Synchronized Dual Bioprinting of Bioinks and Biomaterial Inks as a Translational Strategy for Cartilage Tissue Engineering

Abstract

In this study, we propose a novel synchronized dual bioprinting approach that combines two distinct printing strategies for mechanically and biologically improved substitutes for cartilage tissue engineering. Mechanical stability is achieved using a biomaterial ink (polysaccharide based) for microextrusion printing, whereas the cell-loaded functional bioink (type I collagen based) is bioprinted using cell compatible drop-on-demand (DoD) bioprinting. After an exploratory preselection of bioinks with printability and biocompatible characteristics, articular cartilage substitutes have been dually bioprinted using human knee articular chondrocytes (hKACs) from four independent donors. Our findings showed that dual bioprinting improves the compressive modulus from about 20 kPa (only bioink) to up to 600 kPa (bioink plus biomaterial ink). In addition, the biomaterial ink was found to be well interconnected with the bioink, which is important for future translational research. Chondrogenesis of hKACs was possible in functional bioinks with intermediate collagen concentration among the studied blends and satisfactory printability characteristics. We conclude that dual bioprinting is a promising and advanced strategy for cartilage tissue engineering.

Introduction

Damage and loss of articular cartilage lead to focal lesions that disrupt the distribution of loads across the joint causing acute pain and disability.^1–3 The limited spontaneous repair of articular cartilage injury is well documented as it results from alterations in the structural support for chondrocytes and the avascular nature of cartilage.

Despite efforts done toward engineering cartilage substitutes for clinical applications, current outcomes are still not satisfactory due to the production of mixed repair tissue of variable quantity and quality.^4,5 The choice of materials for cartilage and osteochondral tissue engineering is critical and should ideally be favoring mechanical stability in terms of strength, stiffness, and wear resistance, as well as appealing biological characteristics for tissue remodeling after surgical implantation.

There are two major tissue engineering approaches for cartilage tissue engineering up to now: (1) the fabrication of cell-free mechanically stable scaffolds, which can be repopulated with autologous cells either in vitro before implantation (cells are seeded on the materials' surface, 2D) or in situ after implantation^6,7; and (2) the fabrication of cell-loaded substitutes that are biologically functional, that is, showing maturation standards similar to the native tissue achieved in vitro before implantation (cells are embedded in the material, 3D).^8,9

The effect of using 3D chondrocyte culture for predicting the success of tissue maturation when implanted in vivo has been proved numerous times to be more optimal than its analog experimental setup in 2D.^10–12 In this context, 3D bioprinting has emerged in recent years to fabricate biologically functional materials that resemble native tissues.^5,13,14 Bioprinting applied to cartilage tissue engineering enables the patterning of living cells embedded in specific extracellular matrices in a predefined geometry according to the cartilage defect.

Recent studies on bioprinting of cartilaginous tissues are based either solely on the classic microextrusion of biomaterial inks (resembling the mechanical stiffness of subchondral bone such as, for example, polycaprolactone [PCL] and polyethylene glycol), which were postprinting seeded with cell-loaded materials, or based on the combination of two distinct layers (typically made of collagen and calcium phosphate composites) mimicking on the one hand the cartilage tissue and on the other hand the subchondral bone tissue (Fig. 1a).^6,7,15,16 The main advantage of extruded scaffolds as cartilage substitutes compared to the biphasic approach is the fact that functional cell-loaded materials are incorporated into the scaffold as it is fabricated. However, this benefit comes at the expense of the construct's mechanical properties, since the stiffness of nonprinted forms is higher compared with extruded scaffolds.^9,17

FIG. 1.

Dual bioprinting as an optimized fabrication strategy for articular cartilage tissue engineering. (a) Recent studies showed the fabrication of 90° extruded constructs or casting of biphasic scaffolds.^6,7,16 (b) Dual bioprinting combines the dispensing of supporting materials for improved mechanical strength and stiffness (by microextrusion, I) with functional cell-loaded bioinks (by DoD, II) for optimal in vitro tissue maturation. DoD, drop-on-demand.

