3D Bioprinting Human Chondrocytes with NanocelluloseâAlginate

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3D Bioprinting Human Chondrocytes with NanocelluloseAlginate Bioink for Cartilage Tissue Engineering Applications Kajsa Markstedt, Athanasios Mantas, Ivan Tournier, Héctor Martínez Ávila, Daniel Hägg, and Paul Gatenholm Biomacromolecules, Just Accepted Manuscript • Publication Date (Web): 25 Mar 2015 Downloaded from http://pubs.acs.org on March 27, 2015

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3D Bioprinting Human Chondrocytes with Nanocellulose-Alginate Bioink for Cartilage Tissue Engineering Applications †‡ Kajsa Markstedt, ‡Athanasios Mantas, ‡Ivan Tournier, ‡Héctor Martínez Ávila, ‡Daniel Hägg, † ‡*Paul Gatenholm †Wallenberg Wood Science Center and ‡ Biopolymer Technology, Department of Chemical and Biological Engineering, Chalmers University of Technology, Gothenburg 412 96 KEYWORDS 3D Bioprinting, Cartilage Tissue Engineering, Nanocellulose, Alginate

ABSTRACT The introduction of 3D bioprinting is expected to revolutionize the field of tissue engineering and regenerative medicine. The 3D bioprinter is able to dispense materials while moving in X, Y and Z directions; enabling the engineering of complex structures from the bottom up. In this study a bioink that combines the outstanding shear thinning properties of nanofibrillated cellulose (NFC) with the fast crosslinking ability of alginate was formulated for the 3D bioprinting of living soft tissue with cells. Printability was evaluated with concern to

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Chalmers, Department of Chemical and Biological Engineering, Biopolymer Technology, SE-

412 96 Göteborg, Sweden. Phone: +46317723407. E-mail: [email protected]

ACS Paragon Plus Environment

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printer parameters and shape fidelity. The shear thinning behavior of the tested bioinks enabled printing of both 2D gridlike structures as well as 3D constructs. Furthermore, anatomically shaped cartilage structures, such as a human ear and sheep meniscus, were 3D printed using MRI and CT images as blueprints. Human chondrocytes bioprinted in the non-cytotoxic, nanocellulose-based bioink exhibited a cell viability of 73% and 86% after 1 and 7 days of 3D culture, respectively. Based on these results we can conclude that the nanocellulose-based bioink is a suitable hydrogel for 3D bioprinting with living cells. This study demonstrates the potential use of nanocellulose for 3D bioprinting of living tissues and organs.

INTRODUCTION 3D Bioprinting is an emerging technology expected to revolutionize the field of tissue engineering and regenerative medicine. The 3D bioprinters’ ability to deposit biomaterials with micrometer precision1 in ”cell-friendly conditions” gives it an advantage over other scaffold production technologies which are limited2 with regard to the control of the 3D structure as well as cell distribution. In addition to tissue engineering, the bioprinting technology can also be applied in the development of cell-based sensors3, drug screening models1, 4 and tumor models5. When bioprinting a scaffold and cells simultaneously, the most challenging aspect is the bioinks’ properties. The challenge lies in meeting both the biological requirements and the requirements needed for good printability, the biofabrication window6-7, which often oppose each other. The choice of bioink also determines which bioprinting technology can be used, i.e. dispense the material as drops by using inkjet8; microvalve9 and laser assisted10 techniques or by continuous dispensing, such as microextrusion1. Hydrogels have been pointed out as attractive materials for bioinks6, 11 because they in general are biocompatible, show low cytotoxicity, and their high water content gives them a structural


