阅读说明：本技术 柔性电极及其制造方法 (Flexible electrode and method for manufacturing same ) 是由 劳伦·安妮·科斯特拉 克里斯托弗·K·蒂森 梅丽莎·斯科夫 大卫·雷默 凯尔西·布罗德里克 于 2020-03-06 设计创作，主要内容包括：本文公开的新型柔性电极利用静电纺丝纳米纤维垫的低弯曲刚度来实现手术植入和与外周神经持续双向通信所需的材料特性,而不会危及电子功能。根据本文公开的某些实施例,提供了可植入神经电极探针,其包括具有近端和远端的聚合基片、在基片近端的电极接口；至少一个位于基片远端的神经触点；以及形成在纤维基片上的导电迹线,提供至少一个神经触点和电极接口之间的电通信,其中基片包含聚合纳米纤维的无纺布块。(The novel flexible electrode disclosed herein utilizes the low bending stiffness of electrospun nanofiber mats to achieve the material properties needed for surgical implantation and continuous bidirectional communication with peripheral nerves without compromising electronic function. According to certain embodiments disclosed herein, there is provided an implantable neural electrode probe comprising a polymeric substrate having a proximal end and a distal end, an electrode interface at the proximal end of the substrate; at least one nerve contact located at the distal end of the substrate; and a conductive trace formed on the fibrous substrate providing electrical communication between the at least one nerve contact and the electrode interface, wherein the substrate comprises a non-woven mass of polymeric nanofibers.)

1. A flexible electrode, comprising:

a fibrous substrate comprised of a non-woven block of polymeric nanofibers, said substrate having at least one contact and electrode interface, and

a conductive trace formed on the fibrous substrate, the conductive trace providing electrical communication between the at least one contact and the electrode interface.

2. The flexible electrode of claim 1, wherein the fibrous substrate exhibits an elastic modulus between about 50MPa to about 5 GPa.

3. The flexible electrode of claim 1, wherein the polymeric nanofibers are formed from a plastic material selected from the group consisting of nylon, polycaprolactone, cellulose acetate, polymethyl methacrylate, vinyl alcohol, and polyimide.

4. The flexible electrode of claim 3, wherein the polymeric nanofibers are formed from nylon-6 and/or nylon-6, 12.

5. The flexible electrode of claim 1 further comprising a stabilizing cuff attached to a proximal portion of the fibrous substrate.

6. The flexible electrode of claim 1 further comprising support fibers axially attached to the fiber substrate.

7. The flexible electrode of claim 6, wherein the support fibers are formed of a high strength polymeric material.

8. The flexible electrode of claim 7, wherein the high strength polymeric material comprises high density polyethylene or polyaramid.

9. The flexible electrode of claim 2, wherein the substrate comprises a nanofiber layer comprising polymeric nanofibers and an insulating layer.

10. The flexible electrode of claim 9, wherein the insulating layer comprises a paraxylylene polymer or polydimethylsiloxane.

11. The flexible electrode of claim 1, wherein the substrate comprises a conductive polymer or metal component.

12. The flexible electrode of claim 11, wherein the conductive polymer composition comprises poly (3,4-ethylenedioxythiophene) polystyrene sulfate or gas-phase polymerized poly (3, 4-ethylenedioxythiophene).

13. The flexible electrode of claim 11, wherein the conductive polymer composition comprises conductively doped nanofibers formed from polycaprolactone, cellulose acetate, poly (methyl-methacrylate), or ethylene vinyl alcohol.

14. The flexible electrode of claim 11, wherein the fibrous substrate is coated with a conductive metal composition.

15. The flexible electrode of claim 2, wherein the polymeric nanofiber comprises an insulating coating.

16. The flexible electrode of claim 14, wherein the insulating coating comprises a paraxylylene polymer or polydimethylsiloxane.

17. A manufacturing method of a flexible electrode comprises the following steps:

(a) electrospinning a fibrous substrate layer composed of non-woven fabric blocks of polymeric nanofibers;

(b) depositing a conductive component on or within the fibrous substrate layer; and

(c) electrical traces are lithographically formed on or within the fiber substrate layer.

18. The method of claim 17, further comprising the step of:

(d) coating the substrate layer and electrical traces with a silicone layer; and thereafter

(e) Regions of the silicon layer are etched to form contacts.

19. The method of claim 18, further comprising the step of:

(f) forming the final shape of the electrode.

20. The method of claim 17, wherein step (b) comprises

A photoresist material is coated on the substrate layer and a conductive via pattern is etched.

21. The method of claim 20, wherein step (b) comprises

Adding a conductive material into the channel pattern.

Technical Field

Embodiments disclosed herein relate generally to advanced nanofiber-based material systems employed in electronic applications. According to certain embodiments, the nanofiber-based material systems are particularly useful as electrodes in various applications, such as biomedical applications, which, due to their enhanced flexibility, closely match the surrounding tissue in which the electrode is used.

