Fiber engraving for bioink bioprinting within 3D printed tissue engineering scaffolds

Journal Pre-proof Fiber engraving for bioink bioprinting within 3D printed tissue engineering scaffolds Luis Diaz-Gomez, Maryam E. Elizondo, Gerry L. Koons, Mani Diba, Letitia K. Chim, Elizabeth Cosgriff-Hernandez, Anthony J. Melchiorri, Antonios G. Mikos PII:

S2405-8866(20)30003-8

DOI:

https://doi.org/10.1016/j.bprint.2020.e00076

Reference:

BPRINT 76

To appear in:

Bioprinting

Received Date: 17 September 2019 Revised Date:

22 December 2019

Accepted Date: 2 January 2020

Please cite this article as: L. Diaz-Gomez, M.E. Elizondo, G.L. Koons, M. Diba, L.K. Chim, E. CosgriffHernandez, A.J. Melchiorri, A.G. Mikos, Fiber engraving for bioink bioprinting within 3D printed tissue engineering scaffolds, Bioprinting, https://doi.org/10.1016/j.bprint.2020.e00076. This is a PDF file of an article that has undergone enhancements after acceptance, such as the addition of a cover page and metadata, and formatting for readability, but it is not yet the definitive version of record. This version will undergo additional copyediting, typesetting and review before it is published in its final form, but we are providing this version to give early visibility of the article. Please note that, during the production process, errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain. © 2020 Published by Elsevier B.V.

CREDIT AUTHOR STATEMENT

Luis Diaz-Gomez: Conceptualization, Methodology, Investigation, Writing - Original draft. Maryam E. Elizondo: Investigation, Writing – Review & Editing. Gerry L. Koons: Investigation, Writing – Review & Editing. Mani Diba: Investigation, Writing – Review & Editing. Letitia Chim: Investigation, Writing – Review & Editing. Elizabeth Cosgriff-Hernandez: Supervision, Writing – Review & Editing. Anthony J. Melchiorri: Project administration, Writing – Review & Editing. Antonios G. Mikos: Supervision, Funding acquisition, Writing – Review & Editing.

Fiber engraving for bioink bioprinting within 3D printed tissue engineering scaffolds

Luis Diaz-Gomeza,b,c, Maryam E. Elizondoa,b,c, Gerry L. Koonsa,b,c, Mani Diba,a,b,c, Letitia K. Chima,b,c, Elizabeth Cosgriff-Hernandezd, Anthony J. Melchiorria,b,c, and Antonios G. Mikosa,b,c*

a

Department of Bioengineering, Rice University, 6500 Main Street, Houston, TX 77030, USA;

b

Biomaterials Lab, Rice University, 6500 Main Street, Houston, TX 77030, USA;

c

NIH / NIBIB Center for Engineering Complex Tissues, USA;

d

Department of Biomedical Engineering, University of Texas at Austin, 107 W Dean Keeton

Street, Austin, TX 78712, USA.

*

To whom correspondence may be addressed:

Antonios G. Mikos, PhD Department of Bioengineering, MS-142 BioScience Research Collaborative Rice University 6500 Main Street Houston, TX 77030 E-mail: [email protected] Tel: (713) 348-5355 Fax: (713) 348-4244

Declarations of interest: none

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Abstract In this work, we describe a new 3D printing methodology for the fabrication of multimaterial scaffolds involving the combination of fused deposition modelling and low temperature extrusion of bioinks. A fiber engraving technique was used to create a groove on the surface of a thermoplastic printed fiber using a commercial 3D printer and a low viscosity bioink was deposited into this groove. In contrast to traditional extrusion bioinks that rely on increased viscosity to prevent lateral spreading, this groove creates a defined space for bioink deposition. By physically constraining bioink spreading, a broader range of viscosities can be used. As proof-of-concept, we fabricated and characterized a multimaterial scaffold containing poly(εcaprolactone) (PCL) as the thermoplastic polymer and a gelatin-based bioink. A 7.5 w/v% gelatin methacryloyl (GelMA) bioink loaded with either 5 w/v% poly(lactic-co-glycolic acid) (PLGA) microparticles containing fluorescent albumin or mouse fibroblasts (1x106 cell/mL) was printed at 24ºC. The structure of the composite scaffolds had no significant decrease in porosity or mechanical properties as compared to the PCL control scaffolds, demonstrating the engraving technique did not significantly compromise the mechanical or structural integrity of the scaffold. The encapsulated PLGA microparticles were homogeneously distributed in the GelMA and remained in the scaffolds after incubation in PBS for 24 h at 37ºC. In addition, the viability of the fibroblasts encapsulated in the GelMA bioink and printed in the grooves of the PCL scaffolds was confirmed after 24 h of incubation. Overall, this work provides a new methodology for the preparation of 3D printed scaffolds containing a robust thermoplastic structure in combination with low viscosity bioinks.

Keywords: 3D printing, engraving, tissue engineering, biofabrication, co-printing.

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Abbreviations: gelatin methacryloyl (GelMA), poly(lactic-co-glycolic acid) (PLGA), poly(εcaprolactone) (PCL), fused deposition modelling (FDM), phosphate-buffered saline (PBS), fetal bovine serum (FBS), poly(vinyl alcohol) (PVA).