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similarity to extracellular matrix12-13. Common hydrogels evaluated for 3D bioprinting are the natural polymers collagen14, hyaluronic acid (HA)15, chitosan16 and alginate7. To be suitable for 3D bioprinting, a hydrogel must be viscous enough to keep its shape during printing and must have crosslinking abilities allowing for it to retain the 3D structure after printing. Crosslinking can occur by temperature change14, UV photopolymerization17-19 and ionic crosslinking. Alginates isolated from brown algae have the ability to crosslink at the addition of divalent cations20-21. A common challenge when 3D bioprinting hydrogels is that the printed shapes tend to collapse due to low viscosity. Although the viscosity of alginate can be increased by varying the concentration and molecular weight22 it has not been sufficient for achieving shape fidelity while printing. To increase the structural fidelity hydrogels are often printed in combination with other materials. In such cases the printability of alginate has been improved by the addition of gelatin11, or by printing with a supporting sacrificial polymer23. We and others have previously combined cellulose nanofibrils with alginate to prepare sponges used for adipose tissue24 and for cell encapsulation25. Cellulose nanofibrils, fibrils with one dimension in the nanometer range have gained increased interest as a biomaterial during the last decade. Depending on the cellulose source and its processing conditions cellulose nanofibrils are divided into three main categories; microfibrillated cellulose, nanocrystalline cellulose and bacterial nanocellulose26. Similar for all cellulose nanofibrils are their general properties of cellulose; hydrophilicity and broad chemical modification capacity, combined with properties specific for nanoscale materials due to their high surface area. Cellulose nanofibrils are attractive for biomedical applications due to their good mechanical properties and biocompatibility27. Bacterial nanocellulose fibrils are microscopically similar to collagen fibrils. That is, their width is about 100 nm27, making them similar in size to collagen fibrils28. This similarity makes it


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interesting to use bacterial nanocellulose as a component in scaffolds for soft tissue such as cartilage29-30. Cartilage is an avascular tissue consisting of a small number of chondrocytes (10% to 15%) with limited self-regenerative properties31. The demand for replacing damaged cartilage tissue with alternative approaches such as tissue engineering is therefore high. The aim of this study was to 3D bioprint human chondrocytes using a bioink consisting of cellulose nanofibrils and alginate.

EXPERIMENTAL SECTION MATERIALS Nanofibrillated cellulose (NFC) produced by mechanical refinement and enzymatic treatment as described by Pääkkö et. al32 was provided by Innventia AB (Sweden). Material characterization of the NFC was performed with regard to dry content, sugar composition, fibril size and charge density. The NFC had a dry content of 1.9% with a sugar composition of 95.3% Glucose, 1.7% Mannose and 0.7% Xylose, determined by a Dionex ICS-3000 (USA) system equipped with a CarboPac PA1 (4 × 250 mm). Freeze dried and sterile alginate, SLG100, with a molecular weight of 150-250 kDa and above 60% of α-1-guluronic acid (G) was purchased from FMC Biopolymers (Norway). SLG100 was dissolved to a concentration of 2.5% (w/v) in 4.6% (w/v) D-mannitol aqueous solution. DMannitol (Sigma Aldrich, Sweden) was used to retain the osmolarity33. 90 mM aqueous solution of CaCl2 (Sigma Aldrich, Sweden) acted as the crosslinking solution for the bioink. All Dmannitol and CaCl2 solutions were sterilized by filtering with a 0.1 µm Millex®W filter unit (Merck Millipore). METHODS


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ATOMIC FORCE MICROSCOPY (AFM) Diluted NFC dispersions (0.15 % dry content) were suspended on a glass plate as droplets of 3 mm in diameter and dried under room temperature at 0% relative humidity for 24 hours. The dried films were examined by AFM using a Digital Instrument Dimension 3000 with a type G scanner. The measurement was performed in tapping mode with a standard silicon tip to determine the width of the nanocellulose fibrils. CHARGE DENSITY The charge density of the fibrils was measured as described by Fall et.al34 in fibrillated and centrifuged dispersions by polyelectrolyte titration (PET)35 and the streaming potential was measured by a Stabino particle charge titration analyzer (Microtrac). A NFC dispersion of 7.65 mg in 20 mL of mQ-water was titrated with a polydiallyl dimethyl ammonium chloride (pDADMAC) solution (Mw = 147 kDa, 0.95 µeq/L). At the position where the potential changes sign from minus to plus the added equivalents of pDADMAC have cancelled out all the charges on the fibrils. Thus, the amount of added polymer at this point provides the charge of the fibrils in eq./g. BIOINK PREPARATION The dry content of NFC dispersion was increased from 1.9% to approximately 2.5% by centrifugation (JOUAN CR 3i multifunction, Thermo Scientific) and removal of excess supernatant. The concentrated NFC was mixed intensely and steam sterilized (100 kPa, 121 °C for 20 min) in an autoclave (Varioklav Steam Sterilizer 135T, Thermo Scientific) followed by addition of a 2.5% alginate solution, as described in materials. Four formulations of bioinks were prepared with different proportions of NFC:Alginate; 90:10, 80:20, 70:30, 60:40 as seen in Table 1. All bioinks were stored in sterile conditions at 4 - 8 °C. Table 1. Summary of ink formulations with varied proportions of NFC and Alginate (SLG100). Bioink Formula

