Article pubs.acs.org/Biomac
3D Bioprinting Human Chondrocytes with Nanocellulose−Alginate Bioink for Cartilage Tissue Engineering Applications Kajsa Markstedt,†,‡ Athanasios Mantas,‡ Ivan Tournier,‡ Héctor Martínez Á vila,‡ Daniel Hag̈ g,‡ and Paul Gatenholm*,†,‡ †
Wallenberg Wood Science Center and ‡Biopolymer Technology, Department of Chemical and Biological Engineering, Chalmers University of Technology, Gothenburg 412 96, Sweden 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, which enables 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 cross-linking ability of alginate was formulated for the 3D bioprinting of living soft tissue with cells. Printability was evaluated with concern to 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 noncytotoxic, nanocellulose-based bioink exhibited a cell viability of 73% and 86% after 1 and 7 days of 3D culture, respectively. On the basis of 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.
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INTRODUCTION Three-dimensional (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 sensors,3 drug screening models,1,4 and tumor models.5 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 window,6,7 which often oppose each other. The choice of bioink also determines which bioprinting technology can be used, that is, dispense the material as drops by using inkjet,8 microvalve,9 and laserassisted10 techniques or by continuous dispensing such as microextrusion.1 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 similarity to extracellular matrix.12,13 Common hydrogels evaluated for 3D bioprinting are the natural polymers collagen,14 hyaluronic acid (HA),15 chitosan,16 and alginate.7 To be suitable for 3D bioprinting, a hydrogel must be viscous enough to keep its shape during printing and must have cross© 2015 American Chemical Society
linking abilities allowing for it to retain the 3D structure after printing. Cross-linking can occur by temperature change,14 UV photopolymerization,17−19 and ionic cross-linking. Alginates isolated from brown algae have the ability to cross-link at the addition of divalent cations.20,21 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 weight,22 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 polymer.23 We and others have previously combined cellulose nanofibrils with alginate to prepare sponges used for adipose tissue24 and for cell encapsulation.25 Cellulose nanofibrils, fibrils with one dimension in the nanometer range, have gained increased interest as a biomaterial during the past decade. Depending on the cellulose source and its processing conditions, cellulose nanofibrils are divided into three main categories: microfibrillated cellulose, nanocrystalline cellulose, and bacterial nanocellulose.26 Similar for all cellulose nanofibrils are their general properties of cellulose: hydrophilicity and broad chemical modification capacity, combined with properties Received: December 2, 2014 Revised: March 24, 2015 Published: March 25, 2015 1489
DOI: 10.1021/acs.biomac.5b00188 Biomacromolecules 2015, 16, 1489−1496
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Biomacromolecules specific for nanoscale materials due to their high surface area. Cellulose nanofibrils are attractive for biomedical applications due to their good mechanical properties and biocompatibility.27 Bacterial nanocellulose fibrils are microscopically similar to collagen fibrils. That is, their width is about 100 nm,27 which makes them similar in size to collagen fibrils.28 This similarity makes it interesting to use bacterial nanocellulose as a component in scaffolds for soft tissue such as cartilage.29,30 Cartilage is an avascular tissue consisting of a small number of chondrocytes (10−15%) with limited self-regenerative properties.31 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.
