D. H. Kim thanks the Department of Bioengineering at the University of Washington for the new faculty startup fund. D. H. Kim is also supported by the Perkins Coie Award for Discovery, the Wallace H. Coulter Foundation Translational Research Partnership Award, the Washington State Life Science Discovery Fund, and the American Heart Association Scientist Development Grant (13SDG14560076). J. Macadangdang and A. Jiao thank the support from the NIH Bioengineering Cardiovascular Training Grant Fellowship. &#160;Additional support for this work comes from the National Institutes of Health (NIH) grant R01HL111197 to M. Regnier.

All procedures are conducted at room temperature (~23 &#176;C) unless otherwise noted.
1. Fabrication of Silicon Master
<ol>
	<li>Clean silicon wafer with 100% ethanol or xylene and dry under O2/N2 gas.</li>
	<li>Place silicon wafer in spin-coater at rotation speeds of 2,000-4,000 rpm to produce a 0.3-0.5 &#181;m thick film.</li>
	<li>Pattern the photoresist film with the correct dimensions by using a photolithography system</li>
	<li>Fully immerse the patterned photoresist-coated silicon wafers in an appropriate volume of photoresist developing solution.</li>
	<li>Rinse the developed photoresist-coated silicon wafers with deionized water.</li>
	<li>To form arrays of sub-micron scale ridges with near vertical side walls, deep reactive ion etch the exposed silicon using an etching system.</li>
	<li>Remove the remaining photoresist by placing the silicon wafer in a plasma asher system.</li>
	<li>Cut the silicon wafers with a diamond-tipped cutter into the appropriately sized silicon masters for subsequent replica molding.</li>
</ol>
2. Fabrication of PUA Mold from Silicon Master
NOTE: Volume to be added to silicon master for nanofabrication will vary depending on the area of the nanopatterned master to be replicated as well as the viscosity of the polyurethane acrylate (PUA) solution.
<ol>
	<li>Clean silicon master surface with 100% ethanol or xylene and dry under O2/N2 gas.</li>
	<li>Place silicon master pattern side up in a Petri dish.</li>
	<li>For a silicon master with a 2 cm x 2 cm surface pattern, pipette 40 &#181;l of PUA to the pattern surface.</li>
	<li>Place a sheet of 4 cm x 4 cm transparent polyester (PET) film over the dispensed PUA.</li>
	<li>Press down on the PET sheet and spread the PUA underneath the sheet across the pattern face using a roller or flat edged surface (such as a card) so that the entire pattern is covered by the PUA prepolymer.</li>
	<li>Place silicon master, prepolymer, and PET approximately 10 cm below a 20 Watt (115 V) UV light (&#955; = 365 nm) for 50 sec. To be effective, the UV light wavelength can be anywhere between 310-400 nm. The intensity of the light is 10-15 mW/cm2 at the surface of the substrate.</li>
	<li>After curing, remove PET film slowly with forceps. PUA should attach to the PET film with a negative of the silicon master nanopattern.</li>
	<li>Cure PUA/PET nanopatterns under UV for at least 12 hr prior to use. Overexposure is not an issue.</li>
	<li>To clean silicon masters, place another film of PET on top of the master without the addition of PUA and expose to UV light (&#955; = 365 nm) for 50 sec and remove PET film. This will remove any unreacted monomers.</li>
	<li>Rinse silicon master with 100% ethanol or xylene and dry under O2/N2 gas.</li>
</ol>
3a. Nanopatterning Polyurethane Polymer
<ol>
	<li>Prepare 25 mm diameter circular glass slides by placing in an ozone treatment chamber for 10 min.</li>
	<li>Place ozone-treated glass slides onto small PDMS block for easy handling.</li>
	<li>Apply thin layer of surface adhesion promoter with paintbrush to glass slides. Air dry glass slides for 30 min.</li>
	<li>Place the glass slide on a piece of printer paper.</li>
	<li>Drop dispense 10 &#956;l of polyurethane (PU) pre-polymer (NOA 76) to center of glass slide. Make sure no bubbles are present after addition.</li>
	<li>Place PUA mold, pattern face down, onto the glass slide. Disperse the PU uniformly across the surface of the glass slide by rolling a rubber cylinder roller along the PUA mold. Printer paper will absorb polymer overflow.</li>
	<li>Watt UV lamp. Polymerization time of the PU is dependent on power of UV source.</li>
