The present invention relates to a composition for electrowriting of viable cells, wherein the composition is comprised of a hydrogel comprised of an aqueous medium comprising a proteinaceous biopolymer and a polyethylene glycol or polyethylene oxide. The present invention further relates to a biological scaffold comprised of the hydrogel composition wherein viable cells are arranged and encapsulated in a plurality of hydrogel fibres forming a three dimensional organized porous structure. Furthermore, the present invention relates to a method for production of the biological scaffold by electrowriting and the use of the biological scaffold. The invention furthermore relates to a device for manufacturing a biological scaffold.

Bio printing has become an important tool for fabricating regenerative implants and in vitro cell culture platforms. Bio printing is an emerging technique for the fabrication of biological constructs that can be used in regenerative medicine (RM) and in vitro drug testing. However, in an attempt to better mimic native tissues, there is an increasing demand to engineer structures with finer resolutions and further capture the hierarchical structure and composition of the native extracellular matrix. However, bio printing processes are limited to resolutions on the order of hundreds of micrometres, which hamper the reproduction of intrinsic functions and morphologies of living tissues.

This mimicking of the cellular micro-environment could provide new opportunities for the generation of constructs that can successfully exhibit functions at tissue and organ levels. For example, intercellular interactions at the micro- to nanometre scale, spatiotemporal changes in extracellular matrix structure, and mechanical and topographical cues provided by fibrillar extracellular matrix components are key drivers of cell behaviour, and are likely crucial elements of engineered functional constructs. Until today, the most well-established bio printing processes, i.e. extrusion-based, droplet-based, and light-assisted, with the notable exception of two-photon polymerization, are limited to resolutions close to tens of micrometres, which hamper the reproduction of such cellular microenvironments.

Attempts have been made to achieve high-resolution biological constructs by using non- conventional bio fabrication processes, such as electrospinning and electro-hydrodynamic jetting. Both methods employ electrical fields to create micrometre-scale fibres with encapsulated living cells. Unlike conventional extrusion bio printing methods, here material flow is driven by electrical forces that surpass the surface tension of the liquid ink, allowing for the fabrication of fibres with smaller sizes than the extrusion nozzle diameter. Surprisingly and despite the high electrical inputs used, these methods driven by electrical forces were found to be compatible with processing of living cells, even showing initial steps towards the reconstruction of hierarchical structures embedding cardiomyocytes, or cells from neural lineage. Nevertheless, neither of these techniques could simultaneously meet the requirements to emulate the intrinsic morphologies and local composition of cellular microenvironments, i.e. the three- dimensional (3D) patterning, the deposition of fibres with micron/ sub-micron size diameters and the maintenance of high cell viability.

Despite the fact that electrospinning approaches are compatible with the generation of cell-laden microfibers, they could not be readily applied to organize the fibres into predefined 3D shapes, due to whipping instabilities of the electrified jet. Previous reports on cell electrospinning strategies, disclose cell-laden fibres processed into randomly distributed and disorganized fibrous meshes. With electro-hydrodynamic jetting, on the other hand, structures with more complex three- dimensional patterns could be created, yet fabrication resolution and cell viability were compromised. In addition, previous studies have been carried out with polysaccharide-based hydrogels, such as alginate, which present poor cell adhesive properties and limited mechanical toughness, or synthetic polymers, such as poly(2-ethyl-2-oxazine), which are not compatible with cell encapsulation.

Considering the above, there is a need in the art for biomaterial platforms or biomaterials that are compatible with processing through the application of strong electrical fields, allow for fabrication of complex small-scale three-dimensional geometries, and support high viability of encapsulated cells. Furthermore, there is a need in the art for three dimensional biological scaffolds and the production thereof via the application of strong electrical fields, wherein the scaffolds are comprised of such biomaterials which enables the arrangement of viable cells to form complex organized organ like structures and morphology.

It is an object of the present invention, amongst other objects, to address the above need in the art. The object of present invention, amongst other objects, is met by the present invention as outlined in the appended claims.

Specifically, the above object, amongst other objects, is met, according to a first aspect, by a composition for electrowriting (CEW) of viable cells, wherein the composition is comprised of an aqueous medium comprising at least 10% w/v, preferably at least 15% w/v, more preferably at least 20% w/v of a proteinaceous biopolymer, and between 1 to 10% w/v of a polyethylene glycol or polyethylene oxide, preferably 2 to 8% w/v, more preferably 3 to 6% w/v based on the total volume of the composition. The cell electrowriting (CEW) used herein is an electro-hydrodynamic process that relies on electrical fields to fabricate organized three-dimensional (3D) structures from cell-laden fibres with diameters ranging from 1 to 100 pm. The composition can be used in the CEW process for the production of geometrical biological scaffolds mimicking an organ, wherein the composition is a hydrogel that comprises a protein-based natural-derived biopolymer that are very suitable due to their inherent biocompatibility, bioactive signalling, binding affinity for live cells, and tuneable mechanical properties.