NFC: Alginate (w/w)

NFC (% w/v)

Alginate (% w/v)

Water Content (% w/v)

Ink9010

90:10

2.25

0.25

97.50

Ink8020

80:20

2.00

0.50

97.50


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Ink7030

70:30

1.75

0.75

97.50

Ink6040

60:40

1.50

1.00

97.50

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RHEOLOGY The rheological properties of the bioink and its main components were analysed using the Discovery HR-2 rheometer (TA Instruments, UK) with a peltier plate. Three different setups of rheological measurements were conducted depending on the material being measured: Set-up 1 - Alginate, Set-up 2 - pure NFC and bioink, Set-up 3 - crosslinking of bioink. All measurements were performed at 25 °C and the samples were allowed to reach equilibrium temperature for 60 s prior to each measurement. For Set-up 1 a cone-plate (40 mm, 1.99°) was used and the shear viscosity was measured at shear rates from 0.01s-1 to 1000 s-1.An aluminium plate-plate (20 mm, gap = 300 µm) was used for Set-up 2. Oscillation amplitude sweeps for 0.1 1000 Pa at a frequency of 1 Hz were performed to define the linear viscoelastic region (LVR). From the LVR a stress of 10 Pa was chosen for the oscillation frequency measurements conducted at a frequency range of 10-3 – 103 Hz. Shear viscosity was evaluated by increasing the shear rate from 0.1 to 1000 s-1 at 25 °C. For Set-up 3, the change in moduli while crosslinking the bioink was measured with a serrated plate-plate (8 mm, gap = 1400 µm). Based on oscillation amplitude sweeps for 0.008 – 10 % strain at a frequency of 1 Hz a strain of 0.1% was found to be within the LVR for both non-gelled and gelled bioink. The oscillation frequency measurements were conducted at 0.1 % strain and at a frequency of 1 Hz during 30 minutes. One minute after the measuring was started; 1 ml of 90 mM CaCl2 was dispensed around the measured bioink causing it to gel while simultaneously measuring the storage and loss modulus. 3D BIOPRINTING The bioink was printed using the 3D bioprinter 3D Discovery® from regenHU (Switzerland) seen in Figure 1. The printer head consists of a micro valve, based on electromagnetic jet technology for accurate jetting or contact dispensing, with a 300 µm diameter


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nozzle. The flow rate was controlled by adjusting the dispensing pressure (20-60 kPa), the valve opening time (400-1200 µs), and the dosing distance (0.05- 0.07 mm). The width of the printed lines was controlled by adjusting the parameters mentioned above and the printing speed (10-20 mm/s). Prior to filling the cartridge for each printing, the bioink was stirred with a spatula to reduce settling. Five shapes were 3D printed directly onto glass slides: a small grid (7.2 × 7.2 mm, line spacing 1.2 mm, 6 layers), a large grid (38.5 × 17.7 mm, 6 horizontal lines with 2.5 mm spacing, 10 vertical lines with 3.5 mm spacing, 1 layer), a solid disc (8 mm diameter, 1.5 mm high, line space 0.5 mm, 5 layers) a human ear (22 mm on the long axis) and a sheep meniscus (18 mm on the long axis). These dimensions are those from the CAD models. The dimensions of the 3D printed structures were measured before and after crosslinking the bioink.

