Generating vascular channels within hydrogel constructs using an economical open-source 3D bioprinter and thermoreversible gels

Bioprinting 9 (2018) 7–18

Contents lists available at ScienceDirect

Bioprinting journal homepage: www.elsevier.com/locate/bprint

Generating vascular channels within hydrogel constructs using an economical open-source 3D bioprinter and thermoreversible gels

T

⁎

Ross EB Fitzsimmonsa,b, , Mark S. Aquilinoa, Jasmine Quigleya,b, Oleg Chebotarevb,c, ⁎ Farhang Tarlana,b, Craig A. Simmonsa,b,c, a

Institute of Biomaterials and Biomedical Engineering, University of Toronto, 164 College Street, Toronto, Ontario, Canada M5S 3G9 Translational Biology and Engineering Program, Ted Rogers Centre for Heart Research, University of Toronto, 661 University Ave, 14th floor, Toronto, Ontario, Canada M5G 1M1 c Department of Mechanical and Industrial Engineering, University of Toronto, 5 King's College Road, Toronto, Ontario, Canada M5S 3G8 b

A R T I C L E I N F O

A B S T R A C T

Keywords: Vascularization Pluronic F-127 Gelatin Hyaluronan Bioprinter 3D printing

The advent of 3D bioprinting offers new opportunities to create complex vascular structures within engineered tissues. However, the most suitable sacrificial material for producing branching vascular conduits within hydrogel-based constructs has not yet been resolved. Here, we assess two leading contenders, gelatin and Pluronic F-127, for a number of characteristics relevant to their use as sacrificial materials (printed filament diameter and its variability, toxicity, rheological properties, and compressive moduli). To aid in our assessment and help accelerate the adoption of 3D bioprinting by the biomedical field, we custom-built an inexpensive (< $3000 CAD) 3D bioprinter. This open-source 3D printer was designed to be fabricated in a modular manner with 3D printed/laser-cut components and off-the-shelf electronics to allow for easy assembly, iterative improvements, and customization by future adopters of the design. We found Pluronic F-127 to produce filaments with higher spatial resolution, greater uniformity, and greater elastic modulus than gelatin filaments, and with low toxicity despite being a surfactant, making it particularly suitable for engineering smaller vascular conduits. Notably, the addition of hyaluronan to gelatin increased its viscosity to achieve filament resolutions and print uniformity approaching that with Pluronic F-127. Gelatin-hyaluronan was also more resistant to plastic deformation than Pluronic F-127, and therefore may be advantageous in situations in which the sacrificial material provides structural support. We expect that this work to establish an economical 3D bioprinter and assess sacrificial materials will assist the ongoing development of vascularized tissues and will help accelerate the widespread adoption 3D bioprinting to create engineered tissues.

1. Introduction The development of vasculature for engineered tissue constructs remains a major impediment to the formation of viable tissues for both scientific inquiry and clinical use. One key aspect of this challenge is the generation of branching channels within hydrogel constructs that can then be seeded with endothelial cells to produce vessel-like structures. Favorably, the nascent field of 3D bioprinting is now enabling this technical capability [1,2]. The capacity to design and print complex 3D structures on multiple length scales within a single construct has opened the possibility of generating interconnected and branching vessel systems consisting of both macrovessels (> 1 mm) and small artery/vein-like microvessels (~0.5–1 mm), a feat not possible with tissue engineering methods preceding 3D printing, which have largely

been limited to creating relatively simple planar and tubular macroscopic constructs. In order to print complex 3D branching vessel systems, a number of requirements must be met by both the materials and the 3D printing technology used to create vascular channels. To create perfusable channels within hydrogel constructs, sacrificial materials can be deposited in the desired vascular design during printing and then flushed away once the construct is complete. Such sacrificial materials must be non-toxic and maintain a uniform filament diameter during printing to ensure well-controlled material deposition to accurately and consistently recapitulate the desired vascular architecture. Similarly, the printer itself must have sufficient resolution to print all the necessary channels, including those that will serve as the smaller ~0.5–1 mm small artery/vein-like vessels. (Smaller < 30 µm microvessels

⁎ Corresponding authors at: Translational Biology and Engineering Program, Ted Rogers Centre for Heart Research University of Toronto, 661 University Ave, 14th floor, Toronto, Ontario, Canada M5G 1M1. E-mail addresses: ross.fi[email protected] (R.E. Fitzsimmons), [email protected] (C.A. Simmons).

https://doi.org/10.1016/j.bprint.2018.02.001 Received 17 August 2017; Received in revised form 13 January 2018; Accepted 13 February 2018 2405-8866/ © 2018 Elsevier B.V. All rights reserved.

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 1. 3D Bioprinter Hardware. CAD model and final custom-built 3D bioprinter outfitted with a motorized XYZ stage and five retractable extruder systems in which syringes containing bioinks and/or sacrificial matrices can be inserted (A, B). X/Y movements of the print bed are controlled by stepper motors which rotate driving pulleys outfitted with GT2 belts (C) that connect to pillow blocks on each of the four axis beams (D). These pillow blocks contain linear bearings which allow them to slide along the axis beams, and hence control the movement of the print bed by way of connector beams (E). Vertical movements of the print bed are controlled by three stepper motors within the stage itself (B, E). An inexpensive fixed-temperature heater was constructed using an insulted heater coil (F, G) that maintains syringe contents at ~37 °C (H; error bars = SD, n = 2).

containing a higher cell density, or even cell-only spheroids [3]. While the resolution and ability to print hydrogels with high cell densities make extrusion-based printers particularly attractive for tissue engineering, the cost of commercially available printers is prohibitive to the growth of the field (> $100,000). Less expensive models with fewer features still typically sell for ~$10,000, largely due to the components required to maintain high resolution (~5–10 µm) and precise deposition control (e.g., pneumatics). However, for applications like printing small artery/vein-sized channels (~0.5–1 mm diameter), such high resolution is not required. To that end, we describe here a lower cost ($3000), open-source printer design that can be used for lower resolution applications, like printing perfusable microvessels in tissue constructs. In order to create channels within hydrogel-based constructs, a number of methods have been previously employed. Molten