A composition for use in CEW for biological scaffolds comprising the biomaterial composition should have sufficiently viscosity and moderate electrical conductivity to prevent Rayleigh instabilities and ensure a steady CEW production process without prevent droplet formation. To enhance viscosity the composition comprises between 1 to 10% w/v polyethylene oxide (PEO), or polyethylene glycol (PEG). For example maximum viscosities of 0.8 Pas and 0.4 Pas were obtained for norbomene -modified gelatin (gelNOR) and silk hydrogel systems with 6%w/v and 3.4% w/v PEO, respectively. These PEO concentrations were required to and to thus allow a steady jet formation.

According to a preferred embodiment of the composition, the proteinaceous polymer is one or more selected from the group consisting of a collagen based polymer, gelatin, silk, hyaluronic acid, norbomene-modified gelatin, and silk fibroin, preferable gelatine- or silk fibroin based polymer. The proteinaceous polymer comprises a new class of photo-crosslinkable hydrogel compositions based on proteinaceous polymers, i.e., gelatin and silk fibroin that are compatible with the abovementioned requirements, and thus enables the 3D electrowriting of viable cells on microscale, producing cell-laden microfibers. Gelatin preferred because it is derived from collagen, the main organic constituent of the natural extracellular matrix of mammals. Silk fibroin is preferred due to its unique mechanical properties and potential for cell encapsulation. Importantly, the unique biocompatibility of the gelatin- and silk-based hydrogel systems enable the fabrication of tissue constructs, while maintaining high cell viability (at least 50%, preferably at least 70%). These protein-based polymers with complementary biological and mechanical properties further aids in approaching the functional and structural properties of native cellular microenvironments, using CEW bio fabrication.

According to another preferred embodiment of to the composition, the composition is further comprised of viable cells selected from the group consisting of stem cells, bone marrow-derived mesenchymal stromal cells (MSCs), adipose-derived stromal cells (ASCs), tissue specific cells, cardiomyocytes, neural cells, or a mixture thereof, preferably bone marrow-derived mesenchymal stromal cells (MSCs).

Yet another preferred embodiment of to the composition comprises at least 50%, preferable at least 60%, more preferably at least 70%, most preferably at least 80% viable cells, based on the total percentage of cells present in the composition.

A preferred embodiment of the composition comprises between 1 to 10 ¹³ cells/ml, preferably 10 ² to 10 ¹² cells/ml, more preferably 10 ³ to 10 ¹¹ cells/ml, even more preferably between 10 ⁵ to 10 ¹⁰ cells/ml, most preferably between 10 ⁷ to 10 ⁹ cells/ml. Lower ranges of cell concentration provide an composition that is easier to use in CEW production process and cell printing. However, when the concentration is too low (i.e. less than 1 cell per ml in the composition), cells will die out over extended culture/time in the composition because the cell density is too low. Furthermore, such low densities are difficult to use for tissue engineering. Experiments show that bone marrow- derived mesenchymal stromal cells (MSCs) mixed with the composition in cell concentrations up to 10 ⁸ cells/mL could successfully be processed and cell-laden electro written micro fibers. And more importantly, cell viability analysis of the electro written cells over 1 , 3 and 7 days of culture post-production of the microfiber revealed a high ratio of living cells (>70% at all time points). Present composition preserved the cell viability in the CEW process, i.e. the applied electric field, and the resulting shear stresses within the electro hydrodynamic jet did not induce significant damage to the cells.

Another preferred embodiment of the composition has a gelation time at exposure to visible light of at most 45 seconds, preferably at most 40 seconds, more preferably at most 30 seconds. High crosslinking efficiency within short latency time is of importance to ensure high printing resolution and cell embedding within the electro written fibre, i.e. fast gelation kinetics is important to form stable fibres with a reproducible diameter that can be effectively stacked into ordered 3D structures. This poses a significant challenge for CEW as the composition needs to flow about 10 to 100 times faster within the electro hydrodynamic jet used in CEW than within conventional extrusion-based bio printing. In view of the short latency time, visible light-mediated thiol-ene click reactions and di-tyrosine oxidation photo-chemistries can be used. Hence, norbomene- modified gelatin (gelNOR), which can form a hydrogel network through step growth polymerization in the presence of a multi-functional thiol cross linker, and silk fibroin, which undergoes quick photo-chemical crosslinking due to the presence of tyrosine residues in its protein structure can be used in the composition. For example, to accelerate the gelation kinetics, a two- component photo initiator system, based on a mixture of tris (2,2-bipyridyl) dichlororuthenium(II) hexahydrate and sodium persulfate was used in the fabrication of large tissue constructs with complex geometry.