Figure 1. 3D Bioprinter 3D Discovery® (Switzerland)


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The grids and solid disc were designed in the software BioCAD (regenHU) to generate the process protocol for the 3D bioprinter. For evaluating the resulting line width, large grids (n=10 per bioink) were printed with three alginate solutions (2%, 3% and 4%) and Ink9010. The grids using a dosing distance of 0.07 mm, valve opening time of 1200 µs and printing speed of 20 mm/s. The line width was measured for each grid at 10 different points using the software ImageJ. More complex 3D structures, such as an ear and a sheep meniscus, were printed after converting Stereolithography (STL) files into G-code; the process protocol to be used by the 3D Discovery®. The STL-file of the ear was obtained as described by Nimeskern et. al 36. The STLfile of the sheep meniscus was obtained from a CT-scan using the manual segmentation and 3D reconstruction tools of the software Mimics (Materialise, Belgium). UNCONFINED COMPRESSION Compression testing was performed on casted discs prepared from the four ink formulations seen in Table 1. Using a positive displacement pipette (Microman®, Gilson, USA) and casting units (Q-Gel bio, Lausanne, Switzerland), discs were casted by dispensing 75 µl of bioink onto a casting plate and covering the dispensed drop with a top plate. The casting units with the discs (8 mm × 1.5 mm) were crosslinked in a bath of 90mM CaCl2 (Sigma–Aldrich) for 10 minutes and then immersed in HyClone Basal Medium Eagle (Sigma–Aldrich) for 48 hours at room temperature. After crosslinking and equilibration in medium, the discs had a diameter of 8 mm and a thickness of 1.2 mm in average. Initial dimensions of the discs were measured with a digital caliper. A universal testing machine (Instron Model 5565A, UK) equipped with a 100 N load cell was used for the unconfined compression test, which was performed in wet conditions at room temperature. The strain rate was set to 10%/s and the samples were compressed until 40% compressive strain. Bluehill®


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software (Instron) was used to calculate the compressive stress and strain, compressive stiffness (tangent modulus) at 30% strain and compressive stress at 30% strain for all samples (n=6 per group). CYTOTOXICITY In vitro cytotoxicity testing was performed according to ISO 109935:2009(E), annex C (MTT cytotoxicity test). L929 fibroblasts (ATCC # CCL-1) were cultured in MEM supplemented with 10% FBS and 1% antibiotics/antimycotics and 4 mM L-glutamine (all from Invitrogen). On day 1, solid discs (⌀8 mm) of Ink9010 were printed (30 kPa, printing speed 5 mm/s, dosing distance 0.07 mm, valve opening time 1200 µs) and crosslinked in 90 mM CaCl2 for 10 minutes. Liquid extracts of sterile Ink9010 discs (n=6) were prepared by incubating each sample in 2 ml of complete medium for 24 h under shaking conditions (80 rpm) at 37°C and 5% CO2 under aseptic conditions using sterile cell culture plates. As a positive control, 20% DMSO in complete medium was used and as a negative control, sterile Eppendorf tube lids were placed in 2 ml complete medium. The controls were treated the same as the samples. Ten thousand cells per well were seeded in 96 well plates. On day 2, 100 µl extracts from the samples and controls were added to the wells. As a base line for the test, 100 µl of complete medium was used. At day 3, the medium was removed and replaced with 50 µl of fresh medium and 10 µl of MTS solution (CellTiter 96® Aqueous One Solution Cell Proliferation Assay from Promega). After two hours, absorbance was read at 490 nm using an Epoch plate reader from BioTek. Cell viability was normalized to cells exposed to blank medium and a value above 70 % was considered without cytotoxic potential. 3D BIOPRINTING WITH CHONDROCYTES Human nasoseptal chondrocytes (hNC) were a kind gift from Professor Nicole Rotter at University Medical Center Ulm in Germany. Cells at passage 0 were expanded in culture medium (DMEM/Ham’s F-12 (1:1) medium