Background

Flexible electrodes are applicable in various fields, such as bioengineering (implantable sensors, spinal implants, neural interfaces), wearable sensors, heating elements, etc. The ability to provide patterned electrodes is critical to these functions, but maintaining flexibility is particularly difficult. Neural connections using flexible electronic interfaces are used as a primary example throughout. For example, the development of neural interface technology has made a widespread breakthrough over the past decade. These advances include the most advanced prostheses controlled by neural interface electrodes, and the use of these interface prostheses to achieve sensitization. In sum, these advances have made it possible to improve the quality of life of amputees by replacing a lost limb with an advanced prosthesis. However, as the potential of these techniques continues to evolve, their clinical application is limited by the current lack of long-term stability and precision of the electrodes.

Two major limiting complications of existing neural interface technology are tissue damage and mechanical failure of the device over time. Even the most advanced technologies, such as transverse in-bundle multi-channel electrodes (TIME) and utah's oblique electrode array (USEA), apply high mechanical strain to the surrounding tissue during insertion, resulting in scarring, bleeding, and nerve tissue damage. This is mainly due to the significant mismatch in mechanical properties between the implanted electrode and the tissue in which it is implanted. For example, the elastic modulus of tissue in the Peripheral Nervous System (PNS) is in the range of 0.5-1.0MPa, while the elastic modulus of platinum and silicon (standard electrode interface materials) are 168,000 and 180,000MPa, respectively. This 5-6 order difference in elastic modulus and other mechanical properties leads to various implantation and long-term use problems, such as tissue damage, difficulty in surgical attachment, and relative movement during use. Most critical to performance is that it also leads to glial scarring (or encapsulation) phenomena whereby mechanical injury, cell death and inflammatory responses lead to glial encapsulation of the interface. This inhibits electrical contact with the nerve, eliminating the interface function. Addressing the need for neural interface technologies that can be placed and used over a long period of time to reduce complications would require electrode materials with mechanical properties that more closely match natural tissue. However, these materials must possess both the mechanical properties required for nerve penetration and placement, as well as electrical conductivity and interfacing arrangements for use with existing electrophysiology equipment.

A common problem with neural interface designs and other devices that require electrode flexibility is the material limitations present on each system. Peripheral nerves can undergo strains of up to 10% during normal human movement, a level of deformation that many currently used electrodes cannot withstand without exerting significant pressure on the surrounding delicate nerve tissue. Material property mismatches between the nerve and the implant often result in encapsulation and scarring between the electrode and the nerve fibers, which significantly reduces the signal-to-noise ratio of the electrode and the resulting sensitivity. The use of biological or softened materials provides a better match to the natural tissue properties, but the electrodes still typically contain wire conductive elements that do not soften and retain the material mismatch problem. Mobile or dissolvable carrier systems typically leave voids to fill scar tissue or may degrade into byproducts that lead to swelling, nerve impingement, inflammation, and electrode encapsulation. Biocompatible polymers, such as polyimide, silicone or polytetrafluoroethylene, show great promise for softer insulating layers, but the stiffness values of the materials used for the conductive contacts (e.g., platinum, iridium, or gold) are typically orders of magnitude greater than the stiffness values of neural tissue (600 kPa). Similar bending problems exist with wearable sensors, which may be integrated into textiles or work on the skin, and bending strains of more than 10% are generally expected.

The inability to properly match the material properties of the neural interface device to the native tissue limits the implementation of newly developed neural interface technologies. It would therefore be highly desirable if a neural interface device could be provided that more properly matches the physical properties of the native tissue. Likewise, flexible electrode material systems may also be used in wearable sensors and related devices. Embodiments disclosed herein are directed to providing such a solution.

Disclosure of Invention

To address the five to six order of magnitude difference in elastic modulus associated with currently existing natural tissues, embodiments disclosed herein provide a unique nanofiber-based material interface system. The novel material system disclosed herein takes advantage of the low bending stiffness of electrospun nanofiber mats to achieve the material properties required for surgical implantation and sustained bi-directional communication with human tissue, particularly peripheral nerves, without compromising electronic function. According to certain embodiments, the novel material techniques disclosed herein are particularly useful when embodied as a transverse intra-bundle multi-channel electrode (TIME) system, but may also be applicable to other future neural interface technologies. Similar systems, where the unique electrically patterned nanofiber substrate provides soft tissue-like mechanical properties, are also possible for other surgical implant products or wearable sensors and heating elements.

Unique aspects of the embodiments disclosed herein include:

the nanofiber-based insulating and supporting layers provide high sample flexibility while maintaining sufficient strength to make the device easy to implant and surgically tether.

The nanofiber metal-based or polymer-based conductive layer can reduce modulus mismatch and provide stable mechanical, electrical, and chemical properties during long-term implantation.

Material-based electrode designs are not "device-specific" and therefore can serve as a support platform for a variety of neural interfaces and flexible electronic applications.

As noted above, the neural interface material system of the disclosed embodiments is particularly useful for application of TIME neural interface design, but the disclosed embodiments may be applied in a variety of neural interfaces and biological applications. The novel material systems disclosed herein have inherent flexibility but tensile strength required for in vivo processing and insertion, for example, may find other potential applications in cranial electrode implants, muscle/musculoskeletal electrodes, and wearable surface electrodes. Thus, the flexibility of the novel material system disclosed herein may be able to provide two-way communication between nerves in the body and external electronic devices.