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1. Introduction Bioprinting is a powerful technique to prepare custom tissue engineering scaffolds with high resolution on demand [1,2]. The most explored alternative for the preparation of biodegradable scaffolds is fused deposition modelling (FDM) of thermoplastics [3]. FDM usually results in structures with tunable mechanical properties that can be achieved by incorporating materials with reinforcement properties, such as ceramics [4]. However, the incorporation of cells using FDM must be approached as a post-printing step due to the unaccommodating high temperatures at which FDM printing operates. Static or dynamic seeding is usually used to cellularize these scaffolds, but the high hydrophobicity of most of the synthetic polymers used in FDM often leads to uneven seeding and distribution [5]. This is especially disadvantageous when different cell types need to be spatially organized within specific regions of the scaffolds to mimic the composition of complex tissues [6,7]. The use of cell-laden bioinks, defined as inks containing cells, provides a well-tested strategy to create cell-laden scaffolds for tissue regeneration [8,9]. Bioink printing permits the spatial control of cell deposition within the scaffold. A wide range of natural materials used in bioinks form hydrogels capable of supporting cell encapsulation as well as proliferation and migration after the construct is fabricated and implanted [10–12]. Lack of shape fidelity after printing is one of the main drawbacks of traditional bioinks [13]. The printability of bioinks can be improved by increasing the viscosity of the bioink or decreasing the gelation time to prevent lateral spreading [14,15]. However, the increase in bioink viscosity generates shear stress in the nozzle that can damage cells during extrusion [16,17].

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In addition, the applications of these cell-containing hydrogels are limited due to the poor mechanical properties of the constructs [18]. Despite modifications to increase the mechanical stability and printability of hydrogels, the tensile and compressive strengths of the resulting scaffolds are still significantly limited [19,20]. In order to obtain cell-laden scaffolds with adequate mechanical properties to match those of the targeted tissue, bioinks have been co-printed in parallel with synthetic polymers using commercially available printers with at least two printing heads [21,22]. The combination of extrusion printing and FDM makes it possible to fabricate a polymeric frame with high mechanical properties and introduce cell-laden bioinks to populate the construct with specific cell types and growth factors to promote tissue regeneration [22]. Recently, the use of co-printing to fabricate a mechanically reinforced cartilaginous scaffold by printing parallel strands of poly(ε-caprolactone) (PCL) using FDM and an arginylglycylaspartic acid (RGD)-alginate bioink loaded with mesenchymal stem cells via extrusion has been reported [21]. In another study, poly(lactic acid) (PLA) fibers printed with FDM were co-printed with a gelatin-methacryloyl (GelMA)-based bioink using stereolithography (SLA) to obtain a hierarchical biomimetic bonelike structure [23,24]. These co-printing strategies, however, require hydrogels that can retain their shape after extrusion, which significantly limits the choice of hydrogel materials that can be used in these applications. In addition, printing high viscosity inks adjacent to thermoplastic strands would result in the decrease of scaffold porosity, as well as a low cell viability, as reported previously [21,25,26]. Despite dramatic advances in bioprinting technology and bioinks, the preparation of intricate tissue structures with specific regional characteristics and cell population is limited.

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In this work, we used a commercially available 3D printer to develop a printing methodology to engineer mechanically reinforced cell-laden scaffolds prepared with low viscosity bioinks while maintaining a highly porous and interconnected structure. An engraving printing methodology was developed to create tunable grooves in which low viscosity bioinks can be printed. Printing and engraving reproducibility were assessed, and the structures of the printed, engraved, and filled scaffolds were characterized. Finally, we assessed the suitability of the developed printing methodology to homogeneously incorporate GelMA inks containing poly(lactic-co-glycolic acid) (PLGA) microparticles encapsulating fluorescently-labelled albumin, or fibroblasts while maintaining cell viability.

2. Materials and methods 2.1. Materials Poly(ε-caprolactone) (PCL; nominal molecular weight 50 KDa) was obtained from Polysciences (Warrington, PA). E-Shell® 300 was from EnvisionTEC (Gladbeck, Germany). Gelatin methacryloyl (GelMA) was from Allevi (Philadelphia, PA). Hydroxyapatite, Irgacure 2959 and fluorescein isothiocyanate-bovine serum albumin (FITC-Albumin) were from SigmaAldrich (St. Louis, MO). High glucose Dulbecco modified Eagle medium (DMEM), LGlutamine, penicillin/streptomycin, TrypLE™ Express, and phosphate-buffered saline (PBS) were purchased from Gibco (Waltham, MA). Live/dead viability assay kit was from Thermo Fisher (Waltham, MA). Poly(lactic-co-glycolic acid) (PLGA, 50:50 copolymer ratio) was from Evonik (Essen, Germany). 2.2. Gelatin-based ink preparation

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GelMA ink was prepared from the lyophilized material supplied by the manufacturer. Briefly, 7.5 w/v% GelMA solutions were prepared by dissolving the lyophilized GelMA in sterile PBS for the microparticle encapsulation or high glucose DMEM culture medium for cell encapsulation. These solutions also contained 0.5 w/v% of photoinitiator (Irgacure 2959) to enable UV-induced crosslinking. Inks were stored at 37°C, protected from light, until use.

