioate) for Engineering Neural Tissues From Mouse Embryonic Stem Cells

For tissue engineering applications, the preparation of biodegradable and biocompatible scaffolds is the most desirable but challenging task. &#160;Among the various fabrication methods, electrospinning is the most attractive one due to its simplicity and versatility. Additionally, electrospun nanofibers mimic the size of natural extracellular matrix ensuring additional support for cell survival and growth. This study showed the viability of the fabrication of long fibers spanning a larger deposit area for a novel biodegradable and biocompatible polymer named&#160;poly(glycerol-dodecanoate) (PGD)1 by using a newly designed collector for electrospinning. PGD exhibits unique elastic properties with similar mechanical properties to nerve tissues, thus it is suitable for neural tissue engineering applications. The synthesis and fabrication set-up for making fibrous scaffolding materials was simple, highly reproducible, and inexpensive. In biocompatibility testing, cells derived from mouse embryonic stem cells could adhere to and grow on the electrospun PGD fibers. In summary, this protocol provided a versatile fabrication method for making PGD electrospun fibers to support the growth of mouse embryonic stem cell derived neural lineage cells.

Electrospun Fibrous Scaffolds of Polyglycerol-dodecanedioate for Engineering Neural Tissues From Mouse Embryonic Stem Cells

Synthesis and fabrication of electrospun long fibers spanning a larger deposit area via a newly designed collector from a novel biodegradable polymer named poly(glycerol-dodecanoate) (PGD) was reported. The fibers were able to support the growth of cells derived from mouse pluripotent stem cells.

For tissue engineering applications, the preparation of biodegradable and biocompatible scaffolds is the most desirable but challenging task. &#160;Among ...

Generation and Grafting of Tissue-engineered Vessels in a Mouse Model

The construction of vascular conduits is a fundamental strategy for surgical repair of damaged and injured vessels resulting from cardiovascular diseases. The current protocol presents an efficient and reproducible strategy in which functional tissue engineered vessel grafts can be generated using partially induced pluripotent stem cell (PiPSC) from human fibroblasts. We designed a decellularized vessel scaffold bioreactor, which closely mimics the matrix protein structure and blood flow that exists within a native vessel, for seeding of PiPSC-endothelial cells or smooth muscle cells prior to grafting into mice. This approach was demonstrated to be advantageous because immune-deficient mice engrafted with the PiPSC-derived grafts presented with markedly increased survival rate 3 weeks after surgery. This protocol represents a valuable tool for regenerative medicine, tissue engineering and potentially patient-specific cell-therapy in the near future.

Here, we present a protocol to generate tissue engineered vessel grafts that are functional for grafting into mice by double seeding partially induced pluripotent stem cell (PiPSC) - derived smooth muscle cells and PiPSC - derived endothelial cells on a decellularized vessel scaffold bioreactor.

The construction of vascular conduits is a fundamental strategy for surgical repair of damaged and injured vessels resulting from cardiovascular diseases. ...

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.

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 ...

Synthesis of Thermogelling Poly(N-isopropylacrylamide)-graft-chondroitin Sulfate Composites with Alginate Microparticles for Tissue Engineering

Injectable biomaterials are defined as implantable materials that can be introduced into the body as a liquid and solidify in situ. Such materials offer the clinical advantages of being implanted minimally invasively and easily forming space-filling solids in irregularly shaped defects. Injectable biomaterials have been widely investigated as scaffolds for tissue engineering. However, for the repair of certain load-bearing areas in the body, such as the intervertebral disc, scaffolds should possess adhesive properties. This will minimize the risk of dislocation during motion and ensure intimate contact with the surrounding tissue, providing adequate transmission of forces. Here, we describe the preparation and characterization of a scaffold composed of thermally sensitive poly(N-isopropylacrylamide)-graft-chondroitin sulfate (PNIPAAM-g-CS) and alginate microparticles. The PNIPAAm-g-CS copolymer forms a viscous solution in water at RT, into which alginate particles are suspended to enhance adhesion. Above the lower critical solution temperature (LCST), around 30 &#176;C, the copolymer forms a solid gel around the microparticles. We have adapted standard biomaterials characterization procedures to take into account the reversible phase transition of PNIPAAm-g-CS. Results indicate that the incorporation of 50 or 75 mg/ml&#160;alginate particles into 5% (w/v) PNIPAAm-g-CS solutions quadruple the adhesive tensile strength of PNIPAAm-gCS alone (p&#60;0.05). The incorporation of alginate microparticles also significantly increases swelling capacity of PNIPAAm-g-CS (p&#60;0.05), helping to maintain a space-filling gel within tissue defects. Finally, results of the in vitro toxicology assay kit, 2,3-bis-(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium-5-carboxanilide (XTT) and Live/Dead viability assay indicate that the adhesive is capable of supporting the survival and proliferation of encapsulated Human Embryonic Kidney (HEK) 293 cells over 5 days.

