WO2026020249A1 - Flexible biodegradable and resorbable probe for electrical stimulation or recording electrical signals - Google Patents

Abstract

The present disclosure provides a flexible biodegradable and resorbable brain (or other tissue) stimulation electrode for temporally regulated neuromodulation of neural precursor cells (NPCs), which may also be used for recording electrical signals in the brain or other tissue. Using the cathodally skewed electrochemical window of molybdenum and the volumetric charge transfer properties of a conductive polymer, the electrodes are designed with high charge injection capacity for the delivery of biphasic monopolar stimulation. These electrodes are biocompatible and can deliver an electric field sufficient for NPC activation for at least 7 days post implantation before undergoing resorption in physiological conditions., thereby eliminating the need for surgical extraction. The biodegradable electrode demonstrated its potential to be used for NPC based neural repair strategies.

Description

FLEXIBLE BIODEGRADABLE AND RESORBABLE PROBE FOR ELECTRICAL STIMULATION OR RECORDING ELECTRICAL SIGNALS This international application claims the benefit of priority under Article 8 of the Patent Cooperation Treaty to Luxembourg Patent Application No.100010913, filed on 25 July 2024, the entire contents of which are incorporated herein by reference. FIELD The present disclosure relates to a flexible biodegradable and resorbable probe for electrical stimulation in a subject or recording electrical signals generated by a subject. BACKGROUND Neurological disorders are the leading cause of long-term disability worldwide¹. There are no current treatment options to replace lost cells following injury or disease. The need for novel therapeutic interventions to promote neural repair is clear. One potential strategy that has shown promise involves harnessing the potential of endogenous neural stem and progenitor cells (together termed neural precursor cells, NPCs) ^2,3. Endogenous NPCs are rare cells confined to the subventricular zone (SVZ) lining the walls of lateral ventricles in the mature brain ^3,4. Under homeostatic conditions, resident NPCs give rise to neural cells that contribute to ongoing neurogenesis in vivo. Activation of NPCs are electrosensitive cells that migrate in the direction of electric field and differentiate into neural phenotypes in response to an applied electric field⁵. Brain stimulation involves the delivery of electric fields via electrodes, an approach that has proven effective in activating resident NPCs using ex vivo and in vivo models ^6,7. Transcranial direct-current electric field stimulation (tDCS) has recently been shown to elicit NPCs responses as evidenced by the migration of NSC-derived neuroblasts in a model of ischemia⁸. One challenge associated with tDCS is that it is difficult to control the direction of the electric fields due to low spatial resolution and this can limit the impact of stimulation on NPC migration⁹. In addition, while direct current stimulation is effective in activating NPCs for migration and expanding the NPCs pool^{10– 13}, the unidirectional current flow poses the risk of tissue damage caused by charge accumulation within the tissue. The application of intracranial charge-balanced, monopolar biphasic (BPMP) stimulation overcomes these disadvantages by stimulating the brain with zero net charges injected, thereby reducing the risk of electrode and tissue damage at the interface ^6,14,15. The penetrating intracranial electrodes can also improve the spatial resolution of stimulation, providing a more controlled, localized electric field ¹⁶. A limitation to the current use of platinum electrodes for intracranial BPMP stimulation is the modulus mismatch and recent work has demonstrated that softer cryogel alternatives can deliver electrical stimulation with optimized parameters for NPC activation and reduced neuroinflammation¹⁷. However, the subsequent removal of an implanted device, if necessary following treatment, poses additional risks including surgical intervention and damage to the brain tissue. To address these challenges, the development of bioresorbable electrodes presents a promising alternative. A number of polymer substrates including poly(lactic- co-glycolic acid) (PLGA)¹⁸, polycaprolactone (PCL) ¹⁹ and silk ²⁰ (for example); metals including tungsten (W), molybdenum (Mo) and magnesium (Mg), and dielectrics such as silicon dioxide (SiO2) ^21,22 have demonstrated efficacy for the design of electrodes that degrade from days to weeks, thereby eliminating the need for surgical removal. SUMMARY There is provided a flexible probe, a connector device for use with a pair of flexible probes, and a method for the preparation thereof. Accordingly, in an aspect, there is provided a flexible probe. The flexible probe comprises a biodegradable body having first and second ends, the first end defining a region for electrical connection and the second end defines an electrically conductive site; a bioresorbable metallic electrode extending from the first end to the second end of the body; a biodegradable conductive polymeric coating located on at least a surface of the second end; and an insulating biodegradable polymeric sheath enveloping a portion of the body between the first and second ends. In a particular case, another surface of the second end is coated with one of a conductive polymeric and an insulating polymeric coating. In another case, the other surface of the second end is coated with an insulating polymeric coating. In yet another case, the bioresorbable metallic electrode is made of a metal of _{molybdenum (Mo), tungsten (W), magnesium (Mg), Iron (Fe) and Zinc (Zn), any mixture} of these metals, and alloys of any combination of these metals. In yet another case, the biodegradable conductive polymeric coating is made of any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. In yet another case, the insulating polymeric sheath is made of any of poly(lactic- co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. In yet another case, the probe is configured for insertion or implantation into a tissue, wherein the second end has a needle-like shape. In yet another case, when in use, the probe retains its structural integrity therefore maintaining the stimulation properties of the probes for electrical stimulation for up to 14 days. In yet another case, the electrically conductive site has a thickness in a range from about 20um to about 300 um, and a width in a range from about 100 to about 400 um. In yet another case, the body defines a reservoir proximate to the first end and a delivery channel extending from the reservoir towards an opening proximate to the second end. In yet another case, the delivery channel extends from the reservoir to the opening located at the second end. In yet another case, the probe is configured for surface contact with a tissue wherein the second end has a section having a film-like configuration to adapt to morphology of the surface of the tissue. In yet another case, two or more probes are used for the delivery of an electrical stimulation. In yet another case, the stimulation is a biphasic monopolar current-controlled stimulation. In yet another case, the pair of probes are used for the regulated neuromodulation of neural precursor cells. In yet another case, the neural precursor cells are endogenous brain neural precursor cells. In yet another case, the probes are used for further further deliver of a drug, virus, protein, small molecules, active ingredient or biologically active substance. In yet another case, the probes are used for the recording of an electrical signal, wherein the probe includes two active sites, one site for recording and one site being electrical ground. Accordingly, in an aspect, there is provided an implant. The implant (or connector device) comprises a support; and a pair of probes, the probes being mounted to the support in a parallel orientation to each other with the electrically conductive site of each probe facing each other, the probes being apart from each other by a distance between about 1 mm to about 3 mm. Accordingly, in an aspect, there is provided a method of manufacturing a flexible probe. The method comprises the steps of providing a resorbable or biodegradable substrate; removably securing a mask onto a surface of the substrate, the mask defining a preselected geometry of the probe; depositing on the masked substrate a bioresorbable metallic material which defines an electrode with the preselected geometry; removing the mask; coating a portion of the electrode with an insulating biodegradable or resorbable polymeric material resulting in partially insulated electrode; coating a whole surface of the partially insulated electrode with a layer of patterning polymer; defining an electrically conductive surface of the electrode by removing a section of the layer of patterning polymer; coating at least the conductive surface of the electrode with a layer biodegradable conductive polymer; and removing the layer of patterning polymer from the electrode resulting in a flexible biodegradable and resorbable probe having an electrode with a conductive portion and an insulated portion. I_{n a particular case of this method, the resorbable or biodegradable substrate is} made of any of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. I_{n another case of this method, the bioresorbable metallic electrode is made of} any metal of molybdenum (Mo), tungsten (W), magnesium (MG), Iron (Fe) and Zinc (Zn), any mixture of these metals, and alloys of any combination of these metals. In yet another case of this method, the insulating biodegradable or resorbable polymer is any one of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. In yet another case of this method, the patterning polymer is Parylene. In yet another case of this method, biodegradable conductive polymer is any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. In yet another case of this method, the method further comprises the step of defining a reservoir and a delivery channel within the probe, the reservoir being proximate to an end region of the probe and the delivery channel in open communication with the reservoir and extending from the reservoir towards an opposite end of the probe. Accordingly, in an aspect, there is provided a connector device for use with a first flexible probe having an electrode and a second flexible probe having an electrode and a microfluidic channel, each electrode having a side with an electrical conductive site. The device comprises a first housing portion having an inner wall, the inner wall being configured to receive the first probe with the electrically conductive site facing away from the inner wall; a second housing portion having an inner wall, the inner wall being configured to receive the second probe with the electrically conductive site facing away from the inner wall; a first sheet of conductive material configured to contact a electrical connection region of the electrode of the first probe; a second sheet of conductive material configured to contact the electrical connection region of the electrode of the second probe; a spacer electrically insulating the first sheet of conductive material from the second sheet of conductive material; two openings, each opening jointly defined by a top surface of each housing portions and configured to receive a pin; and a pair of pins, each pin extending through the respective opening and having an end affixed to an end of the corresponding sheet of conductive material, wherein the first and second housing portions are configured to securely mate with each other for defining a cavity enclosing each probe, the corresponding sheet of conductive material and the spacer, such that each probe is resting against the corresponding inner wall and each sheet of conductive material is positioned on opposite sides of the spacer whereby, in the assembled state, the electrical connection region of each probe is pressed against its respective sheet by contact pressure applied by the housing portions and the spacer conductively connecting the electrode of each probe to the corresponding pin, wherein a side surface of the second housing portion defines an opening in fluid communication with the microfluidic channel of the second flexible probe, and wherein a bottom surface of the second housing portion defines two openings, each opening is configured to receive a stimulation site of the electrode of the respective probe, the stimulation sites extending through the openings and projecting outwardly from the second housing portion. In a particular case, the connector further comprises a tubing received within the side opening and extending away from a side surface of the second housing. In an embodiment, the tubing is removably secured to the side opening. In another case, the tubing has an outer diameter corresponding to about the diameter of the side opening defined by the second housing portion thereby an external side wall of the tubing engages with the wall of the side opening. In this embodiment, the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing range between about 20 to about 23-gauges. In yet another case, the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing are about 20 gauges. In yet another case, the connector device further comprises a fluidic channel attachment extending outwardly from a side of the second housing portion and defining an opening concentrically aligned with the side opening of the second housing portion, wherein the opening of the fluidic channel attachment has a diameter greater than the diameter of the side opening of the second housing portion and the tubing is extending outwardly through the opening of the fluidic channel attachment. In yet another case, the opening defined by the fluidic channel attachment has a diameter of about 2 mm. In yet another case, the tubing is bonded to an inner wall of the opening defined by the fluidic channel attachment. In yet another case, the sheet of conductive material is made of copper or silver. In yet another case, the sheet of conductive material is made of copper. In yet another case, each pin has the affixed end conductively bonded to the end of the corresponding sheet of conductive material. In yet another case, the affixed end is soldered to the end of the corresponding sheet of conductive material. In yet another case, each housing portion has a bottom portion having a curvature such that when the housing portions are mating together, a resulting bottom portion of the device has an overall curvature configured to match the curvature of a head of a subject. In yet another case, each housing portion are made of 3-D printing plastic Clear V4^TM. In yet another case, the spacer is shaped as a block. In yet another case, the spacer is made of acrylic. In yet another case, the spacer has a width ranging from about 1.7 mm to about 2.0 mm. In yet another case, the spacer has a width of about 1.8 mm. In yet another case, the connector device further comprises a snap-fit clip system to securely mating both housing portions together wherein one of the housing portion has a pair of cantilevers, each cantilever parallelly extending from an opposite side of the housing portion and having a protrusion at an distal end configured to mate with the side of the other housing portion when the housing portions are engaged with each other. In yet another case, the inner wall of the housing defines a pair of slots, each slot extending inwardly into the inner wall adjacent to a corresponding cantilever, thereby increasing an overall length of the cantilever. In yet another case, the connector device further comprises a pair of pin clips, each pin clip being configured to slot a respective pin into the corresponding opening jointly defined by the top surfaces of the housing portions, the pair of pin clips being located on the top surface of the first housing portion and axially aligned with the two openings. Accordingly, in an aspect, there is provided a device. The device comprises a first neural probe having an electrode with an electrical conductive site, the first neural probe defining a through-opening; a second neural probe having an electrode with an electrical conductive site and a microfluidic channel, the second neural probe defining a through-opening; a central spacer defining a first channel extending from a side surface to an opposing side surface, each probe being secured to one of the opposite sides of the spacer with the electrical conductive site of each probe facing each other and each probe having the opening in alignment with the first channel; a first outer body secured to first probe, the first outer body defining second channel extending from a side surface to an opposing side surface, the second channel being in alignment with the opening of the first probe; and a second outer body having a solid surface secured to the second probe, wherein the opening of each probe is in alignment with the first channel and the second channel for fluid communication between the second channel and the microfluidic channel of the second probe, and wherein the first channel is closed by the solid surface of the second outer body. In a particular case, the device further comprises a tube segment received within the second channel and extending outside the first outer body. In a particular case, the second channel has a diameter greater than the diameter of the first channel. A further understanding of the functional and advantageous aspects of the disclosure can be realized by reference to the following detailed description and drawings. BRIEF DESCRIPTION OF THE DRAWINGS Embodiments disclosed herein will be more fully understood from the following detailed description thereof taken in connection with the accompanying drawings, which form a part of this application, and in which: Figures 1A to G: Design and fabrication of biodegradable electrode for brain stimulation.1A) Design of biodegradable electrodes. The dimensions of each of the cortical penetrating probes are 2 mm in length and 300 μm wide in the cortex. The width of molybdenum probe is 150 μm.1B, 1C) Bright field (scale bar = 1 mm) and enlarged image (c) (scale bar = 0.2 mm) of the stimulation electrode. The uninsulated stimulation site is 350 μm by 1000 μm.1D, 1E) Optical images (scale bar =1 cm) for demonstration of flexibility of the electrode array (d) and probe implants after laser cut (e).1F) Set-up for brain stimulation with current controlled charge-balanced biphasic monopolar stimulation. The optimized waveform for activating NPCs is delivered through the stimulator. The data acquisition (DAQ) unit is connected to record the voltage.1G) Schematic of coronal hemisection showing the placement of biodegradable electrodes in the cortex of the mouse brain. During stimulation, neural precursor cells (NPCs) in the subventricular zone lining the walls of lateral ventricle (LV) proliferate and migrate towards the cathode (-) in response to the applied electric field (EF). +: anode. Figures 2A to 2J: Stimulation capacity of biodegradable Mo electrodes and Mo electrodes with conductive polymer coating.2A) Molybdenum (150 μm width by 2 mm length, 0.003 cm²), gold (150 μm by 2 mm, 0.003 cm²), and platinum (127 μm diameter by 1 mm length, 0.004 cm²) electrodes used for electrochemical characterization.2B) _{top: Electrochemical water window of 150 nm sputter deposited Mo electrodes,} measured in aCSF by recording cyclic voltammetry at a slow scan rate of 20 mV/s. Bottom: Cyclic voltammetry curves and CSC of Mo, Pt and Au electrodes at 200 mV/s, measured in aCSF (n = 3 trials/group, data presented as mean ± S.D.).2C) Voltage transient curves of Mo, Pt and Au in aCSF given charge-balanced, biphasic current pulse with a magnitude of 100 µA and 60 µs pause between the cathodal and anodal charge injection.2D) Corresponding voltage of Au, Mo, MoPH3 electrodes recorded given current-controlled charge balanced biphasic waveform in aCSF.2E) Cyclic voltammetry curves at 200 mV/s of bare Mo, 1 (MoPH1), 3 (MoPH3) and 6 (MoPH6) layers of PEDOT coating on Mo electrodes (350 μm by 1 mm) measured in aCSF.2F) Optical images of PEDOT-coated Mo electrodes, (scale bar = 0.2 mm).2G, 2H) Equivalent Randles circuitry (g) and impedance spectra (h) with phase angle of bare Mo electrode and different layers of PEDOT coatings on Mo electrodes.2I) Voltage transient curves of bare Mo electrode and with 1, 3 and 6 layers of PEDOT coating (PH1, PH3, PH6) on Mo electrodes in aCSF, given a charge-balanced, biphasic current pulse with a magnitude of 200 µA and 60 µs pause between the cathodal and anodal charge injection.2J) Current injection limit and CIC of bare Mo electrode with increasing layers of PEDOT coating in aCSF (see Table 5), calculated based on voltage transient _{curves in Figure 16 (n = 3 trials/group).