Engineering Biocompatible Hydrogel Fibers With Tunable Mechanical Properties For Neural Tissue Engineering

Neural tissue engineering holds a great promise for the treatment of neurodegenerative diseases and peripheral nerve injuries. However, the anisotropic mechanical and electrical properties of the highly aligned neural cells have hindered the development of a faithful in vitro disease models. In this study, a core-shell microuidic extrusion method is implemented to fabricate a cell-laden composite hydrogel ber with tunable mechanical properties. The hybrid hydrogel was formed using a core of GelMA mixed with gelatin seeded with human neuroblastoma cell (SH-SY5Y) and an alginate shell. The composition of the core hydrogel was optimized to support cellular growth and differentiation, yet allow a feasible fabrication of cell-laden bers. The engineered bers were remarkably biocompatible and enabled the formation of highly aligned cellular morphology.


Introduction
The central nervous system (CNS) has a limited capacity to grow new cells to repair the loss or damage of neurons and axonal pathways that accompany conditions such as traumatic brain injury (TBI), stroke, spinal cord injury, and neurodegenerative diseases [1][2][3][4] . Neurological disorders (such as stroke, dementia, meningitis, etc.) are among the leading causes of disability (~ 276 million) and deaths (~ 9 million in 2016) which continue to increase each year 5 . Neurogenesis is a multi-step process of neural stem cell differentiation in the CNS; however, it is restricted to a limited number of areas in the brain (e.g. hippocampus dentate gyrus, etc.), hampering the restoration of the lost neurons 6,7 . Additionally, regeneration of lost axonal pathways in the CNS is insu cient due to the lack of directed guidance as well as inhibitors like glial scar or myelin 8,9 . Conventional regenerative strategies to treat neurological diseases, including gene therapy, stem cell transplantation/engraftment or growth factor delivery, have shown limited success in creating functional tissues due to the low cell engraftment and lack of growth factors localization 10,11 . Animal model, as a traditional alternative, is extensively employed for the regenerative medicine experiments and CNS disorder therapies 12,13 . However, the differences in animal models compared to human body, in addition to the high costs and related ethical issues, have posed serious concerns on the applicability of animal models 14 . Regarding these critical issues, numerous measures have been implemented to develop reliable in vitro disease models 15-17 . Different in vitro techniques such as soft lithography 18 , laser texturing 19 , electrospinning 20 , and electrospraying 21 , have been employed to generate 2D cell-laden constructs. However, these techniques fail to accurately mimic the mechanical properties, architecture and physiology of brain-speci c structure and function 22 . In fact, in 2D culture methods, the functionality of neural cells (including morphology, proliferation, and differentiation) is either limited or completely lost 23,24 . In contrast to the 2D methods, three-dimensional (3D) bioprinted structures have demonstrated new prospective in mimicking the functional native-like tissue microenvironment 25,26 . In 3D bioprinting, cells, biomaterial, and bioactive moieties can be precisely positioned to mimic in vivo conditions 27,28 . This holds a great promise for Page 3/13 future advancements toward having a cost-effective, fast and anatomically accurate representation to address the current issues of CNS disorder therapies 29-31 . Generally, 3D bioprinting techniques can be divided into optical-based and nozzle-based methods 32 . In optical-based methods, the major step is the scaffold photopolymerization where a high energy radiation is exposed to the cell-containing precursors 33 . This can signi cantly reduce the cell viability rate, as compared to nozzle-based methods 34 . Although the optical-based methods have higher printing resolution, the simplicity and the low cost of nozzle-based methods have made the them more prevalent.
Amongst the nozzle-based bioprinting approaches, extrusion bioprinters have the advantage of generating low shear stress on living cells, making them more cytocompatible 34 . Another issue in mimicking neural tissues is the complex microenvironment where neurons, glial cells, blood vessels, pericytes, and extracellular matrix are elaborately bundled 35 . This requires manipulation of different cell types and printing materials simultaneously 36 . For that reason, coaxial extrusion methods can crosslink different biocompatible hydrogels (e.g. core-shell) at the same time through printing a ber-like construct 37,38 . In fact, there are two major superiorities of coaxial extrusion printing over other methods to be used for neural tissue disease models: (i) it provides guidelines for a controlled cell proliferation and axon outgrowth ( ber structure); and (ii) it can delicately position more than one cell type in a 3D tubular structure (core-shell structure).