	<li>Remove the sample from UV light source and carefully peel the PUA mold from PU coated glass slide. Polymerization of the PU is considered complete when the PUA mold peels cleanly away from the sample and the PU glass slide has an iridescent appearance.</li>
	<li>Place finished samples in desiccator for storage for as long as a month.</li>
</ol>
3b. Nanopatterning Poly(Lactide-Co-Glycolide) Hydrogel
<ol>
	<li>Create a flat PDMS mold by vigorously mixing silicone elastomer base and silicone elastomer curing agent in a 10:1 ratio.</li>
	<li>Pour mixed PDMS precursor solution into a Petri dish so that the PDMS precursor reaches 5 mm up the edge of the dish (i.e. so that the PDMS is 5 mm thick).</li>
	<li>Place Petri dish and PDMS precursor in a desiccator for 1 hr to degas.</li>
	<li>Move Petri dish and PDMS precursor into a 65 &#176;C oven for at least 2 hr to cure.</li>
	<li>After PDMS is cured, use a razor to cut the flat PDMS into 3 cm x 3 cm square sections. These will be the flat PDMS molds used later in the patterning process.</li>
	<li>Clean 25 mm diameter circular glass slides by placing in isopropyl alcohol for 30 min in a water sonicator.</li>
	<li>Dry cleaned glass slides under O2/N2 gas.</li>
	<li>Drop dispense 100 &#956;l of PLGA solution (15% w/v PLGA in chloroform) onto glass slide.</li>
	<li>Place a flat PDMS mold on top of the dispensed PLGA to absorb the solvent and obtain a flat PLGA surface. Apply a light pressure (~10 kPa) by placing a 200 g weight on top of the PDMS for 5 min.</li>
	<li>Slowly peel away flat PDMS mold and place the cover glass on a preheated plate (120 &#176;C) for 5 min to remove residual solvent and increase adhesion between the PLGA and the cover glass.</li>
	<li>Place the NP PUA mold on top of the flat PLGA and apply constant pressure (~100 kPa) and heat (120 &#176;C) by placing a 1 kg weight on top of the PUA mold while on the heating plate for 15 min.</li>
	<li>Remove the weight from PLGA cover slide and allow the substrata to cool to room temperature. Do not remove PUA mold until the substrata have been cooled.</li>
	<li>Once the substrata have sufficiently cooled, carefully peel away the NP PUA mold, revealing the NP PLGA substratum. The NP PLGA substratum should have an iridescent appearance.</li>
	<li>Place finished samples in desiccator for storage for as long as a month.</li>
</ol>
4. Cell Seeding and Culture
NOTE: This protocol describes the culture of neonatal rat ventricular myocytes (NRVMs) and H7 human embryonic stem cell-derived cardiomyocytes (hESC-CMs) but other cell sources may be used.
<ol>
	<li>Attach the ANFS coverslips (PU or PLGA) to a 35 mm tissue culture polystyrene dish. Pipette 20 &#181;l of Norland Optical Adhesive (NOA83H) to the bottom of the dish and gently place the ANFS coverslip on top of the NOA. Allow glue to spread out and cover entire coverslip bottom. Cure NOA by exposing dish to UV for 10 min.</li>
	<li>Sterilize ANFS by rinsing with 2 ml of 70% aqueous ethanol solution for 5 min, twice. Remove ethanol by aspiration. Allow ANFS to completely air dry for ~1 hr under the UV sterilization lamp (&#955; = 200-290 nm) in the biological safety cabinet.</li>
	<li>Cellular adhesion is enhanced by coating the ANFS in fibronectin overnight. Dilute fibronectin in DI water to 5 &#181;g/ml. Pipette 2 ml of fibronectin solution into dish. Place in incubator at 37 &#176;C and 5% CO2 overnight (at least 6 hr).</li>
	<li>Obtain NRVMs, hESC-CMs, or other cardiac cells of interest according to previous protocols22.</li>
	<li>Centrifuge cell sample at 1,000 rpm for 3 min to pellet the cells.</li>
	<li>Carefully remove the supernatant by aspiration. Make sure not to disturb the pellet.</li>
	<li>Resuspend cells in appropriate culture media to a concentration of 4.6 x 106 cells/ml.</li>
	<li>Carefully pipette 200 &#181;l of cell suspension onto sterilized ANFS. Make sure cell suspension remains on the coverslip.</li>
	<li>Place cells in incubator at 37 &#176;C and 5% CO2 for 4 hr to allow cells to attach to ANFS.</li>
	<li>Add 2 ml of warm culture media to dish and replace cells in incubator under the same conditions.</li>
	<li>After 24 hr, remove media and wash with 2 ml of DPBS twice to remove excess cells.</li>
	<li>Add 2 ml of warm culture media to dish and replace cells in incubator under the same conditions. Culture the cells to confluence. Replace media every other day.</li>
</ol>

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Authors have nothing to disclose.

