The authors would like to acknowledge the financial support for this project received from the Slovenian Research Agency (grant numbers: P3-0036, and I0-0029), and the Ministry of Science, Education and Sport (grant number: 5442-1/2018/59).

1. Preparation of hydrogels and cross-linking solutions
<ol>
	<li>Briefly, by vigorously mixing, dissolve ALG and CMC powders in ultra-pure water to obtain a total 3 wt% ALG and 3 wt% CMC solution. 
	NOTE: In this work, 5 mL syringes are used for printing; thus, the final amount of material is adjusted to that volume. However, for other extrusion cartridges and printed sample sizes, the amount of prepared material should be scaled accordingly.</li>
	<li>Add 1.5 wt% cellulose nanofibers to the ALG-CMC mixture for additional mechanical reinforcement to reach the desired viscosity, suitable for printing.</li>
	<li>Agitate the hydrogel suspension until homogeneous using an overhead mixer. 
	NOTE: No fibers or bubbles should be present in the hydrogel.</li>
	<li>Prepare 10 mL of 100 mM calcium chloride (CaCl2) solution in ultra-pure water, which is used as the primary cross-linking solution for printing.</li>
	<li>Prepare 10 mL of 5 wt.% CaCl2 solution in ultra-pure water, which is used as a secondary cross-linking solution in post processing of scaffolds. 
	NOTE: Generally, all hydrogel formulations, which are suitable for immediate chemical cross-linking, allow for one-step fabrication of hollow tubes and can be used with this type of core/shell set-up. The printing and cross-linking mechanisms need to be optimized accordingly. Viscosity of the hydrogel will vary depending on the desired composition; however, it can be adjusted with polymer concentrations and the addition of thickening agents (e.g., nanofibers). The ideal viscosity for 3D printing of stable structures is high enough for the extruded filament to retain its shape and for the scaffold to hold its own weight before cross-linking.</li>
</ol>
2. Core/shell printing of perfusable scaffolds
<ol>
	<li>Prior to printing, sterilize the bioprinter by spraying 70% ethanol thoroughly and expose it to UV light for 1 h.</li>
	<li>Turn on the bioprinter and run the controlling software, which comes in a bundle with the 3D printer.</li>
	<li>Perform the homing procedure by pressing the Home icon.</li>
	<li>Using the toolbar command File | Import G-Code, import the generated scaffold g-code.</li>
	<li>Transfer the hydrogel to a sterile 5 mL syringe and place it in one of the extruder mounts of the 3D printer. Via the luer-lock and a short tube, connect it to side luer input of the core/shell nozzle.</li>
	<li>Transfer the cross-linking solution (100 mM CaCl2) to another sterile 5 mL syringe with an attached G27 blunt-end needle and insert it into the top needle holder of the core/shell nozzle. The inner needle should slightly protrude (~1 mm) from the outer core/shell nozzle. Adjust the alignment manually. Insert the second syringe in the extrusion mount.</li>
	<li>To correctly insert the syringes into the mounts (set-up shown in Figure 2), manually control both extrusion mounts by clicking the A and B and Up and Down arrows.</li>
	<li>Before the printing starts, separately extrude the hydrogel and the cross-linking solution to clear all excess air bubbles in the core/shell nozzle and ensure a continuous hydrogel flow.</li>
	<li>Using the Z and Up and Down arrows, manually adjust the distance between the nozzle and printing substrate. It is recommended to use a flat, glass-printing substrate, which has good adhesion. The extrusion nozzle should not be in contact with the substrate to allow for uninterrupted flow of the hydrogel. The optimal distance between the nozzle and substrate (layer height) is typically the same as the width of the outer nozzle diameter, but it is adjusted to the material used and individual printing parameters. Adjust the starting printing height according to individual needs.</li>
	<li>Press the Play button to start the printing process. 
	NOTE: It is recommended to include the printing of a skirt (Figure 3) surrounding the scaffold to ensure the laying of a homogenous hollow filament before printing of the actual scaffold starts. To achieve optimal hydrogel flow with the intention to print optimal scaffolds, vary formulation composition, cross-linking solution composition, and printing parameters (i.e., printing speed, extrusion pressure, printing temperature, distance between the substrate and extrusion nozzle, etc.).</li>