Another preferred embodiment of the composition is further comprised of one or more photo- reactive additives selected from the group consisting of tris(2,2-bipyridyl)dichlororuthenium(II) hexahydrate and/or sodium persulfate. The inclusion of tris(2,2-bipyridyl)dichlororuthenium(II) hexahydrate and sodium persulfate in the composition resulted in a further increase of the conductivity of the hydrogel systems. For sodium persulfate, this increase in conductivity is likely due to the presence of sodium and persulfate ions in the aqueous medium of the composition. The composition of present invention, which may be based on gelatin or silk, can be crosslinked via phenol-phenol coupling induced by exposure to visible light in presence of a photoinitiator or enzymatic mediated crosslinking (e.g. HRP/H ₂ O ₂ , tyrosinase, transglutaminase, sortase, thrombin).

According to another preferred embodiment, the photo-reactive additives are present in the composition at a concentration of between 0.5 to 10 mM, preferably, 1 to 7.5 mM, more preferably 2.5 to 5 mM. Experiments have shown that the presence of the photo-reactive additive determines the window within which the formation of a fluid jet during CEW bio fabrication was still possible. For example, in case levels of sodium persulfate exceeded 10 mM, formation of a continuous jet was no longer observed, likely due to the excess of SPS ionic compounds and the resulting increase in electrical conductivity of the hydrogel.

According to yet another preferred embodiment, the composition is further comprised of one or more cross linkers selected from the group consisting of thiol based cross linker, dithiothreitol, PEG-dithiol, bis-cysteine peptides or other water soluble thiolated compound, preferably PEG- octa-thiol. Gelatin is functionalized with carbic anhydride to achieve norbomene functionalization, mixed with sulfhydryl-containing linkers (e.g., dithiothreitol, PEG-dithiol, bis-cysteine peptides) and crosslinked via thiol-ene photo reaction. In addition to or instead of norbomene, other functional groups can be introduced such as acrylate or methacrylate groups, followed by photocrosslinking or Michael addition reactions.

According to a preferred embodiment, the composition is a hydrogel, preferably a photo responsive hydrogel.

According to yet another preferred embodiment, the composition has an electro conductivity of between 0,1 to 5 mS/cm, preferably 0,5 to 4 mS/cm, more preferably 0,75 to 3 mS/cm, most preferably 1 to 1,5 mS/cm. Experiments have shown that the electrical conductivity of the norbomene-modified gelatin (gelNOR) and silk hydrogels, with and without PEO, was considerably higher than semi-conductive fluids that are considered ideal for steady jet formation under electrical fields. Furthermore, we demonstrate that these high conductivity values did not prevent stable jet formation.

According to a second aspect, a biological scaffold is provided comprised of a composition as mentioned above, wherein the biological scaffold is comprised of viable cells arranged and encapsulated in a plurality of hydrogel fibres forming a three dimensional organized porous structure, wherein the porous structure has a pore size of between 1 to 200 pm, preferably between 5 to 100 pm, more preferably between 10 to 50 pm.

According to a preferred embodiment, the hydrogel fibres have a diameter of between 1 to 100 pm, preferably between 2,5 to 75 pm, more preferably between 4 to 50 pm. The obtained fibre diameters are at least one to two orders of magnitude smaller than the fibre sizes previously reported for extrusion-based bio printing (about 200 pm), and electrohydrojetting processes (>100 pm). The obtained fibre diameters approximate the size of a single cell (animal cell size ~ 10 pm) and are suitable for cell encapsulation.

According to another preferred embodiment, the three dimensional organized porous structure is mimicking the cellular micro-environment and morphology of native tissue and/or an organ.

According to yet another preferred embodiment, the biological scaffold comprises at least 50%, preferable at least 60%, more preferably at least 70%, most preferably at least 80% viable cells.

According to another preferred embodiment, the viable cells are selected from the group consisting of stem cells, bone marrow-derived mesenchymal stromal cells (MSCs), adipose-derived stromal cells (ASCs), tissue specific cells, cardiomyocytes, neural cells, or a mixture thereof, preferably bone marrow-derived mesenchymal stromal cells (MSCs).

According to a preferred embodiment, the biological scaffold has a thickness of at least 10 pm, preferably at least 50 pm, more preferably at least 100 pm, even more preferably at least 200 pm, most preferably at least 500 pm. The biological scaffold is comprised of at least one layer of cells, preferably at least 5 stacked layers of cells, more preferably at least 10 stacked layers of cells, even more preferably at least 25 stacked layers of cells, most preferably at least 40 stacked layers of cells. According to a further aspect, a tissue construct is provided comprising at least one biological scaffold as described above. The composition in combination with the cell-electrowritting also allows for simultaneous printing of multiple hydrogel compositions into multiple biological scaffolds in one single construct.