Functionally mature cardiac tissues are lacking for both in vivo and in vitro applications of cardiac tissue engineering. The CFL nanofabrication methods described here are robust techniques for achieving cellular alignment and influencing macroscopic tissue function due to the scalability of the system. Large areas can easily be patterned and used for cell culture. Macroscopic cellular alignment is essential in cardiac tissue engineering in order to create biomimetic, functional tissue as it influences...

Cardiovascular disease is the leading cause of morbidity and mortality in the world and present a weighty socio-economic burden on an already strained global health system1,17. Cardiac tissue engineering has two distinct goals: (1) to regenerate damaged myocardium after ischemic disease or cardiomyopathy or (2) to create a high fidelity model of the heart for in vitro drug screening or disease modeling.
The heart is a complex organ that must work constantly to supply blood to the body. Densely packed laminar structures of cardiomyocytes and supportive tissues are arranged in helical patterns throughout the heart wall18,19. The heart is also electromechanically coupled20 in a highly coordinated fashion to efficiently eject blood to the body21. Several major hurdles remain to be addressed, however, before nature&#8217;s intricate design can reliably be recapitulated in vitro. First, although robust cardiomyocyte differentiation methods continue to be developed22, hPSC-CMs still exhibit rather immature phenotypes. Their electromechanical properties and morphology most closely match fetal levels23. Second, when kept in traditional culture conditions, both stem cell-derived and primary cardiomyocytes fail to assemble into native, tissue-like structures. Rather, cells become randomly oriented and do not exhibit the banded rod-shaped appearance of adult myocardium24.
The extracellular matrix (ECM) environment with which cells interact plays a significant role in numerous cellular processes11,13,25. The ECM consists of complex, well-defined molecular and topographical cues that significantly influence the structure and function of cells6,26. Within the heart, cellular alignment closely follows the underlying nanometer scale ECM fibers2. The impact of these nanotopographical cues on cell and tissue function, however, is far from completely understood. Preliminary studies of nanometer scale cell-biomaterial interaction indicate the potential importance and impact of sub-micron topographical cues for cell signaling27, adhesion28-30, growth31, and differentiation32,33. However, due to the difficulty in developing reproducible and scalable nanofabricated substrates, such studies could not reproduce the multi-scale cellular effects of the complex in vivo ECM environment. In this protocol, a straightforward and cost-effective nanofabrication technique to produce cell culture scaffolds mimicking native cardiac ECM fiber alignment is described, allowing for a wide range of novel investigations of cardiomyocyte-biomaterial interactions. Understanding how cardiomyocytes interact with the nanoscale ECM environment could allow for the ability to control cellular behavior to more closely mimic native tissue function. Furthermore, cell monolayers are a simplified experimental system compared to 3D structures but still exhibit complex multi-cellular behavior for insightful investigations and functional screening2,34-36. Finally, such scaffolds could be used to improve cellular graft function when implanted into the heart for regenerative purposes37.

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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.

Figure 1 is a schematic overview of the production process for the two fabrication methods. Due to the diffraction of light caused by the nanoscale topography, nanopatterning should result in an iridescent surface to the ANFS. Figure 2 depicts this iridescent surface on a well-patterned 25 mm NP-PU coverslip (Figure 2A) with 800 nm ridge and groove width (Figure 2B). The iridescent appearance of the ANFS will vary slightly depending on the ridge and groo...

Watch this Scientific Journal Video about Capillary Force Lithography for Cardiac Tissue Engineering at JoVE.com

Capillary Force Lithography for Cardiac Tissue Engineering

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 ...

capillary-force-lithography-for-cardiac-tissue-engineering

Research

JoVE Journal

Bioengineering

12.3K 次观看.  University of Washington. 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.