	<li>After printing, carefully remove the substrate with the printed scaffold and pour the secondary cross-linking solution (5 wt.% CaCl2) over the entire scaffold to ensure the cross-linkage over the entire scaffold. Incubate for 1 min at room temperature (RT). 
	NOTE: Make sure that the entire scaffold is submerged in the cross-linking solution. This step is crucial for achieving desired strength properties of the scaffold but will vary depending on the material and cross-linking method used.</li>
	<li>Using a scalpel manually cut the excess skirt material.</li>
	<li>Sterilize the scaffolds under an UV light for 30 min. Carefully flip the scaffold and repeat the sterilization process.</li>
	<li>Carefully detach the scaffold from the substrate by gently pulling it sideways. If the scaffold adheres strongly to the substrate, separate it by inserting a sharp edge between them.</li>
	<li>Transfer the scaffold into a colorless cell culture media (DMEM supplemented with 5 wt.% FBS, 100 U/mL penicillin and 1 mg/mL streptomycin), and incubate them at 37 &#176;C in an atmosphere containing 5 wt.% CO2 for at least 24 h.</li>
</ol>
3. Preparation of endothelial cells and live/dead assay solution
<ol>
	<li>For cell culturing, prepare the advanced DMEM cell culture medium with added phenol red and supplement it with 5 wt.% FBS and 2 mM L-glutamine. Add 100 U/mL penicillin and 1 mg/mL streptomycin.</li>
	<li>Initiate the human umbilical vein endothelial cell (HUVEC) line and passage them in accordance with HUVEC culturing protocols as described47. 
	NOTE: It is recommended to culture the cells in cell culture media with added phenol red for simple visualization during injection of the cells into white translucent scaffolds as described below.</li>
	<li>For cell counting, pipette 100 &#956;L of cells suspended in cell culture media and stain them with 900 &#956;L of 0.1 wt.% trypan blue solution.</li>
	<li>Use an automated cell counter or manual hemocytometer to count and obtain the estimated number of cells in suspension.</li>
	<li>For the live/dead assay, prepare a solution of 4 mM Calcein-AM and 2 mM propidium iodide in sterile PBS. 
	NOTE: The live/dead solution should be prepared directly before conducting the assay.</li>
</ol>
4. Transferring cells into scaffolds
<ol>
	<li>Remove the scaffolds from cell culture media and transfer them into a sufficiently large glass Petri dish.</li>
	<li>Immediately before injecting the cells into the scaffolds, dissociate the HUVECs from the flasks using treatment by 0.25 wt.% trypsin.
	<ol>
		<li>Briefly, dispose cell culture media and incubate the cells with 0.25 wt.% trypsin (~2 mL) for 5 min at 37 &#176;C.</li>
		<li>After incubation add ~3 mL of cell culture media to the trypsinized cells and transfer all detached cells to a centrifuge tube.</li>
		<li>Centrifuge the cells at 200 x g for 5 min and dispose of the supernatant.</li>
		<li>Resuspend the cells in fresh cell culture media.</li>
	</ol>
	</li>
	<li>Count the cells as described previously. Adjust the total cell concentration according to individual needs. In this work, a starting concentration of 340,000 cells/mL is used.</li>
	<li>Resuspend the cells in a sterile syringe with an attached blunt G27 needle.</li>
	<li>Find an entry point and start carefully injecting cells into the scaffolds. The cell suspension flow through a translucent scaffold should be visible. Make sure that the entire scaffold fills up with cell suspension.</li>
	<li>Submerge the scaffolds in cell culture media and incubate them at 37 &#176;C in an atmosphere containing 5 wt.% CO2 for up to 10 days.</li>
	<li>Replenish the cell culture media according to experimental needs.</li>
</ol>
5. Live/dead assay and cell imaging
<ol>
	<li>After incubation, rinse the scaffolds with PBS.</li>
	<li>With a blunt-end needle, carefully inject the previously prepared live/dead solution (4 mM calcein-AM and 2 mM propidium iodide in PBS) into the scaffolds and incubate them in PBS for 30 min at 37 &#176;C. Make sure that the solution flows through the length of the entire scaffold.</li>
	<li>Rinse the scaffolds with PBS.</li>
	<li>Carefully transfer the scaffolds to a glass slide.</li>
	<li>Observe the dyed cells directly in the scaffolds under a fluorescence microscope. 
	NOTE: Viable cells produce green fluorescence and dead cells emit red fluorescence light.</li>
</ol>