According to a further aspect, a method for production of a biological scaffold is provided by electrowriting (CEW) is provided, wherein the method comprises the steps of a) providing the composition as described above, b) applying an electrical force on the composition that surpasses the surface tension of the composition such that a flow of said composition is provided to enable the formation of hydrogel fibres, c) forming a plurality of hydrogel fibres, d) depositing said hydrogen fibres to form an organized porous structure providing a biological scaffold.

To allow controlled patterning of the fibres with the embedded cells by CEW, applied voltage, collector velocity and dispensing pressure are important parameters to control, also to obtain the optimum fibre alignment and fibre diameter. Experiments show that the at a constant voltage, preferably between 0.5 and 5 kV, the constant collector velocity should be between 25 mm/s to 50 mm/s and the applied pressure may not exceed 0.05 bar, as the combination of these parameters allowed the consistent deposition of homogenous fibres, i.e. without a jet break and minimal oscillation of fibre diameter within a jet. Due to the low viscosity of both hydrogel systems, dispensing pressures above 0.05 bar resulted in excessive flow rates and consequent fibre coiling and fibre diameter oscillation.

According to a preferred embodiment, the applied electrical force is between 0.5 to 5 kV, preferably between 1.5 to 4 kV, more preferably between 2.5 to 3 kV. An increase in both fibre straightness and diameter was observed between 2.5 and 3.0 kV, with fibres from both gelatin- and silk hydrogel reaching full straightness (no visual fibre-coiling) at 3.0 kV, therefore a voltage of between 2.5 to 3 kV is preferred. Coiling and reduction of the fibre diameter occurred when voltages between 3.0 and 5.0 kV were applied.

According to another preferred embodiment, the formation of hydrogen fibres is performed at a constant velocity of between 25 mm/s to 50 mm/s and a dispensing pressure of between 0.01 and 0.08 bar, preferably between 0.02 and 0.06 bar. According to yet another preferred embodiment, the hydrogen fibres are deposited in a pattern to form a three dimensional organized porous structure that is mimicking the cellular microenvironment and morphology of native tissue and/or an organ.

According to a preferred embodiment, after deposition, the hydrogel fibres are being exposed to visible light (400 to 450 nm) for at least 30 seconds, preferably at least 1 minute, more preferably at least 2 minutes, resulting in the polymerization of the hydrogen fibres providing said biological scaffold.

According to a further aspect, a device for manufacturing a biological scaffold, preferably as described above, is provided, wherein the device comprises: a holder for receiving a composition for electrowriting, preferably as described above an electro jetting mechanism for creating a electro jet from said composition; a collector for collecting the microfilaments, wherein the collector is within the stable electro jet distance of the electro jetting process; a moving mechanism for moving the collector for creating the three dimensionally ordered biological scaffold.

The electro jetting or electro spinning mechanism is arranged to form a jet or thin structure of a fibre. As mentioned above, a electrical force may thereto be applied is between 0.5 to 5 kV, preferably between 1.5 to 4 kV, more preferably between 2.5 to 3 kV.

In typical electrospinning processes, the initial jet will deform in a swirling structure. The device is preferably arranged to collect the jet prior to the formation of this swirl. The collector for collecting the microfilaments is thereby within the stable electro jet distance of the electro jetting process. This allows forming a predetermined structure, as opposed to a random configuration as is common in electrospinning processes. By moving the moving mechanism in a predetermined manner, a scaffold having a predetermined structure can be formed.

Preferably, the device further comprises a crosslinking mechanism for creating a microfilament from said electro jet. As mentioned above, visible light may be used thereto.

According to a further aspect, a use of the biological scaffold as mentioned above is provided for regenerative medicine applications, drug testing, as organ-on-a-chip model and/or in vitro experiments. The biological scaffold obtained via CEW allows to create microstructure scaffolds that can better resemble cellular microenvironments for regenerative medicine (e.g., muscle fibres, tendons, nerve networks) and organ-on-a chip models. It enable the creation of unique cell culture platforms for applications in developmental biology and drug discovery, because the cellular microenvironment could now be further controlled by the resulting cell-laden fibre size, material composition, scaffold architecture and eventually mechanical properties.

The present invention will be further detailed in the following examples and figures wherein: Figure 1 shows a scheme of the cell electrowriting (CEW) process using a hydrogel composition;

Figure 2 shows the physical-chemical properties of gelatine and silk based hydrogel compositions;

Figure 3 shows the influence of the main cell electrowriting processing parameter;

Figure 4 shows the structural and mechanical stability of CEW hydrogels fibres;

Figure 5 shows the cell electrowriting process of complex shaped cell-laden fibre scaffolds;

Figures 6A-C show a device for manufacturing a scaffold.