2.3. Scaffold printing and engraving Cuboid scaffolds (10×10×5 mm) were printed based on 3D models designed in SketchUp (Trimble; Sunnyvale, CA). Engraved and filled scaffolds were designed as two or three superimposed cuboids of 10×10×5 mm, respectively. All of the 3D models were sliced in 800 µm height layers. PCL in powder form was placed in printer cartridges, heated to 140ºC, and extruded through an 18G metal needle using a 3D-Bioplotter (EnvisionTEC, Gladbeck, Germany). The printing pattern consisted of continuous straight strands of 1.2 mm on-center spacing and layers were printed at 90º angle with respect to the previous layer. The printing process was structured in three steps as summarized in Figure 1: 1) printing of the fibers, 2) engraving, and 3) filling. PCL strands were printed at 140ºC, 4.5 bar and 1.5 mm/s using a needle offset of 0.8 mm. After each layer of PCL was printed, a 20 s delay time was set to allow for the solidification of the extruded polymer. The engraving was performed at 5 mm/s and 0 bars (no material extrusion) using a 22G needle placed in an empty cartridge and heated at 120ºC, and the printing needle offset was decreased to 0.4 mm or 0.3 mm to obtain a depth of the engraving of 100 µm or 200 µm, respectively. Finally, two low-viscosity photocurable inks (GelMA or E-Shell®) were then printed inside the grooves using a 27G needle, and the offset was set back to 0.8 mm. E-Shell® was used as a low viscosity model ink and 0.1 w/v%

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rhodamine B was dispersed in the ink to provide a fluorescent contrast in order to visualize the printed material using epifluorescence microscopy. GelMA has been extensively reported in 3D printing applications for tissue engineering because its excellent biocompatibility, degradability, and printability. Printing conditions for the infill depended on the extruded material and are reported in Table 1. After deposition in the groove, the filler ink was crosslinked using UV light at 5 mW/cm2 for 30 s. The process was repeated for each layer until scaffold fabrication was complete. Microparticle and cell encapsulation studies were carried out in a simple model for demonstration purposes. Scaffolds were printed with three layers (2.4 mm height) where only the second and third layers were engraved following this procedure. PCL scaffolds without engraving were used as controls. In addition to creating composite scaffolds with cell-laden bioinks, the engraved tracks can be backfilled with thermoplastic resins as well to generate complex compositions. To demonstrate the ability to generate composite printed structures, scaffolds containing PCL and HA (30 w/w%; PCL-HA30) were first printed, then engraved, and finally the engraved grooves were filled with PCL. This composition was chosen since the incorporation of HA particles could be used as a contrast agent to evaluate the deposition and structure of the printed materials using microcomputed tomography (microCT)[27–29].

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Material

T (ºC)

Pressure (bar) Speed (mm/s)

Preflow (s)

Postflow (s)

PCL

140

6.5

0.5

1

1

E-Shell®

22

0.3

20

0.2

0.1

GelMA

24

0.4

4.0

0.2

0.2

Figure 1: Scheme for the preparation of multimaterial scaffolds using the engraving printing methodology. First, high temperature (HT) extrusion (FDM) is used to print the PCL fibers using a 18G needle at 140ºC. After each layer is printed, the fibers are engraved using a 22G needle at 120ºC. Finally, the grooves created in the PCL fibers are filled with a bioink extruded at low temperature (LT) and crosslinked using UV light. Captions show the cross-section of a single fiber at each step of the methodology. Table 1: Printing conditions of the filling materials extruded using a 27G needle.

2.4. Validation and characterization of the printing methodology The surfaces of the control, engraved, and filled scaffolds were evaluated by scanning electron microscopy (SEM). Briefly, scaffolds (n = 3) were sputter coated with gold and imaged using scanning electron microscopy (SEM) (Quanta 400 ESEM FEG; FEI, Hillsboro, OR). Samples were imaged under high vacuum at 10 kV and different magnifications. The precision

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of the E-Shell® material printing inside the engraved PCL fibers was further evaluated using epifluorescence microscopy with a Nikon Eclipse Ti2 (Nikon; Tokyo, Japan) using a TRITC filter set. The structure of the scaffolds was also evaluated by microcomputed tomography (microCT). Scaffolds from all experimental groups were scanned using a SkyScan 1272 X-ray microCT (Bruker, Kontich, Belgium). Images were obtained by 180º scanning at 7 µm/pixel resolution with Al 0.25 mm filter at voltage and current settings of 60 kV and 166 mA, respectively. Scan images were reconstructed, resliced, and analyzed using NRecon Reconstruction Software and CTAn software provided with the equipment. The compressive modulus of the control and engraved scaffolds (n = 5; 10 x 10 x 5 mm) was measured in a mechanical testing bench (10 kN load cell; 858 MiniBionixII; MTS, Eden Prairie, MN) and calculated using the TestStar 790.90 mechanical data analysis package included in the manufacturer’s software. Scaffolds were uniaxially compressed perpendicularly to the printing plane at a crosshead speed of 1 mm/min up to 20% strain. Stress-strain plot was calculated from the load vs. displacement data using the initial external dimensions of each sample. 2.5. Microparticle ink preparation, printing, and imaging PLGA particles loaded with fluorescent proteins were fabricated using a water-in-oil-inwater double emulsion solvent evaporation technique as described previously [30,31]. Briefly, 500 mg of PLGA was dissolved in 1.25 mL of dichloromethane at room temperature. Thereafter, 0.120 mL of an 8.33 mg/mL aqueous solution of FITC-albumin was added to the PLGA solution. The first emulsion was carried out using a Qsonica Q125 probe sonicator (Newtown, CT) for 1 min at 50% amplitude. Two mL of 1 w/v% aqueous poly(vinyl alcohol) (PVA) solution were