Regarding the cell proliferation guidelines, a common approach is to use aligned polymer nano bers to fabricate a 3D scaffold as a culture microenvironment 39,40 . The electrospun nano bers can provide a controlled neurite outgrowth and cell alignment replicating the native neuroanatomical architecture 41 . As an example, a 3D scaffold based on a bacterial nanocellulose (BNC) can prove a biocompatible environment for neural cells (neuroblastoma, SH-SY5Y) 42 . The nano bers can also mix with carbon nanotubes to create a conductive 3D composite scaffold in which an exceptional cell (SH-SY5Y) attachment and proliferation along the guidelines can occur 43 . In addition to nano bers, other strategies such as chemical surface treatment, mechanical stimulation, and surface patterning can be implemented for a directional cell spreading and growth 44 . Although these methods can provide a relatively acceptable cell guiding platform, the complexity engaged in simultaneously manipulating of different cell types as well as the higher fabrication costs are still their ongoing challenges.
The ber constructs when fabricated in a core-shell form provides even a broader course of in vitro applications. The structure of coaxial constructs can robustly recapitulate the luminal/tubular con guration of cell bundles of neurovascular/neuromuscular tissues in a 3D environment, rather than a planner organ-on-chip 35,45 . In core-shell constructs, cells located in the core/shell are protected by the outer hydrogel shell that provides both stability and the possibility of cell-cell interactions within the environment 46 . A longitudinal ber-based scaffold fabricated to investigate the drug resistance in Glioblastoma, with human glioma stem cell (GSC23) in the shell and glioma cell line (U118) in the core, showed the superiority of core-shell constructs to mimic the native microenvironment of the neural tissues 47 . A recent study with neural stem cells (core) and Schwann cells (shell) showed an enhancement in cell proliferation and differentiation that can be implemented for spinal cord injury research 48 . In neurodegenerative disease models, however, longitudinal scaffolds do not necessarily provide a native-like microenvironment where neural cells are highly paralleled with the aligned Schwann cells 44,49 . In other words, growing highly aligned neural cells even in the core-shell bers is still a challenge. Therefore, a neural tissue model should be able to align cells in an parallel orientation where the synaptic electrophysiology, plasticity, and neural vesicle tra cking can occur 44 .
Speci cally under the topic of scaffold activated directional neural growth, only a few studies are conducted. As an instant, rat cortical cells and mouse neural stem cell were encapsulated in an ECM (extracellular matrix) protein hydrogel using the coaxial printing method. Despite the extensive investigation on different cell types in this study, the neural outgrowth and directional neural outgrowth is not quantitively measured 50 . Glia cells together with primary cortical neurons are also printed using a coaxial printing strategy that utilizes Gellan gum-RGD as the bioink 37 . Rather than the directed neurite outgrowth of the cells, networking of these types of cells was investigated in that study. Human derived cortical neurons and glial cells (derived from Human-Induced Pluripotent Stem Cells) were used in another bioprinting study 51 . Likewise, the cellular alignment was not the focus in that study.
In addition to CNS related diseases, peripheral nerve injuries and regeneration of neural bers are other important applications of coaxial bioprinting 52 . In fact, the importance of growing neural cells along the desired direction can be realized in the regeneration of the peripheral nervous system 53 . In most of the studies on peripheral neural on-chip models, fabrication of a guiding conduit for directed axonal growth is of interest.
As mentioned above, in vitro cell alignment in neural tissue engineering plays a critical role to faithfully model the neurodegenerative diseases. Herein, we fabricated a core-shell hydrogel scaffold with a superior neural cell alignment within the construct. The hybrid scaffold microenvironment is comprised of gelatin methacryloyl (GelMA)/gelatin and hydrogel as the core, and an alginate as the shell. Human neuroblastoma cells (SH-SY5Y) are cultured in the ber core with a high cell proliferation and more than 85% cell viability. Cell alignment is achieved with a standard deviation of less than 6 in cellular orientation. The microcylindrical 3D scaffold structure is shown to recapitulate the structure of the native extracellular matrix (ECM), as the neural growth is highly aligned with the ber guidelines.