Figure 1 shows a scheme of the cell electrowriting (CEW) process using a hydrogel composition in providing a complex three-dimensional porous structure, i.e. having the morphology of an organ. Visible light-mediated thiol-ene click reactions and di-tyrosine oxidation photo-chemistries were used to photo crosslink the hydrogel composition and produce cell laden microfibers. The CEW process provides optimal cell distribution and provides a resolution as low as between 1 to 100 pm in pore size of the cell laden fibres and can be used in the fabrication of large tissue constructs with complex geometry.

Figure 2 shows the physical-chemical properties of gelatine and silk based hydrogel compositions. A) In situ photorheometry showing the storage modulus (G’) of the compositions as a function of time. Hydrogel samples were irradiated with visible light 30s after the experiment started. Photorheological analysis at different Ru/SPS ratios confirmed gelation of the gelNOR and silk fibroin hydrogel compositions within less than 30 seconds and 1.5 minutes of exposure to visible light, respectively. B) Viscosity changes as a function of the compositions’ PEO concentration. The viscosity of the gelNOR-PEO and silk-PEO blends increased significantly compared to the gelNOR- and silk-only solutions. C) Electrical conductivity of hydrogels as a function of both PEO and photoinitiator concentration. The inclusion of 2/5 mM and 2/10 mM Ru/SPS resulted in a further increase of the conductivity of both hydrogel systems, which are considerably higher than known semi-conductive fluids that have a conductivity of about < 10' ¹¹ mS/cm and are considered ideal for steady jet formation under electrical fields. D) Sol fraction as a function of Ru/SPS photoinitiator concentration. The effect of SPS and PEO on the crosslinking efficiency of the composition was investigated by quantifying the sol fraction upon formation of the polymer network. Both the gelatin and the silk based composition are able to form stable hydrogel fibres at 5mM SPS concentration.

Figures 3A-D show the influence of the main cell electrowriting processing parameters, applied voltage, collector velocity and dispensing pressure, on fibre collection and 3D patterning for both cell-free hydrogel composition. A) Effect of i) voltage, ii) collector velocity and iii) air pressure on fibre diameter. Printability window is represented by background colours. B) Representative microscopic images of the effect of increasing i) voltage and ii) collector velocity on fibre morphology for the gelNOR hydrogel system. The applied voltage was first studied at constant collector velocity (25 mm/s) and applied pressure (0.05 bar), as the combination of these parameters allowed the consistent deposition of homogenous fibres. An increase in both fibre straightness and diameter was observed between 2.5 and 3.0 kV, reaching full straightness (no visual fibre-coiling) at 3.0 kV (Figure 3Ai and Figure 3Bi). By increasing the collector velocity from 25 to 50 mm/s, a reduction in fibre coiling and fibre diameter was observed (Figure 3Aii and Figure 3Bii). The effect of the dispensing pressure was also investigated, yet straight fibre alignment was only observed within a narrow range fibre (Figure 3 Aiii). Due to the low viscosity of both hydrogel systems, dispensing pressures above 0.05 bar resulted in excessive flow rates and consequent fibre coiling and fibre diameter oscillation. C) Print fidelity of fibre hydrogel 3D scaffolds was examined by the patterning capacity of the hydrogel composition, by manufacturing 3D scaffolds with square pore geometries, pore sizes and varying scaffold thicknesses (Figure 3Ci and ii). Scaffolds with precisely configured square pores were obtained. Ci) shows the accuracy of printed pores, Acc pore, as a function of scaffolds pore size. Acc pore was determined using a relative value obtained by the ratio between the design and fabricated pore area. In the case of no deviation between printed and designed pore areas, Acc pore= 1. Cii) shows representative images of printed scaffolds with 1000 pm and 200pm pore size for gelNOR hydrogel. D) Shows the scaffold thickness as a function of the number of stacked hydrogel layers. Di) shows the final scaffold thickness and Dii) representative SEM images of scaffolds with 3 and 10 layers showing perfectly stacked hydrogel fibres at the vertices of squared pores. It was observed that scaffolds with thicknesses of 50 pm and 200 pm for the gelNOR and silk hydrogel systems, respectively, could be fabricated without significantly affecting printing accuracy.