There are two major advantages of using a drop-on-demand (DoD) bioprinting technique for printing cell-loaded bioinks into precise geometrical constructs in direct comparison to microextruded bioinks as presented in other studies.^10,18,19 The first advantage is the potential for printing freeform 3D constructs with complex geometrical features. In contrary, microextruded printing is limited to the extrusion of hydrogel strands in a continuous manner and mostly limited to the printing of 90° flat structures. The second advantage is increased cell viability after bioprinting using a DoD technique instead of microextrusion, which applies uncontrolled shear stress directly to the cells.²⁰ In previous studies we have shown that our cell-loaded bioprinting strategy does not negatively affect cell viability.^21,22

In this study, we propose a novel synchronized dual bioprinting approach that combines two distinct printing strategies for a mechanically and biologically improved cartilage substitute (Fig. 1b). Increased mechanical stability and stiffness are achieved using a biomaterial ink (polysaccharide based) for microextrusion printing, whereas the cell-loaded bioink (type I collagen based) is bioprinted using cell compatible DoD bioprinting. After an exploratory preselection of optimal printing and cell-compatible bioinks, articular cartilage substitutes have been dually bioprinted using human primary chondrocytes.

Experimental Section

Synthesis of bioinks and biomaterial inks

Cell-free biomaterial inks and bioinks were synthesized according to previously published protocols.^{17,20,23–25} Stock solutions of low gelling temperature agarose (Sigma, Hamburg, Germany) of 6 w/v % were produced by dissolving the agarose powder in deionized water and sterilized at 121°C for 15 min. Stock solutions of alginic acid from brown algae (Sigma) of 6 w/v % were prepared by dissolving alginic acid powder in 0.2 μm sterile-filtered deionized water overnight with constant stirring. Biomaterial inks for microextrusion printing were prepared by combining equal parts of agarose and alginic acid solutions using a positive displacement pipette for obtaining the following final blend concentrations: (1) 1.5 w/v % agarose with 3 w/v % alginate, (2) 3 w/v % agarose with 1.5 w/v % alginate, and (3) 3 w/v % agarose with 3 w/v % alginate. Bioinks compliant with chondrocyte encapsulation and differentiation were prepared by mixing 0.3 w/v % type I collagen solution (Biochrom, Berlin, Germany) with 3 w/v % agarose stock solutions at predefined varying concentrations and kept at 37°C in a water bath until being mounted in the prewarmed bioprinter head. Pure 0.3 w/v % collagen working solutions were prepared according to the manufacturer's instructions. Briefly, eight parts of collagen G (Biochrom) were mixed with one part 10-fold DMEM and one part deionized water. The final 0.3 w/v % collagen working solution was neutralized with 2 v/v % 1 M NaOH. The following agarose–collagen blends have been used as bioinks: 1.0 w/v % agarose with 0.1 w/v % collagen and 0.5 w/v % agarose with 0.2 w/v % collagen.

Design, slicing, and printing of bioinks and biomaterial inks

Dually printed samples have been designed with computer-aided design (CAD) software (Inventor, Autodesk, San Rafael). Hollow cylinders with three supporting pins and their negatives have been designed and afterward sliced using a G-code generator for 3D printing (Slic3r, GNU Affero General Public License). The hollow cylinder with three pins was printed with a 3D extruder (Biobot1; Biobots, Philadelphia, PA) that served as the sample's biomaterial ink, whereas its negative part was printed DoD using a custom-made 3D bioprinter,²⁰ which represented the bioink compatible with chondrocyte encapsulation and differentiation. The custom-made 3D bioprinter consisted of a 3-axis robotic platform (ISEL, Eichenzell, Germany) coupled with one print head heated to 34°C. A pressurized air supply (0.5 bar) was mounted to the print head. The printhead included an electromagnetic microvalve with a 300 μm nozzle diameter (Fritz Gyger, Gwatt, Switzerland). Interspersed bioink and biomaterial ink were printed in each layer using two individual printer heads with distinct printing strategies (DoD and microextrusion, Fig. 1). The printing temperature of the biomaterial ink was kept at 30°C, and the cell-loaded bioink was kept at 34°C. These temperatures were selected for being above the gelling temperatures of the bioink and biomaterial ink (above 23°C based on our previously published findings²³) to keep both inks liquid inside the respective reservoir of the printhead during the printing process. As mammalian collagen polymerizes at 37°C, the collagen-based bioinks were kept at temperatures below 37°C (respectively at 34°C). The printing of each layer of the biomaterial ink took ∼30 s, whereas the printing of each layer of the bioink took ∼60 s. Twenty layers of both biomaterial ink and bioink were printed per construct.