The authors declare that they have no competing financial interests.

Nozzle design 
Using the developed core/shell nozzle, integrated into a two-extruder Vitaprint system, hollow, tubular scaffolds were fabricated in a single-step process. To achieve an even thickness of the tube wall through most of the prepared scaffolds, the needle needs to be positioned centrally on the axis of the outer extrusion ring. Standard gauge needles often exhibit a slight, yet significant eccentricity off the axis. Thus, the nozzle body was designed to hold the needle in two places, once at the top (fixing the hub) and once before the final core/shell chamber (fixing the cannula itself), correcting its axial alignment. The precision of the axial alignment increases with the distance between the fixation point. There is, however, a tradeoff between needle length and available nozzle chamber volume. To improve functionality of the set-up further, certain modifications of the nozzle can be implemented: A) a nozzle mount with improved stability, B) additional nozzles for a wider range of needle compatibility, C) a precise adjustment mechanism for needle to nozzle positioning, and D) integrating additional inputs and microfluidic devices for on the fly material preparation.
Hydrogel optimization 
To determine the optimal ALG:CMC ratio, several material iterations were evaluated. Generally, core/shell printing with concentrations above 3 wt.% of both components was rendered impossible, because it did not allow for a continuous hydrogel flow or resulted in clogging of the nozzle. Specifically, ALG concentration above 3 wt.% increased the viscosity excessively and resulted in nozzle clogging, while lower ALG concentrations and higher CMC (&#62;3 wt.%) concentrations slowed down cross-linking times and thus failed to provide sufficient structural support of the scaffold. Core/shell printing was possible with less viscous formulations; however, extruded gel viscosity must be sufficient to sustain long term shape fidelity. In the end, a 1:1 ALG:CMC ratio was proven to be the most suitable choice which confirms a previous study by Maver et al.49. The addition of NFC significantly improved printability and structural rigidity of core/shell printed scaffolds but had no significant effect on cross-linking properties of the material.
Custom applications, optimized for specific cell types and experimental set-ups will require well-tailored scaffolding materials, which will vary in composition and cross-linking mechanisms. The method described in this work is based on an alginate-cellulose mixed polymer solution, which is cross-linked ionically using Ca2+ ions. Alginate itself is a linear polymer of blocks of (1,4)-linked &#946;-d-mannuronate (M) and &#945;-l-guluronate (G) residues that can be reversibly ionically crosslinked by application of Ca2+ and other divalent cations such as Sr2+, Br2+, Mg2+. Nevertheless, the most widely used ion for cross-linking of alginate remains Ca2+ in the form of CaCl2. Ca2+ can also be used in the form of CaSO4 or CaCO3; however, the low solubility of CaSO4 relative to CaCl2 means slower gelation. CaCO3 yields even slower gelation times which can result in weak and inconsistent mechanical properties.
Longer gelation times typically produce a more homogenous construct, however, certain applications, such as core/shell printing requires fast gelation rates50. Mg2+ ions also induce gelation; however, their cross-linking efficiency is about 5x-10x lower, compared to Ca2+, with cross-linking times of 2-3 h. In addition, magnesium ions are more selective towards guluronic units, hence the cross-linking depends more on the chemical composition of the ALG51. In this case, a fast gelation rate is essential to ensure continuous hollow channel formation before the hollow structure can collapse. CaCl2 yields the fastest gelation rate, which is crucial for the direct deposition of hollow filaments. 100 mM CaCl2 was utilized, which adequately stabilized the continuous formation of a hollow filament without causing gel solidification within the nozzle.
Printing and post-processing of scaffolds 
The following steps should be considered during this portion of the process, including 1) ensuring that all solutions and materials including the 3D bioprinter are properly sterilized before printing. 2) When preparing the hydrogel, homogeneity of the material is crucial for continuous printing. Introduction of impurities or air bubbles should be avoided, as they can clog the nozzle and/or disrupt the extrusion. 3) The syringes should be properly connected to the core/shell nozzle via the luer-lock mechanism and correctly inserted in the extruder mounts as seen in Figure 2A,B. 4) Before printing a complex structure, it is recommended to pre-extrude a small portion of the gel and cross-linking solution to clear the excess air bubbles in the core/shell nozzle and ensure a continuous hydrogel flow. This can be incorporated directly into the g-code to improve repeatability. 5) It is helpful to add a skirt surrounding the scaffold to ensure the laying of a homogenous hollow filament before the printing of the scaffold itself starts.
Additionally, 6) to improve adhesion between the printing filament and substrate, it is recommended to use a flat surface with good adhesion (i.e., a glass slide or Petri dish). 7) The extrusion nozzle should not be in direct contact with the substrate to allow for uninterrupted flow of the hydrogel. The initial distance will strongly impact the quality of the print, but thickness of the extruded filament is a good approximation of the initial setting. 8) The starting printing height in the g-code should be adjusted according to individual needs. After the printing parameters are optimized, the scaffold g-code should be imported into Planet CNC software and the printing process started as described in the protocol. 9) To control and optimize hydrogel flow with the intention to print optimal scaffolds, both formulation composition and printing parameters should be varied (i.e., printing speed, extrusion pressure, printing temperature, distance between the substrate and extrusion nozzle, layer height, scaffold size, etc.).
In general, higher flow rates are required to print formulations with higher viscosity. As mentioned, all hydrogel formulations, which are suitable for immediate chemical cross-linking, allow for one-step fabrication of hollow tubes and may be used with the described core/shell set-up. The printing and cross-linking mechanisms need to be optimized accordingly. After printing, all scaffolds were post-processed by secondary cross-linking with 5 wt.% CaCl2 solution, which assured complete cross-linking of the ALG-CMC component and sterilized from both sides under a UV light for at least 30 min. It should be ensured to completely engulf the scaffold with the cross-linking solution and incubate for long enough to complete the cross-linking process. Post-processing will differ based on the material and cross-linking mechanism used, which should be considered beforehand. After post-processing, scaffolds should be removed carefully from the substrate, transferred to cell culture media, and incubated in a controlled atmosphere for at least 24 h before cell seeding. Using a colorless medium will improve visibility of the cell suspension during injection into the scaffolds.
Live/dead assay 
The live/dead solution should be prepared directly before conducting the assay and kept in the dark before conducting the assay, as it contains fluorescence dyes that are prone to bleaching. After the desired incubation time, the cell culture media should be carefully discarded surrounding the scaffolds and rinsed with PBS. Ideally, the same entry point should be used for cell seeding followed by the live/dead assay being injected into the scaffolds.
Importance of results 
Both ALG and CMC have been already used to promote angiogenesis in vitro. Based on its ECM-mimetic features, physical cross-linking, and biocompatibility, ALG has been commonly employed as a component for delivery and controlled release of angiogenic growth factors (e.g., bFGF, HGF, VEGF164, and Ang-1*respectively)52,53,54. Furthermore, in combination with gelatine, CMC has also been used for encapsulating vascular endothelial cells due its rapidly cross-linking capabilities under physiological conditions55. NFC were added to further increase the mechanical stability and shape fidelity of scaffolds. It should be emphasized that the objective was not to enhance the vascularization but to demonstrate the possibility of producing perfusable, hollow ALG-CMC scaffolds, printed in core/shell fashion, which also facilitates the attachment and proliferation of HUVECs. The choice of using an ALG-CMC mixture was based on findings of commonly used, easily accessible, and biocompatible base materials that could enable core/shell printing of hollow channels. Many other materials may be more viable options for enhancing angiogenesis; however, some are not suitable for core/shell printing, as they do not facilitate rapid gelation/cross-linking, which is crucial in this approach.