视频: Capillary Force Lithography for Cardiac Tissue Engineering

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.  ...

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.

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 ...

Plasma Lithography Surface Patterning for Creation of Cell Networks

Systematic manipulation of a cell microenvironment with micro- and nanoscale resolution is often required for deciphering various cellular and molecular phenomena. To address this requirement, we have developed a plasma lithography technique to manipulate the cellular microenvironment by creating a patterned surface with feature sizes ranging from 100 nm to millimeters. The goal of this technique is to be able to study, in a controlled way, the behaviors of individual cells as well as groups of cells and their interactions.
This plasma lithography method is based on selective modification of the surface chemistry on a substrate by means of shielding the contact of low-temperature plasma with a physical mold. This selective shielding leaves a chemical pattern which can guide cell attachment and movement. This pattern, or surface template, can then be used to create networks of cells whose structure can mimic that found in nature and produces a controllable environment for experimental investigations. The technique is well suited to studying biological phenomenon as it produces stable surface patterns on transparent polymeric substrates in a biocompatible manner. The surface patterns last for weeks to months and can thus guide interaction with cells for long time periods which facilitates the study of long-term cellular processes, such as differentiation and adaption. The modification to the surface is primarily chemical in nature and thus does not introduce topographical or physical interference for interpretation of results. It also does not involve any harsh or toxic substances to achieve patterning and is compatible for tissue culture. Furthermore, it can be applied to modify various types of polymeric substrates, which due to the ability to tune their properties are ideal for and are widely used in biological applications. The resolution achievable is also beneficial, as isolation of specific processes such as migration, adhesion, or binding allows for discrete, clear observations at the single to multicell level.
This method has been employed to form diverse networks of different cell types for investigations involving migration, signaling, tissue formation, and the behavior and interactions of neurons arraigned in a network.

A versatile plasma lithography technique has been developed to generate stable surface patterns for guiding cellular attachment. This technique can be applied to create cell networks including those that mimic natural tissues and has been used for studying several, distinct cell types.

Systematic manipulation of a cell microenvironment with micro- and nanoscale resolution is often required for deciphering various cellular and molecular ...

Encapsulation of Cardiomyocytes in a Fibrin Hydrogel for Cardiac Tissue Engineering

Culturing cells in a three dimensional hydrogel environment is an important technique for developing constructs for tissue engineering as well as studying cellular responses under various culture conditions in vitro. The three dimensional environment more closely mimics what the cells observe in vivo due to the application of mechanical and chemical stimuli in all dimensions 1. Three-dimensional hydrogels can either be made from synthetic polymers such as PEG-DA 2 and PLGA 3 or a number of naturally occurring proteins such as collagen 4, hyaluronic acid 5 or fibrin 6,7. Hydrogels created from fibrin, a naturally occurring blood clotting protein, can polymerize to form a mesh that is part of the body's natural wound healing processes 8. Fibrin is cell-degradable and potentially autologous 9, making it an ideal temporary scaffold for tissue engineering.
Here we describe in detail the isolation of neonatal cardiomyocytes from three day old rat pups and the preparation of the cells for encapsulation in fibrin hydrogel constructs for tissue engineering. Neonatal myocytes are a common cell source used for in vitro studies in cardiac tissue formation and engineering 4. Fibrin gel is created by mixing fibrinogen with the enzyme thrombin. Thrombin cleaves fibrinopeptides FpA and FpB from fibrinogen, revealing binding sites that interact with other monomers 10. These interactions cause the monomers to self-assemble into fibers that form the hydrogel mesh. Because the timing of this enzymatic reaction can be adjusted by altering the ratio of thrombin to fibrinogen, or the ratio of calcium to thrombin, one can injection mold constructs with a number of different geometries 11,12. Further we can generate alignment of the resulting tissue by how we constrain the gel during culture 13.
After culturing the engineered cardiac tissue constructs for two weeks under static conditions, the cardiac cells have begun to remodel the construct and can generate a contraction force under electrical pacing conditions 6. As part of this protocol, we also describe methods for analyzing the tissue engineered myocardium after the culture period including functional analysis of the active force generated by the cardiac muscle construct upon electrical stimulation, as well as methods for determining final cell viability (Live-Dead assay) and immunohistological staining to examine the expression and morphology of typical proteins important for contraction (Myosin Heavy Chain or MHC) and cellular coupling (Connexin 43 or Cx43) between myocytes.