(arterioles, capillaries, venules), also necessary for engineering vascularized tissues, are most efficiently and effectively formed through the innate capacity of endothelial cells to undergo microvascular morphogenesis). Finally, a printer must have the capacity to deposit at least two materials (the sacrificial material and tissue material), although being able to print more than two materials is advantageous for generating complex heterogeneous tissues with multiple regions of differing cell/ hydrogel composition. Current bioprinters vary in their methods of deposition and their technical advantages, both of which influence the specific application and price point of the individual printers. For example, microextrusionbased printers (utilizing either pneumatics or mechanical dispensing methods) are more expensive than inkjet-based systems but can achieve a higher printing resolution under optimal conditions (5 µm to millimetres wide vs. ~50 µm wide inkjet droplets) and utilize bioinks 8

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

positional differences; since each pulley is coupled to a 400 step/revolution NEMA 17 axis motor, a single step can result in a displacement of 0.1 mm. Custom pillow blocks are located on each of the four axis beams (Fig. 1D) and contain a linear bearing, which allows it to slide along the axis beam. The two pillow blocks on the same type of axis beam are connected via a rigid beam (i.e. both pillow blocks on the Xaxis beam move in the X-direction and are connected via the X connector beam). These X and Y axis beams are positioned at different heights to ensure the connector beams do not intersect. Movement is conferred to the central block via each of the two connector beams, meaning that when the X pillow blocks are pulled in one direction, the X connector beam moves with them, and pushes the Z-stage central block along the X direction (Fig. 1E). Ultimately, this system allows for X/Y positional control of the Zstage which is mounted atop the central block and is controlled by three NEMA 14 synchronous linear stepper motors that have a single step displacement of 0.003 mm (Fig. 1B, E). Actuating the Z-stage linear stepper motors in unison will drive the print bed downwards or upwards, relative to the central block, which is stationary in Z. The print bed platform was designed to allow the print surface to be easily replaced with alternative printing surfaces, such as surfaces that provide temperature regulation. Additionally, a 3D printed cable chain to the Zstage was included to restrict the movement of the electrical wires to a single plane to prevent wires from becoming caught on moving features of the printer while preserving the range of X/Y motion of the central block.

carbohydrate glass (composed of sucrose, glucose, and dextran) has been used with a modified RepRap Mendel 3D printer to print sacrificial lattices that are dissolved away after they are embedded in hydrogel [4]. Given that the glass is printed in a molten state, hydrogel and embedded cells have to be added to surround the lattice after the printing process is complete in order to preserve cell viability, making this method a suitable option for creating simple homogeneous tissues but a poor option for creating more complex tissues where multiple hydrogel and cell types need to be printed in spatially distinct regions in 3D space. Alternate options exist for sacrificial materials that can be deposited alongside hydrogel-embedded cells during the printing process, such as gelatin and the triblock copolymer PEO (polyethylene oxide)-PPO (polypropylene oxide)-PEO, also known by its trade name of Pluronic F-127 (PF127) [5–7]. Gelatin, a thermoreversible biopolymer consisting of a range of hydrolyzed collagen segments of varying molecular weights, can be printed at a cell compatible temperature of ~37 °C and has already been extensively used in various biomedical applications [8]. The surfactant PF127, on the other hand, demonstrates inverse thermal gelation allowing it to be printed at ambient temperature and then removed at ~4 °C in order to create void vascular channels. PF127 is a surfactant, however, raising the concern of potential cytotoxicity effects on embedded cells. Hence, given the prospective options of gelatin and PF127 and the current lack of any side-by-side comparison of their printability, we aimed to investigate formulations of these two materials for their potential advantages and disadvantages relevant to their use as sacrificial materials in hydrogel-based tissue constructs. Additionally, for sacrificial material formulations found to have poor printability as a result of insufficient viscosity, we also explored using hyaluronan (HA) as a visco-modifying agent. By using our custom-built printer in order to assess the printability of these materials and assessing mechanical properties, we aimed to establish which may be the best option for creating branching vascular channels within engineered tissues.

2.2. Bioprinter design: extruding and heating systems To minimize movement of all five extruders, which otherwise would have required significant torque from the motors, the print heads were designed to be isolated from the XYZ movements executed by the lower part of the chassis. The extruders were also designed to perform rotational movements by way of servo motors to allow each extruder to be vertical while it is the active extruder and slanted backward while it is inactive. This was employed to prevent inadvertent disturbances to the print by inactive extruder tips and to help reduce cell pelleting within the syringe during long print times. The extruding systems (Fig. 1B, E) were designed to hold commercially-available sterile 10 mL syringes to allow for easy replacement, as opposed to using custom-made reservoirs that must be specially fabricated and repeatedly sterilized. Additionally, they allow for monitoring material volume via their transparent graduated wall and can connect to commercially-available Luer-locking tapered nozzle tips with a range of gauge sizes. Extruders were 3D printed with a Makerbot2 using ABS plastic, providing low-cost rigidity and easy customization. They also feature press-fit plates to allow for easy insertion and removal of syringes. A bipolar linear stepper motor is mounted atop each extruder which drives a piston into the syringe reservoir to extrude the desired material. For extruding materials that require heating, we opted for fabricating a custom-made fixed-temperature resistive heater composed of a heating coil wrapped around an insulated canister (Fig. 1F, G), opposed to purchasing a more expensive adjustable temperature heater. The heater was assembled by coiling a heating cable around two canisters, followed by a layer of insulation and electrical tape. The heating cable and insulation were obtained by dismantling a Sunbeam® Moist & Dry Heating Pad. One canister (cut from a 50 mL conical centrifuge tube) was used to surround the syringe barrel, while the other (cut from a 2 mL Eppendorf tube) was used to fit around the nozzle head. By adjusting the amount of insulation and the gap between the canister and syringe, the temperature inside the syringe was tuned to 37 °C to allow for the printing of gelatin. This custom heater could successfully maintain a temperature of 37 °C within the syringe barrel (Fig. 1H).