According to certain embodiments disclosed herein, a flexible electrode is provided that includes a fibrous substrate comprising a plurality of polymeric nanofibers, and a conductive trace formed on the fibrous substrate that provides electrical communication between at least one contact of the electrode and an electrode interface. The fibrous substrate may exhibit flexibility of about 50MPa to about 5GPA and be formed of polymeric nanofibers that may be formed of a plastic material selected from the group consisting of nylon, polycaprolactone, cellulose acetate, poly (methyl-methacrylate), vinyl alcohol, and polyimide). Preferred are nylon-6 (nylon-6) and nylon-6, 12 nanofibers.

A stabilizing cuff may be attached to the proximal portion of the fibrous substrate. To further enhance stability, support fibers formed of a high strength polymer (e.g., ultra-high density polyethylene or polyaramid) may be axially attached to the substrate.

According to certain embodiments, the substrate may include a nanofiber layer including polymeric nanofibers and an insulating layer, for example, a parylene polymer or polydimethylsiloxane.

The fibrous substrate may comprise a conductive polymer component, for example, poly (3,4-ethylenedioxythiophene) polystyrene sulfate and a vapor phase polymerized poly (3,4-ethylenedioxythiophene) metal coating (e.g., gold) deposited on the nanofiber substrate, and/or nanofibers formed from polycaprolactone, cellulose acetate, poly (methyl-methacrylate), and ethylene vinyl alcohol. The polymeric nanofibers may include an insulating coating, such as a paraxylylene polymer or polydimethylsiloxane.

The flexible electrodes disclosed herein may be formed by electrospinning a substrate layer comprised of a nonwoven mass of polymeric nanofibers, depositing a conductive component on or within the nanofiber substrate, and then photolithographically forming electrical traces in the substrate layer. The substrate layer may be coated with a silicon layer and then regions may be etched into the silicon layer to form contacts. The final shape of the electrode may then be formed by removing excess material.

These and other aspects of the present invention will become more apparent upon careful consideration of the following detailed description of the presently preferred exemplary embodiments of the invention.

Drawings

Referring to the drawings wherein:

FIG. 1 is a plan view illustrating a neural interface device, according to an embodiment of the present invention;

FIGS. 2A and 2B are photomicrographs of poly (3,4-ethylenedioxythiophene) polystyrene sulfate (PEDOT: PSS) material and vapor phase polymerized poly (3,4-ethylenedioxythiophene) (VPP PEDOT) material, respectively, which can be used as the conductive component in the substrate of the neural interface device shown in FIG. 1;

FIGS. 3A and 3B are photographs showing a poly (dimethylsiloxane) (PDMS) etching process in a nylon-6 nanofiber mat;

FIG. 4 depicts a sequence according to one embodiment for forming a single layer electrode;

FIG. 5 depicts a sequence according to another embodiment for forming a single layer electrode;

FIGS. 5A-5D are SEM photographs at different magnifications of oblique sections of a nanofiber substrate mat with conductive traces according to one embodiment of the invention;

FIGS. 5E-5G are optical images of a patterned surface on a mat of nanofiber substrate at respectively increased magnification;

FIGS. 5H and 5I are SEM cross-sectional photographs of electrodes fabricated using gold coatings on a nylon-6 nanofiber substrate, showing gold traces on a silicone layer and a silicone-filled nanofiber layer, respectively;

FIG. 6 is a schematic view of the setup of a multi-layer sample of hydrated 1 x Phosphate Buffered Saline (PBS), wherein the area enclosed by the dotted line represents the area of the sample where the tensile test was performed according to the following example;

fig. 7A-7D are bar graphs depicting results of property testing of multilayer samples of nylon-6/PEDOT: PSS + DMSO/nylon-6 and nylon-6/VPP PEDOT samples, immersed in 1 x PBS at 37 ℃ for t 0 weeks (n-3) and t 1 weeks (n-3), where fig. 7A depicts Ultimate Tensile Strength (UTS), fig. 7B depicts percent (%) elongation, fig. 7C depicts elastic modulus (MPa) and fig. 7D depicts toughness (MPa);

fig. 8A-8D are histograms depicting performance test results for nylon-6 and parylene-C coated nylon-6 used to dry and hydrate the test samples, respectively, where the symbol x represents significance when α is 0.05, and where fig. 8A depicts Ultimate Tensile Strength (UTS), fig. 8B depicts percent (%) elongation, fig. 8C depicts elastic modulus (MPa), and fig. 8D depicts toughness (MPa);

FIG. 9 is a graph of resistance (ohms, or ohms) versus frequency (Hz) for various sample materials tested in the following examples;

figure 10 is a bar graph showing bulk conductivity for two PEDOT conductive layer formulations according to the following examples;

FIGS. 11A and 11B are hydrated EIS impedance plots of Zreal (ohms) versus frequency (Hz) at 0 and 1 weeks, respectively, for the hydrated multilayer constructs;

FIG. 12 is an impedance spectrum of a nylon nanofiber sheet with a parylene-C layer and a gold strip electrode;