Arguably, the ultimate objective of tissue engineering (TE) is to produce functional tissues or organs in vitro, which can be used to regenerate or replace injured or diseased parts of the human body1,2,3. Current research in tissue engineering (TE) is focused on individual aspects of the field (scaffolding materials, fabrication procedures, cell sources, etc.)4,5, as well as developing simple in vitro models of tissues and organs that mimic fundamental aspects of their in vivo counterparts. Such models are already useful for many applications, such as drug screening and toxicity studies, especially in cases where conventional 2D cell cultures fail to mimic the dynamic responses of the native tissues6,7,8,9. Three-dimensional in vitro models are usually constructed by combining cells10, physico-chemical cues11, and biologically active molecules12,13 on scaffolds, which are obtained from decellularized tissues or constructed de novo from biological or biocompatible materials14,15,16,17,18.
It is crucial that scaffolds recapitulate the complex 3D microarchitecture and hierarchical structure of the native tissues to enable functionality of the engineered tissues, representative of in vivo tissues19. Despite the significant technological advancement in TE, development of physiologically relevant artificial tissue constructs remains a challenge. Thick tissues ( &#62;200 &#181;m in thickness) are especially problematic, due to limitations such as oxygen and nutrient diffusion20. Progress towards larger tissue constructs has been made; however, the required high proximity of cells to blood vessels in order to transport oxygen and nutrients and promote waste removal must be recapitulated. Vascularization of tissues (or alternatively, fabrication of interconnected 3D vascular networks within tissue constructs) plays a critical role in maintaining cell viability and promoting functions of in vitro engineered tissues, which is more difficult for models in prolonged experiments21,22. Furthermore, the required resolution, structural integrity, and simultaneous biocompatibility has yet to be achieved23.
Several TE approaches have been proposed in an attempt to construct blood vessel-like structures and facilitate vascularization in vitro. Some examples include seeding endothelial cells (also co-cultured with other cells types such as fibroblasts) that self-assemble to generate microvascular networks24, use of vascular progenitor cells and pericytes that promote endothelial cell growth21,25, the delivery of angiogenic growth factors that induce vascularization20,26, using cell sheet technology that allows for control over vascular layering20, and fabrication of highly porous scaffold structures that promote angiogenesis27. The mentioned approaches focus on angiogenesis induction, which generally requires considerable amounts of additional growth factors (e.g., VEGF) and time to form. However, the biggest drawbacks are their limited reproducibility and restricted spatial control over vascular patterning, usually resulting in a random vasculature distribution within the tissue construct that does not necessarily facilitate perfusion.
Additive manufacturing (AM, such as 3D bioprinting) is increasingly involved in the fabrication of 3D constructs using biological or biocompatible materials to create scaffolds suitable for TE. Several AM approaches are being used and developed in parallel (e.g., ink jet- and microextrusion-based methods, different types of lithographic techniques) to produce scaffolds that mimic native tissues in their architecture, biochemistry, and functionality. The individual techniques exhibit certain advantages and disadvantages28, which is why various modifications being explored (e.g., micro-patterning, induced angiogenesis, etc.) to increase the extent to which large, complex, and stable vascular networks can be fabricated22,29,30.
Among these, extrusion bioprinting is the most commonly used method, especially due to the broad range of compatible materials (a generally cell-friendly process28,31,32) as well as exceptional versatility in terms of applications (e.g., embedded and sacrificial printing23,33, fabrication of hollow structures34,35, etc.). The main challenges preoccupying present studies include the transfer from 2D to 3D structures, formation of a dense network of hollow tubes with high spatial resolution, and overall mechanical integrity and shape fidelity during fluid flow in cell culture conditions30.
The most straightforward approach to perfusable tissue is the fabrication of an interconnected network of channels within the construct. The creation of such perfusable channels within a tissue scaffold is expected to solve many of the aforementioned problems, as it immediately allows for nutrient and oxygen diffusion while removing waste products. Therefore, the potential formation of necrotic regions within the construct is avoided36. Such channels may additionally be seeded with endothelial cells (ECs) and serve as artificial blood vessels in 3D tissue models37. In the most elementary sense, a vessel can consist of a hollow channel, soft layer of ECs, and stiff shell. Recently, 3D extrusion of two different materials in a core/shell fashion utilizing co-axial needles for extrusion has gained much interest38,39,40,41, as it allows for fabrication of hollow tubes.
Similar to conventional microextrusion 3D printing, core/shell printing is performed with a co-axial nozzle (e.g., two needles with different diameters aligned on the same axis in a manner, so that the wider needle encloses the narrower one). Thus, two materials can be extruded simultaneously, with one as the central filament or &#8220;inner&#8221; core and a second as the &#8220;outer&#8221; shell41. To date, co-axial bioprinting has been utilized to fabricate structures with solid42, core/shell43, and hollow strands40,44; however, the materials used have not been optimized for both optimal cell viability and mechanical robustness of the printed constructs. As mentioned, the technique provides the possibility to combine biomaterials with different mechanical properties, in which the stiffer one supports the softer one. More importantly, if the scaffold material (e.g., alginate, carboxymethyl cellulose) is extruded as the shell, while the core composed of the cross-linking agent (e.g., calcium chloride) is dispensed from the inner capillary then rinsed out post-printing, it is possible to fabricate a continuous hollow tube in a single step45.
With this in mind, a simple and repeatable one-step method was developed to build well-defined and perfusable scaffolds for the engineering of vascular structures and other tubular tissues. To develop a cost-effective technology, fabrication should ideally be a single-step process. Therefore, a core/shell set-up was adapted and integrated into the 3D bioprinter. The basic design consists of a central nozzle made of metal to avoid deformation during injection, around which a second nozzle of a larger diameter is placed. Such a co-axial nozzle set-up allows for co-extrusion of the two flows and immediate cross-linking of the extruded hydrogel channel. This enables direct fabrication of multilayered hollow filaments, while subsequent cross-linking with higher concentrations of calcium chloride (CaCl2) ensure more permanent stabilization from the outside.
As such, this method allows for concurrent printing of scaffolds and microchannels, in which the hollow hydrogel filaments serve as a scaffold to support the mechanical integrity of 3D constructs and simultaneously act as built-in microchannels to deliver nutrients for cell growth. This protocol provides a detailed procedure of the core/shell 3D bioprinting strategy based on use of a custom-made co-axial nozzle in which hydrogel 3D structures with built-in channels are fabricated by controlling cross-linking to produce hollow filaments, which remain perfusable during cell culture.
The 3D printing set-up used in this work is configured as previously described by Banovi&#263; and Vihar46 and can be divided into three main components: A) a three-axis CNC mechanical set-up with 50 &#956;m positioning accuracy in the X, Y, and Z directions; B) two extruders, adapted for disposable, 5 mL luer-lock syringes, with 1.2 &#956;L voxel resolution; and C) controlling electronics and software.
To facilitate core/shell printing, an appropriate nozzle was developed that can be mounted on one of the extruders (primary extruder, printing the core) and is compatible with G27 blunt-end needles. It also has luer-lock compatibility to connect with the second extruder (printing the shell). The first prototypes were fabricated by inserting a blunt-end G27 needle (inner diameter = 210 &#956;m, outer diameter = 410 &#956;m) into either a G21 needle (inner diameter = 510 &#956;m, outer diameter = 820 &#956;m) or G20 conical tip (inner diameter = 600 &#956;m), then inserting a secondary needle laterally to supply the shell material. However, due to slight bending of the needle shaft, it is not possible to produce a nozzle tip with concentric alignment of the inner and outer needles.
To solve this issue, a new nozzle design was devised that fulfilled the following criteria: 1) it can be manufactured using a 3-axis CNC mill, 2) it can be made from various materials (high performance plastics, such as PEEK or metals), 3) it has luer-lock compatibility for applying shell material, and 4) is compatible for a G27 blunt-end needle and holds it in place at two positions to align the tip with the central axis. A schematic of the nozzle prototype is shown in Figure 1.