2. Materials and methods 2.1. Bioprinter design: chassis and XYZ positioning systems Prior to fabrication of parts, representational CAD models of the chassis, X/Y positioning system, and the extruders were established to serve as a guide for assembling of the printer (Fig. 1A). Design files and a full listing of components and their suppliers is found in Supplemental data. The completed printer (Fig. 1B) includes a 66 cm x 42 cm x 46 cm chassis that was laser cut from 0.25 in. clear acrylic sheets which provides rigidity and clear visualization of the internal components for prototyping purposes. Diaphragms were also included along the back wall of the chassis to help support the weight of the extruders. Several design elements of the chassis, such as interlocking sheets and the T-nut notches for press-fitting support nuts, were borrowed from the Fab@ Home printer design, due to its reported ease of assembly [9]. In the base of the chassis, a Cartesian coordinate system for controlling X/Y positioning of the print bed was employed. This Cartesian system consists of two structural support beams on each of the X and Y axes. In addition, each axis has a pair of driving rods which serve to drive axial movement; one actively turned by a NEMA 17 motor (Fig. 1C), and the other passively driven through a transmission system of pulleys and belts. Each belt of this pulley transmission system is tethered to a set of pillow blocks (Fig. 1D), one on each support beam, to convert rotational movement of the driving rods into linear motion of the Z-stage along the X/Y axes. In order to accomplish this X/Y positioning of the Z-stage, GT2 timing belts with rounded teeth were selected, as the rounded teeth of this belt profile allows for minimal slip when instantaneously reversing the driving direction. As well, 20-tooth driving pulleys were selected to match the GT2 belt profile with a 2 mm pitch. These components are critical in translating the rotational precision of the stepper motor to 9

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 2. Printer operational overview. A CAD model is first converted into G-code using a slicing software, such as Slic3r, which can then be uploaded to the printer interface software (e.g., Pronterface) which relays information to the Duet-Duex microcontroller. The microcontroller then coordinates the actuation of all the X/Y/Z stepper motors and the stepper motors which drive extrusion. The microcontroller also controls the rotational mechanism of the extruders by way of a Maestro USB servo controller. Positional feedback to the Duet-Duex is provided by limit switches on the central block facing in the X/Y/Z directions which signals to the microcontroller when they collide with the X and Y pillow blocks or the underside of the print bed. A power supply is used to power the Maestro, Duet-Duex microcontroller, and all motors.

the print, in case of any displacement. It also allows a print to be paused, adjusted, and resumed, without theoretical loss of the most recent positional information. Moreover, the machine is designed to run in an open-loop and beyond an initial calibration at print start-up no ongoing positional feedback is utilized, as positional information is maintained by the Duet-Duex.

2.3. Bioprinter design: microcontroller and operational overview An open-source Duet v0.6 controller board was selected as the printer controller system. Through pairing it with its attachment board, the Duex, this allowed us to utilize this system to control five axis motors (one in X, one in Y, and three in Z) and five independent extruders. This functionality enables printing up to five materials that can be used to create heterogeneous tissue structures. As summarized in Fig. 2, the Duet's RepRap firmware translates 3D models into functional movements of the extruders and axis motors. Prior to printing, the 3D model is subdivided into stacked planar sections by free “slicer” software, such as Slic3r. Additionally, the software plans a tool path and extrusion criteria for the extruders to physically recreate the 3D model, which can then be exported as a set of G-code commands to be executed by open-source printer interface software, such as Pronterface. The microcontroller then directly actuates all of the X/Y/Z stepper motors necessary to generate the programmed tool path of the print bed through three spatial dimensions. Likewise, it also directly communicates with the servo motors that drive the syringe reservoir pistons of each extruder. However, in order to actuate the servo motors that control the rotation of the extruders, the microcontroller first relays signals to a six-channel Maestro USB servo controller which then communicates with the five motors that regulate the angle of each extruder. Using the Maestro allows the rotation of the extruders to be controlled using only one analog input from the Duet-Duex. A power supply (400 W, 12 V with 33.3 A max, 5 V with 2 A max) is used to power the Maestro, Duet-Duex microcontroller, and all the actuators. In terms of positional feedback, the central block has been equipped with limit switches facing in the X/Y/Z directions, which trigger a signal to the Duet controller when the switch collides with the X and Y pillow blocks, or the Z-stage. This provides the printer with the ability to recognize the limits of its space, and align itself at the beginning of

2.4. Sacrificial material preparation Gelatin solutions were prepared by dissolving porcine type A gelatin (Sigma) in PBS+/+ by heating to 80 °C with repeated agitation. For gelatin-hyaluronan (HA) solutions, dissolved gelatin was combined (1:1) and thoroughly mixed with HA (~1.5–1.8 × 106 Da, Sigma) dissolved in PBS+/+ to produce the desired concentrations of both materials. Pluronic F-127 (Sigma) was dissolved in ice cold PBS+/+ and incubated on a rocker at 4 °C until fully dissolved. For gelatin-alginate tissue ink, dissolved gelatin (in PBS-/-) was combined (1:1) and mixed with dissolved sodium alginate (Sigma) to produce the desired concentrations of both materials. 2.5. Printer testing and filament diameter measurements Examining the basic functionality of the printer was performed by printing water droplets in a defined pattern using each extruder system (design made in Solidworks and converted to G-code by Slic3r) followed by measuring the average distance between the centers of the droplets in the X and Y directions followed by comparing the mean distances to the pre-defined CAD model distances using one-way ANOVA. Graphpad Prism (v7) was used throughout this work for statistical analysis. Proper continuous deposition of material and printing in the Z-direction was confirmed by printing a ‘UofT’ design and simple 30-layer high shapes (cube, cylinder, and star) using 35% PF127. A 10

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

at 80,000 cells per well the day prior to exposure to PF127 (and Triton X-100, which served as a control). Materials were dissolved in PBS+/+ and added to wells at final concentrations ranging from 0% to 10% (w/ v) while maintaining a constant volume of PBS and media for each. Following 24 h of culture, media was removed and replaced with fresh media containing 1 mg/mL methylthiazolydiphenyl-tetrazolium bromide (MTT). After 4 h of incubation with MTT, media was replaced with 1 mL DMSO to dissolve the resulting formazan crystals. Next, 0.1 mL of the DMSO solution was transferred to a 96-well plate and the absorbance was read at 570 nm with a Tecan Sunrise plate reader. The relative cell viability of each condition was determined by normalizing with untreated cells. To examine the viability of HUVEC seeded within hydrogel-based channels, channels were formed by depositing a line of 35% (w/v) PF127 atop a fibrin hydrogel between two affixed needles of a silicone culture chamber, followed by polymerizing another layer of fibrin on top. Fibrin gels were formed by polymerizing 10 mg/mL fibrinogen (Sigma; dissolved in EGM-2 without serum) with 1 unit thrombin (Sigma) per 1 mL fibrinogen. Once the top layer of fibrin had gelled, 4 °C PBS was perfused through the needles to wash away the PF127 in order to form a perfusable linear channel. HUVEC were then seeded in the channels at 7 × 106 cells mL−1 and cultured in EGM-2 under static conditions for 12 h before assessing viability using a LIVE/DEAD staining kit (Molecular Probes). To stain, PBS was first perfused through the channel followed by a 0.5 h incubation with EGM-2 (without serum) containing 0.5 μL calcein AM and 3 μL ethidium homodimer-1 per 1 mL media. Following this incubation, cells were washed once more with PBS and imaged using an inverted fluorescent microscope (Olympus IX71).