FIGS. 13A and 13B are Scanning Electron Microscope (SEM) photographs of a nylon-6 nanofiber and a nylon-6 nanofiber with a parylene-C coating, respectively;

figure 14 is a schematic of a vinyl pattern prepared with PEDT electrodes;

FIG. 15 is a series of plots of electrochemical analyses of PEDOT, PSS and VPP PEDOT materials;

FIGS. 16A and 16B are photographic images of a lithographically processed sample;

FIG. 17 is a schematic of a sequence of photolithography procedures for patterning a nylon-6 nanofiber mat;

FIGS. 18A-18C are photographs showing the results of the sequence shown in FIG. 17, where FIG. 18A shows nylon-6 bands on the wafer, FIG. 18B shows the bands after removal from the substrate, and FIG. 18C is an SEM image of a 50 μm patterned nylon-6 feature;

FIG. 19 shows a plan view illustrating a neural interface device, and an enlarged SEM view of electrode contacts etched thereon, in accordance with an embodiment of the present invention;

FIGS. 20A-20C illustrate, respectively, a neural interface device on a 500 μm scale (FIG. 20A), the device of FIG. 20A in proportion to one cent (FIG. 20B), and a wafer containing a plurality of devices etched thereon (FIG. 20C), in accordance with embodiments of the present invention;

FIGS. 21A-21C are images of a PEDOT: PSS coated silicon wafer (FIG. 21A), completion of photolithography using SU-8 resist and pattern transfer by reactive ion etching (FIG. 21B), and oxygen plasma etching of 6 μm PEDOT (FIG. 21C) after an annealing step and deposition of a 205nm SiN layer;

FIGS. 22A and 22B are bar graphs showing test results for nylon-6 nanofiber bi-layer samples tested in T-peel (T-peel) (FIG. 22A) and single lap joint (FIG. 22B) tests to evaluate the effect of (3-aminopropyl) triethoxysilane (APTES) additives and post-treatment heat treatment;

FIGS. 23A and 23B are bar graphs of direct and indirect cytotoxicity assays, respectively, showing cell viability following direct material exposure;

FIG. 24 is a schematic diagram of a TIME interface prepared for implantation including a neural interface device, according to an embodiment of the present invention.

Detailed Description

Embodiments disclosed herein are based on a new application of electrospun nanofibers as flexible substrates for patterned conductive traces. Thus, these embodiments are better able to mechanically match biological tissue properties, particularly neural tissue in the signature, while maintaining the electrical properties required for bi-directional communication with advanced prosthetic or other machine interfaces. This new material approach disclosed herein takes advantage of the inherent flexibility and biocompatibility of nanofiber mats and demonstrates the potential to utilize these substrates and advanced photolithographic patterning methods to fabricate ultra-flexible materials for biomedical applications.

An implantable neural electrode 10 in accordance with an embodiment of the present invention is depicted in fig. 1. As shown, the electrode 10 includes a substrate 12 formed of a non-woven mass of electrospun nanofibers having an electrical interface 14 (e.g., a TIME interface) at its proximal end and one or more contacts 16 (e.g., SIROF coated contacts) at its distal end. A conductive trace 18 (e.g., gold or PEDOT conductive layer) electrically connects the contact 16 with the electrical interface 14. The nanofiber substrate 12 may extend beyond the contact 16 to form an axially aligned insertion tip 20.

The nanofibers forming the substrate 12 preferably have an average diameter of indefinite length no more than about 100nm to about 750nm, typically between about 300nm to about 500 nm.

The nanofibers forming the substrate 12 may be formed from any suitable polymeric material that is biologically suitable for insertion into mammalian tissue, but is sufficiently flexible to achieve the desired flexibility characteristics. Particularly preferred polymeric materials that can be electrospun to form the substrate 12 include, for example, nylon homopolymers and copolymers, nylon-6 and nylon 6/12, polycaprolactone, and polyimide.

The substrate 12 should also exhibit sufficient flexibility to facilitate implantation and/or wear resistance. In particular, it is preferred that the substrate exhibit an elastic modulus of less than about 160GPA, typically between about 50MPa and about 5 GPA.

Embodiments of the present invention will be described in more detail below.

Constituent parts

A. Insulating nanofibers

The device disclosed herein includes a nanofiber-based interface for bi-directional communication with the peripheral nervous system based on a TIME interface. Since the stiffness of the device is in a cubic proportion to the characteristic dimensions of the system, the use of nanofibers means that the bending stiffness of the nanofibers is exceptionally low, even with high linear tensile strength. Thus, according to the disclosed embodiments, electrospun nanofibers may be utilized to maximize device flexibility and minimize mechanical stress on the peripheral nerve tissue after implantation. At the same time, the linear tensile strength of the material required for insertion and the long-term stability of the device can be maintained.