Figure 4 shows the structural and mechanical stability of CEW hydrogels fibres. Figure 4A shows the swelling of A) gelNOR and B) silk-based cell-free hydrogel fiber before and after Ihour, 1 day, 2 days, 3 days and 7 days incubation in PBS. Figure 4C shows representative microscope images of swollen gelNOR fibers before and after Ih and 2 days of swelling. Figures 4 D-E show representative loading and unloading curves of CEW gelNOR and silk-based cell-free scaffolds measured by nanoindentation. Curves have been averaged over at least 3 measurements. Figure 4F shows the effective stiffness of CEW gelNOR and silk-based cell-free scaffolds and figure 4G compared with scaffolds of same composition obtained by conventional extrusion bioprinting. All nanoindentation experiments were performed on scaffolds after 1-day PBS immersion (equilibrium swelling).

Figures 5A-E show the cell electrowriting process of complex shaped cell-laden fibre scaffolds. Figure 5A and Figure 5B shows cell viability of gelNOR and silk-based cell-laden scaffolds after 1, 3 and 7 days of in vitro culture. Cell electro written scaffolds exhibited high cell viability (>70%) for both hydrogel systems, similar to the conventional extrusion-based bio printing and manual hydrogel casting. These data show that the composition used in the CEW process preserves cell viability, suggesting that the relatively high viscosity of the hydrogel composition, the applied electric field, and the resulting shear stresses within the electro hydrodynamic jet did not induce significant damage to the cells.

Figure 5C) and D) show cell distribution and morphology examined on gelNOR based cell-laden scaffolds and compared with conventional extrusion bio printing. Cell electro written fibres exhibited single cells accurately aligned along the fibre pattern. In contrast, the conventional extrusion bio printed fibres exhibited multiple cells per filament that were randomly aligned along filament pattern. An interesting feature of the composition in the CEW produced cell laden fibres is that each micro fiber (diameter = 2- 10 pm) is significantly smaller than the size of an individual MSC in suspension (= 10-30 pm). Small, submicron features such as grooves, pillars and microfibers are generally known to greatly affect cell morphology via contact guidance and spatial constriction. However, from a morphological point of view, the electro written cells showed only a slight cytoskeletal deformation, yet no significant morphological changes in the nucleus, compared to cells embedded in extrusion-based printed fibres, as observed by measuring cell circularity as a shape descriptor (Figure 5D). Interestingly, the hydrogel compositions, rather than providing a strict confinement for the embedded cells, simply appear to accommodate for the cell size, encapsulating them in a bead-and-necklace-like conformation.

Figure 5E) shows complex shape patterning of representative gelNOR based electro written scaffolds with fluorescent nanoparticles. Cell electro written hexagonal shaped 3D scaffolds exhibited 2 times higher printing accuracy than extruded printed scaffolds. Cell-electrowritting also allowed simultaneous printing of multiple hydrogel compositions in one single construct. In figures 6A-C a device 100 for manufacturing a biological scaffold as described above is shown. As shown in the side view of figure 6A, the device 100 comprises movable stage 1 is movable in the X, Y and Z-direction using linear motors 11 - 13. Arranged above the stage 1 is an electro jetting mechanism 2 which is arranged for creating an electro jet from a nozzle 22 (see figure 6B which shows the electro jetting mechanism in greater detail). The device 100, for instance suitable controller thereof, is arranged to maintain the distance h (see figure 6A) between the movable stage 1 and the nozzle 22 such that the movable stage 1 is within the stable electro jet distance of the electro jetting process. The linear motor 13 for adjusting the position in the Z-direction may be controlled to this end. Using the moving system 10, the stage 1 is moved with respect to the nozzle 22 for forming a predetermined three dimensional structure using electro jetting.

With reference to figure 6C, which is cross-section along plane A in figure 6B, the electro jetting mechanism 2 comprises a holder 23 in the form a luer lock syringe 13 for holding the composition for electro writing. Heating coils 29 are provided to ensure that the composition in the correct temperature range to ensure cell viability and composition characteristics. Upon actuation of the electro jetting mechanism 2, the composition is ejected from the nozzle 22. Also provided in the mechanism is an electrical field generator 25 which is arranged to apply an electrical force on the composition that surpasses the surface tension of the composition such that a flow of said composition is provided to enable the formation of hydrogel fibres from the nozzle 22. In this example, the applied electrical force is between 2.5 to 3 kV.

Also shown in figures 6A and 6B is a curing mechanism generally indicated with 21. The curing mechanism 21 comprises a light source 21a for visible light (400 to 450 nm). The device 100 is arranged to expose the hydrogel fibres after deposition on the stage 1 to visible light (400 to 450 nm) for at least 2 minutes, resulting in the polymerization of the hydrogen fibres providing the biological scaffold.