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supplemented with 10% FBS and 1% pen/strep; Biochrom). After six days of expansion, the cells were trypsinized and counted. Initial cell viability was determined using trypan-blue staining and subsequently hNCs were encapsulated in Ink8020 at a final concentration of 15×106 cells/ml. hNC-laden constructs (4.8 mm × 4.8 mm, line spacing 1.2 mm, layer thickness 0.1 mm, 10 layers) were 3D bioprinted (40 kPa, printing speed5 mm/s, dosing distance 0.05 mm, valve opening time 450 µs) with approximately 3.15×105 cells per construct. After printing, the constructs were crosslinked in 90 mM CaCl2 solution for 10 minutes and rinsed twice in culture medium. As a control for the bioprinting process, non-printed, cell-laden discs (30 µl) were casted as described above and crosslinked as the printed constructs. 3D culture was conducted in standard conditions (37 °C, 5% CO2 and 95% relative humidity), whereat the culture medium was changed three times per week. Non-printed cell-laden hydrogels (n=3) were analyzed after 1 day of culture to examine the effect of the printing process on cell viability. 3D bioprinted cell-laden constructs (n=6) were also analyzed after day 1 and 7 of culture. CELL VIABILITY PRE- AND POST-BIOPRINTING LIVE/DEAD® Cell Imaging Kit (Molecular probes, Life technologies) was utilized to differentiate between live and dead cells during the experiments. Samples were washed in serum-free medium and then incubated in the staining solution for 1 h. After staining, the samples were washed in serum-free medium for 1 h, whereat the medium was changed twice. All incubations were done in standard culture conditions. Images were acquired using an inverted microscope (IX73, Olympus) with FITC and Texas Red filters, digital color camera (XC10) and cellSens imaging software (Olympus). To assess the cell viability in deeper layers of the construct, the software function “instant extended focal imaging” was used, while manually focusing through a depth of 300 µm, to combine the


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in-focus details into a single image. For each time point, five randomly selected images, taken at 10× magnification, were used to evaluate the percentage of cell viability; which was calculated as the average ratio of live to total cells.

RESULTS AND DISCUSSION MATERIAL CHARACTERISATION Atomic Force Microscopy (AFM) imaging of air dried NFC showed that the fibrils had a width of around 20 nm as seen in Figure 2.

Figure 2. AFM of air dried NFC. The width of the nanocellulose fibrils was measured to be around 20 nm. The average calculated value of the NFC charge density was 24 µeq/g. This is a very low charge compared to e.g. carboxymethylated NFC where charges of 515 µeq/g have been reported37 and is presumably appropriate for the mixing with cells. PRINTABILITY Alginates low zero-shear viscosity, seen by the flow curve in Figure 3A, gives pure alginate inks a poor shape fidelity when printing. This was confirmed by measurements of the linewidth in Figure 3B where 2% SLG100 has the highest line width (1.07 mm) when printing compared to 3% SLG100 and 4% SLG100.


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To improve the shape fidelity of the alginate, NFC was used as a main component in the bioink formulation. Compared to alginate, the flow curve of a 2.5 % NFC dispersion in Figure 3A shows that NFC on its own has a high viscosity at low shear rates and is shear thinning. The high viscosity gives excellent shape fidelity when printed; however, the printed shape is destroyed when mechanical force is used as seen in Figure 3, C2.

Figure 3. (A) Flow curves of 2.5% NFC (-), Ink9010 () and alginate solutions: SLG100 4 % (); - SLG100 3 % (); SLG100 2 % (). (B) Line width measurements of 3D printed large grids with alginate inks compared to Ink9010. The photos below the graph show the printed grids and their different line resolutions. (C) Small grid printed with (C1) 3% alginate and (C2) 2.5% NFC. (C3) Small grid of printed and crosslinked Ink9010. Comparing the small grid printed with 3% alginate with the grid printed with 2.5% NFC in Pictures C1 and C2 in Figure 3, illustrates how the low viscosity of alginate reduced the printing resolution and how NFC has been printed nicely but was not gelled. By combining alginate and NFC a bioink with the rheological properties of NFC and the crosslinking ability of alginate was formulated. The successful combination of these two materials is shown by picture C3 in Figure 3: the high printing resolution gives the small grid nicely visible lines and the spatula is able to lift the print since it is gelled by ionic crosslinking of alginate.