One exemplary embodiment of a nanofiber that may be usefully employed as an insulating template for a device embodying the present invention is an electrospun polyamide-6 (nylon-6) nanofiber. Such electrospun nylon-6 nanofibers are advantageous due to the compatibility of nylon-6 polymers with in vivo applications and their high linear tensile strength in nanofiber form. Standard needle-based and scale-up needleless (Elmarco) can be satisfactorily used) An electrospinning process to produce a nonwoven mat of nylon-6 nanofibers. In addition to nylon-6, other polymeric nanofibers can include, for example, nanofibers made from polycaprolactone, nylon-6/12, and polyimide electrospinning.

B. Conductive polymeric layer

For reasons similar to those described above, the implantable devices disclosed herein can also include a nanofiber-based conductive layer. Suitable conductive layers may be formed from nanofibers electrospun from poly (vinyl alcohol), chitosan, cellulose acetate, poly (L-lactic acid), poly (D, L-lactic acid), poly (methyl methacrylate), and polycaprolactone, which may be doped with various conductive carbon fillers (e.g., graphene oxide, carbon nanotubes, and carbon nanoparticles), and core-sheath electrospun nanofibers, in which less doped core nanofiber materials may be used to draw the carbon-doped sheath polymers into a nanofiber structure.

The conductive polymer poly (3,4-ethylenedioxythiophene) (PEDOT) can also be oxidatively polymerized onto the nanofiber template by incorporating the monomer 3,4-Ethylenedioxythiophene (EDOT) into various polymeric nanofiber chemistries (polycaprolactone, cellulose acetate, poly (methyl methacrylate), ethylene vinyl alcohol) and exposing the resulting electrospun mat to an oxidant solution (ferric chloride, ammonium persulfate). If used, the bulk conductivity values required for neural interface applications should be greater than about 100S/cm. Table 1 below shows the various components and end uses of the materials described herein.

TABLE 1

Poly (3,4-ethylenedioxythiophene) (PEDOT) may also be used as a bulk conductive material to improve electrolytic charge transfer between the electrode and neural tissue due to its previously demonstrated incorporation as an electrode coating in neural interface applications. Anionic doping of PEDOT is necessary to obtain high material conductivity; both the conductivity and the hydrolytic stability of the polymer can be adjusted by varying the counter-ion.

Two different deposition methods and anionic doping of PEDOT may also be employed according to embodiments disclosed herein. In this regard, poly (3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT: PSS) is an aqueous deposition product that can be cast as a solid film with bulk conductivity and sufficient electrolytic charge transfer. Gas phase polymerization (VPP) PEDOT, which is formed with anionic p-toluenesulfonate (pTS) and produces a highly flexible conductive nanofiber mat structure, may also be used. For the VPP PEDOT product, the oxidant component is electrospun within the polyvinylpyrrolidone (PVP) nanofiber structure and then exposed to EDOT vapors to initiate polymerization on the nanofiber structure. The PVP nanofiber template was subsequently dissolved in methanol, leaving the nanofiber PEDOT structure (see fig. 2B). Furthermore, the nanofibrous morphology allows electrolyte to penetrate into the conductive layer, increasing the charge transfer capacity by increased surface area compared to solid PEDOT films and possibly promoting cell penetration and adhesion. SEM images of each formulation are shown in fig. 2A and 2B.

Another alternative embodiment of this gas phase polymerization involves casting an aqueous solution of an oxidant component [ ferric tosylate, ferric chloride or Ammonium Persulfate (APS), preferably APS ] within a nanofiber structure (typically nylon-6). The oxidant component was then exposed to EDOT steam, as previously described, resulting in the polymerization of PEDOT directly surrounding the nylon-6 nanofiber structure. This process can be combined with advanced photolithography techniques to focus the polymerization within the patterned electrode template.

C. Additives and coatings

The addition of several additives and coatings serve various beneficial functions in the embodiments disclosed therein and are described in detail below.

Carbon-based dopants-as noted above, carbon-based dopants within the nanofiber structure and concentrated outside the nanofiber structure may be employed.

parylene-C/PDMS-a pad of electrospun nanofibers (e.g., nylon-6) may be coated with parylene-C prior to the photolithographic patterning process to prevent the photolithographic solvent from leaking through the void spaces of the sample, thereby preventing the pattern accuracy from being affected. parylene-C, however, is a moderately hard material (E ═ 2,757MPa), and therefore may affect the mechanical properties of the final nanofiber-based product. Thus, alternatively or additionally, Polydimethylsiloxane (PDMS) (E-1 MPa) can be used as a coating around the nanofiber structure, enabling patterning, insulation and stability of the underlying nanofiber (nylon-6) support layer.

Standard gold traces (thickness 400 nm) can be deposited on top of a parylene-C or PDMS filled nanofiber substrate and then encapsulated in additional insulating material to create an electrode with improved mechanical flexibility (EMod-2.5 GPa) due to the implementation of the nanofiber substrate.

Silane-various silanes, such as (3-glycidoxypropyl) trimethoxysilane (GOPS) and/or (3-aminopropyl) triethoxysilane (APTES), can be used as the insulating and conductive layers to increase adhesion between the layers and improve long-term mechanical stability of the electrode.