Examples

Preparation hydrogel compositions

The following materials were used, Gelatin type A (bloom 180, Roth, Germany), 8- arm polyethylene glycol) thiol (MW: 10,000 g/mol; JenKem Technology, USA); tris(2,2’- bipyridyl)dichlororuthenium(II) hexahydrate (Ru; Sigma-Aldrich, The Netherlands); sodium persulfate (SPS; Sigma-Aldrich, The Netherlands); Poly( ethylene oxide) (PEO; MW: 600,000- 1,000,000 g/mol, Acros Organics, USA); silkworm cocoons (Wildfibres, UK), lithium bromide (Acros Organics, USA). Norbomene -modified gelatin (gelNOR) was synthesized as previously described by Z. Munoz et al., Biomater. Sci. 2014, DOI 10.1039/c4bm000. Briefly, 10% (w/v) porcine gelatin type A (180 bloom; Roth, Germany) was dissolved in phosphate buffered saline (PBS) at 50 °C under constant stirring conditions. 20% (w/v) of carbic anhydride (CA; Acros Organics, Japan) was added to the gelatin solution and the pH was adjusted to 7.5 - 8.0 using 5M sodium hydroxide (NaOH). The reaction was quenched after 24 hours by addition of 3x PBS. After centrifugation at 4000 rpm to remove excess CA, gelNOR was dialyzed against deionized water (MilliQ) at 5 °C for 5 days and water was refreshed 2x per day. Finally, the solution was filter-sterilized and lyophilized. The degree of functionalization was 45% as determined by Fluor aldehyde assay.

Silk fibroin was extracted from Bombyx mori silkworm cocoons as previously described, D. N. Rockwood et al., Nat. Protoc. 2011, DOI 10.1038/nprot.2011.379. Briefly, cut cocoons were boiled in an aqueous solution of 0.02M Na ₂ CO ₃ for 30 min. The degummed fibres were dissolved in a 9.3 M LiBr (Sigma Aldrich) solution at 70°C for Ih, followed by dialysis against water for 48 h, using cellulose dialyzing tubing (MWCO 3.5 kDa, Sigma Aldrich). The resulting 6 % (w/v) silk solution was concentrated to 16 % (w/v) by dialysis against PEG (10 kDa, Sigma Aldrich).

Stock gelNOR solutions were mixed with solutions having different PEO concentrations to obtain a blend with a final gelNOR concentration of 10% and various PEO content (1 - 6%w/v). 10% (w/v) PEO stock solution was prepared by dissolving PEO powder in miliQ. The gelNOR/PEO blends were gently mixed at room temperature, followed by addition of Ru and SPS at various ratios (2:5, 2:10, 2:20). Concentrated silk solutions were prepared with the same PEO concentrations as the gelNOR ink. The silk/PEO blend was mixed at 4 °C, followed by addition of Ru and SPS at the same ratios as the gelNOR.

Rheological measurements hydrogel compositions

Rheological characterization of gelNOR, gelNOR/PEO, silk, silk/PEO, were performed on a Rheometer (Discovery HR-2, TA instruments) fitted with a parallel plate of 20 mm in diameter, a gap distance of 0.5 mm and equipped with a light curing system. To determine the gelation time, the hydrogel compositions with different Ru/SPS ratios were placed between the two plates. In situ photorheometry was performed by using a visible light source with light switched on 30s after initiating the time sweep measurement. All measurements were performed within the linear viscoelastic region, at a strain of 1 % and room temperature. The elastic (G') and viscous modulus (G") were recorded as function of time. To determine the viscosity as function of the shear rate, the hydrogel composition solutions were placed between the parallel plates and data was collected over a shear rate range between 1 to 1000 s-1.

Electrical conductivity measurements hydrogel compositions

Conductivity of the different hydrogel compositions was evaluated using a CDM230 conductivity meter (Radiometer Analytical, France). Calibration was first conducted using NaCl solutions (0.1 and 1 % w/v).

Sol fraction analysis

Crosslinking efficiency was assessed through sol fraction analysis as per previously reported in literature, K. S. Lim, B. S. Schon, N. V. Mekhileri, G. C. J. Brown, C. M. Chia, S. Prabakar, G. J. Hooper, T. B. F. Woodfield, ACS Biomater. Sci. Eng. 2016, DOI 10.1021/acsbiomaterials.6b00149. First, cylindrical gels (0=6 mm x h=2mm) were prepared in custom-made Teflon molds and crosslinked by exposure to visible-light irradiation (wavelength 400 - 450 nm) for 45s. All crosslinked cylinders were weighted for their initial weight (min t=0) and three lyophilized samples to obtain dry weights (mdry t=0). Remaining samples were incubated in PBS at 37 °C, freeze dried and weighted again (mdry). Sol fraction then was then determined as follow, sol fraction = ((min t=0 -mdry) / min t=0).