<table>
	<tbody>
		<tr>
			<td>Alginic acid sodium salt</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>180947</td>
			<td>powder; Mw ~80,000</td>
		</tr>
		<tr>
			<td>ATTC HUV-EC-C [HUVEC]</td>
			<td>LGC Standards (UK)</td>
			<td>ATCC-CRL-1730</td>
			<td>Endothelial Cell Growth Supplement (ECGS) and unidentified factors from bovine pituitary, hypothalamus or whole brain extracts are mitogenic for this line; the cells have a life expectancy of 50 to 60 population doublings.</td>
		</tr>
		<tr>
			<td>Axiovert 40 inverted optical microscope</td>
			<td>Carl Zeiss Microscopy GmbH (Germany)</td>
			<td></td>
			<td>three contrastingtechniques in one objective &#8211; e.g. brightfield,phase contrast and PlasDIC</td>
		</tr>
		<tr>
			<td>Calcium chloride</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>C1016</td>
			<td>anhydrou; granular; &#8804;7.0 mm; &#8805;93.0%</td>
		</tr>
		<tr>
			<td>Cellulose nanofibrils suspension (NFC, 3% (w/v))</td>
			<td>The Process Development Center, University of Maine (Maine, USA)</td>
			<td></td>
			<td>nominal fiber width of 50 nm; lengths of up to several hundred microns</td>
		</tr>
		<tr>
			<td>ELGA Purelab water purification system</td>
			<td>Veolia Water Technologies (UK)</td>
			<td></td>
			<td></td>
		</tr>
		<tr>
			<td>EVOS FL Cell Imaging System</td>
			<td>ThermoFisher Scientific Inc. (Germany)</td>
			<td>AMF4300</td>
			<td>a fully integrated, digital, inverted imaging system for four-color fluorescence and transmitted-light applications</td>
		</tr>
		<tr>
			<td>Gibco Advanced Dulbecco&#8217;s modified Eagle&#8217;s medium (Advance DMEM)</td>
			<td>ThermoFisher Scientific Inc. (Germany)</td>
			<td>12491015</td>
			<td>high glucose; no glutamine; phenol red</td>
		</tr>
		<tr>
			<td>Gibco Dulbecco&#8217;s modified Eagle&#8217;s medium (DMEM)</td>
			<td>ThermoFisher Scientific Inc. (Germany)</td>
			<td>21063029</td>
			<td>high glucose; L-glutamine; HEPES; no phenol red</td>
		</tr>
		<tr>
			<td>Gibco Fetal Bovine Serum (FBS), qualified</td>
			<td>ThermoFisher Scientific Inc. (Germany)</td>
			<td>10270106</td>
			<td>FBS origin: Brazil; 5 % (w/v) FBS</td>
		</tr>
		<tr>
			<td>Hypodermic Sterican needle</td>
			<td>B. Braun Melsungen AG (Germany)</td>
			<td>9180117</td>
			<td>0.40 x 25mm, 27G x 1&#39;&#39;</td>
		</tr>
		<tr>
			<td>L-glutamine</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>G3126</td>
			<td>ReagentPlus&#174;, &#8805;99% (HPLC)</td>
		</tr>
		<tr>
			<td>Live/Dead Cell Double Staining Kit</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>4511</td>
			<td>contains calcein-AM and propidium iodide (PI) solutions; suitable for fluorescence</td>
		</tr>
		<tr>
			<td>Nunc EasYFlask cell culture flasks</td>
			<td>ThermoFisher Scientific Inc. (Germany)</td>
			<td>156367</td>
			<td>Nunclon Delta certified for monolayer formation, cloning efficiency, non-cytotoxic, non-pyrogenic, and sterility; filter caps; culture area of 25 cm2</td>
		</tr>
		<tr>
			<td>Omnifix syringe</td>
			<td>B. Braun Melsungen AG (Germany)</td>
			<td>4617053V</td>
			<td>5 mL Luer Lock</td>
		</tr>
		<tr>
			<td>Penicillin G sodium salt</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>P3032</td>
			<td>powder; BioReagent; suitable for cell culture</td>
		</tr>
		<tr>
			<td>Phosphate buffered saline</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>P4417</td>
			<td>tablet; one tablet dissolved in 200 mL of deionized water yields 0.01 M phosphate buffer, 0.0027 M potassium chloride and 0.137 M sodium chloride, pH 7.4, at 25 &#176;C</td>
		</tr>
		<tr>
			<td>Sodium carboxymethyl cellulose</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>419338</td>
			<td>powder; average Mw ~700,000</td>
		</tr>
		<tr>
			<td>Streptomycin sulfate salt</td>
			<td>Sigma-Aldrich (Germany)</td>
			<td>S9137</td>
			<td>powder; BioReagent; suitable for cell culture</td>
		</tr>
		<tr>
			<td>Ultra-pure water</td>
			<td>Veolia Water Technologies (UK)</td>
			<td></td>
			<td>18.2 m&#8486; cm at 25&#8304;C</td>
		</tr>
		<tr>
			<td>VitaPrint 3D bio-printer</td>
			<td>IRNAS (Slovenia)</td>
			<td></td>
			<td></td>
		</tr>
	</tbody>
</table>