DMSO in the conductive layer-addition of dimethyl sulfoxide (DMSO) to PEDOT: PSS films can be used as a technique to increase conductivity. A commercially available aqueous solution of PEDOT: PSS (Ossil Ltd.) with a surface resistivity of 300 Ω/sq can be used, with an additional 5% (v/v) DMSO being added prior to casting. The conductivity values of PEDOT: PSS (883mS/cm) doped with DMSO can be nearly four times as high as those of undoped PEDOT: PSS (250.76 mS/cm).

STEC enhancers-several stretchability and conductivity (STEC) enhancers can optionally be included in the aqueous PEDOT: PSS film, such as lithium bis (trifluoromethane) sulfonamide, 4- (3-butyl-1-imidazole) -1-butanesulfonic acid trifluoromethanesulfonate, and/or 1-butyl-3-methylimidazolium octyl sulfate. STEC enhancers have been shown to weaken the interaction between PEDOT and PSS to improve the connectivity between the PEDOT domains.

D. Single layer system

The techniques disclosed herein may be embodied in a "single layer" system in which the conductive polymer component is incorporated into the insulating nanofiber component, and the parylene-C insulation is completely replaced with poly (dimethylsiloxane) (PDMS). The addition of a gold or PEDOT conductive layer to the nanofiber substrate will prevent potential delamination and reduce the thickness of the final product. While the embodiments described above use two layers of parylene C for electrical and photolithographic insulation, it has been found that such embodiments may be too stiff (e.g., with an elastic modulus of 3,460 MPa). However, the modulus of elasticity of PDMS is about 1MPa, which will enable further reduction of mechanical mismatch. Fig. 3A and 3B show PDMS incorporation and subsequent etching in a nanofiber mat.

As shown in the production sequence of fig. 4, the single layer electrode can be made by co-electrospinning nylon-6 nanofibers and ferric tosylate (FeTos) oxidant to make VPP PEDOT. After electrospinning the combined nylon/FeToS nanofiber layer (step 1 in fig. 4), a photoresist was applied and etched away from the conductive channel pattern (step 2 in fig. 4). The construction was then exposed to EDOT monomer to polymerize into the conductive channels and the mask was removed (steps 3 and 4 in fig. 4). The low PDMS viscosity allows the spin-on application to close all sides of the PEDOT traces (step 5 in fig. 4) and conductive contacts can be etched in the PDMS (step 6 in fig. 4). The electrode is then patterned into a final shape and rinsed to remove the exposed unreacted oxidant (FeTOS) (step 7 in fig. 4).

Based on the preliminary stability issues of the oxidant component, another electrode production process can be implemented to achieve a fast transition between the oxidant incorporation step and the subsequent polymerization, as shown in fig. 5. In this regard, the nylon-6 nanofiber mat was produced on a chrome-plated silicon wafer, as previously described (step 1 in fig. 5). A photoresist (e.g., S1813 photoresist commercially available from the Shipley Company) can be applied over the entire pad and etched to form the electrode channel pattern (step 2 in fig. 5). Oxidizing agent component (FeTOS, FeCl)₃Or APS) was deposited on top of the electrospun nanofiber mat and excess oxidant was removed using a rotary coater (step 3 in fig. 5). The sample was dried and immediately transferred to a gas phase polymerization chamber with EDOT vapor, and polymerization was induced under vacuum for 5 days (step 4 in fig. 5). The mask can then be removed (step 5 in fig. 5), leaving PEDOT in the channels and PDMS will be deposited around the conductive traces (step 6 in fig. 5). The contact points will be etched (step 7 in fig. 5) and eventually the electrodes will be patterned and removed (step 8 in fig. 5).

Another form of creating nanofiber electrode traces includes depositing titanium (30nm) and gold (100nm) by metal evaporation on a nanofiber substrate. As shown in fig. 5A and 5B, the oblique cross-sectional view of the fiber mat progressively exposes all layers under SEM. The top surface of the fiber mat shown in the upper portion of fig. 5A depicts the metal coated fibers, while the bottom portion of fig. 5B depicts the majority of the photoresist filled mat. At a higher magnification of fig. 5B, it is clear that the uncoated nylon-6 nanofiber layer under the top metal coated fiber can be seen and clearly distinguished. The metal coated fibers appeared brighter, coarser and larger in diameter due to the deposition of 130nm metal. Further investigation of the metal coated surface showed evidence of a uniform coating on and around the top fiber as shown in fig. 5C and 5D. The continuity of the metal and the absence of significant defects on the nanofiber surface seem promising for the fabrication of ultra-flexible nanofiber electrodes. The use of a metallic conductive component will improve the electrical properties compared to PEDOT and its application as a nanofiber coating will improve the mechanical properties compared to the first generation electrodes, where a continuous gold sheet is used as the conductive trace.

After successful validation of the gold deposition, the surface of the metal coating was patterned using a standard TIME electrode trace design. A new layer of photoresist is spin coated on the metal coated nanofiber mat and a pattern is transferred from the photoresist to the metal coating using wet chemical etching. The photoresist was then washed off by soaking in acetone and gentle stirring. Fig. 5E depicts an optical image of the patterned surface and shows that good pattern definition is achieved even in the narrowest trace (6 μm wide) of the design. Fig. 5F and 5G show a very clear contrast between the metal coating traces and the nylon-6 nanofiber mat after metal etching. The nanofibers appear to retain their structure after the photoresist is removed. The preliminary electrical properties of these samples confirm the continuity of the metal coating. If measured 2.5 mm apart using a digital multimeter, the electrodes fabricated using the gold coating on the nylon-6 nanofiber substrate were found to have a conductivity of 1 Ω resistance.