Fabrication of hybrid hydrogel ber
The schematic of the micro uidic setup used to fabricate the hydrogel bers is presented in Fig. 1(A). The system is consisted of two syringe pumps, a co-owing micro uidic chip with two inlets and one outlet, and a calcium chloride (CaCl 2 ) bath. The micro uidic system is made by two concentric metallic needles together creating a core and sheath stream of different uids at the outlet. The outlet of the micro uidic device is connected to the bath with a capillary tube with a 150-µm inner diameter (ID). To fabricate the composite cell-laden hydrogel bers, a solution of sodium alginate (shell) and a solution of GelMA/gelatin containing the neuroblastoma cells (core) are pumped through the micro uidic device (see Material and Method Section for the details) ( Fig. 1(B)). The GelMA/gelatin is pumped to the core channel and the alginate is pumped to the shell channel. Then the fabricated bers are introduced to CaCl 2 bath (see Fig. 2(A)). The ber core is made of GelMA as it provides binding sites for the neural cells 54 . However, wet spinning of GelMA is challenging 55 . Therefore, an alginate shell is formed as a support to physically entrap GelMA monomers in the ber prior to their photocrosslinking 56 . Once the solutions extruded into the CaCl 2 bath, calcium ions replace the sodium and create the calcium alginate hydrogel 57 . The collected bers are exposed to UV light to photocrosslink the GelMA hydrogel ( Fig. 1(B)). The diameter of the core and the thickness of the shell could be tuned simply by controlling the ratio of the ow rates.

Physical property tuning
The setup used for the experiment is shown in Fig. 2(A). The core-shell construct in Fig. 2(B) is fabricated by 3%/2% w/v of GelMA/gelatin ink for the core and 2% w/v alginate ink for the shell. The core-shell regions remained unmixed (Fig. 2(B)). The ow rates of different streams were changed to alter the dimensions of the fabricated bers. Increasing the ow rate ratio between the inner and the outer stream increased the diameter of the printed hydrogel ber (Fig. 2(C)). The mechanical properties of the fabricated bers can be tailored by changing the properties of the used pre-polymers. Mechanical properties of the bers can also be tailored by changing the alginate concentration. The results show that a higher alginate concentration yields larger Young's modulus (Fig. 2(D)). The Young's modulus of the core hydrogel material consisted of GelMA/alginate were measured for different concentrations of GelMA at 1% gelatin. Increasing the concentration of GelMA, the Young's modulus is increased (Fig. 2(E)). The Young's modulus of the 2.5% GelMA and 1 % gelatin solution is about 0.6 kPa, which is around the stiffness of the native brain tissue 58 . Therefore, this concentration is selected as the optimized scaffold for the cell culture microenvironment. Figure 2(F) shows the Young's modulus for different concentrations of GelMA combined with 1% alginate which is higher than the stiffness of the native brain tissue.

Cytotoxicity evaluation
The ability of the fabrication method and the scaffold materials to recapitulate the native microenvironment for cell growth and proliferation is evaluated using different tests for the SH-SY5Y cells (Fig. 3). The lateral shell ow in the core-shell structure con nes the cells to ow within the core region. Figure 3(A) shows the phase contrast images of the cells encapsulated in hydrogel ber at Day 0, 1, and 5 of the fabrication. The viability assessment was performed considering the live/dead assays (Fig. 3(B)).
The live/dead assay results show more than 85% of the cells were alive (appearing as green) at Day 1 and 3. Taken together, the results indicate that the fabricated bers do not induce immediate toxicity and the presence of alginate with the tested concentrations does not reduce the cellular viability. In addition, to evaluate the cells morphology in the 3D structure of hybrid hydrogel, F-actin/DAPI staining was performed for the cells encapsulated in the hydrogel core at Day 1, 5, 7, and 17 ( Fig. 3(C)). These ndings demonstrate cell proliferation in a highly-organized pattern along with the ber guideline. These results indicate the overall biocompatibility of the engineered hydrogel scaffold.