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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 Ink8020 Ink7030 Ink6040
90:10 80:20 70:30 60:40
2.25 2.00 1.75 1.50
0.25 0.50 0.75 1.00
97.50 97.50 97.50 97.50
mm, 1.99°) was used, and the shear viscosity was measured at shear rates from 0.01−1000 s−1. An aluminum plate−plate (20 mm, gap = 300 μm) was used for Set-up 2. Oscillation amplitude sweeps from 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 cross-linking the bioink was measured with a serrated plate−plate (8 mm, gap = 1400 μm). On the basis of 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 nongelled and gelled bioink. The oscillation frequency measurements were conducted at 0.1% strain and at a frequency of 1 Hz for 30 min. 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
EXPERIMENTAL SECTION
Materials. Nanofibrillated cellulose (NFC), produced as described by Päak̈ kö et al.32 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 mm2). 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. D-Mannitol (Sigma-Aldrich, Sweden) was used to retain the osmolarity.33 A 90 mM aqueous solution of CaCl2 (Sigma-Aldrich, Sweden) acted as the cross-linking solution for the bioink. All Dmannitol and CaCl2 solutions were sterilized by filtering with a 0.1 μm MillexW filter unit (Merck Millipore). Methods. 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 h. 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 al.34 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 dimethylammonium 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 canceled 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 the Materials section. Four formulations of bioinks were prepared with different proportions of NFC/alginate: 90:10, 80:20, 70:30, and 60:40 as seen in Table 1. All bioinks were stored in sterile conditions at 4−8 °C. Rheology. The rheological properties of the bioink and its main components were analyzed 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, cross-linking 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
Figure 1. 3D Bioprinter 3D Discovery (Switzerland). head consists of a micro valve, based on electromagnetic jet technology for accurate jetting or contact dispensing, with a 300 μm diameter 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 1490
DOI: 10.1021/acs.biomac.5b00188 Biomacromolecules 2015, 16, 1489−1496
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Biomacromolecules shapes were 3D printed directly onto glass slides: a small grid (7.2 × 7.2 mm2, line spacing 1.2 mm, 6 layers), a large grid (38.5 × 17.7 mm2, 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 cross-linking the bioink. 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 used 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 Gcode, the process protocol to be used by the 3D Discovery. The STLfile of the ear was obtained as described by Nimeskern et al.36 The STL-file 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. By using a positive displacement pipet (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 cross-linked in a bath of 90 mM CaCl2 (Sigma− Aldrich) for 10 min and then immersed in HyClone Basal Medium Eagle (Sigma−Aldrich) for 48 h at room temperature. After crosslinking and equilibration in medium, the discs had a diameter of 8 mm and a thickness of 1.2 mm on 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 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 10993−5:2009(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 cross-linked in 90 mM CaCl2 for 10 min. 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 baseline 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 2 h, 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 supplemented with 10% FBS and 1% pen/strep; Biochrom). After 6 days of expansion, the cells were trypsinized and counted.
Initial cell viability was determined using trypan-blue staining, and subsequently hNC 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 speed 5 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 cross-linked in 90 mM CaCl2 solution for 10 min and rinsed twice in culture medium. As a control for the bioprinting process, nonprinted, cell-laden discs (30 μL) were casted as described above and cross-linked 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. Nonprinted 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 focused through a depth of 300 μm, to combine the 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.
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RESULTS AND DISCUSSION Material Characterization. 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, for example, 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 3, panel A, gives pure alginate inks a poor shape fidelity when printing. This was confirmed by measure1491
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Figure 3. (A) Flow curves of 2.5% NFC (-), Ink9010 (•), and alginate solutions: SLG100 4% (×); SLG100 3% (∗); and 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 cross-linked Ink9010.
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, which indicates 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 min where 90 mM CaCl2 solution was added 1 min after the measurement was started.
ments of the line width in Figure 3, panel B, where 2% SLG100 has the highest line width (1.07 mm) when printing compared to 3% SLG100 and 4% SLG100. 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 3, panel A 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, panel C2.
Comparing the small grid printed with 3% alginate with the grid printed with 2.5% NFC in panels 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 cross-linking ability of alginate was formulated. The successful combination of these two materials is shown by panel 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 cross-linking of alginate. 1492
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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 cross-linking. (1 and 5) Ink6040, (2 and 6) Ink7030, (3 and 7) Ink8020, and (4 and 8) Ink9010. The diameter of the discs was measured before and after cross-linking.