These devices were then filled with a silicone dispersion. Despite the addition of the gold traces, it was determined that silicone successfully penetrated and filled the nanofiber mat. Fig. 5H shows an SEM cross-section of the resulting sample, where the nanofiber-silicone composite can be easily distinguished from the additional silicone coating on top of the sample. The thickness of the top silicone layer was reduced to about 5 μm and in fig. 5I, the gold plated nanofiber traces between the top silicone layer and the silicone filled nylon-6 nanofiber mat were clearly visible. Note that the silicone layer separates from the top of the gold plated surface, which would be addressed by depositing an additional layer of titanium prior to implanting the silicone.

The embodiments disclosed herein will be further understood with reference to the following examples.

Examples

Device characterization

A. Mechanical characteristics

In addition to maintaining adequate electrical performance, it is also important to minimize the mechanical property mismatch between the conductive material and the surrounding neural tissue. For example, the metal wire currently used in conventional devices has an elastic modulus of about 160,000MPa, while the neural tissue is only about 0.6 MPa.

The mechanical properties of the multilayer nylon-6/PEDOT samples using standard tensile testing protocols have been characterized. Before the samples were mounted, the thickness of each sample was measured with a digital thickness gauge, and the width and length of each sample were measured with a digital micrometer. For ease of handling, the samples were mounted in a 1 x 5cm Low Density Polyethylene (LDPE) frame at the top and bottom (fig. 6). Briefly, a Loctite (Loctite) glue was placed on the top and bottom of each sample and the end of each sample was placed in overlapping position with the LDPE framework. The same block of LDPE frame was placed over the lower frame and sample end, sandwiching the sample between the two LDPE. Prior to the tensile test, the samples were cut using an 8.45 mm wide cutting die to ensure uniform force transfer over a consistent area (area outlined in fig. 6) of the 3-layer samples during the test. The prepared sample assembly was loaded into a high precision Instron 5943 tensile testing apparatus with a 500N load cell and pneumatic grips set at 70 psi. The samples were pre-loaded to 0.05MPa to eliminate relaxation and tested for failure at a deformation rate of 5 mm/min. Luna first tensile tests were conducted on samples immersed in 1 x PBS 0 at 37 c for 1 week to study the effect of hydration on tensile properties such as tensile strength, elongation, toughness and elastic modulus. The results are provided in FIGS. 7A-7D.

The mean modulus of elasticity at 0 weeks for the nylon-6/PEDOT: PSS + DMSO sample was 77.77 (+ -1.50) MPa. After 1 week of soaking in 1 × PBS, the average elastic modulus of the samples decreased to 74.56 (+ -13.53) MPa. The ultimate tensile strength of the PEDOT PSS + DMSO samples decreased with longer soak times. Visual evaluation and analysis of the stress-strain curves of these samples indicated that significant failure of each layer occurred during the test protocol. Although the elongation at break decreased slightly after 1 week of soaking, the samples maintained 10% elongation without failure, which is desirable when implanted.

Similar testing of the nylon-6/VPP PEDOT/nylon-6 samples indicated that the mean modulus of elasticity at week 0 and week 1 was 51.91 (+ -6.70) and 57.34 (+ -20.91) MPa, respectively. These values fall within 2 orders of magnitude of native neural tissue. The reduced modulus compared to PEDOT: PSS indicates a flexible nanofiber morphology of VPP PEDOT. Furthermore, the observed elongation of the sample of > 10% further demonstrates the continued investigation of VPP PEDOT as a flexible conductive layer.

As described above, parylene-C additives for lithographic pattern control and electrical layer encapsulation are known to increase the stiffness of the prototype sample. Thus, a tensile test was performed on a nylon-6 nanofiber mat coated with a 2 micron parylene-C layer as compared to an uncoated nylon-6 nanofiber mat (FIGS. 8A-8D).

The elastic modulus of the parylene-C coated nylon-6 nanofibers is significantly greater than that of the dry and hydrated samples of nylon-6. In addition, the initial elongation to failure of the parylene coated nylon-6 was reduced to about 3% upon drying and hydration. A model for the mixed contribution elastic modulus of nylon-6 nanofiber sheets used to analyze parylene-C coatings was developed based on the following formula:

E_Total＝V₁*E₁+V₂*E₂

where E is Young's modulus and V is volume fraction.