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then added to the first emulsion, and sonication was performed for 30 s at 50% amplitude. The resulting emulsion was immediately poured into a hardening solution consisting of 100 mL of 0.3 w/v% aqueous PVA and 100 mL of 2 v/v% aqueous isopropanol. Dichloromethane was evaporated for one hour, and the resulting particles were subsequently collected, washed through three cycles of centrifugation (10,000g for 15 min), and suspended in MilliQ water. The resulting suspension was then filtered through a 40 µm-mesh cell strainer (Fisher Sci; Waltham, MA), flash-frozen in liquid nitrogen, and freeze-dried. The obtained microparticles were stored protected from light at -20ºC until further usage. The obtained microparticles were used to prepare a model ink for drug and bioactive factor encapsulation. The dried microparticles were added to the 7.5 w/v% GelMA ink at 37ºC to obtain a final microparticle concentration of 5 w/v% and were homogeneously dispersed using a vortex. Then, the ink was loaded in a cartridge and placed in the printing head at 23ºC. Scaffolds with 3 layers were printed and engraved as described before and the PLGA microparticle-loaded GelMA ink was printed as filling material in the grooves of the PCL fibers and UV crosslinked at 5 mW/cm2 for 30 s. Then, scaffolds were transferred to 24-well plates and covered with 1 mL sterile filtered PBS. Scaffolds were imaged immediately after printing and again after 24 h of incubation in PBS at 37ºC using an epifluorescence microscope (Eclipse Ti2, Nikon; Tokyo, Japan) and a confocal fluorescent microscope (A1-Rsi, Nikon; Tokyo, Japan). 2.6. Bioink preparation, printing, and cytotoxicity evaluation L929 mouse fibroblasts (ATCC®, CCL-1) were cultured in DMEM supplemented with 10% horse serum, 1% penicillin/streptomycin and 1% L-glutamine in a humidified incubator at 37°C with 5% CO2. Cells were detached from culture flasks at ~80% confluency with TryplE®. Suspended cells were centrifuged at 200g for 5 min to pellet and resuspended in 5 mL of culture

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medium. After cell counting, 5x106 cells were transferred to a 15 mL conical tube and centrifuged as before. The subsequent cell pellet was homogeneously resuspended in 5 mL of 7.5 w/v% GelMA bioink at 37ºC. The cell-laden bioink was poured into a sterile cartridge and placed in a temperature-controlled printing head at 23ºC. The bioink was then printed in the engraved PCL fibers using a 27G needle and crosslinked using UV light at 5 mW/cm2 for 30 s. After printing, the scaffolds were removed from the printing platform and placed in 24-well plates with 1 mL of culture medium and incubated for 24 h in a humidified incubator with 5% CO2 at 37ºC. 2.7. Cytotoxicity assay A live/dead cytotoxicity assay was used to evaluate the effect of encapsulation and printing on cell viability. Briefly, ethidium homodimer-1 and calcein AM stock solutions were added to sterile filtered PBS to obtain working solutions with final concentrations of 4 µM and 2 µM, respectively. Then, culture medium was removed from the wells and scaffolds were soaked in sterile PBS and incubated in 500 µL of the working solution for 30 min. Before observation, the solution was removed, and scaffolds were washed in PBS for 5 min to reduce background signal. Finally, scaffolds were placed upside down on glass slides and cells were observed using confocal fluorescent microscopy (A1-Rsi, Nikon; Tokyo, Japan). 2.8. Statistical analysis Statistics were analyzed using GraphPad Prism (GraphPad, La Jolla, CA). All results are expressed as means ± standard deviation. One-way analysis of variance (ANOVA) and Tukey’s multiple comparison post-test were used. Differences were considered significant for p < 0.05.

3. Results

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3.1. Validation and characterization of the printing methodology

Figure 2: Optical images showing the steps to engrave and fill PCL scaffolds. First, each PCL layer is printed and allowed to solidify. Then, the surface of the PCL fibers is engraved by decreasing the needle offset to the desired depth of the grooves. Finally, GelMA is printed inside the grooves and crosslinked using UV light. Engraved areas are limited by white dashed lines. Scale bar: 1 mm. Image insets show the cross-section reconstruction of a fiber on each step of the engraving method. Scale bar: 400 µm. As shown in Figure 2, the engraving method described in this work allowed for the creation of grooves or channels on the surface of polymeric scaffolds that can be used to incorporate low viscosity inks in a reproducible manner and with variable height. The SEM micrographs and microCT scans of the engraved scaffolds showed that the bulk fiber structure was not altered, and the grooves created on the surface were aligned throughout the center of the fiber (Figure 3). The corners of the (Figures 3A and 3B) also showed the precision of the engraving independently of the depth of the engraving (100 or 200 µm).