} Figures 3A to 3L. Degradation properties of stimulation electrode.3A) Molybdenum (Mo) and sulfur (S) concentration in ppm of DI water and aCSF used to soak MoPH3 electrodes.3B) Change in CSC of PEDOT:PSS-coated Mo electrodes from day 0 to day 7 soaked in aCSF at 37°C. (n = 3 trials/group, data presented as mean ± S.D).3C) Maximum cathodal and anodal electrochemical potential excursions of PEDOT:PSS-coated Mo electrodes in aCSF at 37°C day 7, given a range of injected current magnitudes (n = 3 trials/group, data presented as mean ± S.D.). cat=cathode, an=anode; Ema: max. allowable anodal potential excursion 0.2 V; Emc: max. allowable cathodal potential excursion -1.0 V. Curve fitting and CIC results are presented in Table 7.3D) Change in charge injection capacity of PEDOT-coated Mo electrodes in aCSF at 37°C between day 0 and day 7.3E) Impedance spectra of MoPH3 electrode at day 0 and day 7, soaked in aCSF at 37°C.3F) Voltage transient curves of MoPH3 electrode at day 0 and day 7 in aCSF at 37°C, given charge-balanced, biphasic current pulse with a magnitude of 200 µA and a 60 µs pause between the cathodal and anodal charge injection.3G-3L) Degradation of MoPH3 biodegradable electrodes over 12 weeks in aCSF at 37°C. Inset: bright field microscopic images of the electrode tip (5x). Figures 4A to 4E. Biocompatibility.4A-4B) Photomicrographs of Au and MoPH3 biodegradable electrode implantation sites (dotted lines) in coronal brain sections stained with Iba1+ (green) GFAP+ (red) cells (a) and NeuN+ (cyan) cells (b) after 8 weeks of implantation (scale bar: 1000 µm). Regions for cell quantification are selected 1 mm below the surface of cortex (higher magnification photomicrograph scale bar: 200 µm). Arrows indicate examples of labeled cells.4C-4E) Numbers of Iba1+ (c), GFAP+ (d) and NeuN+ (e) cells in a 200 µm X 200 µm representative area adjacent to the implantation site and 1mm below the cortical surface (dashed square boxes in a,b) (n = 3 mice/group, two-way ANOVA; *p < 0.05). Data presented as mean ± SEM. Figures 5A to 5F: Neurostimulation leads to increased numbers of NPCs in vivo. 5A) Investigation of the expansion of the NPC pool in vivo following electrical stimulation with biodegradable electrode at day 0 and day 7 post-implantation.5B) Schematic of stimulation in the cortex of mice. Two X 2 mm-long electrodes (with 1 mm insulation, shown in blue line) are located 2 mm apart, 0.7- and 2.7-mm lateral to the brain midline (Ml). Ctx: cortex, Dlc: dorsolateral corner, cc: corpus callosum, SVZ: subventricular zone lining of the lateral ventricles, Str: striatum. Dashed lines denote the electrical field distribution.5C) COMSOL simulation reveals a 63.85 V/m electric field delivered to NPCs (Dlc) during cathodal pulse of -200 µA. Electrical conductivity for each brain compartment used for simulation is found in Table 8.5D) Recorded corresponding voltage at the start and end of the 3-hour stimulation given optimized current-controlled stimulation waveform.5E, 5F) Fold change of the number of neurospheres in the contralateral hemisphere (without electrodes) and ipsilateral hemisphere (with electrodes) at 1-hour post-stimulation at day 0 (e) (n = 4 mice/group) and day 7 (f) (n = 3 mice/group) post-electrode implantation (two-way ANOVA (α = 0.05); *p < 0.05). Data presented as mean fold change ± SEM. Figure 6. Degradation mechanism of poly(lactic-co-glycolic acid) (PLGA) in water. Figure 7: Schematic illustration of deposition process for molybdenum electrode patterns. Figure 8: Schematic illustration of PEDOT conductive polymer patterning process and electrode insulation. Figures 9A to 9B: Cathodal charge storage capacity (CSCc) (measured in PBS) _{of 0.5 by 0.5 mm spin coated PEDOT annealed at 130°C and room temperature. The} CSCc of PEDOT annealed at 130°C and room temperature is similar regardless of the scan rate: 1.06 ± 0.017 mC·cm^-2 and 1.04 ± 0.041 mC·cm^-2 respectively at 200 mV·s^-1 (9A); 1.92 ± 0.042 mC·cm^-2 and 2.02 ± 0.106 mC·cm^-2 respectively at 20 mV·s^-1(9B) (n = 3 trials/group, data presented as mean ± S.D.). Figure 10: Front, side and isometric view of electrode connectors, printed with Formlabs 2 SLA printer. The overall dimension of the connector is 6 by 5 by 4.25 mm. Figures 11A to 11D: Example voltage transient curves of Pt (11A), Au (11B) and Mo (11C) electrodes in aCSF given charge-balanced, biphasic current pulse with increasing magnitudes and 60 µs pause between the cathodal and anodal charge injection.11D) Maximum cathodal and anodal electrochemical potential excursions of Mo, Pt and Au in aCSF given a range of injected current magnitudes (n = 3 trials/group, data presented as mean ± S.D.). F_{igures 12A to 12B: 12A) shows Bode and 12B) shows Nyquist impedance} plots of Mo, Pt and Au (inset: equivalent Randles circuit during brain stimulation). Figures 13A to 13B: 13A) Cyclic voltammetry curves at 20 mV/s of bare Mo, 1 (MoPH1), 3 (MoPH3) and 6 (MoPH6) layers of PEDOT coating on Mo electrodes measured in aCSF.13B) Charge storage capacities (CSC) derived from CV curves of bare Mo and different layers of PEDOT coatings on Mo (n = 3 trials/group, data presented as mean ± S.D.). Figure 14: Flowchart of Randles circuitry curve fitting algorithm. Figure 15A to 15E. Curve fitting of equivalent Randles circuit of bare Mo electrodes with increasing layers of PEDOT conductive polymer coating.15A) Example Bode and Nyquist fits of bare Mo electrodes and PEDOT-coated Mo electrodes (inset: equivalent Randles circuit during stimulation).15B-15E) Derived solution resistance (Rs) (15B), charge transfer resistance (Rct) (15C), double layer capacitance (Cdl) (15D) and Warburg coefficient (W) (15E) from curve fitting.10 Figures 16A to 16E. Example voltage transient curves of bare Mo electrodes (16A) (area: 0.0015 cm²) and 1 (16B), 3 (16C) and 6 (16D) layered PEDOT-coated Mo electrodes in aCSF given charge-balanced, biphasic current pulse with increasing magnitudes and 60 µs pause between the cathodal and anodal charge injection.16E) Maximum cathodal and anodal electrochemical potential excursions of bare Mo electrode and different layers of PEDOT coatings on Mo electrodes in aCSF given a range of injected current magnitudes (n = 3 trials/group, data presented as mean ± S.D.). Curve fitting results are presented in Table 5. Figures 17A to 17C: Degradation properties of Mo electrodes and PEDOT coating.17A) Thickness of Mo electrodes measured using optical profiler.17B) Thickness of Mo electrodes in water and aCSF over time at 37 °C.17C) Structural changes of PEDOT:PSS chains in aqueous solutions over time. At first, dissolution of lightly bound hydrophilic PSS-chains occur, followed by PEDOT:PSS chain reorganization and detachment from PLGA surface ³⁸. Figure 18: Voltage transient curve of MoPH6 electrode at day 0 and 7 in aCSF at 37°C, given charge-balanced, biphasic current pulse with a magnitude of 200 µA and 60 µs pause between the cathodal and anodal charge injection. Figure 19: Accelerated degradation the electrode device, characterized by mass loss and pH change of aCSF used to soak the electrode device at 90 °C. The electrode device lost its integrity for mass measurement at 12 hours. The pH of aCSF started to stabilize at around hour 18, which is a sign of complete degradation of the device. Figure 20: Photomicrographs of Au and MoPH3 biodegradable electrode implantation sites in coronal brain sections stained with Iba1+ and GFAP+ cells and NeuN+ cells at 4 weeks post-implantation. Regions for cell quantification are selected 1 mm below the surface of cortex (Scale bar = 1 mm). Figure 21: Stress-strain curves of 50 µm-thick PLGA films tested using 3-point bending set-up. Flexural moduli were derived from 0.05% to 0.1% strain. Figures 22A to 22C: 22A) Mouse under 2% isoflurane anaesthesia with head secured to the stereotactic system. Electrodes are implanted with headcap secured on the skull. Fold change of the number of neurospheres in the contralateral hemisphere (without electrodes) and ipsilateral hemisphere (with electrodes, not stimulated) on day _{7: 22B) (n = 4 mice/group) and day 7; 22C) (n = 3 mice/group) of electrode implantation.} Data presented as mean ± SEM. Figure 23: Biodegradable microfluidic channel fabrication.1) Patterned PDMS is created using preetched silicon wafer with feature depth of 50 μm and width of 50 μm. 2) Compression molding and solvent casting of PLGA pellets creates 50 μm-thick flat and 100 μm-thick patterned PLGA films, respectively.3) Thermal bonding technique allows intimate contact between patterned and flat PLGA films.4) Polymer chains are mobilized, and interdiffused to form a monolithic biodegradable sheet with enclosed microchannels.5) Laser is used to align and cut the sheet into probe shapes with optimized cutting power and speed.6) The finished device probe has a diameter of about 150 μm with the microchannel located in the middle. Figure 24: Microfluidic injection probe assembly.1) PLGA biodegradable probe with delivery channel and stimulating electrodes.2) Surface functionalization of hole- punched PDMS block with plasma activation.3) Biopsy punched 50 μm-thick polyimide film with adhesive side. Surface functionalization of non-adhesive side of PI film with silicon with 3-mercaptopropyltrimethoxysilane (MPTMS). Hydrolysis near alkoxy terminal of MPTMS forms silicon-oxygen network while nucleophilic reaction occurs between PI and mercapto group of MPTMS. Plasma activation of surface treated PI films gives hydroxyl group readily for bonding.4) Permanent, instantaneous room temperature bonding between PI film and PDMS to form injection connector. 5) Alignment, adherence, and lamination of injection connector to the probe.6) Insertion of 28G PTFE tubing to PDMS block for connection between fluid reservoir and the injection probe. Figures 25A to 25D: In vitro injection test with agarose gel.25A) Coomassie blue dye with a concentration of 0.33 μg/μL in DI water is being injected 2 mm deep into the 0.6% agarose gel using a syringe pump at a rate 1 μL/min.25B) Injection of 6 μL of dye-loaded DI water results in no backflow or backtracking along the probe.25C) Dye diffuses into the agarose gel after 1 hour.25D) Diffusive properties as reported by volume change of the dye with respect to time. Figures 26A to 26C: In vivo injection tests.26A) Flow rate-controlled syringe pump is used to deliver Dextran dye (b), aCSF (c) and Brain Derived Neurotrophic Factor (BDNF) (c) at a rate of 1 μL/min. to 2 mm below the skull into the cortex near the SVZ.26B) Injection of 6 μL of Dextran Taxes Red dye with a concentration of 0.33 μg/μL in DI water.26C) Injection of 5 μL of aCSF and 5 μL of BDNF in 0.1% BSA with a concentration of 0.2 μg/μL in aCSF. Figures 27A to 27B: 1 hour after fluorescent Dextran dye injection to cortex. 27A) Injection results in approximately 1.5 mm diffusion span along the anterior- posterior axis of the brain.27B) Injection results in approximately 1 mm diffusion span along the lateral-medial axis. All dye is injected into the cortex. No leakage, backflow or significant backtracking observed. Scale bar: 750 μm. Figures 28A to 28B: Exogenous BDNF is found at the site of the injection.28A) Sections from control mice without any injections.1: No cells in the cortex at the level of the LV.2: endogenous BDNF is highly expressed in the hippocampus. 28B) Coronal sections from mice with aCSF (left) and BDNF (right) injections.3: No cells expressing BDNF are seen in the cortex following aCSF injection.4: Cells expressing BDNF in the cortex where BDNF is injected. Scale bar: 150 µm. Figures 29A to 29D: Degradation and injection test of biodegradable microfluidic delivery device.29A) SEM images of PLGA microfluidic channel cross sections at week 0, 2, 3 and 4 in aCSF at 37 °C. Scale bar: 100 μm.29B) Channel geometry (cross- sectional area) over a period of 4 weeks of degradation. No significant change in channel cross-section area is seen over a period of 4 weeks. Error bars denote standard deviation.29C) Top view of a PLGA film showing the flat surface and porous core at 4 weeks.29D) Injection testing at 10 μL/min. with PLGA microfluidic device at 0, 1, and 2 weeks. F_{igure 30A to 30I: 30A-30B) Microscopic images showing degradation of the} MoPH3 electrode tip in aCSF fluid (a) and 0.5% aCSF agarose gel (b) at 37 °C. Scale bar: 100 µm.30C) Thickness of Mo electrodes in aCSF fluid and 0.5% aCSF agarose gel over time at 37 °C. (30D-30E) Change in charge storage capacity (CSC) (30D) and impedance (at 285 Hz) (30E) of MoPH3 electrodes from day 0 to 7 when soaked in aCSF fluid only or 0.5% aCSF agarose gel at 37 °C. (n = 3 trials/group, data presented as means ± S.D.).30F) Impedance spectra of the MoPH3 electrode at day 0 and day 7, soaked in aCSF fluid or in 0.5% aCSF agarose gel at 37 °C.30G) Maximum cathodal and anodal electrochemical potential excursions of MoPH3 electrodes at day 0 and day 7 in aCSF and 0.5% agarose aCSF gel at 37°C, given a range of injected current magnitudes (n = 3 trials/group, data presented as mean ± S.D.).30H) Voltage transient curves of MoPH3 electrode at day 0 and day 7 in aCSF and 0.5% agarose aCSF gel at 37°C, given charge-balanced, biphasic current pulse with a magnitude of 200 µA and a 60 µs pause between the cathodal and anodal charge injection.30I) Change in current injection limit of MoPH3 electrodes in aCSF fluid and 0.5% agarose aCSF gel at 37°C from day 0 to day 14. Note that the electrochemical measurements of electrodes in aCSF fluid after 2 weeks were noisy and exhibited high variability due to degradation, and therefore these results are not included. Figure 31A to 31E: Variant of the stimulation electrodes and microfluidic channel device (or connector device) and fabrication method thereof.31A) Fabrication of the device.31B) The electrode probe with the microchannel attached to the 22-gauge opening 2-mm thick block.31C-31D) Arrangement of the two acrylic blocks with sealant applied to the 2-mm diameter tubing for preventing leakage and backflow.31E) Injection test with the connector device showing no backflow or leakage. Figure 32A to 32B: Variant of the stimulation electrodes and microfluidic channel device (or connector device), 32A) With dimensions, 32B) Exploded view with components labeled, (Units: mm). Figure 33: Schematic showing the pin and sheet of conductive material with probe. Figure 34: Housing portion with snap fit clips and pin clip features; top view and side view (Units: mm). Figure 35: Housing portion with tubing attachment feature for drug delivery; side view and cross-sectional view taken along line A-A (Units:mm). Figure 36: Cured housing portion with tubing inserted in the fluid channel attachment. Figure 37: Picture of the assembled connector device. Figure 38: Assembly steps for the making of the connector device. Figure 39: Picture showing the droplet formation at the extremity of a probe with micro-fluidic channel. DETAILED DESCRIPTION A detailed description is provided below to facilitate a thorough understanding of the disclosed embodiments and connections thereof. The description is not limited to any particular example included herein. Various embodiments and aspects of the disclosure will be described with reference to the details discussed below. The following description and drawings are illustrative of the disclosure and are not to be construed as limiting the disclosure. Numerous specific details are described to provide a thorough understanding of various embodiments of the present disclosure. The Figures are not to scale. Further, in certain instances, well-known or conventional details are not described in order to provide a concise discussion of embodiments of the present disclosure. As used herein, the