print speed of 1 mm/s and a 25-gauge extruder tip was used for all experiments throughout this work. To confirm the printer could pattern gels embedded with cells while maintaining viability, adipose-derived stromal cells (ASC; purchased from Lonza) were suspended in gelatin methacrylate (GelMA) solution at 2 × 105 cells mL−1 and then printed using a rectilinear line pattern design followed by curing with an Omnicure S2000 UV light source (Lumen Dynamics) at 7.6 mW/cm2 for 90 s. GelMA solution was prepared by dissolving 10% (w/v) gelatin methacrylate (BioBots) and 0.5% (w/v) Igracure 2959 photoinitiator (Ciba Specialty Chemicals) in PBS at 80 °C, followed by cooling to 37 °C prior to adding cells. Once GelMA constructs were UV cured, constructs were cultured in ASC growth media (see Section 2.8 for media components) for 24 h followed by assessment of viability using a LIVE/DEAD staining kit (Molecular Probes). To stain, constructs were washed 3 × with PBS followed by a 0.5 h incubation with ASC media (without serum) containing 0.5 μL calcein AM and 3 μL ethidium homodimer-1 per 1 mL media. Following this incubation, constructs were washed 3 × with PBS and imaged using an inverted fluorescent microscope (Olympus IX-71). The number of live and dead cells per field of view were determined using the ‘Find Maxima’ function of ImageJ. To assess the accuracy and precision of the printer, 35% PF127 was used to print rectilinear line patterns consisting of nine line segments with varying distances set between the lines in the X-axis (CAD model distances: 0.6825, 0.7625, 1.0475, 1.4275, 3.97 mm). Once these designs were printed, the distance between the centers of each line segment were measured in ImageJ and the average distance was compared to the CAD model using t-tests. Percent error was determined as (Abs (CAD model distance – measured line distance)/CAD model distance) x 100%. Percent coefficient of variation was determined as (standard deviation of the measured distance/mean measured distance) x 100%. For filament diameter measurements, rectilinear line segments were printed using each material of interest and the width of each of the nine line segments was measured in ImageJ.

2.9. Printing channels in hydrogel constructs To form hierarchical channels in a simple planar vascular design using 35% PF127 and gelatin-HA (5.25%, 3% w/v), first a CAD model of the desired vascular design was made in Solidworks. Using the desired sacrificial material, this design was then printed atop a ~ 4 mm thick block of alginate-gelatin (4%, 10% w/v) that was formed within a silicone mold at room temperature, followed by encapsulating the sacrificial material with a ~4 mm top layer of alginate-gelatin. The construct was then submerged in a 100 mM CaCl2 bath for ~1 h to crosslink the alginate. Once crosslinked, needles were inserted into either side of the vascular design, and the construct was then heated to 37 °C or chilled to 4 °C, followed by flushing the nascent channels with heated or chilled PBS to remove the gelatin-HA or PF127, respectively.

2.6. Rheometry The viscosities of gelatin (5% and 10% w/v), gelatin-HA (5%, 3% w/v), and PF127 (35% w/v) were analyzed in a shear rate range of 1–100 Hz. During analysis, samples were maintained at the temperature at which they are printed (i.e., 37 °C for gelatin-containing materials and 22 °C for PF127). Measurements were made with a AR2000 rheometer (TA Instruments) under steady state conditions using a 2° 40 mm diameter cone with a 2 min equilibrium step prior to initializing measurements to ensure a correct temperature.

3. Results and discussion 2.7. Compression testing 3.1. Printer testing Materials were prepared for compression testing by using polypropylene molds to produce gel disks ~18 mm high and 12 mm in diameter. Testing was performed using a Test Resources 840 L series mechanical tester (Shakopee, MN, USA) with a 13 mm diameter tester head and a 1.5 kg load cell (WF5). Stress vs. strain curves were generated for each material from 0% to 50% strain at a rate of 0.6% strain per second and the compressive moduli (E) were determined by calculating the tangential slope at a constant stress value of 1 kPa.

To confirm basic printer functionality, we assessed the ability of the printer to create a series of water droplets on the print bed using each of its five extruders (Fig. 3A). To quantitatively compare the prints to the programmed design, the average distance in both the X and Y directions measured between the droplet centers was then compared to the distance present in the CAD model (Fig. 3B). No significant differences were found by one-way ANOVA for any of the extruders implying proper functioning of both the X/Y control systems and the extruding systems of the printer at the length scale tested. Continuous deposition of material (35% w/v PF127) on the print bed in X and Y directions using more complex tool paths was also possible, as demonstrated with the printing of a “U of T” design (Fig. 3C). To confirm proper functioning of the Z-stage, we also successfully printed a number of simple shapes (a cube, a cylinder and a star) 30 layers high using PF127 (Fig. 3D-F). Finally, we confirmed the capacity of the bioprinter to extrude hydrogel-suspended cells in defined patterns while maintaining satisfactory cell viability using the commonly used biomaterial gelatin

2.8. Cell culture and viability assays Human umbilical vein endothelial cells (HUVEC) and adipose-derived stromal cells (ASC) were both purchased from Lonza. HUVEC were cultured in EGM-2 (Lonza) while ASC were cultured in AlphaMEM with 10% FBS, 2 mM additional L-glutamine, 100 U/mL penicillin and 100 μg/mL streptomycin (all Gibco products). Both cell types were used at passage five and trypsinized as per supplier's instructions. For viability assays, HUVEC and ASC were seeded in 24-well plates 11

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 3. Basic functionality testing. The functionality of each extruder and the X/Y control system was tested by depositing a series of water droplets in a defined pattern and comparing the spacing between the centers of the droplets to the CAD model; no significant differences between the print and the model were found using one-way ANOVA (A, B; error bars = SD, n = 4). The printer also performed controlled continuous deposition of material (35% w/v PF127 with blue dye) in X and Y directions using a more complex tool path, as demonstrated with the printing of a “U of T” design (C). The basic functionality of the Z-stage was tested by printing multi-layered structures of various shapes (a cube, cylinder, and star, each 30 layers high) composed of PF127 (D-F). Gelatin methacrylate (GelMA) was used to confirm the capacity of the printer to extrude hydrogel-suspended cells in defined patterns while maintaining cell viability. Using a 25-gauge extruder tip, rectilinear line patterns of GelMA (10% w/v) could be printed (G) and the viability of GelMA-encapsulated adipose-derived stromal cells (ASCs) within the resulting constructs was assessed by LIVE/DEAD staining 24 h after printing and UV curing (H).