Cellular alignment
Cellular alignment is a critical consideration for in vitro neural tissue model systems. We observed that the hybrid hydrogel system signi cantly promotes cellular alignment along the bers. Figure 4 shows the qualitative and quantitative results of the cellular orientation within the two-week periods after encapsulation. Fluorescent staining of neuroblastoma F-actin laments indicates the alignment of the cells with respect to the ber orientation (Fig. 4(A)). The orientation of F-actin lament was quanti ed and measured with respect to the direction of the channel. The results showed that after 7 days, actin laments were oriented along the channel with the standard deviation of less than 10 degrees. The orientation was improved at Day 14 with the standard deviation of less than 6 degrees (Fig. 4(B)). The cell culture density within the 3D tubular structure was estimated at Day 7 and 14 per unit volume of the channel to compare the seeding density and those of the cells on the following days. The results showed that the seeding density was 15 million/ml; whereas the density of cells after 7 and 14 days of culture was increased to 27 and 50 million/ml, respectively ( Fig. 4(C)).

Numerical simulation
Here, the shear stress occurring on the cells, and the cells con nement within the ber's core during the extrusion are numerically simulated using COMSOL ® Multiphysics. The geometry of two coaxial tubes with the inner diameters of 0.838 mm and 0.603 mm, and the length of 12 mm was generated. The Navier-Stokes equation and a two-phase discrete phase model (DPM) were used to describe the uid ow and the distribution of cells inside the tube, respectively. The simulations were conducted at different shell ow rates of 2, 3, and 4 ml min − 1 for the constant core ow rate of 2 ml min − 1 . Figure 5 shows the cross-section view for the cells distribution and the shear stress on the cells surface at different shell ow rates (2, 3, and 4 ml/min) and distances from the nozzle head in the channel. Initially, the cells are distributed within the radius of 0 < < 0.3 mm. After passing through the channel, they concentrate within the region with 0 < < 0.17 mm. The pressure gradient arisen from the shell ow pushes the cells towards the center of the core stream. Figure 6 shows the shear stress distribution at different radius for three shell ow rates of 2, 3, and 4 ml/min along the channel (the channel side view). At rst, a sudden jump in the shear stress pro le is observed. However, after passing a critical length, the shear stress stays approximately constant. We speculate that the directed axonal outgrowth is obtained due to three phenomena: 1) the spatial con nement of the cells in the bers and the stiffness of the core that applies mechanical stimulation to maintain the cells in a concentric distribution; 2) the alignment of polymeric sheets due to the ow induced shear stress during the extrusion process; and 3) the alignment of unmixed GelMA domains. GelMA and alginate do not mix homogeneously due to their dissimilar physical properties such as density, and thermal sensitive viscosity; since alginate hydrogel does not provide a cell friendly matrix, cells align toward the elongated GelMA networks (see Fig. 1). It has been shown that the size of hydrogel constructs play a key role in the morphological growth of the cells 59,60 . In smaller GelMA features, the cells can sense the boundaries of the GelMA microenvironment so that they align themselves within the boundaries. Here, due to the stiffness difference between the core and the shell, cells can align with the ber's axis.