Analysis of the contact surface between bioinks and biomaterial inks

Due to the fact that two distinct printing strategies with two different printing materials have been used for the same cartilage substitute, it was necessary to evaluate the contact surface between both materials after printing. Thus, casted 5 × 10 mm cylinders composed of both material types were prepared with a single contact surface for scanning electron microscopy (SEM). Samples were fixed in 2 v/v % glutaraldehyde and 0.1 M Sorenson's buffer (pH 7.4) for 24 h. Thereafter, samples were dehydrated in acetone and critical point dried (E-300 Critical Point Dryer, Polaron Equipment, London, United Kingdom). Before imaging samples were coated with a thin gold layer. Imaging was done with a SEM (ESEM XL 30 FEG, FEI, Philips, Eindhoven, The Netherlands).

Mechanical characterization of dually printed samples

For evaluating the mechanical stiffness of printed samples, a universal testing machine was used (MiniZwick 2.5, Zwick, Ulm, Germany). Samples were compressed at a cross-head speed of 4 mm/min until rupture, as previously described.²³ Compressive tangent moduli were calculated from stress/strain curves at defined strains between 5% and 20%.

Pellet and cell-loaded hydrogel culture

Human primary knee articular chondrocytes (hKAC) were isolated from knee joints of patients undergoing total knee arthroplasty upon written consent and allowance from the local ethics committee (EK 305/13). Briefly, articular cartilage was separated from the bone and cut in 2 × 2 mm² pieces. Only macroscopically healthy cartilage removed from femoral condyles was used. The pieces of cartilage were enzymatically digested with 50 mg/mL collagenase (Roche, Mannheim, Germany) dissolved in GlutaMAX medium (Life Technologies, Carlsbad, CA) at 37°C and 5% CO₂ under constant stirring. After 24 h incubation, the digested cartilage solution was filtered using a 100 μm nylon sieve (BD Falcon, Franklin Lakes, NJ), washed with PBS, and centrifuged thrice at 1500 rpm for 10 min. The cells were resuspended in 1 mL medium (GlutaMAX; Life Technologies) supplemented with 10% FCS (PAN, Aidenbach, Germany), 1% gentamicin (Life Technologies), 1% penicillin and streptomycin (PAA, Cölbe), and 0.5% fungizone amphotericin B (Life Technologies). Medium was changed twice a week, and hKACs were expanded in culture until use for pellet and cell-loaded bioprinting experiments (Supplementary Fig. S1). Pellet cultures of hKACs were prepared by seeding 250,000 cells per well (conic 96-well plates), centrifuging at 300 rcf for 5 min, and adding either expansion medium (controls) or chondrogenic differentiation medium (Stem Cell Technologies). Pellets were cultured at 37°C and 5% CO₂ for 21 days. Cell-hydrogel cultures were prepared by embedding 2 × 10⁶ cells per mL of hydrogel both for controls and chondrogenic samples. Selected hydrogels for hKAC culture had the following final concentrations: (1) 0.3 w/v % collagen, (2) 1 w/v % agarose with 0.1 w/v % collagen, and (3) 0.5 w/v % agarose with 0.2 w/v % collagen. All three tested cell-loaded hydrogels (bioinks) had the same initial cell density (2 × 10⁶ cells/mL). After preparation of cell-loaded hydrogel samples, an incubation time of 21 days took place at 37°C and 5% CO₂ with medium changes thrice a week. Selected cell-loaded 0.5 w/v % agarose with 0.2% collagen hydrogels for bioprinting experiments was prepared in the same way as pilot cultures, mounted in the printer head of the 3D bioprinter, and printed DoD as the bioink. After printing, samples were incubated for 21 days in chondrogenic medium (Stem Cell Technologies) with medium changes thrice a week.