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 4, panel A. This is in good agreement with the shear thinning properties previously observed for nanocelluloses.32,38 The flow curves of the four different bioinks are all very similar, which indicates 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 cross-linking 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 4, panel B show that the storage modulus G′ and loss modulus G″ increased for higher proportions of NFC in the bioink formulation. The tan δ values, Figure 4, panel C, 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, which indicate that the inks were more gel-like than liquid. The gelling behavior of the bioinks has been studied by measuring the loss modulus and storage modulus over time while cross-linking with CaCl2. The gelling curves for Ink7030 and Ink8020 in Figure 4, panels D and E show a fast increase in modulus at t = 1 min and gradually become linear. The gelling curve of Ink9010 follows the trend of the other inks at the start but decreases after 5 min, which indicates that the gelled structure is being disrupted. Ink6040 shows a 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 cross-linking. The unconfined compression test revealed the mechanical properties of all groups. The stress strain curves of all groups can be seen in Figure 5, panel A. 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 nonlinear 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 5, panel B, 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 4, panels D and E. This indicates that a high alginate concentration impairs the mechanical properties. Regarding the compressive stress, Figure 5, panel C, all groups show a similar behavior, where Ink9010 displays the lowest value as expected. To evaluate the change in shape and size after cross-linking printed solid discs, the four bioink formulations were studied directly after printing and after 10 min of cross-linking. All discs were printed with the same process protocol, and the variety in size before cross-linking (9−10 mm) is mainly due to the properties of the different inks. A large change in the shape is seen in Figure 5, panel D 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 cross-linking, while at the same time swells at the center as seen in pictures 5 and 6 taken from the side in Figure 5, panel D. Printed discs of Ink8020 and Ink9010 do not show the same change in shape after cross-linking; the diameter only 1493
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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 (hNC) were used for 3D bioprinting of gridded constructs. On the basis of 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 hNC 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 7, panel A). 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 independent t test determined a significantly higher viability in hNC before embedding (95.3 ± 0.1%) compared to after embedding in bioink and cross-linking 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 nonprinted group after 1 day of culture (Figure 7, panels A−C), which indicates that the preparation and mixing were 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 nonprinted and 3D bioprinted constructs, which indicates a successful mixing of the cells in the bioink. The decrease in cell viability, after embedding the hNC and cross-linking the constructs, could be attributed to the shear forces applied to the cells during mixing with the bioink or to the long cross-linking process. Although there was a loss of cells after the embedding and bioprinting processes, the cells were able to recover after 7 days of culture, as shown by the higher cell viability compared to day 1 (Figures 7, panels D 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.
decreases 0.5 mm, and the swelling is barely noticeable when the disc is examined from the side in pictures 7 and 8 in Figure 5, panel D. Taking all results into account, rheology, compression, and shape deformation, Ink8020 would be the optimal ink formulation. It shows both a good gelling while cross-linking and shows little tendency of shape deformation. 3D Bioprinting Bioink. Because of the high shape fidelity of the formulated ink, the whole print with all its layers could be dispensed before cross-linking the structure with CaCl2. After 10 s, the effect of cross-linking could be seen. The printing process showing a cross-linked small grid acting as a stiff gel is depicted in Figure 6, panels A−C.
Figure 6. (A) 3D printed small grids (7.2 × 7.2 mm2) with Ink8020 after cross-linking. (B) The shape of the grid deforms while squeezing, and (C) it is restored after squeezing. (D) 3D printed human ear and (E and F) sheep meniscus with Ink8020. Side view (E) and top view (F) of meniscus.
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 6, panels D−F. Even with printing times of up to 20 min 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 culture.39 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 noncytotoxic.
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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 pressureessential conditions when 3D bioprinting with living cells. Before crosslinking, the properties of NFC were dominating the bioink and making it shear thinning with high shape fidelity and increased printing resolution as compared to 3D printing with alginate. After cross-linking, 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 1494
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Figure 7. (A) Viability of hNC before and after 3D bioprinting. Representative images showing dead (red) and live (green) cells (B) before and (C) after bioprinting hNC in Ink8020 and 3D culture for 1 day. (D and E) Representative images (at 4× and 10× magnifications) showing dead and live cells in 3D bioprinted constructs after 7 days of culture.
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 hNC 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.
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of the Project E!8355 CELLINK. The Knut and Alice Wallenberg Foundation is gratefully acknowledged for sponsoring the Wallenberg Wood Science Center.