analyzed by LIVE/DEAD staining (Fig. 3H). In order to evaluate the precision and accuracy of the printer (repeatability and fidelity to a specified CAD model, respectively), a number of PF127 rectilinear line patterns were printed with varying spacing between the lines to facilitate assessment of the printer's capabilities at multiple length scales (Fig. 4A). The average distance between the midpoints of the lines for each print type was then compared

methacrylate (GelMA). Using a 25-gauge syringe tip, rectilinear line patterns of GelMA (10% w/v) could successfully be produced (Fig. 3G) with a mean filament width of 1.136 mm and a coefficient of variation (CV) of 42%. To assess viability, adipose-derived stromal cells (ASCs) were printed within GelMA followed by UV curing to form solid gel constructs. Favorably, after 24 h of culture within printed GelMA constructs encapsulated ASCs exhibited 83 ± 7.7% (SD) viability when 12

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 4. Fine resolution testing. PF127 rectilinear line patterns deposited on the print bed as per a CAD model with the specified spacing between lines (A). Quantification of the distances between the midpoints of each line segment compared to the CAD model (B; error bars = SD, n = 2; significant differences determined using t-tests; * = p < 0.05, ** = p < 0.01). Accuracy conveyed as mean differences from CAD model vs. CAD model line spacing, and then normalized to CAD model spacing as absolute percent error to adjust for the relative size of the print design (C, D; error bars = SD). Precision conveyed as standard deviation of the measured spacing (i.e., error bar values in B) vs. CAD model line spacing (E), and normalized to the mean measured spacing as percent coefficient of variation (F).

could potentially adapt their CAD model by scaling it up or down by the appropriate amount to adjust for this error to better recapitulate the desired construct dimensions. With regard to precision, as the spacing between lines decreased from 3.97 mm to 0.683 mm the standard deviation of the measured spacings increased from < 0.01 mm to 0.035 mm (Fig. 4E). This was also apparent when data was normalized to the size of the design as the coefficient of variation increased from 0.23% to 5.68% as the line spacing decreased from 3.97 mm to 0.683 mm, with a prominent increase just below 1 mm (Fig. 4F). This result could potentially be attributed to increased sensitivity to random variation in the printing process at smaller length scales, such as minor vibrations affecting the extruder tip. For comparison, higher end printers (> $100,000) such as RegenHU's Biofactory and EnvisionTEC's 3D Bioplotter have precisions

to the respective CAD model (with line spacings varying from 0.6825 to 3.97 mm). A trend was noted that the larger the designed width, the larger its difference from the CAD model, with significant differences found for the designs with lines spacing of 0.7625 mm and upwards (p < 0.05 using t-tests) (Fig. 4B). While differences to the CAD model were noted to generally increase from − 0.069 mm to + 0.193 mm as the size of the design increased (Fig. 4B, C), when normalized for the size of the design as an absolute percent error, however, a trend of decreasing error was noted from ~12% to ~5% as line spacing increased from < 1 mm to > 1.4 mm (Fig. 4D). This was expected since differences from the CAD model become increasingly negligible as the size of the print increases. Regardless of print size (though likely of greater concern at smaller scales given the higher percent error), if this error is deemed unsatisfactory for a particular application, the user

13

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 5. Sacrificial material filament measurements. Deposited rectilinear line patterns of PF127 (P), gelatin (G), and gelatin-hyaluronan (G+H; HA at 3% w/v) resulting from a single-pass of an extruder outfitted with a 25-gauge tip at a print speed of 1 mm/s (A). Quantification of the widths of the deposited materials from two trials (triangles and squares) at the specified w/v concentrations (B). Each point represents a measurement of one of the nine segments composing a rectilinear line pattern, while lines indicate the mean width of the deposited filaments for each condition. Note that the 0 mm measurements of one of the trials using 10% (w/v) gelatin was the result of clogging in the extruder tip resulting in no material being deposited.

which forms a solid gel, we reasoned that increasing the viscosity of gelatin without increasing its ability to undergo gelation may be of benefit. Hence, we explored the option of adding the visco-modifying agent, hyaluronan (HA), to gelatin to increase viscosity, as an alternative to increasing gelatin concentration which had also promoted gelation. Favorably, HA has been used extensively as an injectable biomaterial owing to its biocompatibility and shear-thinning properties [10,11]. In assessing numerous formulations of gelatin-HA, an optimal formulation of 5.25% (w/v) gelatin with 3% (w/v) HA produced filaments with the lowest width of all gelatin-only and gelatin-HA formulations with a mean width of 0.75 mm and a CV of 7%. Similar to gelatin alone, gelatin-HA formulations with higher concentrations of 7.5% and 10% (w/v) gelatin resulted in poorer uniformity of filament thickness (CVs of 32% and 19%, respectively), likely as a result of intermittent clogging due to premature gelation in the extruder tip prior to deposition. Hence, overall, in cases where vascular channels need to be printed with high uniformity and at the smallest diameter possible, PF127 would be the most appropriate choice, followed by Gel-HA (5.25%, 3% w/v).

of 5 µm and 1 µm, respectively, while lower end printers (~$10,000) such as the Biobot 1 and Seraph Robotics’ Scientist 3D Printer have precisions of 10 µm and 5–10 µm (as reported by company websites). Favorably, at its most precise (feature spacing of > 1 mm), our printer has a similar precision of ~10 µm, whereas if feature spacing is < 1 mm precision declines. 3.2. Printability and filament diameter measurements Next, we assessed the printability of sacrificial materials, in terms of width and uniformity of the deposited filament, as both qualities are critical in discerning the suitability of a material for consistently creating vascular designs of suitable dimensions to mimic sub-millimeter arteries and veins. To do so, we printed rectilinear line patterns with a 25-gauge extruder tip and measured the thickness of the lines produced by a number of material formulations (Fig. 5A, B). PF127 (35% w/v) was found to have a mean width of 0.4136 mm and a coefficient of variation (CV) of 16.2%. In contrast, gelatin (2.5–7.5% w/ v) filaments were wider (0.98–1.1 mm mean widths) and more variable (44.3–53.3% CV). We attribute this poor control over filament width to insufficient viscosity of the gelatin solution at these lower concentrations. At a higher concentration of 10% (w/v), deposited gelatin filaments favorably had reduced variability in their width with a CV of 12.9% and a mean width of 0.69 mm, but also tended to transiently obstruct the extruder tip, on occasion preventing printing. The variability in gelatin filament thickness and the issues with transient clogging noted here may have contributed to the less uniform appearance of vascular channels in prior work using gelatin compared those produced using PF127 [5–7]. Motivated by the otherwise favorable characteristics of gelatin (i.e., biocompatible, thermoreversible gelation within cell-compatible temperatures), we explored whether the printability of gelatin could be improved upon. Given that PF127 does not suffer clogging issues due to it being of a paste-like consistency at its maximal solidity, unlike gelatin