Chondrocyte viability after DoD bioprinting

Cell viability was assessed by live/dead staining immediately after, 1, and 7 days after bioprinting. Live/dead staining (5% fluorescein diacetate (FDA) and 5% propidium iodide (PI) in Ringer's solution, all from Sigma) was added to each sample, and samples were imaged with a fluorescent microscope (Zeiss Imager.M2m, Germany). Cells from three independent donors (n = 3) were used for the viability test (Supplementary Fig. S2).

Histology and immunohistochemistry of articular cartilage, chondrocyte pellets, and cell-loaded hydrogels

After the culture periods, pellets, cell-loaded hydrogels, and printed samples were fixed in 4 v/v % paraformaldehyde (PFA) for either 30 min (pellets) or 1 h (hydrogels). Human articular cartilage was cut in small pieces and fixed overnight in osteosoft (Merck Millipore, Burlington, VT). Thereafter, samples were washed once with PBS. Pellets were embedded in 3 w/v % agarose and transferred to paraffin embedding cassettes, and cell-loaded hydrogels and cartilage samples were transferred directly to cassettes. All samples were submerged in 70% ethanol until automated dehydration has started. After this overnight step, samples were embedded in paraffin, sliced with 5 μm thickness, and stained histologically or with specific antibodies. The histological slices were stained with hematoxylin and eosin for morphological evaluation as described elsewhere²¹ and with toluidine blue staining for detection of proteoglycans. Briefly, toluidine blue stainings were done after routine slide deparaffinization, washing with distilled water, and stained with acidic Toluidine Blue solution (pH 2.0–2.5) for 2 min. Then, slides were washed thrice with distilled water, transferred to ascending alcohol series, and mounted with vitro clud. Immunocytochemical stainings were applied after fixation with PFA using aggrecan as primary antibody (ACAN, 1:100, mouse monoclonal; Abcam, Cambridge, United Kingdom) and Alexa 594 as secondary antibody (goat anti-mouse IgG; Abcam). Cell distribution inside the functional hydrogel blends was visualized with phalloidin actin (Alexa 488, 1:50; Life Technologies, Carlsbad, CA).

Analysis of GAG production

After incubation, pellet and hKAC-loaded samples were collected from culturing vessels, transferred to 1.5 mL reaction tubes, and incubated with 125 μg/mL papain digestion buffer (Sigma-Aldrich, St. Louis, MO) at 60°C for 24 h. Enzymatic reaction was stopped by increasing the incubation temperature to 70°C for 10 min. Samples were homogenized, and 100 μL were transferred to new reaction tubes. The quantitative detection of total glycosaminoglycans (GAGs) was done using a colorimetric assay according to the manufacturer's instructions (Blyscan, Biocolor, Carrickfergus, United Kingdom). Briefly, 1 mL Blyscan dye reagent was added to each sample and incubated for 30 min at room temperature. Thereafter, reaction tubes were centrifuged at 12,000 rpm for 10 min, and excess liquid was carefully drained. Five hundred microliters of dissociation reagent were added to each tube; samples were vortexed and centrifuged at 12,000 rpm for 5 min to remove air bubbles. The colorimetric intensity of samples was measured at 656 nm, and sGAG concentrations were calculated from the standard curve. To normalize the GAG content to the DNA, a DNA Quantification Kit was used according to the manufacturer's instructions (Quant-iT PicoGreen ds DNA Assay Kit; Life Technologies). Final GAG and DNA concentrations were calculated in ng/μL.