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.

The aim of this work was to develop an easy-to-manufacture core/shell nozzle with luer-compatibility for core/shell printing of woodpile structures. In addition, a straightforward and repeatable one-step printing protocol was described, which is simple to modify and accommodates a broad range of materials and different chemical cross-linking mechanisms to build well-defined and perfusable scaffolds for engineering of vascular and other tubular tissue structures.
Core/shell nozzle 
The nozzle is composed of a G27 blunt-end needle (for printing the inner axial filament) and the nozzle body, which holds the needle in place and creates an outer nozzle for the shell filament with a connection port for material input. A schematic is shown in Figure 1. Two 5 mL syringes, which are placed into the individual extruders and provide the core and shell materials. Tubing connects the nozzle body with the syringe, providing the shell material.
The complete assembly of the core/shell nozzle and syringe set-up is shown in Figure 2. The first functional prototype of the nozzle body was manufactured by CNC milling a block of polyoxymethylene (POM). Compatibility and sealing between the element, a G27 needle and tubing was tested by installing into Vitaprint. No leakage of the shell material was observed in the nozzle, luer-lock connector, or between body and needle. The nozzle body fits tightly with the needle hub (luer connector), ensuring synchronized motion of the whole nozzle with the extruder.
Scaffold building 
The most straightforward approach to fabricating scaffolds is by depositing materials layer by layer where printing direction is changed every consecutive layer, typically at an angle of 90&#176;. Due to the visco-elastic and hygroscopic properties of the hydrogels, retaining shape fidelity of the printed structures remains challenging. The main purpose of this protocol section is 3D printing of perfusable core/shell scaffolds using two polymers, which have previously shown promising results for scaffold building (ALG and CMC) with the addition of cellulose nanofibers (NFC) for increased mechanical stability. Both ALG and CMC are negatively charged, water-soluble linear copolymers48 and both contain carboxyl groups, which can be cross-linked with the addition of divalent cations. Ca2+ particles form ionic bonds with two functional groups simultaneously, forming connections between polymer chains, increasing gel rigidity.
Printing of perfusable scaffolds 
The aim of this process is to 3D-print simple woodpile scaffold structures, which span over several layers and retain their shape fidelity as well as perfusability until becoming fully cross-linked. This requires even coextrusion of core and shell without crossover or retraction movements, which can interrupt the flow. Therefore, typical CAD modelling and slicing methods are less suitable. In this work, a manually designed g-code is used, and a python-based g-code generator was developed for fast g-code preparation.
The scaffolds were structured in a woodpile grid shape and built on a flat glass surface. The fabrication was performed layer-by-layer, depositing the crossing lines of each succeeding layer in an angle 90&#176; to the previous one. Additionally, each succeeding layer was 2% narrower in the X and Y directions, ensuring continuous filament support by preceding layers. The distance between filaments (macropore size) can be precisely designed in the g-code by taking into account the outer diameter of the extruded filament (0.8 mm) and the distance between gridlines (3 mm).
Scaffolds were required to fulfill key inclusion criteria for further consideration. First, scaffolds with at least 4 layers in height were required to retain their structural integrity and geometry (e.g., macropore size) during printing, to be eligible for further development. Second, scaffolds needed to remain perfusable (stable microchannels) even after being incubated for 7 days in cell culture media at 37 &#176;C. At appropriate time intervals (1, 2, 5, and 7 days) scaffolds were taken out of the culture media and tested to see if they were still perfusable. In Figure 4A, the cross-section of a freshly printed and post-processed scaffold displays a clearly visible hollow channel inside the filament. In Figure 4B, it is clear that even after 72 h incubation in cell culture media at 37 &#176;C, the filament retains the hollow structure through the entire scaffold length.
Several formulations were printable, remained structurally stable, preserved the printed geometry, and remained perfusable; however, a single one was chosen for further testing (i.e., 3 wt.% ALG + 3 wt.% CMC + 1.5 wt.% NFC), which allowed printing of perfusable scaffolds with up to 10 layers. The printing process with the custom nozzle is shown in Figure 5A. Core/shell printing was possible with less viscous formulations; however, gels with higher concentrations did not allow continuous flow through the nozzle. For primary cross-linking (core material, delivered during printing) 100 mM CaCl2 was utilized, which adequately stabilized the continuous formation of a hollow filament, without causing gel solidification within the nozzle. Post-printing, the scaffolds were soaked in a 5 wt.% CaCl2 solution to completely cross-link the hydrogel for long-term shape fidelity. A finished scaffold sample is pictured in Figure 5B. During the optimization, process of the formulation an artificial dye was used in the core solution, to allow visual evaluation and examining the quality of the extruded filament. The dye was not used for the fabrication of the final scaffolds, which were prepared for cell seeding and cultivation.
Live/dead assay 
After incubation, a live/dead assay was used to visualize the ECs and distinguish between the living (green) and the dead (red) cells within the incubated scaffold. This served two main purposes: A) to determine whether the scaffolds provide a biocompatible environment to promote growth and adhesion without exhibiting harmful effects on the cell, and B) to visualize the structural integrity of tubular structures and their internal channel system in more detail.