Discussion
As it was shown in Fig. 4 (C), the cellular alignment together with the cellular density was increased after 14 days. The high density of the cells can be attributed to the highly oriented actin laments observed on Days 7 and 14, as the higher density of the cells causes the elongation of the cells in the channel direction.
Although the use of neuroblastoma cells showed an excellent cell alignment and directional neural growth, these types of cells are not the best model for neural stem cells. Considering this fact, more investigations in needed to adjust the properties of the current system, including geometrical and chemical properties of the bers, so it can prove its performance to recapitulate the in vivo environment.
Regarding the numerical simulations, it was shown that the system goes under a sudden jump in the shear stress, which after passing a critical length, it stays approximately constant. This suggests that increasing the shell ow rate delays the stress stabilization (longer critical length). As both Figs. 5 and 6 represent, the lowest shear stress value on the cells surface happens in the center of the channel and it increases gradually toward the channel walls. Moreover, the rise in shell ow rate causes the cells to experience higher shear stress.
The simulation results explain the stress-induced elongation of the polymer sheets due to ow extrusion force. Crosslinking of the hybrid hydrogel prevents the reorganization of these elongated structures. Thus, we speculated that the encapsulated cells were following this organization. The shear stress is higher close to the channel walls and therefore, elongation of the GelMA domains is higher in those regions.

Conclusion
A cell-laden composite hydrogel ber constructed with tunable mechanical properties was fabricated. A micro uidic system was used to fabricate core-shell bers with a cell-laden core and shell from a biocompatible cell-laden hybrid hydrogel of GelMA/gelatin. Despite many other studies on the application of coaxial bioprinting, the focus of this study was on the quanti cation of the cellular alignment and tunability of the scaffold constructs speci cally for neural cells. The hydrogel bers showed cell attachment, proliferation, and excellent growth within the ber's core axis. The bers directed the axonal outgrowth and generated a highly aligned cell pattern.

Fabrication of hydrogel bers
To fabricate the hydrogel bers, a co-ow device was utilized by assembling two needles of different sizes. The needles were chosen in a way that the outer dimeter of one is less than the inner diameter of the other one, where there would be a space between the needles for liquid to ow through. Here we used needle sizes of 18G and 20G to fabricate the device. The head of the needles were trimmed and attened using sandpaper. An inlet (focusing ow inlet) was made in the plastic chamber of the bigger needle. The needles were then assembled together and glued. A transparent chemical resistant clear PVC tubing was then connected to the outlet and placed in a 2% calcium chloride solution in water.
As it is illustrated in Fig. 1, hydrogel bers were fabricated using inner phase of 3% GelMA mixed with 2% gelatin (different concentrations were made), and the focusing ow phase of 2% sodium alginate. The two phases were injected into the inlets of the micro uidic device. The ow rates of the solutions were nely controlled by syringe pumps (PHD 2000, Harvard Apparatus) to create bers with different sizes. The hydrogel bers were collected in a Petri dish containing 2% calcium chloride in water to fully crosslink the hydrogel. The bers were then exposed to UV light (OmniCure s200, USA) for 30 seconds to crosslink the GelMA hydrogel in the bers. The UV intensity was set at 850 mW, and the distance between the tip of the ber optic and the Petri dish was set at 8 cm.
SH-SY5Y cells encapsulated in pre-described hybrid hydrogel using the following steps; after detaching cells (by trypsin), they were resuspended in cell culture media with the density of 15 M mL − 1 . The cells solution was added to the GelMA and gelatin hydrogel solution. Subsequently, GelMA/glatin containing the cell solution was mixed with PI solutions.

Viability assay (Live/Dead)
To evaluate the in vitro biocompatibility of the obtained hydrogel bers, SH-SY5Y encapsulated in the hybrid hydrogel bers were cut into smaller segments (1 cm) and placed in a 12-well plate. The viability test was performed by incubating the cell-laden samples for 15 minutes at 37°C with a mixture of 2 µl ml − 1 ethidium homodimer-1 (EthD-1, red, dead cells, Invitrogen) and 0.5 µl ml − 1 calcein AM (live cells, green, Invitrogen). Fluorescence Microscopy was performed using a Zeiss Observer D1 microscope of an X-Cite Series 120Q uorescence source.

Actin/DAPI staining
Samples were xed in 4% paraformaldehyde (Electron Microscopy Sciences), washed in DPBS and subsequently stained with phalloidin (life Technologies, labeled with Alexa F 594) and DAPI as described in the manufacturer's manual. Microscopy was performed as described above.