terms, “comprises” and “comprising” are to be construed as being inclusive and open ended, and not exclusive. Specifically, when used in the specification and claims, the terms, “comprises” and “comprising” and variations thereof mean the specified features, steps or components are included. These terms are not to be interpreted to exclude the presence of other features, steps or components. As used herein, the term “exemplary” means “serving as an example, instance, or illustration,” and should not be construed as preferred or advantageous over other configurations disclosed herein. As used herein, the terms “about” and “approximately”, when used in conjunction with ranges of dimensions of particles, compositions of mixtures or other physical properties or characteristics, are meant to cover slight variations that may exist in the upper and lower limits of the ranges of dimensions so as to not exclude embodiments where on average most of the dimensions are satisfied but where statistically dimensions may exist outside this region. It is not the intention to exclude embodiments such as these from the present disclosure. Unless otherwise specified, the terms “about” and “approximately” mean plus or minus 25 percent or less. It is to be understood that unless otherwise specified, any specified range or group is as a shorthand way of referring to each and every member of a range or group individually, as well as each and every possible sub-range or sub-group encompassed therein and similarly with respect to any sub-ranges or sub-groups therein. Unless otherwise specified, the present disclosure relates to and explicitly incorporates each and every specific member and combination of sub-ranges or sub-groups. As used herein, the term "on the order of", when used in conjunction with a quantity or parameter, refers to a range spanning approximately one tenth to ten times the stated quantity or parameter. According to an embodiment as shown in Figure 1A to 1E, a flexible probe (10) is provided. The probe (10) has a biodegradable body (101) having a first end (102) and a second end (103), a bioresorbable metallic electrode (20) extending from the first end (102) to the second end (103) of the body (101), a coating (40) made of biodegradable conductive polymeric located on at least a surface of the second end (103), and an insulating biodegradable polymeric sheath (30) enveloping a portion of the body (101) between the first end (102) and second end (103). As shown in Figure 1A, the first end (102) has a region for electrical connection (201) and the second end (103) has at lease a surface defining an electrically conductive site (202) According to an embodiment, the other surface of the second end (103) of the probe (10) may be coated with a conductive polymeric coating or an insulating polymeric coating. Alternatively, the surface may be coated with an insulating polymeric coating. According to an embodiment, the bioresorbable metallic electrode (20) may be made of metal such as molybdenum (Mo), tungsten (W), magnesium (Mg), Iron (Fe) and Zinc (Zn), any mixture of these metals or alloys made of any combination of these metals. The biodegradable conductive polymeric coating (40) may be made of poly(3,4- ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) or any derivatives thereof. The insulating polymeric sheath (30) may be made of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper or cellulose based biodegradable materials. According to an embodiment, the electrically conductive site (202) has a thickness in a range from about 20µm to about 300 µm, and a width in a range from about 100 µm to about 400 µm. According to an embodiment, as best shown in Figure 1A to 1G, the probe (10) is configured/shaped for insertion or implantation into a tissue (such as brain tissue 5) of a subject (2), wherein the second end (103) has a needle-like shape. According to an embodiment, the probe (10) is configured for surface contact with a tissue (such as brain tissue (5)) of a subject (2) wherein the second end (103) has a section having a film-like configuration to adapt to morphology of the surface of the tissue (5). According to an embodiment, as best seen in Figure 24 and Figure 25, the body (101) has a reservoir (105) proximate to the first end (102) and a delivery channel (106) extending from the reservoir (105) towards an opening (107) proximate to the second end (103). Alternatively, the delivery channel (106) extends from the reservoir (105) to the opening (107) located at the second end (103). According to an embodiment, the probe (1), when in use, retains its structural integrity such that the probe (10) can maintain its stimulation properties for electrical stimulation for up to 14 days. According to an embodiment, the two or more probes (10) are used for the delivery of an electrical stimulation. During use, the probes (10) produce a biphasic monopolar current-controlled stimulation. According to an embodiment, two probes (10) are used for the regulated neuromodulation of neural precursor cells (504). According to another embodiment, the neural precursor cells (504) are endogenous brain neural precursor cells. According to an embodiment, the probes (10) are used not only to provide electrical stimulation but also for the delivery of a drug, virus, protein, small molecules, active ingredient or biologically active substance. According to an embodiment, the probes (10) may be also used to record an electrical signal. In this situation, the probe (10) has two active sites, one site for recording and one site being electrical ground. According to an embodiment, as shown in Figure 10, an implant (or connector device) (60) is provided. The implant (60) has a support (601), a base (603) with two openings (602) and a pair of probes (10). The probes (10) have their penetrating portion extending through the openings (602) outside the support (601). The electrical connection region (201) of each probe (10) extends upwardly outside the support (601). As shown in Figure 10, the base (603) has a curvature to rest on the curvature of the head’s subject when in use. The probes (10) are mounted to the support (601), with a majority of the electrodes (20) residing within the support (601), in a parallel orientation to each other with the electrically conductive site (202) of each probe facing each other (the conductive site (202) being part of the stimulation site (206). The probes (10) may be apart from each other by a distance between about 1 mm to about 3 mm. According to an embodiment, a method of manufacturing a flexible probe (10) is provided. The method comprises the steps of: providing a resorbable or biodegradable substrate; removably securing a mask onto a surface of the substrate, the mask defining a preselected geometry of the probe; depositing on the masked substrate a bioresorbable metallic material which defines an electrode with the preselected geometry; removing the mask; coating a portion of the electrode with an insulating biodegradable or resorbable polymeric material resulting in partially insulated electrode; coating a whole surface of the partially insulated electrode with a layer of patterning polymer; defining an electrically conductive surface of the electrode by removing a section of the layer of patterning polymer; coating at least the conductive surface of the electrode with a layer biodegradable conductive polymer; and removing the layer of patterning polymer from the electrode resulting in a flexible biodegradable and resorbable probe having an electrode with a conductive portion and an insulated portion. According to an embodiment, as shown in Figure 24, the method further comprises the step of defining a reservoir (105) and a delivery channel (106) within the probe (10). The reservoir (105) is located proximate to the end region (102) of the probe and the delivery channel (106) is in open communication with the reservoir (105) and extending from the reservoir (105) towards an opposite end (103) of the probe (10). According to an embodiment, the resorbable or biodegradable substrate used for the manufacturing of the flexible probe (10) may be made of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, _{rice paper or cellulose based biodegradable materials. The bioresorbable metallic} electrode used for the manufacturing of the flexible probe (10) may be made of molybdenum (Mo), tungsten (W), magnesium (MG), Iron (Fe) or Zinc (Zn). The material may also be any mixture of these metals or alloys of any combination of these metals. The insulating biodegradable or resorbable polymer used for the manufacturing of the flexible probe (10) may be poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper or cellulose based biodegradable materials. The patterning polymer used for the manufacturing of the _{flexible probe (10) is Parylene. The biodegradable conductive polymer used for the} manufacturing of the flexible probe (10) may be poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) or any derivatives thereof. Materials and Methods Electrode fabrication The electrode fabrication procedure is divided into 4 parts: electrode substrate fabrication, electrode deposition, conductive polymer patterning and device insulation. _{A. PLGA substrate fabrication} The PLGA substrates were fabricated with compression molding using flat PDMS surfaces to ensure smoothness. Briefly, a clean, 4-inch silicon wafer was firstly treated with trichloro(1H,1H,2H,2H-perfluorooctyl)silane (Sigma Aldrich, Canada) via vapour phase silanization (3 hours in a sealed glass chamber heated to 180 degrees Celsius) to increase the hydrophobicity of the surface. Two blocks of PDMS (Sylgard 184, base 10: 1 curing agent) with dimensions 7.5 cm X 5 cm X 0.5 cm was then cured on top of surface treated silicon wafer using laser-cut acrylic sheets. Two PDMS blocks, cleaned with IPA and DI water, and dried under nitrogen air, were then used to sandwich PLGA pellets for compression molding at 150 degrees Celsius at 1000 psi using a laser-cut 50-μm-thick polyimide mold. The 5 cm X 2.5 cm X 50 μm PLGA film was then carefully taken out from the compression mold.50 μm was selected to be the thickness of the electrodes to prevent buckling during insertion. Buckling of electrode penetration To ensure smooth penetration to the brain, the electrodes need to be stiff enough to avoid buckling during insertion. The insertion force must be lower than the Euler’s critical buckling force (Fcritical) which is dependent on the elastic modulus of the probe material (E), area moment of inertia (I), effective length factor (K) and the unsupported length of the electrode during insertion (L): ^ 2 EI F critical _{^ 2 (1)} Using equation 1, the can be determined. The minimum critical buckling force must be higher than the force required for a neural probe to penetrate the cerebral cortex, which ranges from 321 to 464 μN, as, measured by previous studies. The flexural modulus for PLGA substates was measured to be 3.261 ± 1.066 GPa, comparable with the reported modulus measured by tensile test and nanoindentation test⁴⁹, and AFM⁵⁰. An effective length factor of 2.1 is chosen based on the assumption that both translation and rotation of the electrode are fixed at the connector and are free at the probe tip. Substituting the length of electrode L = 2 mm, the area moment of inertia is calculated to be 2.7432 e-19 m⁴. Then, the required thickness can be calculated based on both Ix and Iy axis through the centre of mass to be: ^ b = 0.077 μm (2) ^ b = 21.11 μm (3) Hence, the thickness of the electrodes must be at least 104 μm to facilitate penetration to cortex. The electrodes are designed and fabricated to have a thickness of 0.002’’ (50.8 μm) to achieve a critical buckling force of 6.97 mN that exceeds the force required for cortex penetration of about 500 μN ⁵¹. _{B. Mo electrode deposition (Figure 12)} PVA were used to improve adhesion between electrode and PLGA substrate. To prepare the PVA solution, 1 wt% of PVA (Mw 130,000, 99+% hydrolyzed) was dissolved in DI water heated at 90 degrees Celsius. To introduce the adhesive layer, the hydrophobic surface of PLGA films were firstly modified with plasma activation, followed by spin coating of the 1% PVA solution at 2000 rpm for 30 s. After coating, the PVA- coated PLGA substrates were annealed at room temperature overnight. To deposit Mo to PLGA substrates, a laser-cut metal stencil (Stencils Unlimited) was used as mask for electrode patterning on the 50-μm thick PLGA substrates. Neodymium magnets were used to ensure zero-gap between the metal mask and substrate.300 nm Mo was deposited (AJA International ATC Orion 5 Sputter Deposition System) at a rate of 1 Å/s. After deposition, magnets and metal masks were carefully removed, yielding Mo electrode patterns on PLGA substrates. C. Conductive polymer patterning (Figure 13) Parylene-C patterning was used to create conductive polymer patterns on PLGA substrates. Briefly, 10 μm of parylene-C was deposited (SCS 20120 Parylene Coater) to the PLGA substrate with Mo electrode patterns. Parylene-C was then patterned with CO2 laser cutter (Universal Laser System VLS3.75; power 2%, speed 100%) such that it cut through the parylene-C layer with minimal disturbance to the underlying PLGA film. Parylene-C in areas designated for conductive polymer coating were then peeled off with ultra-fine tweezer under a microscope, followed by 10 min. of corona activation of the sample for improving the wettability of conductive polymer solutions during coating. To prepare PEDOT solution for spin coating, 5 v/v% ethylene glycol (Sigma), 0.5 μl/mL of dodecyl benzene sulfonic acid (DBSA, Sigma) and 1 wt% of 3- glycidoxypropyltrimethoxy-saline (GOPS, Sigma) was added to PEDOT:PSS aqueous dispersion (PH1000, Clevios^TM) to improve conductivity, surface wettability and promote crosslinking respectively. This formulation has been shown to yield highest conductivity compared with pristine PEDOT:PSS solution and post-treatments with acid (Table 1). PEDOT solution was then spin coated on PLGA film at 1000 rpm to give 180 nm thickness after annealing ³⁴. After single or multiple layers of spin coating, PEDOT- coated samples were then annealed at room temperature for 24 hours before peeling off the sacrificial Parylene-C layer. Figure 14 shows that the charge storage capacity of PEDOT annealed at room temperature for 24 hours is comparable to that of annealed at 130 °C for 1 hour. Table 1: Electrical conductivity of annealed PEDOT:PSS films after different post- treatments. P_{rocessing Method Treatment Thickness Conductivity} Literature (S/cm) Cast mold at 130 C, None 10 μm 0.6 to 0.9 Order of 1 anneal for 15 min. S/cm Cast mold at 130 C, MSA ~ 5 μm ~880 >2500 anneal for 15 min. S/cm Spin coat at 500 rpm EG + DBSA + ~210 nm 1500 to 2000 1330 to for 30s, followed by GOPS 5000 S/cm annealing at 130 C for 15 min. D. Insulation and probe fabrication (Figure 13) For insulation layer, a 25-μm thick PLGA film was fabricated by solvent casting. Briefly, 0.041 g PLGA pellets were dissolved in acetone to 50 mg/mL and solvent casted with stereolithography (SLA) printed mold clamped securely on top of flat PDMS surface. The PLGA film was then laser cut and adhered to the bare Mo electrode part of the device for insulation (Figure 13). The electrode patterns were then aligned, and laser cut to fabricate 350 μm diameter neural probe for electrical stimulation. Device fabrication and assembly For device assembly, the laser-cut probes were assembled onto a 3-D printed connector for connection to the neurostimulator. The connectors were originally designed in SolidWorks 2021 (Dassault Systèmes SE, France) and produced using Form 2 desktop SLA 3D printer (Formlabs Inc, USA) with Clear FLGPCL04 photopolymer resin. The laser-cut probes were attached to the 7 mm X 4 mm X 5 mm connector with polyimide adhesive. M1.5 screws were installed to lead the connection from the sputtered Mo electrode to the tip of machine pin. A 25-μm diameter copper wire was soldered to the tip of machine pin and wrapped around the head of the screws. A probe length of 2 mm was ensured from the bottom of the connector to the tip of the probe (Figure 15A-15E). Characterization Electrochemical characterization Cyclic voltammetry and electrochemical impedance spectroscopy analyses were conducted using electrochemical analyzer (CHI6054E, CH Instruments, USA) to determine the charge storage capacity (CSC) and interfacial impedance, respectively. To calculate the charge injection capacity (CIC), stimulus generator (STG4002-1.6mA) was used to perform current injections while the corresponding voltage transients were recorded with National Instrument 6255-USB Analog Data Acquisition (DAQ) box. Three-electrode configuration was adapted for all electrochemical characterizations (reference electrode RE: Ag/AgCl; counter electrode CE: platinum; working electrode WE: the electrode of interest Au, Mo, MoPH1-6). Artificial cerebrospinal fluid (aCSF; comprised of NaCl, 1M KCl, MgCl2, 155 mM NaHCO3, 1M glucose, 108 mM CaCl2, penicillin-streptomycin and dH2O) was used for the electrolyte in all in vitro measurements. The cathodal CSC, determined using cyclic voltammetry with sweep rate ν was calculated by (4) where GSA is the geometric surface contact area between the electrodes and aCSF, V1 and V₂ are the water reduction and oxidation potential, respectively, and I(V) is the current density. The CIC was determined by applying biphasic, charge-balanced, cathodal-first alternative current pulses across the working and counter electrode while measuring the corresponding voltage across the working and reference electrode. A very large resistor was connected from working to counter electrode to dissipate residual charge to ensure the charge is well balanced. To ensure zero bias potential, current magnitude was modified on the order of 100 nA on the anodal phase (ian.). The cathodal (Emc) and anodal potential excursion (Ema) were recorded by taking the voltage 10 μs after the end of cathodal and anodal current pulse, respectively. CIC was calculated from the cathodal (Qcat.) charge delivered per unit area at which either Emc or Ema reaches the water reduction or oxidation potential, respectively as the current magnitude increases: (5) where icat. is the current pulse phase