An interfacial self-assembling bioink for the manufacturing of capillary-like structures with tuneable and anisotropic permeability (2024)

We conducted experiments and simulations to characterize bioink shape fidelity as determined by the diffusion-reaction mechanism triggered upon ELK1-in-GO printing. First, an 18 µl aqueous solution of ELK1 (2 wt%) was manually injected into an aqueous suspension of GO (0.10 wt%, 100 µl) in a 96-well plate to generate a vertical tube. The process was video recorded (figure S1) and imaged at different time-points to investigate the kinetics of tube wall formation (figure 2(a)) and movie S2). Separately, a 'transport of diluted species' COMSOL model simulating similar experimental conditions was generated and used to validate the experimental results. Both experimental and simulation results were in agreement and revealed that the ELK1-GO tube exhibited a similar diameter as that of the starting ELK1 injected drop and the ELK1-GO tube wall thickness did not increase after ∼60 s of interaction (figure 2(a), figure S5). These results demonstrate that the initial volume of the ELK1 injection determined the outside dimension of the final printed structure (figure 2(a)) and that this dimension was not affected by the growth of the ELK1-GO interfacial tube wall as a result of its diffusion-reaction process.

The printability of the ELK1-GO bioink was characterized using a commercially available 3D printer (3D Discovery^™, RegenHU, Switzerland) with adjustable nozzle heights and diameters, dispensing pressures, and printing speeds [31]. By fixing nozzle height and diameter, we assessed the relation between dispensing pressure and printing speed. The fidelity of fabricated ELK1-GO tubes was assessed by optically and systematically measuring tube diameters at different locations [32], which revealed increased printing fidelity with increasing printing speed (figures 2(b) and (c)). Larger diameter differences within a single tube led to larger variances, which resulted in rugged tube walls (figure 2(b), printing speed: 10 mm s⁻¹). By varying the dispensing pressure of the ELK1 solution between 0.02 and 0.10 MPa and the printing speed (PS) between 10 and 100 mm s⁻¹ (nozzle height and diameter were fixed, figure 2(b), table), we found that tubes were generated when the dispensing pressure was between 0.04 and 0.08 MPa. When printing within this pressure range and printing speeds above 30 mm s⁻¹, variances in tube wall geometry within a single tube were under 5%, which reflected high fidelity tubes with smooth walls (figure 2(b)). Utilizing a dispensing pressure of 0.06 MPa and printing speed of 30 mm s⁻¹, we were able to print tubes of up to 12 cm in length and 2 mm in diameter. It is important to mention, however, that the limitation on tube length was not lack of tube integrity but rather the size of the plate used to hold the GO solution during printing.

Printing fidelity was also characterized by varying the concentration of the GO suspension (figure 2(c)). In this case, lower concentrations of GO (0.05 wt%) generated tubes with lower printing fidelity and thinner tube walls (∼40 µm), while higher GO concentrations (0.10 and 0.15 wt%) resulted in tubes with higher printing fidelity and thicker walls (∼60 µm) (figures 2(d) and (e)). We speculate that this difference in wall thickness results from interfacial membranes forming with varying levels of permeability as a result of having different concentrations of GO. We have demonstrated that upon initial ELK1-GO co-assembly, a diffusion barrier is formed, which defines its downstream diffusion-reaction mechanism that generates the distinctive multilayer membrane [29]. We reason that lower concentration of GO (0.05 wt%) leads to a more permeable diffusion barrier, which allows for higher and faster ELK1 diffusion and consequently thinner multi-layered membranes. In contrast, higher concentration of GO (0.15 wt%) upon co-assembly results in the formation of a less permeable diffusion barrier, which results in thicker membranes.

Having characterized printing reproducibility and fidelity, we then focused on exploiting the co-assembling nature of the bioink to control ELK1-GO tube wall porosity by varying the concentration of GO. Tubes were first co-assembled with varying GO concentrations (0.05 wt%, 0.10 wt%, and 0.15 wt%) and cut to extract flat membranes (the tube wall) (figure S1(a)). The luminal (inner) and abluminal (outer) sides of these membranes were then examined by SEM following standard preparation protocols [29]. As expected, increased GO concentrations led to decreased porosity, which was easily noticeable on both inner and outer surfaces of the membranes (figure 3(a)). Upon closer inspection, significant differences in porosity between the inner and outer surfaces were observed on all membranes, independently of GO concentration (figure 3(b)), which was also confirmed by histological sectioning (figure 3(d)). These results demonstrated the possibility to modulate the ELK1-GO tube wall pore sizes from 0.5 to 1.2 µm on the inner surface and 1.6–4.0 µm on the outer one by modifying GO concentrations (figure 3(c)).