3.3. Rheometry of sacrificial materials To investigate the association between filament diameter/uniformity and viscosity we measured the viscosity of several sacrificial materials (Fig. 6). As expected, 5% (w/v) and 10% (w/v) gelatin, which had larger filament widths and greater variability, also had significantly lower viscosities compared to 35% (w/v) PF127 and Gelatin-HA (5%, 3% w/v) (p < 0.05 for viscosities measured at 1–31.6 Hz, Tukey tests following two-way ANOVA). Measured at 37 °C from 1 to 100 Hz, gelatin had viscosities with a range of 0.018–0.007 Pa·s for 5% (w/v) gelatin, and 0.032–0.033 Pa·s for 10% (w/v) gelatin. Such values are comparable to previous viscosity measurements of ~0.01 Pa·s from 1 to 12 Hz at a similar concentration of 7% (w/v) gelatin [12]. As anticipated, given their reduced filament widths and reduced variability, 14

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

PF127 and gelatin-HA had higher viscosities from 1 to 100 Hz with a range of 619.4–9.8 Pa·s for 35% PF127, and 196.3–3.7 Pa·s for gelatinHA (5%, 3% w/v). Of note, PF127 also had significantly higher viscosities than gelatin-HA from 1 to 17.8 Hz (p < 0.05, Tukey tests), which may help explain why it has a comparatively smaller filament width. This range of values for gelatin-HA seem consistent with previous measurements in that gelatin-HA (7%, 1% w/v) with a lower HA content has been found to have a range of ~8–0.8 Pa·s from 1 to 100 Hz which is of intermediate viscosity between our own measurements of pure gelatin and gelatin-HA (5%, 3% w/v) [12]. Also similar to our own findings, Restylane SubQ (a partially crosslinked 2% (w/v) HA-based dermal filler) has been measured to have a viscosity of 198.4 Pa·s at 0.7 Hz, which is similar to our own value of 196.3 Pa·s at 1 Hz for gelatin-HA (5%, 3% w/v) [13,14]. Also comparably, non-crosslinked 2.5% HA (w/v) (an HA concentration slightly lower than what we measured) was found to have a viscosity of ~45 Pa·s at 1 Hz [15]. Notably, pronounced shear thinning behavior was observed for both PF127 and gelatin-HA; this combination of pronounced shear thinning and high viscosity appears ideal for a printable material as it allows for extrusion and yet filament diameter maintains relatively constant compared to lower viscosity materials. Overall, these results suggest uniform filaments can be generated if the selected material has a sufficient viscosity of ~200–4 Pa·s from 1 to 100 Hz, and that materials with a viscosity of ~0.01 Pa·s from 1 to 100 Hz are prone to poor

Fig. 6. Rheometry of sacrificial materials. Viscosity vs. shear rate of 35% (w/v) PF127, gelatin-HA (5%, 3% w/v), 10% (w/v) gelatin and 5% (w/v) gelatin at the temperatures at which the materials are printed (error bars = SD, n = 2). Measurements were made with a AR2000 rheometer under steady conditions using a 2° 40 mm diameter cone. Tukey tests following a two-way ANOVA indicated no significant (NS) differences between 5% (w/v) and 10% (w/v) gelatin at any shear rate. * = NS difference between gelatin-HA and PF127 at specified shear rates, ** = NS difference between gelatin-HA and either gelatinonly material, § = NS difference among any of the four materials at specified shear rates. Where not otherwise indicated, data points at the same shear rates are significant (p < 0.05).

Fig. 7. Compression testing of sacrificial materials. Compressive moduli (determined at a stress value of 1 kPa from data collected from 0% to 50% strain) and yield strains of the specified sacrificial materials (A, B; error bars = SD, n = 4). Representative examples of stress-strain curves from 0% to 50% strain (C). Asterisks indicate significant differences determined using Tukey tests. * = p < 0.05, **** = p < 0.0001.

15

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 8. Toxicity testing of pluronic F-127. MTT viability assays of adipose-derived stromal cells (ASC) (A), and human umbilical vein endothelial cells (HUVEC) (B), after exposure to varying concentrations of PF127 and Triton X-100 for 24 h. Triton X-100 was included as a control for the assay. All data is normalized as a percentage of the absorbance at 570 nm of the untreated condition (error bars = SD, n = 3). Asterisks indicate a significant difference from the 0% condition by Dunnett's multiple comparison test. To assess the viability of HUVEC seeded within PF127-formed channels, a line of PF127 was deposited atop a fibrin gel between two needles followed by polymerizing a second layer of fibrin on top. After washing away the PF127 with 4 °C PBS to form a fibrin-based channel (C), HUVEC were then seeded and their viability assessed by LIVE/DEAD staining 12 h later (D). **** = p < 0.0001.

a comparable difference between 5% (w/w) gelatin and 15% (w/w) gelatin (concentrations similar to 5% (w/v) and 15% (w/v) based on our own tests), with moduli of ~13 kPa and 81 kPa, respectively [16]. Hence, overall, PF127 would be superior to gelatin-containing materials for supporting construct overhangs during the printing process due to its relatively higher compressive modulus, and if gelatin is to be used, using a higher concentration would be advantageous for this application. However, it also bares mentioning that the capacity of gelatin to solidify into an elastic gel may be favorable over the paste-like consistency of PF127 in some instances. In scenarios where external and/or internal structural reinforcement of nascent tissue constructs is of importance during handling or manipulation (e.g., transferring into crosslinking solution after printing), gelatin would likely be superior to PF127 by merit of its capacity to resist plastic deformation in its gel phase. Being of a paste-like consistency at ambient temperature, PF127 exhibited plastic deformation during compression with a yield strain of ~1.7–2.5% and a yield stress of 1.71 ± 0.3 kPa (SD), unlike gelatin and gelatin-HA which did not exhibit plastic deformation over the same strain range tested of 0–50% (Fig. 7B, C).