Statistical analysis

Data from the mechanical compression tests and biological characterization were analyzed using Student's t-test, and statistical significance was defined as *p < 0.05 and **p < 0.01.

Results

CAD and slicing software for synchronized dual bioprinting

The proof-of-concept study of dual printing was done using two distinct printing strategies. On the one hand, more viscous and stiffer cell-free biomaterial inks were extruded with a 27 G needle in a cylindrical shape as the contour of the scaffold and in smaller three-pin combinations useful for improved mechanical stiffness after future implantation. On the other hand, cell-loaded collagen-based bioinks were printed using an electromagnetic microvalve under a constant pressure of 0.5 bar. Stereolithographic (STL) data were derived from CAD cylindrical constructs as shown in Figure 2. Each utilized printing strategy had a corresponding CAD, being the DoD design the negative counterpart of the microextrusion design. It was feasible to slice STL files for extrusion printing using Slic3r; however, STL files used for DoD printing were sliced using a custom-made Java-based slicing software. Using DoD for cell-loaded materials has proved to be more spatially precise than using a microextrusion technique.²⁶ In addition, DoD printing did not interfere with previously extruded material, as it would have been the case if the whole construct was solely extruded (the movement and touch of the syringe during printing could have destroyed the bottom printed layers). Finally, it was feasible to dually print the distinct materials using two printing strategies and printer heads, as shown in Figure 2, without losing printing precision compared to singularly printed materials.

FIG. 2.

CAD and slicing software for synchronized dual bioprinting. Each printing strategy (microextrusion and DoD) needed a distinct slicing software for optimized printing precision. Both supporting and functional materials were printed individually (single printing) or in combination (dual printing). The diameter of the dually printed scaffold was 16 mm. CAD, computer-aided design.

Bioink–biomaterial ink interface after dual bioprinting

As the printing of two distinct materials could have originated additional challenges regarding the adhesion and connectivity between both printed parts, scanning electronic micrographs at the interface site were analyzed (Fig. 3). It was possible to observe an interconnection between the polysaccharide-based biomaterial ink with the collagen-based functional bioink. Interestingly, at the interface line collagen fibers were elongated toward the polysaccharide phase and vice versa.

FIG. 3.

The interface between supporting material (combination of agarose and alginate) and functional bioink (combination of agarose and type I collagen). Collagen fibers elongated toward the supporting material phase, whereas polysaccharide chains extended toward the bioink.

Mechanical properties of dually bioprinted cartilage substitutes

The aim of this study was on the one hand to improve the biological functionality of printed scaffolds as cartilage substitutes by including chondrocyte-loaded hydrogels during bioprinting (bioinks) and on the other hand to improve the mechanical stiffness of the final construct by combining DoD printing with microextrusion printing of stiffer cell-free materials. For this purpose, combinations of agarose and alginate hydrogels were printed in proof-of-concept cylindrical shapes and compressed using a universal testing machine (Fig. 4). It was noted that by increasing the solid concentration of biomaterial inks it was possible to fine-tune the mechanical stiffness of the whole dually bioprinted construct. The highest compressive modulus value, corresponding to 3 w/v % agarose with 3 w/v % alginate, was in the same range as native human articular cartilage (∼700 kPa).²⁷

FIG. 4.

Compressive tangent moduli of printed scaffolds. Bioprinted bioink (functional material) scaffolds are significantly less stiff than dually bioprinted scaffolds. Higher solid concentration of agarose and alginate blends resulted in improved mechanical stiffness of dually bioprinted scaffolds (1.5%AG-3%ALG+Bioink <3%AG-1.5%ALG+Bioink <3%AG-3%ALG). Dually bioprinted samples of 3%AG-3%ALG+Bioink showed the highest compressive tangent modulus among the tested samples at a compressive strain of 20% (p < 0.05 in comparison with 3%AG-1.5%ALG+Bioink for n = 3). This peaking result is in the same range as the compressive stiffness of native articular cartilage.²⁷ * Indicates statistical significance, p < 0.05.