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Figure 3: (A-D): SEM micrographs showing the surface of the PCL scaffolds engraved with grooves of depths 100 µm (PCL 100 µm; A, C) and 200 µm (PCL 200 µm; B, D). The fidelity of the engraving can be observed in both straight fibers and the corners as well as in the layers below. (E): MicroCT reconstruction of a PCL scaffold with 200 µm engraving depth. Scale bar: 500 µm (A-D) and 250 µm (E).

The incorporation of the filling material in the engraved fibers was evaluated using both the low viscosity, photocurable E-Shell® ink and the thermoplastic PCL filler. A fluorescent dye, rhodamine B, was dispersed in the E-Shell® ink to visualize the deposition of the ink inside the grooves created on the PCL fibers using an epifluorescence microscope. The SEM and epifluorescence micrographs (Figure 4) confirmed the precision of the deposition of E-Shell® inside the engraved PCL fibers, including in the curved turns of the fibers. The deposited material appeared to fill the complete volume of the grooves without any detrimental artifact such as material overfill and subsequent spreading outside of the engraved grooves.

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Figure 4: Epifluorescence images (A, B) and SEM (C, D) micrographs of the PCL 200 µm and PCL 200 µm-E-Shell® fibers. The high fidelity of the low viscosity ink printing allowed for the precise deposition of E-Shell® inside the 200 µm grooves created on the PCL fibers, including inside the curves at the edges of the scaffold. Scale bar: 500 µm (A, B and C) and 250 µm (D). Similarly, PCL-HA30 scaffolds were engraved and filled with PCL extruded at 140ºC and analyzed using SEM and microCT to evaluate the effect of thermoplastic filling in a thermoplastic engraved scaffold. (Figure 5). PCL-HA30 was selected as the fiber material since the incorporation of ceramics can be used as a contrast agent to highlight the differences in the composition of the fiber and the filling material. The deposition of PCL within the grooves of the PCL-HA30 scaffolds was clearly observed using microCT. The sliced reconstructions showed

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that the PCL is well distributed along the groove. PCL filling using a 22G needle resulted in the deposition of a continuous strand of 400 µm in diameter along the groove previously created in the PCL-HA30 fiber.

Figure 5: SEM micrographs (A, B) and microCT reconstructions (C, D) of a PCL-HA30 scaffold, engraved with 200 µm depth grooves, and filled with PCL. A and B show different areas of the surface of a single scaffold layer obtained using SEM. C and D show microCT reconstructions of two different layers of a single scaffold. The precision of the deposition was evident and resulted in the complete and continuous filling of the 200 µm grooves created on the PCL-HA30 fibers independently of the area or the layer of the scaffold. Scale bar: 500 µm.

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MicroCT scans were then used to evaluate the fiber diameter, pore size, porosity, and engraving depth. PCL scaffolds composed of 6 layers were printed and used as controls. Similarly, PCL scaffolds with 100 µm (PCL 100 µm) or 200 µm (PCL 200 µm) engraving depth were printed and characterized to evaluate the effect of the engraving on the final structure of the scaffold. Finally, PCL 200 µm-E-Shell® scaffolds were also characterized to evaluate the effect of the filling process in the overall structural properties of the resulting scaffolds. As shown in Table 2, the fiber diameter and the pore size of the controls, engraved and filled scaffolds were statistically similar (p ˃ 0.05). The engraving depth measured for both 100 and 200 µm needle offset engravings resulted in measured depths of 95 ± 27 and 215 ± 23 µm, respectively. The effect of filling the engraved PCL fibers with a low viscosity ink such as E-Shell® was also evaluated. The results showed that the pore size and fiber diameter of the scaffolds were not significantly changed by the filling of the grooves with the E-Shell® ink compared to the PCL control scaffolds. Furthermore, the porosity of the scaffolds was not significantly affected by the incorporation of the E-Shell® compared to PCL controls. Table 2: Structural characterization of PCL controls; PCL 100 µm and PCL 200 µm engraved scaffolds; and PCL 200 µm-E-Shell® scaffolds was assessed by taking 30 random measurements of each scaffold per group (n = 5 scaffolds in each group). Fiber diameter

Pore size

Porosity

Engraving depth

Engraving width

(µm)

(µm)

(%)

(µm)

(µm)

PCL control

1028 ± 38

910 ± 55

47 ± 6

-

-

PCL 100 µm

1030 ± 22

924 ± 33

50 ± 8

95 ± 27

465 ± 23

PCL 200 µm

1044 ± 23

921 ± 59

52 ± 7

215 ± 23

485 ± 34

1003 ± 51

925 ± 37

48 ± 6

203 ± 15

469 ± 14

PCL 200 µm-E-Shell

®

17

The mechanical properties of the PCL controls and the PCL 200 µm scaffolds were evaluated in order to characterize the effect of engraving on the overall mechanical strength of the scaffolds. The compressive tests performed on the PCL controls and the 200 µm engraved scaffolds showed that the creation of the grooves on the surface of the PCL scaffolds had no significant effect on the compressive modulus of the scaffolds (Figure 6). The compressive modulus of the PCL control scaffolds was 70 ± 13 MPa. Similarly, PCL engraved scaffolds showed a compressive modulus of 61 ± 11 MPa.