The results of the live/dead assay are shown in Figure 6. In the presence of intracellular esterases, the plasma membrane permeable Calcein-AM is converted to Calcein, emitting green fluorescence light in live cells. On the other hand, apoptotic cells are visualized by membrane impermeable propidium iodide, which fluoresces in red when intercalated in the DNA double helix. Live/dead images and bright-field pictures of scaffolds were combined to help visualize the cells inside the hollow channels. The staining solution was injected, and the assay performed directly in the 3D printed scaffolds after scaffold-cell incubation for 48 h.
It should be noted that a relatively small seeding density of ECs was used (340,000 cells/mL), as this study served only as a proof-of-concept for core/shell printing of perfusable scaffolds. The most significant conclusion of the live/dead assay is that even after 48 h, no dead cells (red) were observed, proving that neither the scaffold material itself, nor its degradation products exhibited toxic effects. Furthermore, the ECs did in fact adhere and remain attached inside the scaffolds and seemed to form evenly distributed agglomerates when grown inside the channels. This suggests that the described fabrication method and scaffold formulation provide a suitable framework for building in vivo, relevant, tubular tissue morphologies. Besides mimicking cell-ECM interactions and tight cell-cell communication in all three spatial dimensions, complex tissue engineering also requires constant cell exposure to fresh medium to sustain their viability. This in turn can be achieved by a dense channel network under continuous perfusion, which warrants further investigation in future work, and will require optimizing material and growth parameters to facilitate long-term tissue engineering of vasculature.
<img alt="Figure 1" class="xfigimg" src="/files/ftp_upload/59951/59951fig1.jpg" /> 
Figure 1: Core/shell nozzle prototype. (A) The overall design and the main components of the nozzle body prototype are shown. The nozzle is completed by inserting a blunt-end G27 needle through the top. The top and bottom needle holders immobilize and realign the needle with the nozzle axis, ensuring that the tip extends from the nozzle through the center.&#160; To connect the nozzle with &#8220;shell&#8221; material extruded from the secondary syringe, tubing with a luer-lock connector is attached to the lateral input. From here, the material is forwarded to the nozzle through a narrow channel. The fabrication of the mentioned channel requires drilling in two positions, producing holes that need to be capped after manufacturing). (B) Shown is a close-up of the nozzle with an inserted G27 needle, extending out of the nozzle. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig1large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 2" class="xfigimg" src="/files/ftp_upload/59951/59951fig2.jpg" /> 
Figure 2: Final core/shell set-up. (A) Shown is the completed core/shell nozzle with correctly attached syringes containing the hydrogel (right), building the &#8220;shell&#8221; and cross-linking solution (left) extruded as the &#8220;core&#8221;. (B) Shown is the core/shell set-up installed into the Vitaprint system with two extruders. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig2large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 3" class="xfigimg" src="/files/ftp_upload/59951/59951fig3.jpg" /> 
Figure 3: G-code of the tubular scaffold. Here a screenshot of the printer software&#160; is shown, specifically the path preview (A) and raw g-code of the first layer (B). The g-code is a set of instructions with target coordinates in absolute spatial directions (X, Y, Z), as well as extrusion (A,B) in relative directions. The G command determines the type of instruction, whereas G1 represents linear movement towards the target coordinates and G92 determines the initial starting position. In addition, the feed-rate of following commands is determined with instruction F in mm/min. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig3large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 4" class="xfigimg" src="/files/ftp_upload/59951/59951fig4.jpg" /> 
Figure 4: Cross-section of a scaffold strand with a hollow interior. (A) Shown is the cross-section slice of the freshly printed and post-processed scaffold. (B) Shown is the cross-sectional slice of the scaffold after being incubated in cell culture media for 72 h. While the nozzle shape defines the extrusion of a tube with a round cross-section, the filament appears to be somewhat flattened on deposition. The internal channel, however, stays intact and retains its form during incubation. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig4large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 5" class="xfigimg" src="/files/ftp_upload/59951/59951fig5.jpg" /> 
Figure 5: Core/shell printing of scaffolds. Here, fabrication (A) of a three-layered hollow-tube scaffold and its final form (B) are shown. For improved visualization, the cross-linking solution in the core was stained with a red dye. The formulation exhibits sufficient mechanical stability to retain scaffold stability, even if thicker structures (up to 10 layers, data not shown) are fabricated. The outer dimensions of the final structure were approximately 27 mm x 27 mm x 3.5 mm. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig5large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 6" class="xfigimg" src="/files/ftp_upload/59951/59951fig6.jpg" /> 
Figure 6: Live/dead assay performed directly in the scaffolds. A suspension of HUVECs was injected into the inner scaffold channel, incubated for 48 h, and treated with live/dead dye. Viable HUVECs emit green fluorescence light, which is represented in the bright spots of the image. Dead cells emit green fluorescence light; however, none are visible in the observed scaffold. The distribution of the cells also signifies the shape and retained perfusion capabilities of the channels. To a minor extent, the live/dead assay solution also stained the scaffolding material, producing light fluorescence under the microscope. <a href="https://www.jove.com/files/ftp_upload/59951/59951fig6large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>