Choice of bioink for bioprinting and chondrogenic tissue maturation

Recently, we have shown the effect of the 3D matrix on cell shape and fate.^17,23 For a successful chondrogenic differentiation it was necessary to test different combinations of collagen-based materials as functional parts of the printed scaffold. The first screening qualitative analyses were performed on toluidine stained histological slices (Fig. 5). It was possible to observe a marked proteoglycan staining (purple color) in pellet cultures (positive control) followed by cultures in pure type I collagen hydrogels (COL +++). The success of chondrogenic differentiation was supported by a quantitative GAG assay normalized to the DNA content in each sample. The highest agarose concentration in the agarose-collagen blend (COL +) resulted in the least noticeable proteoglycan staining and lowest measured GAG/DNA ratio. Interestingly, intermediate agarose and collagen concentration in the blend (COL ++) showed in-between chondrogenic differentiation levels among the tested hydrogel blends (n = 4). For this reason, COL ++ was the chosen blend as functional material (bioink) for further dual bioprinting experiments.

FIG. 5.

Chondrogenic differentiation potential in hydrogel blends with varying mixing ratios. (a) After 21 days in culture, KAC pellets showed marked proteoglycan expression (purple color). Followed by the pellet cultures (positive controls), KAC-loaded collagen gels were the samples with the highest proteoglycan expression. (b) In addition, quantitative GAG expression was the highest in pure collagen hydrogels (COL +++), followed by intermediate agarose-collagen blends (COL ++) and blends with the highest agarose concentration (COL +). Data presented as mean ± standard deviation using primary hKACs from four independent donors (n = 4). Statistical significance was obtained using Student's t-test, and it is marked as *p < 0.05 and **p < 0.01. hKAC, human knee articular chondrocyte.

Chondrogenic differentiation and in vitro maturation of dually bioprinted samples

After selecting the most suitable functional material as bioink for KAC encapsulation and differentiation (COL ++), dual bioprinting experiments were performed with KAC. After printing and incubating in vitro for 21 days, cell-loaded specimens were histologically and biochemically evaluated. The cells contained in the bioprinted samples were homogeneously distributed as shown by phalloidin actin staining and imaged with two-photon microscopy (Fig. 6 ACTB). In addition, the presence of aggrecan (ACAN) was confirmed histologically (Fig. 6 ACAN), and the production of proteoglycans was observed by Toluidine Blue staining at the interface between the supporting and functional dually bioprinted materials (Fig. 6 Toluidine Blue). In addition, GAG expression was equivalent in samples with casted and bioprinted functional bioinks containing the cells (Fig. 6, graph). This result is of importance because it shows that the dual bioprinting process does not negatively affect chondrogenesis. The stability of dually bioprinted samples was also confirmed by microcomputed tomography (μCT, Fig. 6), where it was possible to notice a preserved initial printed structure (marked by brighter and darker tones for supporting and functional printed materials, respectively) as proposed in Figure 2.

FIG. 6.

Dually bioprinted chondrocyte-loaded samples fabricated with the selected COL ++ (intermediate agarose-collagen blend) as functional bioink (bioprinted by DoD) and 1.5% agarose with 3% alginate blends as supporting material (3D printed by microextrusion). After 21 days of incubation in CIM, it was possible to observe the presence of collagen type I fibers in the functional material by SHG imaging, positive actin staining for chondrocytes distributed homogeneously inside the functional material (COL ++), and positive ACAN and toluidine blue stainings after immunohistological and histological preparation of the samples. Quantitative GAG expression was equivalent for casted and bioprinted functional materials, therefore showing that the dual bioprinting process is not affecting negatively the chondrogenesis of cells after bioprinting (p > 0.05 using primary hKACs from four independent donors and printing experiments, n = 4). Dually bioprinted samples were additionally imaged by μCT to confirm the stability of the constructs (brighter areas correspond to the supporting material, whereas darker areas correspond to the functional material). CIM, chondrogenic induction medium; SHG, second harmonic generation.