Synthesis of Thermogelling PolyN-isopropylacrylamide-graft-chondroitin Sulfate Composites with Alginate Microparticles for Tissue Engineering

An injectable tissue engineering scaffold composed of poly(N-isopropylacrylamide)-graft-chondroitin sulfate (PNIPAAm-g-CS)-containing alginate microparticles was prepared. The adhesive strength, swelling properties and in vitro biocompatibility are analyzed in this study. The characterization techniques developed here may be applicable to other thermogelling systems.

Injectable biomaterials are defined as implantable materials that can be introduced into the body as a liquid and solidify in situ. Such materials offer ...

A Multi-Cue Bioreactor to Evaluate the Inflammatory and Regenerative Capacity of Biomaterials under Flow and Stretch

The use of resorbable biomaterials to induce regeneration directly in the body is an attractive strategy from a translational perspective. Such materials induce an inflammatory response upon implantation, which is the driver of subsequent resorption of the material and the regeneration of new tissue. This strategy, also known as in situ tissue engineering, is pursued to obtain cardiovascular replacements such as tissue-engineered vascular grafts. Both the inflammatory and the regenerative processes are determined by the local biomechanical cues on the scaffold (i.e., stretch and shear stress). Here, we describe in detail the use of a custom-developed bioreactor that uniquely enables the decoupling of stretch and shear stress on a tubular scaffold. This allows for the systematic and standardized evaluation of the inflammatory and regenerative capacity of tubular scaffolds under the influence of well-controlled mechanical loads, which we demonstrate on the basis of a dynamic co-culture experiment using human macrophages and myofibroblasts. The key practical steps in this approach&#8212;the construction and setting up of the bioreactor, preparation of the scaffolds and cell seeding, application and maintenance of stretch and shear flow, and sample harvesting for analysis&#8212;are discussed in detail.

The goal of this protocol is to execute a dynamic co-culture of human macrophages and myofibroblasts in tubular electrospun scaffolds to investigate material-driven tissue regeneration, using a bioreactor which enables the decoupling of shear stress and cyclic stretch.

The use of resorbable biomaterials to induce regeneration directly in the body is an attractive strategy from a translational perspective. Such materials ...

Engineering

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 ...

Nondestructive Monitoring of Degradable Scaffold-Based Tissue-Engineered Blood Vessel Development Using Optical Coherence Tomography

Engineered vascular grafts with structural and mechanical properties similar to natural blood vessels are expected to meet the growing demand for arterial bypass. Characterization of the growth dynamics and remodeling process of degradable polymer scaffold-based tissue-engineered blood vessels (TEBVs) with pulsatile stimulation is crucial for vascular tissue engineering. Optical imaging techniques stand out as powerful tools for monitoring vascularization of engineered tissue enabling high-resolution imaging in real-time culture. This paper demonstrates a nondestructive and fast real-time imaging strategy to monitor the growth and remodeling of TEBVs in long-term culture by using optical coherence tomography (OCT). Geometric morphology is evaluated, including vascular remodeling process, wall thickness, and comparison of TEBV thickness in different culture time points and presence of pulsatile stimulation. Finally, OCT provides practical possibilities for real-time observation of the degradation of polymer in the reconstructing tissues under pulsatile stimulation or not and in each vessel segment, by compared with the assessment of polymer degradation using scanning electron microscopic(SEM) and polarized microscope.

A step by step protocol for nondestructive and long-period monitoring the process of vascular remodeling and scaffold degradation in real-time culture of biodegradable polymeric scaffold-based tissue-engineered blood vessels with pulsatile stimulation using optical coherence tomography is described here.

Engineered vascular grafts with structural and mechanical properties similar to natural blood vessels are expected to meet the growing demand for arterial ...

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 ...