at which the working electrode polarization potential reaches the reduction potential. To obtain the impedance spectra and Nyquist diagrams, a 0.5 V sine wave with frequency ranging from 10⁵ to 10⁰ Hz was applied with 12 measurements per decade. The impedance spectra and Nyquist plots were fitted using MATLAB. Curve fitting of the equivalent Randles circuit The impedance data were fitted using equivalent Randles circuit presented in Figure 13. All data were separated into Bode and Nyquist results, which included magnitude and frequency, and real and imaginary values respectively. The Randles circuit models the curves, and finding the parameters was the purpose for this curve fitting as shown below in Eq.6 which represents the equivalent impedance ^_^^. Formula for the Warburg impedance ^_^ is shown in Eq.7. _{Z R ct} ^ _{Z W} (6) e_{q ^ R s ^ ^ ^} By substituting Eq.7 into Eq.6, the equivalent impedance was broken down into real and imaginary components. This is shown below in Eq.8 and Eq.9 respectively. This was used to curve fit using a combination of Bode and Nyquist. The plots were re- organized into frequency vs real ^_^, and frequency vs imaginary ^_^ as two different curves. ^ + ^_{^ ^^} The MATLAB, iterated 20 times (Figure 14). This function was bases off a preliminary guess of the coefficients, and then optimized. Therefore, it was important that the preliminary guess was within magnitude of the real component (ie.1A or 1mA). This was then related for the imaginary component, using the result of the real component fitting previous as initial guesses. In order to better mesh between the two, the result of the imaginary fitting was used to fit into the real component curve and imaginary component again one iteration each before obtaining the final result. Along with every iteration above, the coefficient of determination was calculated to monitor the fitting. Thickness characterization Optical profiler was used to measure the thickness of Mo during degradation.50 nm Mo patterns were sputtered on glass slides and were soaked in DI water or aCSF solutions at 37^oC and were dried under nitrogen prior to thickness measurement at each time point. Conductivity characterization Conductivity of PEDOT:PSS conductive polymer coatings and sputtered Mo during degradation were measured using 4-probe with Keithley 2400 source meter. Mo concentration measurement Inductively coupled plasma atomic emission spectroscopy (ICP-OES, Thermo Scientific) was used to measure the concentration in ppm of Mo contents in solutions. A wavelength of 202.03 was used and a standard curve was generated at 1, 3, 5 and 10 ppm for calibration. PLGA modulus measurement 3-point bending test was performed on 5.08-mm long, 6-mm wide PLGA films using thermomechanical analyzer (TMA Q400) at 37 ^oC. A ramp force of 0.002 N/min. was applied, and flexural stress and strain was recorded for modulus calculation as shown in Figure 21. Electrode implantation surgery All animal work was approved by the University of Toronto Animal Care Committee in accordance with institutional guidelines (protocol no.20011279). Male C57/BL6 mice at 6 weeks of age were implanted with either conductive polymer-Mo- based biodegradable electrodes or gold electrodes for biocompatibility testing and electrical stimulation. Surgeries were performed on anaesthetized mice induced with 5% and maintained at 1.5~2.5% isoflurane (Fresenius Kabi) via inhalation. An incision on the shave head was made to expose the skull, and subsequently disinfecting it with 70% ethanol, povidone-iodine, and ethanol. Eye lubricant (Dura tears) was applied. Mice were then secured on a stereotactic apparatus and placed on a 37°C thermal pad for thermal regulation were injected with 1.5ml of Ringer’s lactate solution intravenously and 0.2mg/kg of meloxicam subcutaneously. A 0.018’’ diameter drill bit (07289 #77, Kyocera Group) attached to a dental drill (P/N 8177, David Kopf Instruments) was utilized for all holes drilled for electrode insertions. For biocompatibility experiments, one hole was drilled at 0.8 mm rostral, -2.0 mm lateral relative to bregma. For stimulation experiments, two holes were drilled using the same set-up at locations 0.8 mm rostral, -0.7 mm and -2.7 mm lateral relative to bregma. Reverse action forceps were then used to lower electrodes into the cortex, and the connectors were secured with Insta-cure+cyanoacrylate glue (BSI-106, 14.2 g, Bob smith industries, CA, USA) before skin was sutured. Mice were then monitored until fully awake in a clean cage under a heat lamp. Meloxicam (2.0 mg/kg) was administered subcutaneously daily for the first 3 days post-surgery. Neurosphere Assay Adult male C57BL/6 mice of 6-8 weeks of age (Charles River) were anesthetized with isoflurane (Fresenius Kabi) before cervical dislocation. The SVZ was dissected, and cells were dissociated and plated in the neurosphere assay ^11,44,45. In brief, tissue was submerged in enzyme mix containing hyaluronidase (0.83 mg/mL, Millipore- Sigma), trypsin (1.33 mg/mL, Millipore-Sigma), and kynurenic acid (0.13 mg/mL, Millipore-Sigma) for dissociation at 37°C for 25 minutes. It was then centrifuged at 1500 revolutions per minute (RPM) for 5 minutes before resuspending cells in trypsin inhibitor (0.67 mg/mL, Worthington Biochemical Corporation). Following trituration, cells were then centrifuged for 5 minutes at 1500 RPM. Following resuspension in supplemented serum-free media (S.SFM; comprised of 10X Dulbecco’s modified Eagle’s medium/F12, 30 % glucose, 7.5 % NaHCO3, 1 M Hepes buffer, l-glutamine, hormone mix, penicillin and streptomycin, epidermal growth factor (20 ng/mL; Millipore-Sigma), basic fibroblast growth factor (20 ng/mL; Millipore-Sigma), and heparin (2 μg/mL, Millipore Sigma)), cells were triturated and centrifuged at 1500 RPM for 3 minutes. Cells were then resuspended in 1mL of S.SFM before determining the number of viable cells with a hemocytometer and plated at clonal density (10 cells/µl) ⁴³. The numbers of neurospheres with diameters >80 µm were counted after 7 days in vitro. Immunohistochemistry I_{n vivo biocompatibility studies were performed to assess the neuronal response} and neuroinflammatory response elicited by Au and biodegradable electrodes. Mice were transcardially perfused using a microprocessor pump drive (Masterflex model 7518-10) with a flow rate of 15 ml/min with 1xPBS and 4% paraformaldehyde (PFA) in 1xPBS at either 4 or 8 weeks after electrode implantation. Electrodes were carefully removed prior to brains being extracted. Brains were cryoprotected in 4% PFA overnight then transferred to 30% sucrose in 1xPBS for 48 hours prior to sectioning (HM535 NX, ThermoFisher Scientific). Coronal brain sections of 20 μm in thickness were collected on Superfrost Plus glass slides (Fisher Scientific) and were subsequently stored at - 20°C until processing. Slides with an implantation site were selected for immunohistochemistry staining, following protocols adapted from previous studies ^46–48. Sections were rehydrated with 1xPBS for 5 minutes then permeabilized with 0.3% Triton-X 100 for 20 minutes. Blocking solution containing 10% normal goat serum (NGS) in 0.3% Triton-X and 1% bovine serum albumin (BSA) was placed on the slides for 1 hour at room temperature. Sections were then incubated in primary antibodies with chicken monoclonal purified IgG glial fibrillary acidic protein (GFAP; 1:500; Abcam, AB4674), and rabbit monoclonal purified IgG ionized calcium binding adaptor molecule 1 (Iba1; 1:400; Wako, LEH3120), or rabbit polyclonal anti-NeuN for neurons (1:200, Sigma-Aldrich, ABN78) at 4 °C overnight. Prior to adding secondary antibodies, 1x PBS in Tween was used to wash the slides 3 times for 5 minutes each. Slides were incubated with goat anti-chicken IgG 568 (1:400, Invitrogen A11041), goat anti-rabbit IgG 488 (1:400, Invitrogen A11001), or goat anti-rabbit IgG 647 (1:400, A21245), DAPI (1:5000, Vector, H-1, 200) for 1 hour before washing with 1xPBS for 3 times for 5 minutes each. A glass coverslip was mounted with DAKO mounting medium to cover the sections. Imaging of Iba1+, GFAP+, NeuN+ cells was performed with a fluorescent inverted microscope (Axio Observer D1, Zeiss, NY, USA).40,000 μm² areas at 1 mm deep in the parenchyma and 25 μm away from the electrode track were selected for cell counting to avoid the high background autofluorescence surrounding the implant site. All counts were performed by a blinded observer. Electrical stimulation For in vivo stimulation, MoPH3 biodegradable electrodes were implanted at 0.8 mm rostral, -0.7 mm and -2.7 mm lateral relative to bregma. Charge-balanced biphasic current-controlled stimulation was applied using the STG4002-1.6mA simulator. To account for residual charge in the brain tissue caused by imperfect charge balance of the stimulator, a large resistor (1 MOhm) was connected in parallel between the cathode and anode to dissipate the residual charge, while not influencing the current being delivered to the tissue. The anodal current magnitude is also adjusted with a minute current magnitude of less than 0.2 μA to maintain zero interpulse potential. Charge- balanced biphasic stimulation parameters applied had a cathodal amplitude of 200 μA with a pulse width of 500 μs, followed by anodal amplitude of 50 μA with 2000 μs then a silencing period (interpulse interval) of 1000 μs. Stimulations were performed for 1 hour for neurosphere assay experiments on day 0 (the time of implantation) or day 7 post- implantation. Animals were sacrificed via cervical dislocation for neurosphere assay or perfusion for immunohistochemistry. Modelling of dissolution rate of Mo The thickness of sputtered Mo (h0) is much smaller than the width and length of the electrode and interconnect pattern, and hence can be modelled using a simple 1-D reactive diffusion model with reactive diffusion equation: (10) where D is the and w (z, t) is the water concentration at time t and location z with z = 0 at the bottom of Mo. By assuming constant water concentration at z = h0 (w|z = h0 = w0) and zero water flux at z = 0 (∂w/∂t|z = 0 = 0), Equation 5 can be solved and the rate of dissolution can be estimated as where tc is the MMo and MH2O are the molar masses of Mo and water, respectively, ^ is the density of Mo and q = 1 representing one water molecule reacts with one Mo atom in the dissolution reaction. (7) Table 2: Parameters used for calculating the dissolution rate of Mo outlined in “Curve fitting of the equivalent Randles circuit”. _{Parameter Symbol Value} Reaction constant of Mo k 1.21×10^-5 s^-1 Diffusivity of water D 2 ×10^-16 cm² s^-1 Molar mass of water MH2O 18 g mol^-1 Molar mass of molybdenum MMo 96 g mol^-1 Water concentration of Mo at t w0 1 g cm^-3 = 0 D_{ensity of molybdenum ^ 10.22 g cm-3} By Initial thickness of Mo h0 300 ×10^-4 cm substituting the parameters in Table 2, the dissolution rate in pure water is determined: 1 ^{^} ^ _{= ^^^} ^^ ^ _{× 2 × 10} ^ ^ 1 ^^^{^} _{× 96} ^ ^^^ _{1.21 × 10^^} 1 ^ _{× (300 × 10} ^{^^}^^)^{^} ^ _{= 2.5671 ×} ^ _{= 22.18 ^^/^^^} Statistical analysis Statistical analyses were carried out using Origin and GraphPad Prism 9 (GraphPad Software). Cellular analyses and, neurosphere analysis were analyzed with two-way ANOVA. For all analyses, p < 0.05 was considered as statistically significant. Results and Discussion Design and fabrication of biodegradable electrode To design a biodegradable probe, the probe substrate, metal interconnect, insulation and stimulation electrodes were selected based on the degradation rate and biocompatibility of materials. The degradation rate of the material reflects the time that the material will stay in the body and the biocompatibility is critical to ensure that the end-products will not have detrimental effects on cell survival or exacerbate neuroinflammation. The aim was to provide neurostimulation to activate NPCs for 1 week based on previous studies showing that 7 days of NPC activation is sufficient to promote neural repair following injury ^23–27. Poly(lactic-co-glycolic) acid (PLGA), approved by the U.S. Food and Drug Administration (FDA) and Health Canada, was chosen for both the substrate and insulation layer due to its flexibility and well- established biocompatibility and biodegradability ^28,29. (Figure 1A-1E). The degradation time of PLGA can also be tuned based on its monomer ratios of glycolic and lactic acid²⁹ (Table 3, Figure 6). Table 3: Degradation time of poly(lactic-co-glycolic acid) (PLGA) in water. The ratio between lactide and glycolide monomers in PLGA affect the degradation time in water²⁹. Polymer Degradation Time (weeks) Poly(D,L-lactide- 5-6 co-glycolide) 85/15 Poly(D,L-lactide- 4-5 co-glycolide) 75/25 Poly(D,L-lactide- 1-2 co-glycolide) 50/50 Molybdenum (Mo) was chosen for the interconnects due to its relatively slow dissolution rate of 20 nanometers (nm) per day compared with other bioresorbable metals^21,22 (Figure 1A-1B, Table 4). The conductive polymer poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) was used as a coating on Mo for the stimulation site of the electrode to improve the charge injection capacities and reduce interfacial impedance of the electrodes during stimulation (Figure 1A-1C)³⁰. While PEDOT:PSS cannot be completely resorbed, it is biocompatible with partial degradation properties such that it can disintegrate from the biodegradable matrix when present in small amounts¹⁹. Table 4: Dissolution rates and mechanisms of bioresorbable metals^21,22. Bioresorbable Dissolution rate Mechanism Metals in PBS (nm/day) Mo 20 (PBS) 2Mo + 2H2O + 3O2 → 2H2MoO4 W 150 (PBS) 2W + 2H2O + 3O2 → 2H2WO4 Mg 1700 (DI) Mg + 2H2O → Mg(OH)2 + H2 Zn 3500 (PBS) Zn + 2H2O ^ Zn(OH)2 + H2 To prevent damage of brain tissue, electrode geometry and contact area within the brain tissue was carefully designed. As described, the maximum allowable charge injection per phase is 4 nanocoulombs (nC) for microelectrodes smaller than 10,000 µm² and 30 μC/cm² for macroelectrodes greater than 100,000 µm²³¹. It has previously been shown that the optimized stimulation paradigm for NPC activation⁶ (Figure 1F-1G) comprises a cathodal charge injection of 0.1 μC/phase (-200 μA magnitude and 500 μs pulse width) (Figure 1F), hence a macroelectrode iwas designed with the stimulating site of 1000 µm by 350 µm to ensure sufficient contact area with neural tissue for safe charge delivery (Figure 1A-1C). Fabrication of the biodegradable neural probes was not feasible using standard photolithography protocols due to the heat sensitivity of PLGA substrates³² therefore a laser-based approach was developed to pattern the interconnects and stimulating electrodes (Figures 7-9A, 9B). The electrode implant design consists of two parallel microfabricated neural probes which are inserted into the cerebral cortex of the brain. (Figure 1A-1C, 1F). Each probe is 2 mm long, 350 µm wide and 50 µm thick and comprised of a 300 nm-thick, 150 μm-wide Mo interconnect on the surface of the PLGA probe (Figure 1A). The 1000 μm Mo interconnect located closer to the skull is insulated with a 25 μm thick PLGA film, while the bottom 1000 μm of the probe is the stimulating site coated with PEDOT:PSS (Figure 1B-1C). The two neural probes were mounted on a 3-D printed connector adhered to the surface of skull and serves as an interface between the tissue and stimulator (Figure 1F, Figure 10). Electrochemical properties and stimulation capacity of molybdenum Mo demonstrated its suitability to be used as a stimulating electrode for delivering current-controlled BPMP stimulation based on its superior electrochemical properties including the cathodally skewed electrochemical window with high charge injection capacity (CIC).The electrochemical window of bare sputtered Mo (Figure 2A) was determined to range from -1.0 V (reduction potential) to 0.2 V (oxidation potential) (Figure 2B), as compared to -0.6 to 0.6 V for conventional electrodes such as Pt and Au³³. Cyclic voltammetry (CV) measurements based on this electrochemical window reveals a cathodal charge storage capacity (CSC) of 2.50 ± 0.06 mC/cm² for Mo. These values are comparable to conventional platinum (Pt) wires and sputter gold (Au) electrodes (Figure 2B) however, Mo is less likely to cause undesired hydrogen gas formation during charge injection compared with Pt and Au at cathodal potentials of less than -0.6V. As a result, the negatively skewed electrochemical window of Mo (Figure 2B) is beneficial for biphasic charge-balanced current controlled brain stimulation where the desired stimulation waveform is more cathodally favored. While CSC represents the amount of charge that can be stored in the double layer during current injection, CIC directly reports the maximum amount of charge that can be injected without damaging the electrodes. By determining the maximum potential excursions from voltage transient curves measured in artificial cerebrospinal fluid (aCSF) (Figure 2C, Figure 11), the cathodal CIC is calculated to be 0.22 mC/cm² with a current injection limit (CIL) of 1.35 mA which is much higher than Pt wires (CIC = 0.013 mC/cm²) and sputtered Au (CIC = 0.036 mC/cm²) with similar areas. The impedance spectra demonstrate a slightly lower charge transfer resistance at low frequencies for Mo compared with Pt and Au (Figure 12) which is likely attributed to the formation of oxides on the surface of Mo during anodal charge injection thereby facilitating charge storage and transfer. At higher frequencies from 10³ to 10⁵ Hz, the impedance of Mo is slightly higher than that of Pt and Au, but the slope and phase angles are all close to 0, indicating Mo, Pt and Au have similar ionic resistances in aCSF. Taken together, these findings reveal that Mo can deliver the desired cathodal current of 200 µA to the brain tissue with a lower polarization potential compared to Au (Figure 2D) and Pt, providing a safer alternative for brain stimulation. Conductive polymer coating on molybdenum The conductive polymer PEDOT:PSS is a promising material for use in neural interfaces, attributed to its good biocompatibility, high CSC and the capacity to spin coat multiple ultra-thin nanometer-thick layers of the material ^30,34. Due to PEDOT:PSS’s capacity to absorb water and uptake ions, the entire volume of the material participates in the charge transfer process during stimulation ³⁴. Hence, Mo