Discussion

For the first time it was shown that it is possible to dually bioprint 3D constructs with improved mechanical and biological characteristics provided by the presence of collagen for cartilage tissue engineering. The most relevant steps of this work were (1) the synchronized printing of two distinct materials providing both mechanical and biological enrichment to the scaffold, (2) successful chondrogenic differentiation in printable hydrogel blends that were used to carry the chondrocytes, and (3) improved mechanical stiffness and chondrogenic differentiation after bioprinting compared to control samples only, including cell-carrying bioinks.

Recent studies have shown hydrogel-based fabricated constructs as articular cartilage substitutes with great mechanical properties.^28,29 However, cell-carrying materials printed alongside cell-free supporting materials like PCL were confined to the space in between printed rods of the supporting material. As a consequence, nutrients were scarce to the cells, and the overall stiffness of printed constructs was in a range above the stiffness of native articular cartilage tissue.²⁸ It is well accepted that articular cartilage substitutes with mismatching mechanical properties to the adjacent tissue can lead to tissue degradation.³⁰ Thus, dually bioprinted cartilage constructs are an alternative to the former scaffolds, since their stiffness can be tuned in a range between ∼100 and 600 kPa. This finding may be used as an asset for translational purposes, due to the fact that articular cartilage injuries aren't limited to one specific area of the joint. Since the mechanical properties of articular cartilage in the knee joint slightly vary, depending if the region of interest is covered by the meniscus or not, the mechanical stiffness of such dually bioprinted substitutes could be adjusted depending on the patient's needs.³¹ Moreover, the advantage of our proposed dual bioprinting strategy compared to other similar studies, such as presented by Melchels et al., is that it utilizes biomaterial inks with thermal and ionic polymerization mechanisms, which are more cell friendly to the nearby cell-loaded bioinks compared to other gelling mechanisms like photo crosslinking (e.g., for supporting hydrogels like poloxamer 407).¹⁸ In our present study, we showed that human KACs produce GAGs in an equivalent manner after the bioprinting process compared to casted controls, indicating that cells are not only surviving the printing but also able to maintain their chondrogenic phenotype.

A further key finding of this study was the interconnecting interface between bioink and biomaterial ink using dual bioprinting, contrarily to the typical nonadherent interface between PCL and other bioinks.¹⁰ Considering that this study proposes an optimized dual bioprinting method that is based on two distinct strategies, which in turn are dedicated to print materials with varying rheological properties, it was necessary to show that two individual materials were well connected at their interface. Scanning electron micrographs at the interface between both materials showed that collagen fibers (bioink) elongated toward the polysaccharide-based biomaterial ink and vice versa. These results are encouraging for future in vivo testing, which will require stable 3D constructs for ectopic and later entopic surgical implantations.

The selection of the optimal bioink for this work has required on the one hand that this material is printable with a DoD nozzle and on the other hand that it supports chondrogenesis. We have shown in recent studies that agarose-collagen type I blends are suitable bioink candidates for bioprinting and tissue engineering applications and that agarose-containing gels are printable as 3D constructs.^17,23 Thereby, hKAC chondrogenesis has been tested in three bioinks with varying concentrations: COL +++ (0.3 w/v % collagen, nonprintable control blend for chondrogenesis with the highest collagen concentration), COL ++ (0.5 w/v % agarose with 0.2 w/v % collagen, blend with intermediate concentration of collagen and intermediate printability characteristics), and COL + (1.0 w/v % agarose with 0.1 w/v % collagen, blend with the lowest collagen concentration and best printability characteristics). Our results have shown a significant increase in GAG production in the three blends cultured in chondrogenic induction medium (CIM) compared to control groups cultured in expansion medium. Moreover, GAG production was the highest in the blends with the highest collagen concentration cultured in CIM, followed by intermediate blends and blends with the lowest collagen concentration. In addition, proteoglycan histological staining with Toluidine Blue showed a similar trend for chondrogenesis in the blends. Altogether, preliminary evaluating studies for chondrogenesis have shown that there was a compromise in the hKAC chondrogenic potential when combining agarose with collagen type I hydrogels serving as bioink. For these reasons, further dual bioprinting experiments were performed with COL ++ as bioink, for proving its advantages with regard to both printability and aiding hKAC chondrogenesis.