Cell electrowriting (CEW)

Hydrogel fabrication was performed with an in-house built CEW device. Briefly, hydrogel was loaded in a temperature controlled printhead (temp range 0 - 120 °C) connected to a high precision air pressure (0.01-2 bar, VPPE-3-1-1/8-2-010-E1 557771, Festo). The hydrogel was electrified with a high voltage power supply (Heinzinger LNC 30000-2pos, 0-30 kV) and the electrified hydrogel fibre was collected in a computer controlled high-precision XY stage (LG-motion, UK). The CEW device was protected by an acrylic box to ensure stable environmental conditions. All the fabrication experiments were performed at room temperature using glass syringes (3 ml, 25 G needle nozzle). Hydrogel jet formation as well as fibre diameter and morphology were investigated at increasing voltage [2, 5] kV, for a constant velocity 25 mm/s and pressure 0.05bar; at increasing collection velocity [5, 50] mm/s, for a constant voltage 2.5kV (gelNOR), 3kV (silk) and pressure 0.05 bar; and at increasing pressure [0.05, 0.3] bar, for a constant collection velocity of 25 mm/s and voltage 2.5kV (gelNOR), 3.0kV (silk). Print fidelity and layer stacking experiments were performed at a constant collector velocity [25, 30] mm/s, voltage 2.5 (gelNOR) and 3.0 (gelNOR), and air pressure 0.05bar. All experiments were conducted at a constant collector distance of 5 mm and constructs were irradiated with visible light (400 - 450 nm) during jet collection and after printing for approx. 5min. Squared- and hexagonal-shaped pore fabrication experiments were performed with same CEW parameters. Jet and scaffold fabrication were monitored by digital USB microscopes.

Extrusion-based bio printing:

Extrusion printing of gelNOR and silk hydrogels was achieved by pneumatic dispensing with a Cellink Inkredible bio printer (Cellink, Sweden). Extrusion was performed at room temperature, with a 25 G needle tip, collector velocity of lOmm/s and an applied pressure of 0.9 bar. After printing constructs were irradiated with visible light (400 - 450 nm) for approx. 5 min.

Fibre and scaffold imaging

An Olympos BX51 fluorescent microscope (Olympos, The Netherlands) was used using a TRITC filter to obtain high magnification images of printed hydrogel fibres. Fluorescence microscopy for fibre quality analysis was evaluated through imaging of the natural fluorescence exhibited by gelatin and silk polymers. Stereomicroscopy images of fabricated 3D constructs were acquired with an Olympus stereomicroscope (Olympus Soft Imaging Solutions GmbH, The Netherlands). Accuracy of hydrogel 3D constructs was also analysed with SEM (Phenom Pro, Phenom-World, The Netherlands) at an acceleration voltage of 5-10kV. Prior SEM imaging samples were gold plated (2 nm) using a Q150R rotary-pumped sputter (Quorum Technologies, UK). Print fidelity was quantified using a relative value, Accpore, expressed by the following index, Acc pore= Afab / Ades. Afab is the fabricated pore area and Ades is the designed pore area. Pore areas were quantified using Image J.

Cell isolation and culture

Equine tissue samples and cells were obtained from deceased horse donors, donated to science by their owner and according to the guidelines of the Institutional Animal Ethical Committee. Equine - derived bone marrow-derived MSCs were isolated as previously described, V. H. M. Mouser et al., Bio fabrication 2017, DOI 10.1088/1758-5090/aa6265.

After isolation, cells were cultured in MSC expansion medium consisting of DMEM + GlutaMAX (Gibco, 31966, The Netherlands) supplemented with fetal bovine serum (FBS, 10% v/v, Gibco, 10270, The Netherlands), penicillin/streptomycin (1%, Gibco, The Netherlands), 1-ascorbic acid-2- phosphate (0.2 x 10-3 M, Sigma Aldrich, The Netherlands), and basic fibroblast growth factor (Ing/mL, PeproTech, United Kingdom), medium was refreshed twice per week. Cells were expanded until passage 4 and used at a density of 10 ⁸ cells/ml for both cell electrowriting and extrusion-based bio printing. Cell viability

Cell viability in printed fibres was quantified through a LIVE/DEAD assay (Calcein AM, ethidium homodimer-1, Life Technologies, The Netherlands) using a confocal microscope (Leica SPX8, Leica Systems, The Netherlands) for imaging after 1, 3 and 7 days in culture (n = 5). Cast cell-laden hydrogels used as control groups were prepared using Teflon molds (0=6 mm x h=2mm).

Immunofluorescent staining

At day 1 and 7 after cell electrowriting and extrusion bio printing, cell-laden constructs were fixed using 4% neutral buffered formalin. Pollowing Triton-X membrane permeabilization, samples were stained with phalloidin and DAPI. Images were captured using a Leica SP8X confocal microscope and cell shape was quantified using ImageJ.

The present invention is not limited to the embodiment shown, but extends also to other embodiments falling within the scope of the appended claims.