Together with the previous results of printing fidelity and tuneability of geometry and porosity, we can conclude that higher GO concentration and printing speed (50 mm s⁻¹ < PS < 70 mm s⁻¹) can be used to fabricate higher fidelity tubes with smooth and thicker (∼60 µm) fully formed tube walls but less porosity after 60 s of co-assembly. By contrast, lower GO concentration and printing speed (30 mm s⁻¹ < PS < 50 mm s⁻¹) would generate lower fidelity tubes with rugged and thinner (∼40 µm) tube walls exhibiting higher levels of porosity. Importantly, upon ELK1-GO co-assembly, the thickness of the resulting interfacial structures depends on the time-dependent diffusion-reaction process of formation (figure 2(a)). Consequently, given the high ELK1-GO affinity, it is possible to stop the co-assembly process at different time points, resulting in interfacial membranes of different thicknesses. We found that it was possible to stop the co-assembly process after just 15 s by removing the fabricated tube from the GO suspension, resulting in tube wall thicknesses down to ∼2 µm when printing at >70 mm s⁻¹. Furthermore, tube wall thicknesses were observed to reach a maximum of  ∼20 µm in thickness (figure 2(c)) after ∼60 s of co-assembly, which did not seem to change after longer periods of time. We speculate that these thin tube walls result from the low ELK1 volumes released upon printing at fast printing speeds (>70 mm s⁻¹), offering an opportunity to reproducibly fabricate with resolutions down to a few microns in size (figures 1(c) and (d)). In summary, these results demonstrate how bioprinting fidelity and resolution can be tuned through both printing parameters from the top-down and self-assembling events taking place from the bottom-up.

The possibility to control tube wall porosity opens an opportunity to fabricate membranes, capillaries, and other structures with tuneable permeability that could be designed to recreate that of biological tissues. For example, proper vascular permeability is of outmost importance in many biological processes such as regulating diffusion of nutrients and waste products, serving as a semipermeable barrier between blood contents and tissues to screen toxic molecules, and regulating the microenvironment of the ECM [33]. Furthermore, diffusion across vascular and lymphatic tissues is vital to enable appropriate filtration, protection, and delivery of key solutes, cells, and factors. The capacity to develop vascular structures or fluidic devices with tuneable permeability is critical to recreate and optimize biofunctionality. To assess the effect of tube wall porosity on its permeability, we designed a simple capillary-based permeability testing device where the membrane (cut from the ELK1-GO tube) can be positioned in the middle, leaving two free chambers on either side (figure S2(a)). The chambers can be filled with any solution, enabling diffusion of solutes from one side to the other (figure S2(a)). We tested the permeability of the membrane to fluorescein sodium salt and 20 kDa FITC-dextran in water solution to recreate physiological charge and size barriers, respectively. These two functions are normally observed in vascular tissue [33]. Using epifluorescence microscopy imaging, permeability constants were calculated based on the fluorescence intensity profiles recorded on the chamber where the solutes diffused to (figure 4(a)). Specifically, exponential curves describing the trend of the average intensity within the chamber as a function of time were calculated for each ELK1-GO membrane and the permeability constant measured based on a diffusion constant using a fitting plot [34] (figure 4(b)).

To confirm these results, we established a 'transport of diluted species' COMSOL model to predict the permeability of sodium salt through the membrane (figure 4(a)). Similar to the experimental setup, sodium salt exhibited faster diffusion than 20 kDa FITC-dextran (figure 4(c)) with a permeability constant of ∼5 × 10⁻⁵cm s⁻¹. Furthermore, as expected, we found that the permeability constants decreased with increasing concentration of GO for both fluorescein sodium salt (from 1 × 10⁻⁵ cm s⁻¹ to 5 × 10⁻⁵ cm s⁻¹) and 20 kDa FITC-dextran (from 2 × 10⁻⁶ cm s⁻¹ to 58 × 10⁻⁶ cm s⁻¹) (figure 4(c)).

These results demonstrate the potential of the ELK1-GO bioink to fabricate membranous and tubular structures while enabling regulation of permeability. By simply modulating GO concentration, it is possible to fabricate complex structures made from thin membranes with permeability constants comparable to those of the native retina (∼2 × 10⁻⁵ cm s⁻¹) [35], lymphatic-vessels (∼6.4 × 10⁻⁶cm s⁻¹) [36], an hUVEC-containing Transwell (∼5 × 10⁻⁶ cm s⁻¹) [37], and organ-on-chip devices of the blood-brain-barrier (∼5 × 10⁻⁵ cm s⁻¹) [38].