was combined with PEDOT:PSS to further enhance the stimulation capacity enabling a higher CIC and lower interfacial impedance. The effects of PEDOT:PSS coating thickness on the electrochemical and the degradation properties were determined with the goal of delivering the desired current with minimal probe thickness. Cyclic voltammetry curves at 200 mV/s (Figure 2E) and 20 mV/s (Figure 13) show that the cathodal CSC increases with more layers of PEDOT:PSS coating. The CSC of MoPH1, MoPH3 and MoPH6 (1, 3 and 6 layers of coating, Figure 2F) is 4.74 ± 0.76, 7.15 ± 1.02, and 10.04 ± 0.82 mC/cm², respectively. The impedance spectra also demonstrate lower interfacial impedance at low frequencies with thicker PEDOT:PSS coatings (Figure 2G-2H). Near-zero phase angle at high frequencies indicates resistance-dominated impedance. Based on equivalent Randles circuit analysis (section “Curve fitting of the equivalent Randles circuit”, Figure 14), it can be concluded that PEDOT:PSS coating reduces the solution resistance, charge transfer resistance and Warburg diffusion-based impedance, while increasing the double layer capacitance (Figure 15A-15E). With increasing layers of PEDOT coating, the solution resistance did not show significant differences which could be explained by a similar area of contact with the electrolyte. For the other components in the Randles circuit, double layer capacitance, charge transfer resistance and Warburg impedance, 3 and 6 layers of PEDOT:PSS coating show increased capacitance and reduced charge transfer and diffusion-based impedance compared with 1 layer of coating. This is attributed to the fact that with thicker PEDOT:PSS coating, the bulk conductive polymer allows volumetric electron-ion interaction site during charge injection, thereby enhancing charge carrier properties and facilitating charge diffusion and storage. CIC measured through voltage transient curves (Figure 2I) reveal that with more layers of PEDOT:PSS coating, the potential excursions reach the water reduction (-1.0V) and oxidation potential (0.2V) at a slower rate (Figure 16A-16E, Table 5), indicating that more charges can be injected with a thicker coating without irreversible Faradic reactions. As shown in Figure 2J, the PEDOT:PSS-coated Mo electrodes exhibit a high cathodal CIC of 0.22, 0.33, and 0.75 mC cm² with 1, 3 and 6 layers of coating respectively. Together, these results indicate that with 1 layer of PEDOT:PSS coating, Mo electrodes have a 10-fold higher CIC compared to conventional Pt and Au electrodes for brain stimulation^30,33,34. _{Table 5: Calculation of current injection limit (CIL) and charge injection capacity (CIC) of} PEDOT-coated Mo electrodes from Figure 16e. Curve fit (y = mx + b) R² Current Area Charge injection limit (cm²) injection (µA) capacity (µC cm^-2) Mo Anodal V = 7.74854E-4 i + _{0.936 83.56 0.0018 23.21} 0.00576 C_{athodal V = -0.00129 i – 0.03499 0.984 748.07 0.0018 207.80} MoPH1 Anodal V = 4.28255E-4 i + _{0.986 463.21 0.0036 64.33} 0.00163 C_{athodal V = -5.45524E-4 i – 0.992 1559.29 0.0036 216.57} 0.14937 MoPH3 Anodal V = 2.84942E-4 i + _{0.966 679.68 0.0036 94.40} 0.00633 C_{athodal V = -3.78827E-4 i – 0.977 2376.84 0.0036 330.12} 0.09959 MoPH6 Anodal V = 1.27398E-4 i – _{0.972 1579.70 0.0036 219.40} 0.00125 C_{athodal V = -1.85393E-4 i + 0.956 5427.28 0.0036 753.79} 0.00618 Next, the potential difference due to double layer capacitance and charge transfer resistance was assessed at the electrode-aCSF double layer given a stimulation waveform with -200 μA cathodal pulse (Figure 2D). While replacing Au with Mo electrodes, the polarization potential can be reduced from 1.74 to 1.24 V. Mo electrodes with 3 layers of conductive polymer PEDOT:PSS coating can further reduce this potential more than 27 times down to only 0.06 V, an indication of capacitive charging with minimal Faradic charge injections to ensure stimulation safety without causing electrode or tissue damage. Electrochemical and degradation properties of conductive polymer-coated molybdenum in physiological conditions The degradation properties of the electrodes in physiological conditions were assessed to determine whether the required stimulation parameters could be delivered during resorption. The rate of Mo dissolution was first determined by measuring the probe thickness using an optical profiler (Figure 17). In deionized (DI) water, the dissolution rate was found to be 21.1 nm/day, which closely matched our calculated rate of 22.08 nm/day based on a kinetics model as outlined in section Modelling of dissolution rate of Mo^21,22,35. Notably, and as predicted, the dissolution rate was higher in aCSF at 31.0 nm/day due to the presence of ions in aCSF such as Cl- that has been shown to accelerate the dissolution reaction ^21,36. To accommodate the higher rate of degradation, 300 nm of Mo was sputtered on to PLGA substrate with the goal of maintaining the stimulation capacity of the electrode for the desired period of 7 days. Mo electrodes were analysed with inductively coupled plasma atomic emission spectroscopy (ICP-OES) and Mo was detected after 1, 2 and 3 weeks of soaking in both DI water and aCSF. (Figure 3A, Figure 17, Table 6). ICP-OES analysis over 3 weeks also demonstrates higher dissolution rate of Mo in aCSF than DI water as seen in higher concentration of Mo observed in week 1 and 2 of incubation in physiological condition. Table 6: Mo concentration before and after 2-week incubation of sputtered Mo in DI water and aCSF, assessed using inductively coupled plasma atomic emission spectroscopy (ICP-OES) at wavelength of 202.03. S_{olution Mo concentration} (ppm) DI water – Day 0 0.016 DI water – Day 14 46.282 DI water – Day 14, diluted 1:20 in Milli-Q 2.049 Water aCSF – Day 0 0.072 aCSF – Day 14, diluted 1:20 in Milli-Q 2.653 water The thickness of the PEDOT conductive polymer coating was then optimized to achieve the desired stimulation capacity over a 7-day period to deliver the optimized cathodal stimulation current of 200 µA for NPC activation. Since conductive polymers cannot be fully resorbed by body fluids, the goal is to use minimum conductive polymer coating possible while maintaining stimulation properties. As the conductive polymer network is being fully exposed to aqueous conditions, water molecules weaken the hydrogen bonding between PEDOT oligomers and hygroscopic polystyrene sulfonate (PSS) chains, resulting in partial dissolution of PSS chains followed by chain reorganization which contributes to the loss of electrochemical properties required for stimulation ^37,38 (Figure 17C). Dissolution of PSS chains is evidenced by the increase in sulfur content in aCSF from the first to third week of electrode soaking soaking (Figure 3A). In addition to PSS dissolution, the uptake of water modules was also reported to expand the distance between PEDOT oligomers and reducing the electrical conductivity³⁹. As shown in Figure 3B, by 7 days, the change in the CSC was reduced by half irrespective of the numbers of PEDOT coatings (1, 3, or 6) while the CIC reduction occurred at different rates (Figure 3C-3D, Table 7). Table 7: Calculation of current injection limit (CIL) and charge injection capacity (CIC) of PEDOT-coated Mo electrodes after 1 week in aCSF from Figure 4. Curve fit (y = mx R² Current Area Charge + b) injection (cm²) injection limit (µA) capacity (µC cm^-2) M_{oPH1 Anodal V = 0.1024 i + 0.950 1.94 0.0036 0.27} 8.5591E-4 C_{athodal V = -1.11315 i + 0.972 0.9 0.0036 0.13} 0.00113 M_{oPH3 Anodal V = 0.00228 i + 0.949 87.33 0.0036 12.13} 8.78245E-4 C_{athodal V = -0.00243 i + 0.977 411.97 0.0036 57.22} 0.00108 _{MoPH6 Anodal V = 0.00102 i + 0.994 195.15 0.0036 27.10} 9.45339E-4 C_{athodal V = -0.00132 i + 0.993 758.5 0.0036 105.35} 0.00122 With 1 layer of coating, the current injection limit (CIL) dropped from 748 µA at day 0 to approaching zero after 1 week of incubation in aCSF. However, less of a CIL reduction was observed with 3 and 6 layers of coating the CIL maintained 411 and 759 µA, respectively (Figure D, well above the maximum allowable current magnitude of - 210 μA without inducing tissue damage (30 μC/cm² with contact area of 0.0035 cm²). The impedance of Mo electrodes was then characterized with 3 layers of conductive polymer coating at day 7 (Figure 3E). An increase in the electrode resistance was observed as indicated by the rise in the high frequency impedance and the large instantaneous voltage drop after the end of cathodal pulse (Figure 3F, Figure 18), however, the increase in cathodal potential excursion (Emc) remained within the water reduction potential of -1.0 V with a 200 µA magnitude current pulse, indicating stimulation with MoPH3 electrodes with optimized parameters would not lead to gas or chemical byproducts formation during the first week of degradation. The stable low frequency impedance also indicates there are minimal impedance changes of both charge transfer resistance and Warburg impedance at the electrode- tissue interface. Further, the close-to-zero phase angle at above 100 Hz demonstrates resistance-dominated impedance 7 days after exposure to physiological conditions (Figure 3E). With the expectation that the degradation rate in vivo in the brain will be slower than aCSF in a solid-phase hydrogel-like environment, we examined the degradation performance of the electrodes in 0.5% aCSF agarose gel. As predicted, the rate of degradation was about two-thirds of that observed in aCSF fluid conditions (Figure 30). Hence, with the goal of delivering the optimized stimulation current of -200 µA with minimal conductive polymer residues, Mo electrodes with 3 layers of PEDOT:PSS coating (MoPH3) proved sufficient for delivering stimulation within the safety threshold for one week and thus was employed for subsequent experiments. Physical degradation of MoPH3 was further characterized in physiological conditions (37°C in aCSF) through optical and microscopic images (Figure 3G-3J). The degradation of the fabricated electrode started with the dissolution of Mo in the first week (Figure 3H). After maintaining the electrochemical properties required for stimulation during the first week, Mo undergo near-complete dissolution at the end of second week. (Figure 3I) With the dissolution of Mo, water uptake of the conductive polymer network and the initiation of PLGA substrate degradation, the PEDOT:PSS coating began to separate from the hydrophobic PLGA surface by the fourth week (Figure 3J). The detachment of the conductive polymer from the substrate leads to chain reorganization and subsequent disintegration which, together with bulk dissolution of PLGA polymer, results in degradation of the entire device (implantable probe portion) within a few months (Figure 3K-3L), as evident by the mass loss of the device and stabilization of pH in the surrounding environment in an accelerated degradation test at 90 °C in aCSF (Figure 19). In vivo biocompatibility study on MoPH3 electrodes The biocompatibility of the MoPH3 biodegradable electrodes was assessed in vivo by comparing outcomes with the FDA-approved Au electrodes with the same dimensions. Adult mice were implanted with 2 mm long MoPH3 probes or Au probes into the motor cortex. At 4- and 8-weeks post-implantation, in the absence of stimulation, brains were harvested and immunostained for the presence of microglia (Iba1+), reactive astrocytes (GFAP+), neurons (NeuN+) and cell nuclei (DAPI+) in the ipsilateral (implanted) and contralateral (control) cortices (Figure 4A-4B, Figure 20). The numbers of Iba1+/DAPI+, GFAP+/DAPI+ and NeuN+/DAPI+ cells were assessed. As predicted, irrespective of the electrode material, the numbers of neuroinflammatory cells (microglia and reactive astrocytes) were increased in the ipsilateral hemisphere compared to the contralateral hemisphere at both 4 and 8 weeks post-implantation (Figure 4C-4D). No significant difference in neuroinflammatory marker expression was observed between MoPH3 and Au implants at 4 weeks and 8 weeks post implantation (Figure 4C-4D) in the ipsilateral cortices. Further, while the numbers of NeuN+ cells were reduced in the ipsilateral cortex around the implantation site compared to the contralateral hemisphere in both the Au and MoPH3 implanted brains, the numbers of NeuN+ cells were not significantly different between Au and MoPH3 implanted cortices after 4 weeks (p = 0.23). Notably, significantly more surviving neurons were observed at 8 weeks post-implant in the MoPH3 implanted brains (p = 0.0280) (Figure 4E). These results suggest that MoPH3 biodegradable electrodes are biocompatible and do not lead to an enhanced neuroinflammatory response compared to conventional Au electrodes. Similar effects were reported with probes made with SU-8 and silicon where the extent of glial activation is comparable with conventional high modulus electrodes such as silicon, yet greater neuronal survival is seen in close proximity to the implant site ^40–42. This suggests that glial activation may depend more on the mechanical trauma rather than the material properties and higher neuronal survival near the implant site may be attributed to MoPH3 electrode biodegradability and/or better mechanical coupling with the soft brain tissue, as the flexural modulus of PLGA was measured to be 3.261 ± 1.066 GPa (Figure 21), which is >1 order of magnitude lower than Au electrodes (79 GPa). The promising biocompatibility of the MoPH3 electrodes, along with its stimulating capacity, makes it an appealing candidate to be used as a temporary neurostimulation tool for NPC activation. Neurostimulation with MoPH3 electrodes activates neural precursor cells in vivo The ability of MoPH3 electrodes to activate NPCs was then evaluated. One hour of charge-balanced current-controlled stimulation was delivered to the adult mouse brain on the day of electrode implantation (day 0) (Figure 5A-5B, Figure 21A-2B) with an optimized stimulation waveform of 200 µA cathodal current ⁶ (Figure 1H). Based on COMSOL Multiphysics software simulation, this results in a 63.85 V/m electric field being delivered to NPCs in the dorsolateral corner of the SVZ (Figure 5C, Table 8). To demonstrate consistency of electric field delivery, stimulation was applied for 3 hours and a stable voltage waveform at the beginning and the end of the stimulation period was recorded. (Figure 5D). The marginal voltage increase (less than 0.2V) observed was likely attributed to possible oxidation of molybdenum during anodal current delivery. Table 8: Electrical conductivity and dielectric properties used in COMSOL modelling^52,53 Compartment Electrical conductivity [S/m] C_{SF (SVZ) 1.8} Matter) Corpus 0.18 Callosum (White Matter) Striatum (Grey 0.3 Matter) To evaluate NPC activation, mice were sacrificed and the SVZ was microdissected and cells were plated in the NPC colony forming “neurosphere” assay. A single neurosphere is clonally derived from a neural stem cell hence the numbers of neurospheres reflects the size of the NPC pool⁴³. The electrical stimulation resulted in a 3.08 ± 1.28 -fold increase in the number of neurospheres isolated from the stimulated ipsilateral hemisphere (Figure 5E, Figures 22B, 22C) compared to the contralateral (unstimulated) hemisphere and sham controls (implanted and unstimulated, ipsilateral hemisphere). This expansion in the size of the NPC pool is similar to previous studies using Pt electrodes and the same stimulation parameters⁶. Investigation was conducted to determine if electrical stimulation with the MoPH3 electrode could activate NPCs during degradation in the first week following implantation. The same stimulation parameters were delivered at day 7 post- implantation and the neurosphere assay was performed following stimulation. As shown in Figure 5F, stimulation at 7 days post-implantation resulted in a significant expansion of the size of the NPC pool revealing that the degradation of the electrode during the first week of implantation did not affect its ability to deliver electrical stimulation and elicit a biological response. These findings reveal the efficacy of MoPH3 biodegradable electrodes for NPC activation. Development of biodegradable stimulation electrodes that incorporate microfluidic material delivery system According to an embodiment, the stimulating electrode may have a microfluidic _{delivery platform incorporated in the electrode to enable the intracranial delivery of} drugs virus, protein, small molecules, active ingredient or biologically active substance to the brain. The microfluidic channels may be designed with biodegradable substrate materials that are integrated with the stimulation electrodes of the present invention. The study revealed that the bulk degradation properties of PLGA material does not affect the microchannel geometry and injection properties. Device fabrication The fabrication of microfluidic delivery system is divided into two parts, the micropatterned channel (Figure 23), and the attachment of injection interface (Figure 24). The microchannel is fabricated by thermal bonding of two soft lithographed PLGA films to form an enclosed channel for fluid passage. As shown in Figure 23, the bottom flat layer and the top microchannel layers were prepared by hot embossing and solvent casting using flat and patterned PDMS slabs respectively, which were cured on top of pre-etched silicon wafers. To create the channel with an inlet, the flat PLGA film was then hole punched with 2 mm-diameter biopsy punch before aligning and thermally bonding with the patterned film above its glass transition temperature of 50 °C for intimate contact between the layers through polymer-chain interdiffusion [26]. The thermally bonded device was then aligned, and laser cut into probe shapes with optimized cutting speed and power such that the probe diameter can be reduced to ~150 μm. To attach the injection connector/interface, the substrate of probe (10) was laminated with adhesive layers of thin-polyimide (PI) films followed by bonding a PDMS block (108). As illustrated in Figure 24, the surface of PDMS block (108) with a hole (108a) punched through it and the non-adhesive side of a polyimide film were first activated with a hydroxyl group. To surface treat the polyimide film, its surface was functionalized with silicon using 2 v/v% 3-mercaptopropyltrimethoxysilane (MPTMS). Hydrolysis near the alkoxy