Watch this Scientific Journal Video about Core/shell Printing Scaffolds For Tissue Engineering Of Tubular Structures at JoVE.com

Core/shell Printing Scaffolds For Tissue Engineering Of Tubular Structures

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 ...

coreshell-printing-scaffolds-for-tissue-engineering-tubular

Research

JoVE Journal

Bioengineering

9.2K Views.  University of Maribor. 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.

Video: Core/shell Printing Scaffolds For Tissue Engineering Of Tubular Structures

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 ...

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 ...

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

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

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

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

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

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 ...

Self-reporting Scaffolds for 3-Dimensional Cell Culture

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

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

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

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 ...

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 ...

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. ...

Epithelial Cell Repopulation and Preparation of Rodent Extracellular Matrix Scaffolds for Renal Tissue Development

This protocol details the generation of acellular, yet biofunctional, renal extracellular matrix (ECM) scaffolds that are useful as small-scale model substrates for organ-scale tissue development. Sprague Dawley rat kidneys are cannulated by inserting a catheter into the renal artery and perfused with a series of low-concentration detergents (Triton X-100 and sodium dodecyl sulfate (SDS)) over 26 hr to derive intact, whole-kidney scaffolds with intact perfusable vasculature, glomeruli, and renal tubules. Following decellularization, the renal scaffold is placed inside a custom-designed perfusion bioreactor vessel, and the catheterized renal artery is connected to a perfusion circuit consisting of: a peristaltic pump; tubing; and optional probes for pH, dissolved oxygen, and pressure. After sterilizing the scaffold with peracetic acid and ethanol, and balancing the pH (7.4), the kidney scaffold is prepared for seeding via perfusion of culture medium within a large-capacity incubator maintained at 37 &#176;C and 5% CO2. Forty million renal cortical tubular epithelial (RCTE) cells are injected through the renal artery, and rapidly perfused through the scaffold under high flow (25 ml/min) and pressure (~230 mmHg) for 15 min before reducing the flow to a physiological rate (4 ml/min). RCTE cells primarily populate the tubular ECM niche within the renal cortex, proliferate, and form tubular epithelial structures over seven days of perfusion culture. A 44 &#181;M resazurin solution in culture medium is perfused through the kidney for 1 hr during medium exchanges to provide a fluorometric, redox-based metabolic assessment of cell viability and proliferation during tubulogenesis. The kidney perfusion bioreactor permits non-invasive sampling of medium for biochemical assessment, and multiple inlet ports allow alternative retrograde seeding through the renal vein or ureter. These protocols can be used to recellularize kidney scaffolds with a variety of cell types, including vascular endothelial, tubular epithelial, and stromal fibroblasts, for rapid evaluation within this system.

This protocol describes decellularization of Sprague Dawley rat kidneys by antegrade perfusion of detergents through the vasculature, producing acellular renal extracellular matrices that serve as templates for repopulation with human renal epithelial cells. Recellularization and use of the resazurin perfusion assay to monitor growth is performed within specially-designed perfusion bioreactors.

This protocol details the generation of acellular, yet biofunctional, renal extracellular matrix (ECM) scaffolds that are useful as small-scale model ...

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 ...