uniformity during printing and should be avoided if stringent control of filament diameter is of importance for the application. In the future, the filament uniformity of prospective materials with viscosities between these values will have to be further investigated to better establish the lower viscosity threshold requisite to produce the desired uniformity. 3.4. Compression testing of sacrificial materials Next, we measured the compressive modulus of several sacrificial materials to investigate their suitability for serving as external support material for tissue constructs with overhangs that require mechanical support during printing. A higher compressive modulus of the support material would be beneficial to better avoid outright collapse of features, or even to avoid subtle mechanical creeping of the construct during fabrication which would increasingly interfere with accurate printing as the tissue construct compresses under its own weight. To assess the compressive modulus of the various sacrificial materials, gels were compressed from 0% to 50% strain and the compressive moduli (E) were determined at a constant stress value of 1 kPa (Fig. 7A). We found that the compressive modulus of 35% (w/v) PF127 (102.2 ± 20.76 kPa (SD)) was significantly higher using Tukey tests than that of 5.25% (w/v) gelatin (9.79 ± 0.14 kPa; p < 0.0001), gelatin-HA (5.25%, 3% w/v) (11.81 ± 0.81 kPa; p < 0.0001), and 20% (w/v) gelatin (73.24 ± 9.63 kPa; p < 0.05). Additionally, 20% (w/v) gelatin was found to have a greater modulus than both 5.25% (w/v) gelatin and gelatin-HA (5.25%, 3% w/v) (both p < 0.0001, Tukey tests). Corroborating our own findings, prior measurements have noted

3.5. Toxicity testing of pluronic F-127 While gelatin is of biological origin and is composed of hydrolyzed peptide chains of predominantly type I collagen, PF127 is synthesized to contain a central hydrophobic polypropylene oxide block and two hydrophilic polyethylene oxide blocks which together impart surfactant 16

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Fig. 9. Fabricating perfusable channels. Strategy for fabricating channels within hydrogel-based constructs using sacrificial materials (A) and preliminary testing of 3D printing hierarchical vascular structures on the print bed using 35% (w/v) PF127 and Gelatin-HA (5.25%, 3% w/v) (B, C). 3D printing of an example hydrogel construct composed of a tissue material, gelatin-alginate, using a heated extruder (D, E). Demonstration of perfusion (before and after) with red dye in PBS after crosslinking constructs in a 100 mM calcium chloride bath and expelling 35% (w/v) PF127 and Gelatin-HA (5.25%, 3% w/v) from alginate-gelatin hydrogel constructs using either 4 °C or 37 °C PBS, respectively (F; G, macroscopic view of construct with channels made using Gelatin-HA).

low as 0.01% (w/v) (p < 0.0001 by Dunnett's multiple comparison test), ASCs treated with PF127 did not show a significant decrease in viability even at concentrations as high as 10% (w/v) (Fig. 8A). Likewise, when we next assessed PF127 toxicity at 5% and 10% (w/v) using HUVEC no decrease in viability was detected (Fig. 8B). Additionally, to confirm low toxicity in the context of seeding cells within hydrogelbased channels formed using PF127, we formed channels using 35% (w/v) PF127 within the commonly used biomaterial, fibrin (Fig. 8C). After washing away the PF127 with 4 °C PBS, fibrin-based channels were then seeded with HUVEC at 7 × 106 cells mL−1. Favorably, HUVEC within the channels were viable when imaged 12 h after cell seeding as indicated by LIVE/DEAD staining (Fig. 8D). These results for both cell types imply PF127 is indeed a reasonably satisfactory material with regard to cytotoxicity. That said, in addition to the vasculaturerelated cell types tested here, it may be prudent to also assess the toxicity of PF127 on the specific parenchymal cell types to be used in any future applications. Moreover, further investigation into the other

properties as a result of an amphiphilic structure. This characteristic of PF127 has been exploited for its miscibility-enhancing abilities in a number of industrial applications and pharmaceutical delivery systems [17,18]. Hence, while PF127 successfully produced the smallest filament width and with the least variability, one possible disadvantage of PF127 is that it is a surfactant and residual material remaining within channels after removal could potentially have cytotoxic effects. In order to address this main concern of PF127, we used an MTT (methylthiazolydiphenyl-tetrazolium bromide) viability assay to assess its toxicity on endothelial cells (ECs) and adipose-derived stromal cells (ASCs), which are commonly used as a vascular-supporting cell type to maintain EC viability and encourage capillary outgrowth in engineered tissues [19,20]. ASCs were exposed to both PF127 and the surfactant Triton X-100 (used as a control) for 24 h at 0.01–10% (w/v); 10% was selected as a maximum since PF127 undergoes a sol-to-gel transition at concentrations higher than ~15% (w/v) at 37 °C [21]. While Triton X100 resulted in a significant reduction in viability at concentrations as

17

Bioprinting 9 (2018) 7–18

R.E. Fitzsimmons et al.

Toronto) for their expertise and technical assistance with rheometry. This project is financially supported by Canadian Institutes of Health Research (CIHR) operating grants. RF was financially supported by a CIHR (MOP-102721, RMF-111624 and MOP-130481) Banting and Best Doctoral Scholarship. The authors report no conflicts of interests.

subtler effects of PF127 on cellular activity would also be of merit; for example, PF127 has been shown to affect the activity of lipoprotein lipases [22]. Further expanding our understanding of the bioactivity of PF127 with specific cell types would help resolve if any risks of using PF127 for a particular application may outweigh its many benefits as a sacrificial material.

Appendix A. Supporting information 3.6. Fabricating perfusable channels Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bprint.2018.02.001.

Lastly, we demonstrated the methods by which the most promising sacrificial materials, PF127 (35% w/v) and gelatin-HA (5.25%, 3% w/ v), could be used by the bioprinter to create hydrogel constructs with perfusable hierarchical channels (Fig. 9A). A CAD model of a simple vascular structure was designed in Solidworks and the printing of this pattern was then tested with both materials on the bare print bed. Consistent with the satisfactory printing fidelity both materials showed in our earlier material optimization testing with rectilinear line patterns, both materials could successfully print the desired vasculature design (Fig. 9B, C). After this preliminary test, both sacrificial materials were printed on hydrogels of a tissue-scaffold material composed of gelatin-alginate, followed by embedding in another layer of gelatinalginate (gelatin provides immediate gelation allowing for printing of 3D shapes (Fig. 9D, E), while alginate provides permanent gelation at 37 °C once crosslinked). After submerging the construct in a calcium chloride bath (100 mM) to polymerize the alginate, the sacrificial materials were expelled from the constructs by perfusing either 4 °C or 37 °C PBS for PF127 and gelatin-HA, respectively. As seen with the perfused red dye, both sacrificial materials could successfully create perfusable channels within the hydrogel constructs (Fig. 9F, G). While the approach implemented in this work to generate planar vascular designs may be useful for relatively thin constructs (e.g., skin and other epithelial tissues), future work should be focused on optimizing the conditions necessary to create vertical channels in order to create perfusable constructs of thicker dimensions. By further adapting these methods for heterogeneous vascular structures with varying diameters of vessels, it is anticipated that more complex and physiologically-similar vasculature could be formed within engineered tissues to permit media and/or blood perfusion.