Figure 6: Compressive modulus of control PCL scaffolds and PCL 200 µm scaffolds. Data reported as means ± standard deviation (n = 5 scaffolds). 3.2. Bioink Printing First, a 7.5 w/v% GelMA ink loaded with 5 w/v% FITC-albumin-loaded PLGA microparticles was printed in the grooves of engraved PCL fibers to show the effect of printing and incubation on the microparticle distribution (Figure 7 A-B and Figure S1). This ink was

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chosen as a model to demonstrate the feasibility for including various compounds or biological factors within the engraved scaffolds. Background signal from the GelMA was evident in the micrographs obtained using both confocal and epifluorescence microscopy. The results showed that the GelMA ink was accurately printed inside the PCL fiber grooves without observing any significant overfilling or spreading over the scaffold. The microparticles were homogeneously distributed inside the hydrogel without significant aggregation. The maintenance of the structure of the printed scaffolds was also evaluated after incubation in PBS for 24 h (Figure S2). The results showed that the ink and the microparticles were still retained inside the grooves of the PCL fibers, without any apparent disruption of the hydrogel structure. Finally, L929 mouse fibroblasts were encapsulated in a GelMA hydrogel ink and printed in the grooves of engraved PCL fibers. This cell type was chosen for the present study as they are used for preliminary cytotoxicity evaluation for a wide range of biomaterials, and are recommended by ISO Standard 10993-5 [32]. After 24 h of incubation, cell viability was evaluated by means of the live/dead staining, containing calcein AM and ethidium homodimer-1. Cell viability was demonstrated by the presence of a large number of green-stained cells (live) compared to the number of red-stained cells (dead) present in the GelMA deposited in the grooves (Figure 7 C-D).

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Figure 7: (A, B): Confocal micrographs from two different PCL 200 µm scaffolds filled with a 7.5 w/v% GelMA ink containing FITC-albumin-loaded PLGA microparticles. The homogeneous distribution of the microparticles and the accuracy of the printing inside the grooves created on the PCL fibers were revealed. Engraving was outlined by white dashed lines. (C, D): Merged images from confocal microscopy of live/dead staining, revealing cell viability after 24 h of cell culture in both a straight fiber (left) and in the curve on the edge of the layer (right). Engraved regions are outlined by a white dashed line. Scale bar: 50 µm. 4. Discussion In this study, we developed a new 3D printing methodology to prepare tissue engineering scaffolds composed of a mechanically robust PCL structure and incorporating low viscosity biomaterials in manner that maintains the scaffolds’ highly interconnected porous structure. This

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novel methodology involves the engraving of the upper surface of the PCL fibers to create a continuous groove along the direction of the fiber that serves as a mold for the subsequent printing of low viscosity biomaterial inks or bioinks. As a layer-by-layer printing methodology, this process allows for the deposition of different material compositions within the same construct without compromising the mechanical properties or the porosity of the scaffolds. Furthermore, the viscosity of the filling material does not need to meet normal printability requirements (e.g., post-printing maintenance of the fiber structure) since the filling material is deposited and contained inside the grooves [33]. This is a major advantage of the engraving technique compared to methods involving parallel printing of hydrogels and thermoplastic materials, or pure hydrogel scaffolds that require a higher viscosity during the printing process to maintain the fiber structure and prevent the spreading of the printed material. Another advantage of the engraving method is that the process of printing, engraving and filling can be carried out using a commercial bioprinter equipped with exchangeable heads. In this work, 4 heads were used sequentially: main fiber material head, engraving head, filling material head and, finally, UV light head. The main limitation of this technique is the resolution of the grooves, since we used commercially available 22G hollow needles for the engraving step. The use of smaller needles (24G and 27G) was also evaluated (data not shown). However, the mechanical weakness of the thin and hollow needles often resulted in the bending of the needle tip and the failure of the engraving. This limitation can be potentially overcome in future studies by using solid tips that can withstand stronger forces, enabling the smaller dimensions of engraving for polymeric fibers. Similarly, we found that the depth of the engraving is limited to 200 µm since greater engraving depths usually lead to the bending of the engraving needle. However, this limitation could

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potentially be overcome in future studies by employing consecutive engraving steps of up to 200 µm each until the desired groove depth is reached. In order to create a cell-laden scaffold with sufficient mechanical properties to fulfill the requirements of the host tissue, conventional co-printing techniques rely on the coaxial or parallel printing of alternate fibers of thermoplastics and cell-laden hydrogels [21,34]. These methods can result in the fabrication of mechanically reinforced cellularized scaffolds, but there are some drawbacks that must be overcome to ensure the success of the tissue regeneration process. For instance, the coaxial printing of thermoplastics and cell-laden hydrogels is highly influenced by the heat transfer between the fused deposited material, such as PCL, and the hydrogel thus resulting in a decrease in the cell viability [8]. Also, the shear-stress observed during the printing process often results in a marked decrease in the viability of the cells encapsulated in the scaffolds [17]. As a proof of concept, we used E-Shell®, a commercially available methacrylate-based resin, as filling material which has a viscosity of 339.8 mPa·s (according to the product´s datasheet). As a comparison, extrusion printing requires a minimum viscosity of 30x107 mPa·s to obtain reproducible constructs [33]. The low viscosity of E-Shell® makes it impossible to use this material for extrusion printing as it has virtually no structural retention following extrusion and can only be used in laser-based printing methods, such as SLA. The engraving printing method enabled us to include E-Shell® in the grooves of the PCL and PCL-HA30 fibers at room temperature without the need of including sacrificial materials, intricate printing setups, or postprocessing steps [35]. The engraving method also successfully incorporated a microparticle-loaded hydrogel in the PCL scaffold. FITC-albumin-loaded PLGA microparticles (used as models for drug and