terminal of MPTMS forms a silicon-oxygen network while nucleophilic reaction occurs between the substrate and the mercapto group of MPTMS. In this way, plasma activation of both the PDMS and PI films gives hydroxyl groups to allow permanent, instantaneous room temperature bonding [27, 28]. The adhesive side of the PI film was then used to laminate the device (10) and 28G PTFE (Teflon) tubing (109) was inserted to the PDMS block (108) to connect the fluid reservoir (105) with the injection probe (20b). Injection tests A series of injection tests were performed using the devise. The first injection test was performed in 0.6% agarose phantom as a surrogate for brain tissue. Coomassie blue dye with a low molecular weight of 825.97 g/mol was used to monitor the injection in real-time. Dye-loaded DI water was injected 2 mm deep into the gel using a stereotactic system. The probe was inserted 3 mm deep into the gel and after 30 seconds the probe was retracted to 2 mm and 5 μl of dye was injected via a syringe pump at a rate of 1 μl/min for 5 min [29]. The probe was then retracted slowly, over a period of 2 minutes. As shown in Figure 25, no leakage or backtracking can be observed, demonstrating the reliability of injection device using the proposed fabrication protocol. After injection, Figure 25D shows the dye diffusion in the agarose gel after 1 hour. We next performed in vivo injections with the device. Fluorescent red Dextran with a molecular weight of 10 kDa, which is on the same order of magnitude as BDNF (27 kDa), was used. The dye-loaded DI water was injected into cortical mouse brain tissue above the SVZ (AP+0.6; L+2.0, relative to Bregma) with similar injection procedure as described (Figure 26A). The concentration of the dye is 0.33 μg/μL, similar to the concentration of BDNF injection reported previously [18]. Mice were sacrificed at 1 hour after the injection, brains were removed and snap frozen using liquid nitrogen chilled isopentane (Figure 26B). The frozen brain was cryosectioned at 50 μm thickness at -16C and sections were placed on slides for imaging. Figure 27 demonstrates successful dye injection in the cortex. The diffusion was analyzed by examining the intensity of TxRed along each axis. It was found that the dye diffused evenly in the cortex along the anterior-posterior and medial-lateral axis for distances of 1.5 mm and 1 mm, respectively. After the reliability of the newly designed biodegradable injection probe in vitro and in vivo has been demonstrated, BDNF protein (PeproTech #450-02) was injected into the brain and its diffusion and cellular uptake was assessed using immunohistochemistry [30]. BDNF was prepared by reconstituting in aCSF to obtain a final concentration of 0.2 μg/μL.0.1% BSA was added to BDNF-aCSF solution to maintain bioactivity. Figure 26C shows the experimental groups: (1) Insertion of the probe but no injection (control), (2) aCSF injection into the left hemisphere (sham control) and (3) BDNF injection to the right hemisphere of the same brain. With similar injection procedure described above, the mice were perfused 1 hour after injection and the brain was extracted and fixed for visualization of BDNF. Tissue was cryosectioned at 20 μm thick sections and placed on slides. Sections containing SVZ and hippocampus were stained with primary antibody of BDNF rabbit monoclonal (1:250; abcam, ab213323) and secondary antibody of goat anti rabbit IgG Alexa Fluor 488 (1:400, abcam, ab150077). Endogenous BDNF is abundant in the hippocampus and Consistent with what is observed in Figure 28A showing strong expression of BDNF in the hippocampus. Exogenous BDNF injected near the SVZ was imaged as shown in Figure 28B. The cortex of the BDNF injected hemisphere demonstrated stronger BDNF expression than the contralateral aCSF injected hemisphere. Degradation properties microfluidic channel Channel and fluid flow properties were assessed during the degradation period of PLGA to examine the delivery parameters over time. To examine the effects of degradation on channel geometry and injection properties, channels were soaked in aCSF at 37 °C. Figure 29A shows the preliminary studies with SEM images of the channels over a period of 4 weeks. It can be seen that the film thickness and channel cross-sectional area does not change over time (Figure 29B), despite more pores being observed as soaking time increases. The surface of the film is smooth compared to the porous core (Figure 29C), suggesting that the fluid will not diffuse to the PLGA bulk as it is travelling through the microfluidic channel. This phenomenon can be explained by the auto-catalysis degradation mechanism of PLGA. As the material is submerged in aCSF, water absorption starts, and ester bond cleavage begins. The lactic and glycolic acid can diffuse away quickly from the surface, and they are trapped within the polymer, leading to catalyzed degradation in the polymer core. Given the channel geometry is stable over the degradation time, it is anticipated that fluid injection will be sustained over time. Figure 29D shows preliminary results of the channels, before and after 1 week of soaking in aCSF. No significant difference is seen when dye-load DI water is injected into aCSF solution at a rate of 10 μL/min. Variant of the connector device According to an embodiment, a variant of the stimulation electrodes and microfluidic channel device (or connector device) is provided (Figure 31A). The device (200) comprises two neural probes (210 and 211) assembled in a sandwiched assembly between two outer bodies (320, 330) and a central spacer (310). The first probe (210) has an electrode with an electrical conductive site and the second probe (211) has an electrode with an electrical conductive site and a microfluidic channel. Both probes (210 and 211) have a through-opening (210a, 211a). The spacer (310) has a first channel (311) extending from a side surface to an opposing side surface. Each probe (210 and 211) is secured to one of the opposite sides of the spacer (310) with the electrically conductive site of each probe facing each other and having its opening (210a, 210b) in alignment with the first channel (311). The first outer body (330) has second channel (331) extending from a side surface to an opposing side surface and is secured to first probe (210) with the second channel (331) being in alignment with the opening (210a) of the first probe (210). The second outer body (320) has solid surfaces with no opening or channel. The outer body (320) is secured to the second probe (211). As shown in Figure 31A, the openings (210a and 211a) are in alignment with the first channel (311) and second channel (331) providing fluid communication between the second channel (331) and the microfluidic channel of the second probe (211). Since the outer body (320) as a solid surface secured to the probe 210, the solid surface of the outer body (320) closes the first channel (311) As shown in Figure 31A, the inlet (331) has a diameter greater than the diameter of the opening (311). The body (320) also provides structural stability to the assembly/device (200). As best seen in Figure 31E, a short tubing segment (340) is received within the through- type inlet (331) of the of first outer body (330). The tubing (340) may be used for the facilitating the delivery of a drug, virus, protein, small molecules, active ingredient or biologically active substance. According to an embodiment, the central spacer (310) and the outer bodies (320 and 330) may have the shape of a block. The blocks may be made of acrylic. They may be laser-cut and dimensioned to about 7mm X about 5mm. The central block (310) may have a thickness of about 5 mm, and the two outer blocks (320 and 330) may have a thickness of about 1.5 mm. The opening of the outer block (330) may have a diameter of about 2 mm, and the opening (311) of the central block (310) may have a diameter of about 22 gauge. A sealant may be applied to the tubing (340) to prevent any leak between the outer wall of the tubing (340) and the wall of the inlet (331). According to another embodiment, the fabrication method of the stimulation electrodes and microfluidic channel device comprises the following steps: Electrode fabrication 1. PLGA pellets were dissolved in acetone, and spin coated on cleaned glass substrates. The coated glass substrate was baked at 60C or left at room temperature for solvent (acetone) evaporation 2. Attachment of laser-cut metal mask, followed by attachment of magnets to minimize gap between metal mask and PLGA substrate. 3. Deposition of 300 nm Mo, then remove magnets and metal mask. Microchannel fabrication 4. Surface activation of hydrophobic PLGA film, followed by spin coat PEDOT:PSS conductive polymer 5. Use photolithography and dry etching techniques to create 50-um wide, 50-um deep features on Si wafer. This wafer will be used as negative mold for PDMS cast. The mold was salinized using trichloro(1H,1H,2H,2H-perfluorooctyl)silane to increase the hydrophobicity of the surface. 6. Cast of PDMS (Sylgard 18410: 1 curing agent) on to mold, followed by peel off once cured. This PDMS will be used as positive mold for PLGA microchannel fabrication 7. Solvent cast (PLGA pellets in acetone) of PLGA on PDMS mold. PLGA film was peeled off after solvent evaporation. 8. A biopsy punch was used to create a hole aligning the “U” shaped electrode. Then, the electrode part (flat) and microchannel part (patterned) were bonded through thermal bonding, with the electrode aligning exactly with the microchannel. In some cases, the two PLGA films were bonded together by compression during which the there were still solvents left in PLGA that facilitates PLGA polymer chain interdiffusion. 9. Bonded device can be laser cut into probe shape with electrical connection (“U” shape part). Device assembly 10. Three acrylic blocks were used to form the device connectors. 11. Two electrode probes (1 with and 1 without microchannel) were first aligned to the two sides of a 2-mm thick block with a 22-gauge opening. 12. A 1.5-mm thick block with a 2-mm diameter opening was then attached for tube insertion. 13. Another 1.5-mm thick block was attached to the other side of the probe to provide stability. 14. A 2-mm diameter tubing was inserted in the 2-mm diameter opening. 15. A sealant was applied to the tubing for preventing any leakage or backflow. Variant of the connector device According to an embodiment, as shown in Figure 32A, Figure 32B and Figure 38, a connector device (100) for use with two flexible probes (10a ,10b) is provided. The first flexible probe (10a) has an electrode (20) and the second flexible probe (10b) has an electrode (20) and a microfluidic channel (106). Each electrode (20) has a side with an electrically conductive site (202). The connector device (100) comprises a first housing portion (130), a second housing portion (140), a first sheet (150a) of conductive material, a second sheet (150b) of conductive material, a spacer (170), two openings (180), each opening (180) is configured to receive a pin (160) which is extending through the respective opening (180). As shown in Figure 32B, the first housing portion (130) has an inner wall (131), which is configured to receive the first probe (10a) with the electrically conductive site (202) facing away from the inner wall (131). The second housing portion (140) has an inner wall (141) (best seen in Figure 35), which is configured to receive the second probe (10b) with the electrically conductive site (202) facing away from the inner wall (141). The first sheet (150a) of conductive material is configured to contact the electrical connection region (201) of the electrode (20) of the first probe (10a) and the second sheet (150b) of conductive material is configured to contact the electrical connection region (201) of the electrode (20) of the second probe (10b) (Figure 33). As shown in Figure 32A and Figure 32B, the openings (180) at the top surface (132) of the device (100) are defined when both halves (130,140) are joined together, each half (130,140) having two semi circular recesses (133) on its top surface (132). As shown in Figure 33, each pin (160) has an end (161) affixed to an end (151) of its corresponding sheet (150a,) of conductive material. As shown in Figure 32B and Figure 38, the first housing portion (130) and the second housing portion (140) are configured to securely mate with each other to form the outside of the connector device (100) defining a cavity enclosing each probe (10a, 10b), the corresponding sheet (150a, 150b) of conductive material and the spacer (170) such that each probe (10a, 10b) is resting against the corresponding inner wall (131, 141) and each sheet (150a, 150b) of conductive material is positioned on opposite sides of the spacer (170), whereby, in the assembled state, each probe (10a, 10b) is pressed against its respective sheet (150a, 150b) by contact pressure applied by the housing portions (130, 140) and the spacer (170) conductively connecting the electrode (20) of each probe (10a, 10b) to the respective pin (160a, 160b). The second housing portion (140) has an opening (143) on its side surface (144). When the device (100) is assembled, the opening (143) is in fluid communication with the microfluidic channel (106) of the second flexible probe (10b) (not shown). As shown in Figure 35, the bottom surface (149) of the second housing portion (140) has two openings (190). Each opening (190) is configured to receive a stimulation site (206) of the respective probe (10a, 10b). The stimulation sites (206) extend through the openings (190) and project outwardly from the second housing portion (140) (as best seen in Figure 32A, Figure 37 and Figure 38). The design of the connector device (100) allows the device (100) to have two sections: 1) an electrical stimulation section and 2) a drug delivery integration section. This allows the device’s two stimulation modes to be isolated and independently used. According to an embodiment, the connector device (100) may have a dimension of 8.8 mm X 6.4mm X 4.1mm According to an embodiment, as shown in Figure 37, and Figure 38, the connector device (100) also has a tubing (145) received within the side opening (143) and extending away from the side surface (144) of the second housing (140). According to an embodiment, the tubing (145) may be removably secured to the side opening (143). The tubing (145) may have an outer diameter corresponding to about the diameter of the side opening (143). This geometry and sizing allow the external side wall of the tubing (145) to engage with the wall of the side opening (143). The diameter of the side opening (143) and the outer diameter of the tubing (145) may range between about 20 to about 23-gauges. Alternatively, the diameter of the side opening (143) and the outer diameter of the tubing (145) are about 20 gauges. As shown in Figure 32A, Figure 32B, Figure 35 and Figure 38, a fluidic channel attachment (146) extends outwardly from the side (144) of the second housing portion (140) and has an opening (147) concentrically aligned with the side opening (143). This opening (147) has a diameter greater than the diameter of the side opening (143) and the tubing (145) extends outwardly through the opening (147). The tubing (145) may be bonded to the inner wall of the opening (147) with a sealant (148). According to an embodiment, the opening (147) may have a diameter of about 2 mm. According to an embodiment, the sheet (150a, 150b) of conductive material may be made of copper or silver. Alternatively, the sheet (150a, 150b) of conductive material may be made of copper. According to an embodiment, each pin (160) may have the affixed end (161) conductively bonded to the end (151) of the corresponding sheet (150a, 150b) of conductive material. In another embodiment, the affixed end (161) may be soldered to the end (151) of the corresponding sheet (150a, 150b) of conductive material. According to an embodiment, as shown in Figure 32A, Figure 34, Figure 35 and Figure 38, each housing portion (130 and 140) has a bottom portion (149) having a curvature such that when the housing portions (130 and 140) are mating together, the bottom portion of the device (100) may have an overall curvature configured to match the curvature of a head of a subject. The device curvature design allows the device to be snuggly-placed the subject’s head (for example a mouse’s head) According to an embodiment, each housing portion (130 and 140) may be made of 3-D printing plastic Clear V4^TM. According to an embodiment, the spacer (170) is shaped as a block (as shown in Figure 32B) and may be made of acrylic. The spacer (170) may have a width ranging from about 1.7 mm to about 2.0 mm. Alternatively, the spacer (170) may have a width of about 1.8 mm. As mentioned above, the first and second housing portions (130,140) are configured to securely mate with each other. According to an embodiment, the connector device (100) has a snap-fit clip system to securely mating both housing portions (130 and 140) together. As shown in Figure 32A, Figure 32B, Figure 34 and Figure 38) The housing portion (130) has a pair of cantilevers (134) extending from its sides (135) in a parallel orientation. Each cantilever (134) has a protrusion (136) at its distal end. The protrusion (136) is configured to engage the side surface (144) of the housing portion (140) (as best shown in Figure 32A, Figure 32B and Figure 38). During assembly, the cantilever (134) flexes to allow insertion, and the protrusion (136) snaps into engagement with the side surface (144) to lock both housing portions (130, 140) together. In an embodiment, as shown in Figure 34, the inner wall (131) of the housing (130) has a pair of slots (137). Each slot (137) extends inwardly into the inner wall (131) adjacent to a corresponding cantilever (134). This configuration increases the overall length of the cantilever (134). According to an embodiment, as shown in Figure 32A, Figure 32B, Figure 34 and Figure 38, the device (100) also has a pair of pin clips (138). Each pin clip (138) is configured to slot a respective pin (160) into the corresponding opening (180) jointly defined by the top surfaces (132) of the housing portions (130 and 140). The pair of pin clips (160are located on the top surface (132) of the first housing portion (130) and and each is axially aligned with the opening (180). Fabrication of the connector device The connector device was made in two sections: 1) an electrical stimulation section and 2) a drug delivery integration section. This design configuration allowed the device’s two stimulation modes to be isolated and independently used. To allow clean separation between the stimulator and electrode systems, the pins (160) were kept as a part of the connector device (100) and used between the stimulator port and the flexible probes (110a, 100b) (Figure 32A, Figure 32B, Figure 37 and Figure 38). There are two major methods of electrical conduction: means of fusion and pressure. Means of fusion include methods of heating, such as welding, soldering and brazing connections. Since the electrodes (20) of the flexible probes (110a, 100b) are not solderable and cannot be heated, pressure was selected as the electrical conduction method, which includes clamping and compressed connections. Pressure conduction utilizes surface contact for the electrical connection, which depends on three parameters: the contact pressure, the contact area and the material. Contact pressure deforms asperities and increases the real contact area. Since resistance Ω is inversely proportional to area A and conductance G as shown in Eq.1, an increase in the area decreases the resistance and increases the conductance. Secondly, materials have various conductivities - a material property reflecting on the ability to conduct an electrical current. Silver and Copper are the two most conductive easily available materials. Copper has a conductivity of about 58 MS*m−1 at room temperature (pure copper). Ω = (ρ×L)/A=1/G Eq.1 The design of the connector device was developed with a focus on contact area and pressure. In order to maximize the contact area, the pin (160) and electrode portion (20) of the flexible probe (110a, 110b) were made into contact by a piece of copper sheet (150a, 150b), with one end of the copper sheet (150a, 150b) soldered to the pin (160) as shown in Figure 33. The contact pressure was applied between the electrode portion (20) and the metal sheet (150a, 150b) by a three-component system. With two housing portions (130 and140) spaced between a spacer block (170), the pressure between the spacer block (170) and housing portions (130 and 140) sandwiched each electrode portion (20) of the probes (110a, 110b) and its corresponding copper sheet (150a, 150b) together for electrical conduction (Figure 32B). To reduce the connector size, a snap-fit design was selected to secure the two housing portions (130 and 140) together. Due to geometry constraints and sizing requirements, a cantilevered uniform cross-sectional design was chosen. A pair of cantilevers (134) extend from the wall as shown in Figure 34. Slots (137b) were made on the wall in order to increase the clip length while minimizing the overall connector dimension. Additional pin clips (138) were integrated in order to allow the pin (160) to slot into the connector device (100) after soldering (Figure 32A, Figure 32B, Figure 34 and Figure 38). These features were designed on the same housing portion (130) (as shown in Figure 34), which allowed the additional features of the fluidic channel attachment (146) to be added on the other portion (140) without changing the design shown at this stage. This design enabled tuning of resistance between the copper sheet (150a, 150b) and electrode (20) by adjusting the spacer block (170) width between the housing portions (130 and 140). In order to allow fluidic ejection, an inlet into the housing portion (140) was required, with the outlet being the micro-fluidic channel (116) in the flexible probe (110b). A tubing with a diameter of 21-gauge was used to allow the tubing (145) to be pressed-fitted in the inlet (143) and the 3D-printing plastic, Clear V4^TM was used due to machining limitations. Additionally, to prevent leakage caused by unnecessary components, a single block design was selected. This design allowed the tubing (145) to be detached from the housing portion (145) when the drug is not being delivered. The attachment (146) was built on the snap-fit housing portion (140) where a 2-mm diameter hole (147) was printed on the housing portion (140) as shown in Figure 35. The goal was to maintain mostly the dimension of the housing portion (140), while building a system with no backflow or leakage. This was achieved by an additional larger-diameter opening (147) that provided additional surface area for a sealant (DOWSIL^TM 736) to latch on without increasing the original thickness of the housing (1 mm). Since there are differences in the accuracy of 3D-printing and laser-cutting, the hole diameters for the opening (143) ranging between 20 and 23-gauges were tested and a hole size of 20- gauge was determined sufficient for this connector device, where the tubing (145) was able to be fitted into the hole (143) with no gap between the external wall of the tubing (145) to the wall of the opening (143). For the purpose of simplification in testing and limited access to micro-fabrication facilities, the connector device was tested using two single-modal probes. The electrical conduction was reported using the resistance measured between the machined pins and the left probe (having only an electrode), and the drug delivery was tested for continuous flow ejection from the right probe (having only the microfluid channel). The final mass of the connector design was measured to be 0.39 g. The housing portions (130 and 140) were printed using the SLA-based 3D- printer, Form2, with a resolution of 25 μm using the resin Clear V4^TM. The spacing blocks were laser cut from a 3 mm thick acrylic sheet to a width of 1.8 mm. In order to maximize pressure in between the electrode (20) and copper sheet (150a, 150b), block widths ranging from 1.7 to 2.0 mm were assembled to test.1.8 mm was the largest width that allowed assembly without housing failure (cracking) and therefore was used in this final design. A 28-gauge tubing (Hamilton #20928) was fitted through the right housing portion (140), with sealant (DOWSIL^TM 736) applied into the larger-diameter opening (147) using a syringe with 21-gauge extruder. A 27-gauge pin (1000) was left in during the curing of 7 days in order to ensure that the tubing stays centered in the right housing as demonstrated in Figure 36. The pin (1000) was replaced with a needle after curing, and this configuration allowed the injecting pipe/tubing to be detached from the housing when not used. The inner wall (141) of the housing was filed slightly and cleaned with isopropyl alcohol to make sure the smoothness of the wall (141) and remove debris before the probe (10b) was taped using a double-faced tape, Arcare® 90106NB (step 1). This adhesive was used throughout the assembly process. In order to simply the fabrication process of the copper pieces (150a and 150b), the following manufacturing method was used. Due to the thickness of copper sheet (0.0035in), the sheet may be cut into shape simply using scissors. The specific dimensions and shape of the piece were not critical (150a and 150b), provided that it fully covered the electrode (20) being stimulated, and that the two pieces of copper sheets (150a and 150b) do not come in contact when soldered and assembled into the device (100). A hole was punched into each sheet using a sharp small drill bit, until a machined pin can be comfortably slid in with some resistance. Each piece of copper sheet (150a and 150b) and its respective pin (160) were then soldered together, with the solder forming a thin layer between the pin and the sheet. Each pin (160) was then cut off from the surface of the respective copper sheet (150a and 150b) and filed until smooth. Each piece (150a and 150b) was then bent. The pin (160) of the first copper- pin soldered piece was clip-fitted into the left housing portion (130) following the non- channeled electrode probe (10a). The copper piece (150a) was further secured down to the spacer block (170) using a piece of double-faced tape. The pin (160) of the other copper-pin soldered piece was then clipped in, before the right housing-probe assembly was snapped to the left housing portion (140) to form the final assembled device (100) as shown in Figure 37. The connector device (100) required a total of six steps as shown in Figure 38: 1) securing the probe having only an electrode to the inner wall of to the first housing portion; 2) securing the probe having the electrode and microfluidic channel to the inner wall of the second housing, 3) securing the first copper sheet and pin assembly against the first probe inside the first housing portion; 4) placing the spacer against the first copper sheet; 5) securing the second copper sheet and pin assembly against the spacer inside the first housing portion (therefore setting up the electrical stimulation section), and 6) securing the second housing portion to the first housing portion (therefore setting up the drug delivery integration section). Prior the assembly of the connector device, each pin is secured to its respective sheet of conductive material and a tubing is secured to the opening of the fluidic channel attachment. Three assembled devices s were tested to determine the resistance between the machined pin and connected left probe, which showed successful electrical connection with an average resistance of. about 5 Ω These assembled devices were tested to simulate for drug injection using diluted Coomassie blue dye using a flow velocity of 5 µL/min in order to speed up the observation. Droplet formations were initially found in all as demonstrated in Figure 39. Conclusion In an embodiment, the present disclosure provides a flexible probe. The flexible probe comprises a biodegradable body having first and second ends, the first end defining a region for electrical connection and the second end defines an electrically conductive site; a bioresorbable metallic electrode extending from the first end to the second end of the body; a biodegradable conductive polymeric coating located on at least a surface of the second end; and an insulating biodegradable polymeric sheath enveloping a portion of the body between the first and second ends. In an embodiment, another surface of the second end is coated with one of a conductive polymeric and an insulating polymeric coating. In an embodiment, the other surface of the second end is coated with an insulating polymeric coating. In an embodiment, the bioresorbable metallic electrode is made of a metal of _{molybdenum (Mo), tungsten (W), magnesium (Mg), Iron (Fe) and Zinc (Zn), any mixture} of these metals, and alloys of any combination of these metals. In an embodiment, the biodegradable conductive polymeric coating is made of any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. In an embodiment, the insulating polymeric sheath is made of any of poly(lactic- co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. In an embodiment, the probe is configured for insertion or implantation into a tissue, wherein the second end has a needle-like shape. In an embodiment, when in use, the probe retains its structural integrity therefore maintaining the stimulation properties of the probes for electrical stimulation for up to 14 days. In an embodiment, the electrically conductive site has a thickness in a range from about 20um to about 300 um, and a width in a range from about 100 to about 400 um. In an embodiment, the body defines a reservoir proximate to the first end and a delivery channel extending from the reservoir towards an opening proximate to the second end. In an embodiment, the delivery channel extends from the reservoir to the opening located at the second end. In an embodiment, the probe is configured for surface contact with a tissue wherein the second end has a section having a film-like configuration to adapt to morphology of the surface of the tissue. In an embodiment, two or more probes are used for the delivery of an electrical stimulation. In an embodiment, the stimulation is a biphasic monopolar current-controlled stimulation. In an embodiment, the pair of probes are used for the regulated neuromodulation of neural precursor cells. In an embodiment, the neural precursor cells are endogenous brain neural precursor cells. In an embodiment, the probes are used for further further deliver of a drug, virus, protein, small molecules, active ingredient or biologically active substance. In an embodiment, the probes are used for the recording of an electrical signal, wherein the probe includes two active sites, one site for recording and one site being electrical ground. In an embodiment, there is provided an implant. The implant comprises a support; and a pair of probes, the probes being mounted to the support in a parallel orientation to each other with the electrically conductive site of each probe facing each other, the probes being apart from each other by a distance between about 1 mm to about 3 mm. In an embodiment, there is provided a method of manufacturing a flexible probe. The method comprises the steps of providing a resorbable or biodegradable substrate; removably securing a mask onto a surface of the substrate, the mask defining a preselected geometry of the probe; depositing on the masked substrate a bioresorbable metallic material which defines an electrode with the preselected geometry; removing the mask; coating a portion of the electrode with an insulating biodegradable or resorbable polymeric material resulting in partially insulated electrode; coating a whole surface of the partially insulated electrode with a layer of patterning polymer; defining an electrically conductive surface of the electrode by removing a section of the layer of patterning polymer; coating at least the conductive surface of the electrode with a layer biodegradable conductive polymer; and removing the layer of patterning polymer from the electrode resulting in a flexible biodegradable and resorbable probe having an electrode with a conductive portion and an insulated portion. I_{n an embodiment of this method, the resorbable or biodegradable substrate is} made of any of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. I_{n an embodiment of this method, the bioresorbable metallic electrode is made of} any metal of molybdenum (Mo), tungsten (W), magnesium (MG), Iron (Fe) and Zinc (Zn), any mixture of these metals, and alloys of any combination of these metals. In an embodiment of this method, the insulating biodegradable or resorbable polymer is any one of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. In an embodiment of this method, the patterning polymer is Parylene. In an embodiment of this method, biodegradable conductive polymer is any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. In an embodiment of this method, the method further comprises the step of defining a reservoir and a delivery channel within the probe, the reservoir being proximate to an end region of the probe and the delivery channel in open communication with the reservoir and extending from the reservoir towards an opposite end of the probe. In an embodiment, there is provided a connector device for use with a first flexible probe having an electrode and a second flexible probe having an electrode and a microfluidic channel, each electrode having a side with an electrical conductive site. The device comprises a first housing portion having an inner wall, the inner wall being configured to receive the first probe with the electrically conductive site facing away from the inner wall; a second housing portion having an inner wall, the inner wall being configured to receive the second probe with the electrically conductive site facing away from the inner wall; a first sheet of conductive material configured to contact a electrical connection region of the electrode of the first probe; a second sheet of conductive material configured to contact the electrical connection region of the electrode of the second probe; a spacer electrically insulating the first sheet of conductive material from the second sheet of conductive material; two openings, each opening jointly defined by a top surface of each housing portions and configured to receive a pin; and a pair of pins, each pin extending through the respective opening and having an end affixed to an end of the corresponding sheet of conductive material, wherein the first and second housing portions are configured to securely mate with each other for defining a cavity enclosing each probe, the corresponding sheet of conductive material and the spacer, such that each probe is resting against the corresponding inner wall and each sheet of conductive material is positioned on opposite sides of the spacer whereby, in the assembled state, the electrical connection region of each probe is pressed against its respective sheet by contact pressure applied by the housing portions and the spacer conductively connecting the electrode of each probe to the corresponding pin, wherein a side surface of the second housing portion defines an opening in fluid communication with the microfluidic channel of the second flexible probe, and wherein a bottom surface of the second housing portion defines two openings, each opening is configured to receive a stimulation site of the electrode of the respective probe, the stimulation sites extending through the openings and projecting outwardly from the second housing portion. In an embodiment, the connector further comprises a tubing received within the side opening and extending away from a side surface of the second housing. In an embodiment, the tubing is removably secured to the side opening. In an embodiment, the tubing has an outer diameter corresponding to about the diameter of the side opening defined by the second housing portion thereby an external side wall of the tubing engages with the wall of the side opening. In this embodiment, the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing range between about 20 to about 23-gauges. In an embodiment, the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing are about 20 gauges. In an embodiment, the connector device further comprises a fluidic channel attachment extending outwardly from a side of the second housing portion and defining an opening concentrically