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CHARACTERIZATION OF BIOINKS Four bioinks were formulated and studied to find the most suitable proportions of NFC:Alginate with regard to rheological properties, shape fidelity and mechanical strength. Rheological studies of the different bioinks (Ink6040, Ink7030, Ink8020 and Ink9010) showed that all bioinks were shear thinning with viscosity flow curves very similar to the curves of pure NFC in Figure 4A. This is in good agreement with the shear thinning properties previously observed for nanocelluloses32,38. The flow curves of the four different bioinks are all very similar indicating that for the tested ratios the viscosity is highly dependent on the dry content of the bioink, rather than the proportions of NFC and Alginate. Thus, the crosslinking properties can be controlled by varying the proportion of alginate to NFC without influencing the viscosity and hence the printability. The measurements of frequency oscillation in Figure 4B shows that the storage modulus G' and loss modulus G'' increased for higher proportions of NFC in the bioink formulation. The tan δ values, Figure 4C, calculated from G’/G’’ were used to evaluate how gel-like the bioinks were. All bioinks had tan δ values below 1 at the measured frequencies indicating that the inks were more gel-like than liquid.


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Figure 4. Rheological data for the different bioink formulations: - Ink9010; - Ink8020; Ink7030; - Ink6040. (A) Flow curves of the different bioink formulations. All bioinks show very similar viscosities and shear thinning properties. Ink6040 shows a slightly lower viscosity indicating that the alginate starts to have an influence on the viscosity. (B) Storage modulus, G’(closed symbols) and loss modulus, G’’ (open symbols) of the bioink formulations as a function of frequency. (C) Tan δ of the bioink formulations. (D) Storage modulus G’ and (E) loss modulus G’’ measured over 30 minutes where 90 mM CaCl2 solution was added 1 minute after starting the measurement. The gelling behavior of the bioinks has been studied by measuring the loss modulus and storage modulus over time while crosslinking with CaCl2. The gelling curves for Ink7030 and Ink8020 in Figure 4D and Figure 4E show a fast increase in modulus at t=1 min and gradually becomes linear. The gelling curve of Ink9010 follows the trend of the other inks at the start but decreases after 5 minutes indicating that the gelled structure is being disrupted. Ink6040 shows a


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lower modulus than what is expected and does not follow the trend of increased modulus with increasing alginate concentration. Casted discs were evaluated to determine their strength after crosslinking. The unconfined compression test revealed the mechanical properties of all groups. The stress strain curves of all groups can be seen in Figure 5A. Ink6040 shows a higher mechanical strength than Ink9010, while Ink7030 is closer to Ink6040 as expected. Ink8020 displays a more linear stress-strain behavior than all other groups. Since almost all groups revealed a non-linear stress-strain behavior as being viscoelastic materials, the compressive stiffness (tangent modulus) and compressive stress at 30% strain are important for the mechanical characterization. Regarding the compressive stiffness, shown in Figure 5B, Ink9010 shows a noticeably lower value than Ink7030, which could be a consequence of Ink9010 having the lowest alginate concentration. Ink6040 shows a lower value of compressive stiffness than Ink7030 similar to the results from the rheology in Figure 4D & E. This indicates that a high alginate concentration impairs the mechanical properties. Regarding the compressive stress, Figure 5C, all groups show a similar behavior where Ink9010 displays the lowest value as expected.


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Figure 5. (A) Stress-strain curves of casted discs from - Ink9010; - Ink8020; - Ink7030; - Ink6040 under unconfined compression test. (B) Compressive stiffness at 30 % strain and (C) compressive stress at 30 % strain of casted discs from the four bioink formulations. (D) Photographs of 3D printed solid discs from the four bioink formulations (1-4) before and (5-8) after crosslinking. (1 & 5) Ink6040, (2 & 6) Ink7030, (3 & 7) Ink8020 and (4 & 8) Ink9010. The diameter of the discs was measured before and after crosslinking. To evaluate the change in shape and size after crosslinking printed solid discs, the four bioink formulations were studied directly after printing and after 10 minutes of crosslinking. All discs were printed with the same process protocol and the variety in size before crosslinking (9-10 mm) is mainly due to the properties of the different inks. A large change in the shape is seen in Figure 5D for Ink6040 and Ink7030 compared to the other bioinks. The diameter of discs printed with Ink6040 and Ink7030 decreases by 1 mm compared to the initial size before crosslinking while at the same time swells at the center as seen in pictures 5 and 6 taken from the side in Figure 5D. Printed discs of Ink8020 and Ink9010 do not show the same change in shape after