Following the preparatory studies on dual bioprintability and chondrogenic potential in functional materials, dually bioprinted constructs were prepared using hKACs from four independent donors and incubated for 21 days. Three main findings were considered in this experimental setup. First, chondrocytes were homogeneously distributed in the dually bioprinted samples after the incubation period, as shown by two-photon microscopy and phalloidin actin staining. The chondrocytes maintained their original round shape inside the functional materials, which is important for their chondrogenic differentiation and to avoid that they form fibrotic tissue (e.g., in case chondrocytes would elongate inside the hydrogel blend).³⁰ Second, hKACs in casted and dually printed constructs showed equivalent GAG expression, thereby demonstrating that the printing process is not negatively influencing their chondrogenic differentiation. This finding is a major aspect of this work, since dually bioprinting is an important and necessary tool for precise spatial positioning of cells and materials for cartilage tissue engineering applications. Recent studies have shown that the bioprinting process can strongly affect cell viability, and therefore, it is important to control the shear stress levels at the printing nozzle.²⁰ Third, dually bioprinted constructs showed a stable and preserved initial printed shape after the incubation time as shown by μCT imaging. Together with the scanning electron micrographs that revealed a good interconnectivity between the two printed materials, μCT images showed maintained appearance of the constructs after the printing process and incubation period. As previously justified, these encouraging results are ideal for future in vivo testing experiments.

In future studies, dual bioprinting will be used for creating more complex articular cartilage substitutes with matching sizes and shapes to that of the original defects measured in the patients. Therefore, future studies will focus on improving the printing resolution, for example, by incorporating different printing scales (centimeter, millimeter, and micrometer) within the same bioprinting platform. Furthermore, the bioprinted constructs will be cultured in dynamic bioreactors with controlled levels of oxygen to provide ideal conditions for the production of cartilaginous-rich matrix in mechanically stable constructs.

Conclusions

In this study, we introduced a synchronized dual bioprinting approach that combines two distinct printing strategies for a mechanically and biologically improved cartilage substitute. After an exploratory preselection of optimal printing and cell-compatible bioinks, articular cartilage substitutes have been dually bioprinted using primary chondrocytes from four independent donors. Our findings revealed that dual bioprinting improved the stiffness from about 20 kPa (only bioink) to up to 600 kPa (bioink plus biomaterial ink). In addition, the biomaterial ink was found to be well interconnected with the bioink, which is important for future translational research. Chondrogenesis of hKACs was possible in bioinks with intermediate collagen concentration among the studied blends and satisfactory printability characteristics. Altogether, dual bioprinting is a promising strategy for cartilage tissue engineering and can be further optimized using either improved supporting printable materials or smartly tuned functional materials.

Footnotes

Acknowledgments

This research project is supported by the START-Program of the Faculty of Medicine, RWTH Aachen. In addition, this work was jointly supported by the Core Facilities “Two-Photon Imaging” and “Immunohistochemistry” of the Interdisciplinary Center for Clinical Research (IZKF) Aachen within the Faculty of Medicine at RWTH Aachen University. The authors thank Sophie Lecouturier from the Department of Orthopedics and Roswitha Davtalab from the Department of Dental Materials and Biomaterials Research at the RWTH Aachen University Hospital for isolating knee articular chondrocytes used for the bioprinting experiments.

Author Disclosure Statement

No competing financial interests exist.

Supplementary Material

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