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ABBREVIATIONS NFC, nanofibrillated cellulose; pDADMAC, polydiallyl dimethylammonium chloride; LVR, linear viscoelastic region; STL, stereolithography; hNC, human nasoseptal chondrocytes; G′, storage modulus; G″, loss modulus
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REFERENCES
(1) Derby, B. Science (Washington, DC, U. S.) 2012, 338, 921−926. (2) Nakamura, M.; Iwanaga, S.; Henmi, C.; Arai, K.; Nishiyama, Y. Biofabrication 2010, 2, 014110. (3) Falconnet, D.; Csucs, G.; Michelle Grandin, H.; Textor, M. Biomaterials 2006, 27, 3044−3063. (4) Chang, R.; Nam, J.; Sun, W. Tissue Eng., Part C 2008, 14, 157− 166. (5) Hutmacher, D. W. Nat. Mater. 2010, 9, 90−93. (6) Malda, J.; Visser, J.; Melchels, F. P.; Jungst, T.; Hennink, W. E.; Dhert, W. J.; Groll, J.; Hutmacher, D. W. Adv. Mater. (Weinheim, Ger.) 2013, 25, 5011−5028. (7) Khalil, S.; Sun, W. J. Biomech. Eng. 2009, 131, 111002−111002. (8) Xu, C.; Chai, W.; Huang, Y.; Markwald, R. R. Biotechnol. Bioeng. 2012, 109, 3152−3160. (9) Lee, W.; Debasitis, J. C.; Lee, V. K.; Lee, J.-H.; Fischer, K.; Edminster, K.; Park, J.-K.; Yoo, S.-S. Biomaterials 2009, 30, 1587− 1595. (10) Gruene, M.; Pflaum, M.; Deiwick, A.; Koch, L.; Schlie, S.; Unger, C.; Wilhelmi, M.; Haverich, A.; Chichkov, B. N. Biofabrication 2011, 3, 015005. (11) Chung, J. H. Y.; Naficy, S.; Yue, Z.; Kapsa, R.; Quigley, A.; Moulton, S. E.; Wallace, G. G. Biomater. Sci. 2013, 1, 763. (12) Frampton, J. P.; Hynd, M. R.; Shuler, M. L.; Shain, W. Biomed. Mater. (Bristol, U. K.) 2011, 6, 015002. (13) Guillotin, B.; Guillemot, F. Trends Biotechnol. 2011, 29, 183− 190. (14) Lee, V.; Singh, G.; Trasatti, J. P.; Bjornsson, C.; Xu, X.; Tran, T. N.; Yoo, S. S.; Dai, G.; Karande, P. Tissue Eng., Part C 2014, 20, 473− 484. (15) Song, S. J.; Choi, J.; Park, Y. D.; Lee, J. J.; Hong, S. Y.; Sun, K. Artif. Organs 2010, 34, 1044−1048. (16) Almeida, C. R.; Serra, T.; Oliveira, M. I.; Planell, J. A.; Barbosa, M. A.; Navarro, M. Acta Biomater. 2014, 10, 613−622.
AUTHOR INFORMATION
Corresponding Author
*Phone: +46317723407. E-mail:
[email protected]. Author Contributions
The manuscript was written through contributions of all authors. Development of bioinks, rheology, and cross-linking studies were carried out by K.M. A.M. performed the mechanical testing. Evaluation of printer parameters and line width measurements was carried out by I.T. and A.M. The cell studies were carried out by D.H. and H.M.A. All authors have given approval to the final version of the manuscript. Notes
The authors declare no competing financial interest.