In our previous study, we demonstrated how the ELK1-GO self-assembling system can be incorporated within liquid-in-liquid printing to biofabricate tubular structures comprising cells exhibiting high viability and growth for at least 7 d [29]. To demonstrate the functionality and biological relevance of the bioink, we conducted similar permeability tests but after endothelialisation of the lumen of the ELK1-GO tubes. This setup enabled us to better account for the localization of the cells and their effect on tube-wall permeability. Different densities (5 ×10³, 1 ×10⁴, 5 ×10⁴, or 1 ×10⁵ ml⁻¹) of hUVECs were seeded on the luminal side of membranes (cut from ELK1-GO tubes) (0.10 wt% GO) placed on Petri dishes (figure S2(b), top). We chose the method of seeding hUVECs after tube formation instead of directly co-assembling with cells to more precisely account for the localization of the cells, their interaction, and the resulting effect on permeability. As expected, after 48 h of culture, membranes exhibited different hUVEC densities including ∼30%, 50%, 80%, and >95%, as measured by confocal microscopy imaging (figure 5(a), confocal images). Then, these membranes were inserted into the permeability test device with hUVECs facing one of the chambers (figure S2(b), bottom). Then, 40 kDa FITC-dextran [39] in phosphate-buffered saline (PBS) was flowed into the chamber facing the cells while PBS alone was flowed into the other. Here, we used 40 kDa FITC-dextran instead of the 20 kDa FITC-dextran as it is widely used to test biomaterial permeability [39]. The results demonstrated that membrane permeability tended to decrease with increased cell density, with membranes without cells (control) exhibiting the highest permeability constant (2.2 ×10⁻⁶cm s⁻¹) while those with the highest cell density exhibited the lowest permeability constant (6.5 ×10⁻⁷cm s⁻¹) (figures 5(b) and (c)). In conclusion, the presence of hUVECs growing on the ELK1-GO tube wall significantly influenced the transport of solutes in a way that is dependent on the level of hUVEC confluency. We speculate that lower cell densities led to lower number of endothelial cell junctions, which in turn led to higher permeabilities. These results further support the possibility to tune structure permeability and, in particular, to do it through the presence of cells. For example, the permeability constant of ELK1-GO-hUVEC membranes (∼1 ×10⁻⁶ cm s⁻¹) with confluence levels above 50% were comparable with that of tumour micro vessels (1.8 ×10⁻⁶ cm s⁻¹) [40].

The possibility to rapidly and controllably engineer living capillary- and vascular-like structures remains a major unmet challenge in biofabrication, tissue engineering, and in vitro models. Feng et al have demonstrated how liquid-in-liquid printing can be used to fabricate perfusable tubular structures [24] but the system relies on aqueous nanoclay dispersions and solutions of H2N-PDMS-NH2 in toluene, which prevents the use of cells and recreation of key anatomical and physiological features. On the other hand, perfusable channels within bulk gel structures have been reported by Skylar-Scott et al [41], but these structures fail to recapitulate capillary features such as an outside surface and thin wall thicknesses, which consequently affect biomechanics and diffusion kinetics. Our system aims to overcome some of these challenges by providing a facile way to manufacture biohybrid capillary-like structures with physiological properties.

To further demonstrate the potential of the self-assembling ELK1-GO bioink for the biofabrication of complex structures with innovative properties, we printed tubes but now exhibiting different porosities and corresponding permeabilities within the same structure along the tube axis (figure 6(c), illustrator). These types of structures are observed in biological systems such as renal tubules [42] and arteriovenous networks [43]. To our knowledge, these structures have not been fabricated in a single printing process. We first designed an experimental setup that enabled the formation of a MilliQ water solution made with different concentrations of GO contiguously increasing from 0.15 wt% on the left side, 0.10 wt% on the middle, and 0.05 wt% on the right side (figure 6(a)). The GO was stained with Rhodamin B prior to printing of the ELK1 to enable GO visualization in the tube wall. In this setup, we found that the gradient of the GO in solution was maintained for at least 3 h, which was confirmed by a 'transport of diluted species' COMSOL model (figure 6(a)) and facilitated the printing process. Upon extrusion of the ELK1 solution into the GO solution from the lower to the higher concentration of GO, the tube formed immediately at the interface of both liquids with the tube wall progressively assembling with increasing concentrations of GO (figure 6(b)). The resulting tubular structure was then characterized via epifluorescence microscopy, which revealed a tube wall with increasing content of GO and consequently a porosity gradient (figure 6(c), red). To test the effect of this porosity gradient on the permeability of the overall tube, 20 kDa FITC-dextran solution was perfused through and, as expected, its diffusion progressively decrease along the tube direction as the GO concentration increased (figures 6(c), green and (d)). These kinds of anisotropic structures are prevalent in biological systems such as in renal tubules [42] and arteriovenous networks [43]. Our bioink enables the fabrication of such anisotropic vascular structures in a simple manner.