We describe the isolation of neonatal cardiomyocytes and the preparation of the cells for encapsulation in fibrin hydrogel constructs for tissue engineering. We describe methods for analyzing the tissue engineered myocardium after the culture period including active force generated upon electrical stimulation and cell viability and immunohistological staining.

Culturing cells in a three dimensional hydrogel environment is an important technique for developing constructs for tissue engineering as well as studying ...

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 ...

Tissue Engineering of a Human 3D in vitro Tumor Test System

Cancer is one of the leading causes of death worldwide. Current therapeutic strategies are predominantly developed in 2D culture systems, which inadequately reflect physiological conditions in vivo. Biological 3D matrices provide cells an environment in which cells can self-organize, allowing the study of tissue organization and cell differentiation. Such scaffolds can be seeded with a mixture of different cell types to study direct 3D cell-cell-interactions. To mimic the 3D complexity of cancer tumors, our group has developed a 3D in vitro tumor test system.	
	Our 3D tissue test system models the in vivo situation of malignant peripheral nerve sheath tumors (MPNSTs), which we established with our decellularized porcine jejunal segment derived biological vascularized scaffold (BioVaSc). In our model, we reseeded a modified BioVaSc matrix with primary fibroblasts, microvascular endothelial cells (mvECs) and the S462 tumor cell line. For static culture, the vascular structure of the BioVaSc is removed and the remaining scaffold is cut open on one side (Small Intestinal Submucosa SIS-Muc). The resulting matrix is then fixed between two metal rings (cell crowns).	
	Another option is to culture the cell-seeded SIS-Muc in a flow bioreactor system that exposes the cells to shear stress. Here, the bioreactor is connected to a peristaltic pump in a self-constructed incubator. A computer regulates the arterial oxygen and nutrient supply via parameters such as blood pressure, temperature, and flow rate. This setup allows for a dynamic culture with either pressure-regulated pulsatile or constant flow.	
	In this study, we could successfully establish both a static and dynamic 3D culture system for MPNSTs. The ability to model cancer tumors in a more natural 3D environment will enable the discovery, testing, and validation of future pharmaceuticals in a human-like model.

Tissue Engineering of a Human 3D in vitro Tumor Test System

Methods to create human 3D tumor tissues as test systems are described. These technologies are based on a decellularized Biological Vascularized Scaffold (BioVaSc), primary human cells and a tumor cell line, which can be cultured under static as well as under dynamic conditions in a flow bioreactor.

Cancer is one of the leading causes of death worldwide.  Current therapeutic strategies are predominantly developed in 2D culture  systems, which ...

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. ...

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 ...

A Protocol for Decellularizing Mouse Cochleae for Inner Ear Tissue Engineering

In mammals, mechanosensory hair cells that facilitate hearing lack the ability to regenerate, which has limited treatments for hearing loss. Current regenerative medicine strategies have focused on transplanting stem cells or genetic manipulation of surrounding support cells in the inner ear to encourage replacement of damaged stem cells to correct hearing loss. Yet, the extracellular matrix (ECM) may play a vital role in inducing and maintaining function of hair cells, and has not been well investigated. Using the cochlear ECM as a scaffold to grow adult stem cells may provide unique insights into how the composition and architecture of the extracellular environment aids cells in sustaining hearing function. Here we present a method for isolating and decellularizing cochleae from mice to use as scaffolds accepting perfused adult stem cells. In the current protocol, cochleae are isolated from euthanized mice, decellularized, and decalcified. Afterward, human Wharton&#39;s jelly cells (hWJCs) that were isolated from the umbilical cord were carefully perfused into each cochlea. The cochleae were used as bioreactors, and cells were cultured for 30 days before undergoing processing for analysis. Decellularized cochleae retained identifiable extracellular structures, but did not reveal the presence of cells or noticeable fragments of DNA. Cells perfused into the cochlea invaded most of the interior and exterior of the cochlea and grew without incident over a duration of 30 days. Thus, the current method can be used to study how cochlear ECM affects cell development and behavior.

The goal of this protocol is to demonstrate an effective method to decellularize and decalcify mouse cochleae for utilization as scaffolds for tissue engineering applications.

In mammals, mechanosensory hair cells that facilitate hearing lack the ability to regenerate, which has limited treatments for hearing loss. Current ...