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crosslinking; the diameter only decreases 0.5 mm and the swelling is barely noticeable when examining the disc from the side in pictures 7 and 8 in Figure 5D. Taking all results into account; rheology, compression and shape deformation, Ink8020 would be the optimal ink formulation. It shows both a good gelling while crosslinking and shows little tendency of shape deformation. 3D BIOPRINTING BIOINK Due to the high shape fidelity of the formulated ink, the whole print with all its layers could be dispensed before crosslinking the structure with CaCl2. After 10 seconds the effect of crosslinking the printed structure could be seen. The printing process showing a crosslinked small grid acting as a stiff gel is depicted in Figure 6A-C.

Figure 6. (A) 3D printed small grids (7.2 × 7.2mm) with Ink8020 after crosslinking. (B) The shape of the grid deforms while squeezing and (C) it is restored after squeezing. (D) 3D printed human ear and (E & F) sheep meniscus with Ink8020. Side view (E) and top view (F) of meniscus.


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Once the printing of 2D structures was controlled, larger constructs were printed. Shapes resembling cartilage tissues such as an ear, and a meniscus were successfully printed, Figure 6DF. Even with printing times of up to 20 minutes for these larger constructs, the prints did not collapse or lose their shape during the printing process due to the viscosity of the ink. CYTOTOXICITY ANALYSIS To investigate any potential harmful effects of the nanocellulose-based bioink on cell viability, an indirect cytotoxicity test was performed in accordance with ISO standard 10993-5:2009(E). Ink9010 was chosen since it contained the highest amount of NFC, and previous research has shown that alginate is compatible with cell culture39. The cytotoxicity analysis showed no indications of potential harmful effects of the bioink. There was no significant difference between the negative control (105.7 ± 6.3%) and Ink9010 (105.6 ± 2.3%). Moreover, the positive control (20 % DMSO) resulted in complete cell death (3.1 ± 0.8%). Since cell viability was well above 70%, the cytotoxic potential of the evaluated bioink was classified as non-cytotoxic. This suggests that this bioink has excellent potential as a hydrogel biomaterial for tissue engineering and 3D bioprinting. 3D BIOPRINTING WITH CHONDROCYTES As a proof of concept, human nasoseptal chondrocytes were used for 3D bioprinting of gridded constructs. Based on the evaluation of the four ink ratios with regard to material properties, printability and gelling, Ink8020 was chosen to test the influence of this bioink on cellular responses in a short-term 3D culture. Viability of the hNCs was analyzed prior to embedding in the bioink by using the trypan-blue exclusion method. This test revealed an initial cell viability of 95.3 ± 0.1% (Figure 7A). Furthermore, cell viability in the bioink constructs was analyzed before and after printing using a Live/Dead cell-imaging assay, which is used to differentiate between live and dead cells. Live cells emit green fluorescence in the cytoplasm, whereas nuclei of dead cells emit red fluorescence. Analysis using


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independent t-test determined a significantly higher viability in hNCs before embedding (95.3 ± 0.1%) compared to after embedding in bioink and crosslinking the constructs (69.9 ± 13.3%, before printing group) at the p< 0.05 level. No significant changes in cell viability were observed in the printed constructs (72.8 ± 6.0%, after printing group) compared to the non-printed group after 1 day of culture (Figure 7A-C) indicating that the preparation and mixing was what caused a decrease in viability. Moreover, a significant increase in cell viability was detected in the printed constructs after 7 days of 3D culture (85.7 ± 1.9%) compared to day 1. A homogeneous cell distribution was observed in the non-printed and 3D bioprinted constructs, indicating a successful mixing of the cells in the bioink. The decrease in cell viability, after embedding the hNCs and crosslinking the constructs, could be attributed to the shear forces applied to the cells during mixing with the bioink or to the long crosslinking process. Although there was a loss of cells after the embedding and bioprinting processes, the cells were able to recover after seven days of culture, as shown by the higher cell viability compared to day 1 (Figures 7D and E). The dead cells were most likely flushed out of the bioink matrix, since fewer dead cells were observed at day 7. These results combined with the cytotoxicity analysis clearly demonstrate that this bioink is a biocompatible hydrogel well suited for 3D bioprinting of living cells. In this work small constructs were 3D bioprinted with cells. However, the upcoming challenge lies in the bioprinting of large anatomical-sized constructs that maintain a high cell viability. To overcome this challenge, large constructs could be 3D printed with an open inner structure to improve the supply of nutrients to embedded cells.