aligned with the side opening of the second housing portion, wherein the opening of the fluidic channel attachment has a diameter greater than the diameter of the side opening of the second housing portion and the tubing is extending outwardly through the opening of the fluidic channel attachment. In an embodiment, the opening defined by the fluidic channel attachment has a diameter of about 2 mm. In an embodiment, the tubing is bonded to an inner wall of the opening defined by the fluidic channel attachment. In an embodiment, the sheet of conductive material is made of copper or silver. In an embodiment, the sheet of conductive material is made of copper. In an embodiment, each pin has the affixed end conductively bonded to the end of the corresponding sheet of conductive material. In an embodiment, the affixed end is soldered to the end of the corresponding sheet of conductive material. In an embodiment, each housing portion has a bottom portion having a curvature such that when the housing portions are mating together, a resulting bottom portion of the device has an overall curvature configured to match the curvature of a head of a subject. In an embodiment, each housing portion are made of 3-D printing plastic Clear V4^TM. In an embodiment, the spacer is shaped as a block. In an embodiment, the spacer is made of acrylic. In an embodiment, the spacer has a width ranging from about 1.7 mm to about 2.0 mm. In an embodiment, the spacer has a width of about 1.8 mm. In an embodiment, the connector device further comprises a snap-fit clip system to securely mating both housing portions together wherein one of the housing portion has a pair of cantilevers, each cantilever parallelly extending from an opposite side of the housing portion and having a protrusion at an distal end configured to mate with the side of the other housing portion when the housing portions are engaged with each other. In an embodiment, the inner wall of the housing defines a pair of slots, each slot extending inwardly into the inner wall adjacent to a corresponding cantilever, thereby increasing an overall length of the cantilever. In an embodiment, the connector device further comprises a pair of pin clips, each pin clip being configured to slot a respective pin into the corresponding opening jointly defined by the top surfaces of the housing portions, the pair of pin clips being located on the top surface of the first housing portion and axially aligned with the two openings. In an embodiment, there is provided a device. The device comprises a first neural probe having an electrode with an electrical conductive site, the first neural probe defining a through-opening; a second neural probe having an electrode with an electrical conductive site and a microfluidic channel, the second neural probe defining a through-opening; a central spacer defining a first channel extending from a side surface to an opposing side surface, each probe being secured to one of the opposite sides of the spacer with the electrical conductive site of each probe facing each other and each probe having the opening in alignment with the first channel; a first outer body secured to first probe, the first outer body defining second channel extending from a side surface to an opposing side surface, the second channel being in alignment with the opening of the first probe; and a second outer body having a solid surface secured to the second probe, wherein the opening of each probe is in alignment with the first channel and the second channel for fluid communication between the second channel and the microfluidic channel of the second probe, and wherein the first channel is closed by the solid surface of the second outer body. In an embodiment, the device further comprises a tube segment received within the second channel and extending outside the first outer body. In an embodiment, the second channel has a diameter greater than the diameter of the first channel.

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Claims

WE CLAIM: 1. A flexible probe, comprising: a biodegradable body having first and second ends, said first end defining a region for electrical connection and said second end defines an electrically conductive site; a bioresorbable metallic electrode extending from said first end to said second end of the body; a biodegradable conductive polymeric coating located on at least a surface of said second end; and an insulating biodegradable polymeric sheath enveloping a portion of the body between the first and second ends. 2. The probe according to claim 1, wherein another surface of said second end is coated with one of a conductive polymeric and an insulating polymeric coating. 3. The probe according to claim 2, wherein the other surface of said second end is coated with an insulating polymeric coating. 4. The probe according to any one of claims 1 to 3, wherein said bioresorbable metallic electrode is made of a metal of molybdenum (Mo), tungsten (W), magnesium _{(Mg), Iron (Fe) and Zinc (Zn), any mixture of these metals, and alloys of any} combination of these metals.

5. The probe according to any one of claims 1 to 4, wherein said biodegradable conductive polymeric coating is made of any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. 6. The probe according to any one of claims 1 to 5, wherein said insulating polymeric sheath is made of any of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. 7. The probe according to any one of claims 1 to 6, wherein said probe is configured for insertion or implantation into a tissue, wherein said second end has a needle-like shape. _{8. The probe according to any one of claims 1 to 7, wherein, when in use, the} probe retains its structural integrity therefore maintaining the stimulation properties of the probes for electrical stimulation for up to 14 days. 9. The probe according to anyone of claims 1 to 8 wherein the electrically conductive site has a thickness in a range from about 20um to about 300 um, and a width in a range from about 100 to about 400 um.

10. The probe according to anyone of claims 1 to 9, wherein the body defines a reservoir proximate to said first end and a delivery channel extending from said reservoir towards an opening proximate to said second end. 11. The probe according to claim 10, wherein the delivery channel extends from said reservoir to said opening located at said second end. 12. The probe according to any one of claims 1 to 6, wherein said probe is configured for surface contact with a tissue wherein said second end has a section having a film-like configuration to adapt to morphology of the surface of the tissue. 13. The use of two or more probes according to any one of claims 1 to 12 for the delivery of an electrical stimulation. 14. The use of a pair of probes according to claim 13, wherein the stimulation is a biphasic monopolar current-controlled stimulation. 15. The use of the pair of probes according to claim 14 for the regulated neuromodulation of neural precursor cells. 16. The use of the pairs of probes according to claim 15 wherein the neural precursor cells are endogenous brain neural precursor cells.

17. The use of the probes of claims according to anyone of claims 13 to 16 further comprising the delivery of a drug, virus, protein, small molecules, active ingredient or biologically active substance. 18. The use of at least one of the probes according to any one of claims 1 to 12 for the recording of an electrical signal, wherein the probe includes two active sites, one site for recording and one site being electrical ground. 19. An implant comprising: a support; and a pair of probes according to anyone of claims 1 to 11, the probes being mounted to the support in a parallel orientation to each other with the electrically conductive site of each probe facing each other, the probes being apart from each other by a distance between about 1 mm to about 3 mm. 20. A method of manufacturing a flexible probe, comprising the steps of: providing a resorbable or biodegradable substrate; removably securing a mask onto a surface of said substrate, said mask defining a preselected geometry of the probe; depositing on the masked substrate a bioresorbable metallic material which defines an electrode with said preselected geometry; removing said mask; coating a portion of said electrode with an insulating biodegradable or resorbable polymeric material resulting in partially insulated electrode; coating a whole surface of said partially insulated electrode with a layer of patterning polymer; defining an electrically conductive surface of said electrode by removing a section of said layer of patterning polymer; coating at least said conductive surface of said electrode with a layer biodegradable conductive polymer; and removing said layer of patterning polymer from the electrode resulting in a flexible biodegradable and resorbable probe having an electrode with a conductive portion and an insulated portion. _{21. The method of claim 20 wherein the resorbable or biodegradable substrate is} made of any of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. _{22. The method of claim 20 or 21, wherein the bioresorbable metallic electrode is} made of any metal of molybdenum (Mo), tungsten (W), magnesium (MG), Iron (Fe) and Zinc (Zn), any mixture of these metals, and alloys of any combination of these metals. 23. The method according to any one of claims 20 to 22, wherein the insulating biodegradable or resorbable polymer is any one of poly(lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), polycaprolactone (PCL), poly(ethylene glycol), silk, rice paper and cellulose based biodegradable materials. 24. The method according to any one of claims 20 to 23, wherein the patterning polymer is Parylene. 25. The method according to any one of claims 20 to 24, wherein biodegradable conductive polymer is any one of poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT;PSS), polyaniline (PANI), polypyrrole (PPy) and any derivatives thereof. 26. The method according to any one of claims 20 to 25, further comprising the step of defining a reservoir and a delivery channel within the probe, the reservoir being proximate to an end region of the probe and the delivery channel in open communication with the reservoir and extending from the reservoir towards an opposite end of the probe. 27. A connector device for use with a first flexible probe having an electrode and a second flexible probe having an electrode and a microfluidic channel, each electrode having a side with an electrical conductive site, the device comprising: a first housing portion having an inner wall, the inner wall being configured to receive the first probe with the electrically conductive site facing away from the inner wall; a second housing portion having an inner wall, the inner wall being configured to receive the second probe with the electrically conductive site facing away from the inner wall; a first sheet of conductive material configured to contact a electrical connection region of the electrode of the first probe; a second sheet of conductive material configured to contact the electrical connection region of the electrode of the second probe; a spacer electrically insulating the first sheet of conductive material from the second sheet of conductive material; two openings, each opening jointly defined by a top surface of each housing portions and configured to receive a pin; and a pair of pins, each pin extending through the respective opening and having an end affixed to an end of the corresponding sheet of conductive material, wherein the first and second housing portions are configured to securely mate with each other for defining a cavity enclosing each probe, the corresponding sheet of conductive material and the spacer, such that each probe is resting against the corresponding inner wall and each sheet of conductive material is positioned on opposite sides of the spacer whereby, in the assembled state, the electrical connection region of each probe is pressed against its respective sheet by contact pressure applied by the housing portions and the spacer conductively connecting the electrode of each probe to the corresponding pin, wherein a side surface of the second housing portion defines an opening in fluid communication with the microfluidic channel of the second flexible probe, and wherein a bottom surface of the second housing portion defines two openings, each opening is configured to receive a stimulation site of the electrode of the respective probe, the stimulation sites extending through the openings and projecting outwardly from the second housing portion. 28. The connector device according to claim 27, further comprising a tubing received within the side opening and extending away from a side surface of the second housing. 29. The connector device according to claim 28, wherein the tubing is removably secured to the side opening. 30. The connector device according to claim 28 or 29, wherein the tubing has an outer diameter corresponding to about the diameter of the side opening defined by the second housing portion thereby an external side wall of the tubing engages with the wall of the side opening. 31. The connector device according to claim 30, wherein the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing range between about 20 to about 23-gauges. 32. The connector device according to claim 31, wherein the diameter of the side opening defined by the second housing portion and the outer diameter of the tubing are about 20 gauges.

33. The connector device according to any one of claims 28 to 32, further comprising a fluidic channel attachment extending outwardly from a side of the second housing portion and defining an opening concentrically aligned with the side opening of the second housing portion, wherein the opening of the fluidic channel attachment has a diameter greater than the diameter of the side opening of the second housing portion and the tubing is extending outwardly through the opening of the fluidic channel attachment. 34. The connector device according to claim 33, wherein the opening defined by the fluidic channel attachment has a diameter of about 2 mm. 35. The connector device according to claim 33 or 34, wherein the tubing is bonded to an inner wall of the opening defined by the fluidic channel attachment. 36. The connector device according to any one of claims 27 to 35, wherein the sheet of conductive material is made of copper or silver. 37. The connector device according to claim 36, wherein the sheet of conductive material is made of copper. 38. The connector device according to any one of claims 27 to 37, wherein each pin has the affixed end conductively bonded to the end of the corresponding sheet of conductive material.

39. The connector device according to claim 38, wherein the affixed end is soldered to the end of the corresponding sheet of conductive material. 40. The connector device according to any one of claims 27 to 39, wherein each housing portion has a bottom portion having a curvature such that when the housing portions are mating together, a resulting bottom portion of the device has an overall curvature configured to match the curvature of a head of a subject. 41. The connector device according to any one of claims 27 to 40, wherein each housing portion are made of 3-D printing plastic Clear V4^TM. 42. The connector device according to any one of claims 27 to 41, wherein the spacer is shaped as a block. 43. The connector device according to claim 42, wherein the spacer is made of acrylic. 44. The connector device according to claim 42 or 43, wherein the spacer has a width ranging from about 1.7 mm to about 2.0 mm. 45. The connector device according to claim 44, wherein the spacer has a width of about 1.8 mm.

46. The connector device according to any one of claims 27 to 45, further comprising a snap-fit clip system to securely mating both housing portions together wherein one of the housing portion has a pair of cantilevers, each cantilever parallelly extending from an opposite side of the housing portion and having a protrusion at an distal end configured to mate with the side of the other housing portion when the housing portions are engaged with each other. 47. The connector device according to claim 46, wherein the inner wall of the housing defines a pair of slots, each slot extending inwardly into the inner wall adjacent to a corresponding cantilever, thereby increasing an overall length of the cantilever. 48. The connector device according to any one of claims 27 to 47, further comprising a pair of pin clips, each pin clip being configured to slot a respective pin into the corresponding opening jointly defined by the top surfaces of the housing portions, the pair of pin clips being located on the top surface of the first housing portion and axially aligned with the two openings. 49. A device comprising: a first neural probe having an electrode with an electrical conductive site, the first neural probe defining a through-opening; a second neural probe having an electrode with an electrical conductive site and a microfluidic channel, the second neural probe defining a through-opening; a central spacer defining a first channel extending from a side surface to an opposing side surface, each probe being secured to one of the opposite sides of the spacer with the electrical conductive site of each probe facing each other and each probe having the opening in alignment with the first channel; a first outer body secured to first probe, the first outer body defining second channel extending from a side surface to an opposing side surface, the second channel being in alignment with the opening of the first probe; and a second outer body having a solid surface secured to the second probe, wherein the opening of each probe is in alignment with the first channel and the second channel for fluid communication between the second channel and the microfluidic channel of the second probe, and wherein the first channel is closed by the solid surface of the second outer body. 50. The device according to claim 49, further comprising a tube segment received within the second channel and extending outside the first outer body. 51. The device according to claim 48 or 49, wherein the second channel has a diameter greater than the diameter of the first channel.