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growth factor loaded microparticles) were successfully encapsulated in the GelMA ink and printed in the grooves of the PCL fibers. The results showed the homogeneous distribution of the microparticles throughout the hydrogel without significant aggregation. Furthermore, the microparticles were retained in the hydrogel for at least 24 h in PBS at 37ºC, suggesting the stability of the construct. This technique might be of interest in drug delivery systems. More specifically, the encapsulation and sustained release of labile biomolecules, especially growth factors, is still a challenge in the regenerative medicine field and bioprinting [36–39]. The main drawback of the previously reported 3D printed scaffolds for controlled release of bioactive molecules is that the release profile is limited by the degradation rate of the scaffolds [40]. Nonetheless, future investigation is necessary to comprehensively evaluate the capacity of the current technology for drug delivery applications. Traditional drug delivery systems incorporate bioactive molecules or encapsulated molecules in the polymeric matrix resulting in systems with a poor compromise between the delivery rate and the mechanical properties [41]. However, drug-containing hydrogels incorporated within the engraved fibers could have release properties independent of the degradation of the polymeric framework. Furthermore, the high temperatures needed to print thermoplastics such as PCL or PLGA prohibit the incorporation of thermolabile molecules, such as growth factors. However, the engraving method allows for the incorporation of bioactive molecules and microparticles under mild conditions of temperature and shear stress. Furthermore, the release can be controlled exclusively by modifying the matrix where the molecules are encapsulated while maintaining the mechanical properties required for the tissue which the scaffold is to be implanted.

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The maintenance of the porous structure and interconnectivity when co-printing materials are used remains a main challenge. The deposition of parallel fibers leads to a decrease in the porosity and interconnectivity of the scaffolds [42]. Such porosity and pore interconnectivity are crucial to cell and tissue infiltration and exchange of nutrients, waste products and degradation byproducts [42–44]. The analysis of the structure of the engraved and filled scaffolds did not show any decrease in the porosity of the scaffolds due to the incorporation of the filling ink. As shown in the mechanical characterization, the compressive properties of the scaffolds were not significantly affected by the engraving process and were in the range of human trabecular bone (50-150 MPa) [45]. Furthermore, the reduced contact area between the PCL portions of fibers arising from the engraving did not result in a decrease in the integrity of the scaffolds. Despite the forces applied to the scaffold during the engraving process, no disruption or delamination in the structure was observed. Parallel printing of multiple materials using high viscosity hydrogels as cell carriers also presents a major drawback related to the shear stress submitted to the cells during the printing process, which is already widely described in the literature [17,46]. Specifically, this issue arises from the need of a printable bioink that can maintain the fiber structure during the printing process until crosslinking, which is typically addressed using high viscosity inks. As a result, when printing, these inks experience high shear stress at the extrusion needle which results in a marked decrease of cell viability [16,47]. On the contrary, if the viscosity of the ink is too low, the spreading of the material after printing makes reproducible and precise printing impossible [48]. However, the engraving method described here overcomes these drawbacks. The engraved grooves serve as a cast where the bioinks can be deposited regardless of their viscosity.

24

The viability of the encapsulated cells after printing was assessed using a Live/Dead staining, which is one of the main approaches commonly used to evaluate the effect of printing on cell survival [16]. The representative images of the staining analysis showed a large number of viable fibroblast cells relative to the number of dead cells were successfully encapsulated in the low viscosity GelMA bioink and printed through a 27G. These results are in agreement with previous studies describing the increase in the cell viability at low bioink viscosities [49]. Furthermore, despite the low concentration of GelMA used in this study, a homogeneous hydrogel was achieved by crosslinking with UV light after printing and remained inside the PCL engraved fibers for at least 24 h of immersion in PBS or culture medium. Despite the small volume of the grooves (0.1 µL per mm length of groove), it was possible to incorporate over 5000 cells per layer using GelMA bioink containing a cell concentration of 106 cells/mL. Since the strands are deposited at a 90º angle with respect to the previous layer, the majority of the hydrogel surface area will be in direct contact with the cell culture medium, allowing the growth and survival of the cells. Moreover, due to the low viscosity of the 7.5 w/v% GelMA bioink, the density of encapsulated cells in the bioink might be significantly increased at later points of culture [50]. The amount of bioink that can be incorporated in the grooves of scaffolds produced using this technique is limited by the engraved volume of the strands. The total cell number incorporated in the constructs is lower compared to methods involving parallel printing of hydrogels and thermoplastic materials, or pure hydrogel scaffolds prepared with GelMA bioinks. Nevertheless, the engraving method enables the preparation of cellularized scaffolds with mechanical properties over three orders of magnitude higher than biodegradable hydrogel scaffolds reported previously [51]. Furthermore, engraving has no effect on the pore size of the

25

scaffolds, which might be required to ensure the scaffolds are compatible with cell infiltration for osteochondral applications [52,53]. Since the developed technology is a layer-by-layer method, it is possible to use the engraving method to precisely control the distribution of cells and growth factors while printing. For instance, we envision the development of a biodegradable scaffold for osteochondral tissue regeneration mimicking the structure and composition of the different regions of the bonecartilage interface, including their cellular and extracellular content by altering the bioink used in each layer. The engraving process may provide a highly innovative method for complex, multicellular tissue regeneration by integrating regional bioactive factors into biomimetic hierarchical architectures.