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ACKNOWLEDGMENTS 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 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 is greatly acknowledged for financial support 1495
DOI: 10.1021/acs.biomac.5b00188 Biomacromolecules 2015, 16, 1489−1496
Article
Biomacromolecules (17) Pescosolido, L.; Schuurman, W.; Malda, J.; Matricardi, P.; Alhaique, F.; Coviello, T.; van Weeren, P. R.; Dhert, W. J.; Hennink, W. E.; Vermonden, T. Biomacromolecules 2011, 12, 1831−1838. (18) Kolesky, D. B.; Truby, R. L.; Gladman, A. S.; Busbee, T. A.; Homan, K. A.; Lewis, J. A. Adv. Mater. (Weinheim, Ger.) 2014, 26, 3124−3130. (19) Bertassoni, L. E.; Cardoso, J. C.; Manoharan, V.; Cristino, A. L.; Bhise, N. S.; Araujo, W. A.; Zorlutuna, P.; Vrana, N. E.; Ghaemmaghami, A. M.; Dokmeci, M. R.; Khademhosseini, A. Biofabrication 2014, 6, 024105. (20) Lee, K. Y.; Mooney, D. J. Prog. Polym. Sci. 2012, 37, 106−126. (21) Augst, A. D.; Kong, H. J.; Mooney, D. J. Macromol. Biosci. 2006, 6, 623−633. (22) Kong, H.-J.; Lee, K. Y.; Mooney, D. J. Polymer 2002, 43, 6239− 6246. (23) Lee, J. S.; Hong, J. M.; Jung, J. W.; Shim, J. H.; Oh, J. H.; Cho, D. W. Biofabrication 2014, 6, 024103. (24) Krontiras, P.; Gatenholm, P.; Hagg, D. A. J. Biomed. Mater. Res., Part B 2015, 103, 195−203. (25) Park, M.; Lee, D.; Hyun, J. Carbohydr. Polym. 2015, 116, 223− 228. (26) Klemm, D.; Kramer, F.; Moritz, S.; Lindström, T.; Ankerfors, M.; Gray, D.; Dorris, A. Angew. Chem., Int. Ed. 2011, 50, 5438. (27) Bäckdahl, H.; Helenius, G.; Bodin, A.; Nannmark, U.; Johansson, B. R.; Risberg, B.; Gatenholm, P. Biomaterials 2006, 27, 2141−2149. (28) Starborg, T.; Kalson, N. S.; Lu, Y.; Mironov, A.; Cootes, T. F.; Holmes, D. F.; Kadler, K. E. Nat. Protoc. 2013, 8, 1433−1448. (29) Svensson, A.; Nicklasson, E.; Harrah, T.; Panilaitis, B.; Kaplan, D. L.; Brittberg, M.; Gatenholm, P. Biomaterials 2005, 26, 419−431. (30) Ahrem, H.; Pretzel, D.; Endres, M.; Conrad, D.; Courseau, J.; Müller, H.; Jaeger, R.; Kaps, C.; Klemm, D. O.; Kinne, R. W. Acta Biomater. 2014, 10, 1341−1353. (31) Yamaoka, H.; Asato, H.; Ogasawara, T.; Nishizawa, S.; Takahashi, T.; Nakatsuka, T.; Koshima, I.; Nakamura, K.; Kawaguchi, H.; Chung, U. I.; Takato, T.; Hoshi, K. J. Biomed. Mater. Res., Part A 2006, 78A, 1−11. (32) Paakko, M.; Ankerfors, M.; Kosonen, H.; Nykanen, A.; Ahola, S.; Osterberg, M.; Ruokolainen, J.; Laine, J.; Larsson, P. T.; Ikkala, O.; Lindstrom, T. Biomacromolecules 2007, 8, 1934−1941. (33) Olderoy, M. O.; Lilledahl, M. B.; Beckwith, M. S.; Melvik, J. E.; Reinholt, F.; Sikorski, P.; Brinchmann, J. E. PLoS One 2014, 9, e91662. (34) Fall, A. B.; Burman, A.; Wagberg, L. Nord. Pulp Pap. Res. J. 2014, 29, 176−184. (35) Wågberg, L.; Winter, L.; Ö dberg, L.; Lindström, T. Colloids Surf. 1987, 27, 163−173. (36) Nimeskern, L.; Martinez Avila, H.; Sundberg, J.; Gatenholm, P.; Muller, R.; Stok, K. S. J. Mech. Behav. Biomed. Mater. 2013, 22, 12−21. (37) Wagberg, L.; Decher, G.; Norgren, M.; Lindstrom, T.; Ankerfors, M.; Axnas, K. Langmuir 2008, 24, 784−795. (38) Nechyporchuk, O.; Belgacem, M. N.; Pignon, F. Carbohydr. Polym. 2014, 112, 432−439. (39) Rouillard, A. D.; Berglund, C. M.; Lee, J. Y.; Polacheck, W. J.; Tsui, Y.; Bonassar, L. J.; Kirby, B. J. Tissue Eng., Part C 2011, 17, 173− 179.
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DOI: 10.1021/acs.biomac.5b00188 Biomacromolecules 2015, 16, 1489−1496