References [1] V. Mironov, R.P. Visconti, V. Kasyanov, G. Forgacs, C.J. Drake, R.R. Markwald, Organ printing: tissue spheroids as building blocks, Biomaterials 30 (2009) 2164–2174. [2] R.P. Visconti, V. Kasyanov, C. Gentile, J. Zhang, R.R. Markwald, V. Mironov, Towards organ printing: engineering an intra-organ branched vascular tree, Exp. Opin. Biol. Ther. 10 (2010) 409–420. [3] S.V. Murphy, A. Atala, 3D bioprinting of tissues and organs, Nat. Biotechnol. 32 (2014) 773–785. [4] J.S. Miller, K.R. Stevens, M.T. Yang, B.M. Baker, D.-H.T. Nguyen, D.M. Cohen, et al., Rapid casting of patterned vascular networks for perfusable engineered three-dimensional tissues, Nat. Mater. 11 (2012) 768–774. [5] V.K. Lee, D.Y. Kim, H. Ngo, Y. Lee, L. Seo, S.-S. Yoo, et al., Creating perfused functional vascular channels using 3D bio-printing technology, Biomaterials 35 (2014) 8092–8102. [6] D.B. Kolesky, R.L. Truby, A.S. Gladman, T.A. Busbee, K.A. Homan, J.A. Lewis, 3D bioprinting of vascularized, heterogeneous cell-laden tissue constructs, Adv. Mater. 26 (2014) 3124–3130. [7] D.B. Kolesky, K.A. Homan, M.A. Skylar-Scott, J.A. Lewis, Three-dimensional bioprinting of thick vascularized tissues, Proc. Natl. Acad. Sci. 113 (2016) 3179–3184. [8] K. Su, C. Wang, Recent advances in the use of gelatin in biomedical research, Biotechnol. Lett. 37 (2015) 2139–2145. [9] E. Malone, H. Lipson, Fab@Home: the personal desktop fabricator kit, Rapid Prototyp. J. 13 (2007) 245–255. [10] J. Necas, L. Bartosikova, P. Brauner, J. Kolar, Hyaluronic acid (hyaluronan): a review, Vet. Med. 53 (2008) 397–411. [11] A. Fakhari, C. Berkland, Applications and emerging trends of hyaluronic acid in tissue engineering, as a dermal filler and in osteoarthritis treatment, Acta Biomater. 9 (2013) 7081–7092. [12] J. Picard, S. Giraudier, V. Larreta-Garde, Controlled remodeling of a protein-polysaccharide mixed gel: examples of gelatin-hyaluronic acid mixtures, Soft Matter 5 (2009) 4198–4205. [13] H. Sundaram, B. Voigts, K. Beer, M. Meland, Comparison of the rheological properties of viscosity and elasticity in two categories of soft tissue fillers: calcium hydroxylapatite and hyaluronic acid, Dermatol. Surg. 36 (2010) 1859–1865. [14] A. Verpaele, A. Strand, Restylane SubQ, a non-animal stabilized hyaluronic acid gel for soft tissue augmentation of the mid- and lower face, Aesthet. Surg. J. 26 (2006) S10–S17. [15] R.J. Wendling, A.M. Christensen, A.D. Quast, S.K. Atzet, B.K. Mann, Effect of carboxymethylation on the rheological properties of hyaluronan, PLoS One 11 (2016) e0162849. [16] A. Karimi, M. Navidbakhsh, Material properties in unconfined compression of gelatin hydrogel for skin tissue engineering applications, Biomed. Eng. -Biomed. Technol. 59 (2014) 479–486. [17] M. Fahad, M.A. Khan, M. Gilbert, Evaluation of thermal gelation of F-127 in a nonaqueous solvent and its suitability as a support material for additive manufacturing, Adv. Mater. Res. 911 (2014) 226–231. [18] J.J. Escobar-Chavez, M. Lopez-Cervantes, A. Naik, Y.N. Kalia, D. QuintanarGuerrero, A. Ganem-Quintanar, Applications of thermoreversible pluronic F-127 gels in pharmaceutical formulations, J. Pharm. Pharm. Sci. 9 (2006) 339–358. [19] S. Merfeld-Clauss, N. Gollahalli, K.L. March, D.O. Traktuev, Adipose tissue progenitor cells directly interact with endothelial cells to induce vascular network formation, Tissue Eng. Part A. 16 (2010) 2953–2966. [20] S. Kachgal, A.J. Putnam, Mesenchymal stem cells from adipose and bone marrow promote angiogenesis via distinct cytokine and protease expression mechanisms, Angiogenesis 14 (2011) 47–59. [21] H. Geng, H. Song, J. Qi, D. Cui, Sustained release of VEGF from PLGA nanoparticles embedded thermo-sensitive hydrogel in full-thickness porcine bladder acellular matrix, Nanoscale Res. Lett. 6 (2011) 312. [22] T.P. Johnston, W.K. Palmer, Mechanism of poloxamer 407-induced hypertriglyceridemia in the rat, Biochem. Pharmacol. 46 (1993) 1037–1042.

4. Conclusions In conclusion, we found that PF127 is generally superior to gelatin as a sacrificial material for creating vascularized tissues by merit of its filament uniformity during printing and its greater compressive modulus; that said, we also noted that the uniformity of gelatin could be improved upon through the addition of hyaluronan and that gelatinbased sacrificial materials have a higher compressive yield strain that PF127. We anticipate that the experimental framework used here may be useful in assessing additives for improving upon these materials and exploring potential alternatives in the future. We expect that this work on assessing sacrificial materials and establishing an economical 3D bioprinter will assist the ongoing development of vascularized tissues and will help accelerate the widespread adoption 3D bioprinting to create such engineered tissues. Acknowledgements The authors gratefully acknowledge Prof. David James and Mr. Philip Hoang (Mechanical and Industrial Engineering, University of

18