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A

B

⃰

100

Viability of hNCs (%)

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⃰

80 60 40

C

20 0

Before printing Before Before embedding printing Before embedding

D

After printing (day 1)

After printing (day 7)

After printing (day 1) After printing (day 7)

E

Figure 7. (A) Viability of human nasoseptal chondrocytes (hNC) before and after 3D bioprinting. Representative images showing dead (red) and live (green) cells (B) before and (C) after bioprinting hNCs in Ink8020 and 3D culture for one day. (D and E) Representative images (at 4× and 10× magnifications) showing dead and live cells in 3D bioprinted constructs after seven days of culture. CONCLUSIONS This study reports on a novel bioink composed of nanocellulose and alginate. Several structures representing cartilage tissue were 3D printed with high fidelity and stability. The rheological properties of the bioink provided excellent printability at room temperature and low pressure - essential conditions when 3D bioprinting with living cells. Before crosslinking, the properties of NFC were dominating the bioink making it shear thinning with high shape fidelity and increased printing resolution as compared to 3D printing with alginate. After crosslinking,


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the properties of alginate were dominating with increased storage modulus correlating with higher amounts of alginate in the bioink. When rheology, compression testing and shape fidelity were weighted together, the bioink Ink8020 had the optimal composition for 3D bioprinting. However, different applications may require different mechanical properties of the material, and the different compositions of bioink can therefore have different usages. The results from the indirect cytotoxicity test and cell viability analysis of the 3D printed bioink with hNCs showed that the tested bioink is biocompatible and a suitable material for cell culture. Our results demonstrate the potential use of nanocellulose for 3D bioprinting of living tissues and organs, particularly for diagnostics. We conclude that a bioink composed of nanofibrillated cellulose and alginate is a suitable hydrogel for 3D bioprinting with living cells for growth of cartilage tissue.

AUTHOR INFORMATION Author Contributions The manuscript was written through contributions of all authors. Development of bioinks, rheology and crosslinking studies have been carried out by K.M. A.M. has performed the mechanical testing. Evaluation of printer parameters and line width measurements has been carried out by I.T. and A.M. The cell studies have been carried out by D.H. and H.M.A. All authors have given approval to the final version of the manuscript.

ACKNOWLEDGMENT We would like to acknowledge Åsa Blademo and Tom Lindström at Innventia AB (Stockholm, Sweden) for providing nanofibrillated cellulose used in this study. Also we would like to thank Oruc Koklukaya, Fibre and Polymer Technology, KTH Royal Institute of


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Technology for help during the charge density measurements. We acknowledge the Centre for Cellular Imaging at the Sahlgrenska Academy, University of Gothenburg for the use of the microscope equipment. We appreciate the kind gift of nasoseptal chondrocytes from Professor Nicole Rotter at the Department of Otorhinolaryngology, University Medical Center Ulm, Germany. EU program Eureka and Vinnova, Sweden are greatly acknowledged for financial support of the Project E!8355 CELLINK. The Knut and Alice Wallenberg Foundation is gratefully acknowledged for sponsoring the Wallenberg Wood Science Center.

ABBREVIATIONS NFC, nanofibrillated cellulose; pDADMAC, polydiallyl dimethyl ammonium chloride; LVR, linear viscoelastic region; G’, Storage Modulus; G’’, Loss Modulus; STL, Stereolithography

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TABLE OF CONTENTS GRAPHIC


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3D Bioprinting Human Chondrocytes with NanocelluloseâAlginate

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