5. Conclusions A new methodology is reported involving the engraving of printed fibers to exert spatial control over construct microenvironment, cellular organization, and bioactive factor presentation. Specifically, a fiber engraving method allows for high-precision bioprinting of microparticleloaded bioinks as well as low viscosity bioinks encapsulating viable cell populations spatially within 3D printed tissue engineering scaffolds. In summary, this work provides a robust printing methodology for the incorporation of high and low viscosity inks in thermoplastic scaffolds without adversely affecting the mechanical or structural properties of the constructs.

Acknowledgements

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This work was supported by the National Institutes of Health (P41 EB023833). L.D.-G. acknowledges Consellería de Cultura, Educación e Ordenación Universitaria for a Postdoctoral fellowship (Xunta de Galicia, ED481B 2017/063). G.L.K. was supported by the Robert and Janice McNair Foundation MD/PhD Student Scholar Program. M.D. was supported by Rubicon postdoctoral fellowship from the Netherlands Organization for Scientific Research (Project No. 019.182EN.004). References [1]

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Figure captions

Figure 1: Scheme for the preparation of multimaterial scaffolds using the engraving printing methodology. First, high temperature (HT) extrusion (FDM) is used to print the PCL fibers using a 18G needle at 140ºC. After each layer is printed, the fibers are engraved using a 22G needle at 120ºC. Finally, the grooves created in the PCL fibers are filled with a bioink extruded at low temperature (LT) and crosslinked using UV light. Captions show the crosssection of a single fiber at each step of the methodology.

Figure 2: Optical images showing the steps to engrave and fill PCL scaffolds. First, each PCL layer is printed and allowed to solidify. Then, the surface of the PCL fibers is engraved by decreasing the needle offset to the desired depth of the grooves. Finally, GelMA is printed inside the grooves and crosslinked using UV light. Engraved areas are limited by white dashed lines. Scale bar: 1 mm. Image insets show the cross-section reconstruction of a fiber on each step of the engraving method. Scale bar: 400 µm.

Figure 3: (A-D): SEM micrographs showing the surface of the PCL scaffolds engraved with grooves of depths 100 µm (PCL 100 µm; A, C) and 200 µm (PCL 200 µm; B, D). The fidelity of the engraving can be observed in both straight fibers and the corners as well as in the layers below. (E): MicroCT reconstruction of a PCL scaffold with 200 µm engraving depth. Scale bar: 500 µm (A-D) and 250 µm (E).

Figure 4: Epifluorescence images (A, B) and SEM (C, D) micrographs of the PCL 200 µm and PCL 200 µm-E-Shell® fibers. The high fidelity of the low viscosity ink printing allowed for the precise deposition of E-Shell® inside the 200 µm grooves created on the PCL fibers, including inside the curves at the edges of the scaffold. Scale bar: 500 µm (A, B and C) and 250 µm (D).

Figure 5: SEM micrographs (A, B) and microCT reconstructions (C, D) of a PCL-HA30 scaffold, engraved with 200 µm depth grooves, and filled with PCL. A and B show different areas of the surface of a single scaffold layer obtained using SEM. C and D show microCT reconstructions of two different layers of a single scaffold. The precision of the deposition was evident and resulted in the complete and continuous filling of the 200 µm grooves created on the PCL-HA30 fibers independently of the area or the layer of the scaffold. Scale bar: 500 µm.

Figure 6: Compressive modulus of control PCL scaffolds and PCL 200 µm scaffolds. Data reported as means ± standard deviation (n = 5 scaffolds).

Figure 7: (A, B): Confocal micrographs from two different PCL 200 µm scaffolds filled with a 7.5 w/v% GelMA ink containing FITC-albumin-loaded PLGA microparticles. The homogeneous distribution of the microparticles and the accuracy of the printing inside the grooves created on the PCL fibers were revealed. Engraving was outlined by white dashed lines. (C, D): Merged images from confocal microscopy of live/dead staining, revealing cell viability

after 24 h of cell culture in both a straight fiber (left) and in the curve on the edge of the layer (right). Engraved regions are outlined by a white dashed line. Scale bar: 50 µm.

Disclosure statement

Luis Diaz-Gomez, Maryam E. Elizondo, Gerry L. Koons, Mani Diba, Letitia K. Chim, Elizabeth Cosgriff-Hernandez, Anthony J. Melchiorri, and Antonios G. Mikos, authors of the manuscript “Fiber engraving for bioink bioprinting within 3D printed tissue engineering scaffolds” submitted to the journal Bioprinting, declare no conflict of interest.