Since the length and width of the sheet are uniform, the volume fraction term can be reduced to a thickness fraction. This model was created to estimate the stretching effect of the thinnest possible parylene-layer on nylon. Using this model, the measurements were compared initially with literature values for the elastic modulus of parylene-C. From the data collected, the parylene-C elastic modulus was calculated to be 3,460MPa, while the literature value was 2,757 MPa. This indicates that the model will allow reasonably accurate estimation of the modulus for samples containing thinner parylene-C layers. In this regard, it is assumed that the 1 μm parylene-C layer may be the thinnest layer capable of providing sufficient electrical and photolithographic insulation. This thickness was input to the model to predict a hydrated elastic modulus of 692 MPa. This analysis confirms that the presence of parylene-C makes the structure stiffer and less elastic (although still significantly less stiff than the metal electrodes currently used). By comparison, the simulated 1 μm amorphous silicon carbide layer provided a predicted elastic modulus of 41,989MPa, much greater than that required for such an electrode employed in the embodiments disclosed herein.

A. Description of electrical characteristics

The following table summarizes the electrical property analysis of the devices according to embodiments disclosed herein:

type of analysis	Measuring properties
		Four point probe analysis	Bulk material conductivity
Cyclic voltammetry test	Material stability, charge transfer
		Electrochemical impedance spectroscopy	Impedance and phase angle of electrolysis
Equivalent circuit model fitting	Charge transfer characteristics

Characterization of bulk materials

Material systems have been electronically characterized using Electrochemical Impedance Spectroscopy (EIS) to ensure that the insulating chemistry is actually insulating and the conducting chemistry achieves the conductivity profile required for bi-directional communication with neural tissue. The resistivity of the nanofibers was measured to evaluate the conductivity change imparted by varying the bulk polymer, dopant chemistry, loading density, sample thickness and manufacturing process. For such characterization, samples were hydrated with 10-150. mu.S/cm standard solution and placed at 1cm using either a Gamry Interface 1000 electrochemical workstation or a Biologic SP200 system²Gold interdigitates the top of the electrode sensor. The EIS system then applies a sinusoidal voltage across the electrodes to measure the response current wave and evaluate the ohmic resistance and capacitance of the material. In that<In the range of 1Hz to 1MHz, a curve of measured impedance values can be generated to determine the conductivity of the sample in the region of interest. The impedance value was measured at an applied frequency of 1Hz to 1MHz and was measured at 10cm²Was applied with an alternating voltage of 10mV and a curve as shown in fig. 9 was generated. These resistance measurements can be converted to bulk material conductivity values (mS/cm) using gold interdigitated electrode (IDE) calibration curves.

Due to the limitations of the EIS system in characterizing highly conductive samples, the Jandel Four Point Probe (FPP) system is also used for conductive layer characterization. The four-point probe operates by inputting and then monitoring a dc constant current signal through four equidistant microcontact points. Current passes between the outer points, while the potential drop between the inner points is related to the material resistance. Assuming that the thickness of the layer to be measured is less than 5mm and the sample diameter is greater than or equal to 40 mm, the following equation:

where "V" is the applied voltage and "I" is the applied current, which determines the material resistance (Rs) Ω/sq for a particular case. The material resistivity (Ω cm) was then determined by multiplying the sample resistance by the sample thickness, and the conductivity (S/cm) was determined correlatively. The measurement is performed by placing the sample on a slide under the probe and lowering the probe head until it contacts the sample. A current, typically in the range of 10-100 mua, is selected to adequately measure the voltage drop across the sample.

The conductivity of both the PEDOT conductive layer formulation and the nylon-6 nanofiber mat was measured by a four point probe. Although the conductivity of nylon was too low to be detected, both PEDOT formulations achieved average conductivities in excess of 100S/cm; far in excess of the calculated 35S/cm required for use as an electrode material (fig. 10).

Hydrolytic encapsulation capacity of nano fiber

To investigate the ability of a nylon-6 nanofiber mat to effectively encapsulate and insulate conductive components, EIS measurements of IDE-based thick (16.33 μm) and thin (6.33 μm) hydrated (1 week in 1 × PBS) nylon-6 nanofiber mats were obtained. These data appear in fig. 11A and 11B. Although the impedance spectra of the two nylon layers at t-0 indicate some encapsulation, the impedance of both nylon-6 build layers decreased to below that of PBS after 1 week of immersion in 1 x PBS.

These results show that while nylon-6 provided partial insulation for the electrodes at the 0 week time point, the electrical encapsulation was insufficient and dropped to essentially zero at the 1 week measurement. Based on these results, it was determined that the nanofiber mat required an additional insulating component. A relatively thin (1-2 μm) layer of parylene-C was therefore added as an insulating additive. As shown by the impedance spectrum shown in fig. 12, parylene-C successfully isolated the nanofiber layer from electrolyte permeation as shown by the offset between the red and black dashed lines. SEM images of nylon nanofibers before and after parylene-C deposition are also shown in fig. 13A and 13B, respectively. Beneficially, the layer application also improves lithographic patterning by filling voids within the nanofiber material and preventing photoresist bleed.

Description of electrolytic Charge transfer characteristics

To further characterize the flexible neural electrode constructed in accordance with fig. 1, a 3-electrode Electrochemical Impedance Spectroscopy (EIS) cell was configured to measure the electrolytic interface characteristics. The cell used the PEDOT conductive layer of the electrode as the working electrode, a large platinum mesh as the counter (or auxiliary) electrode, and a BASi Ag/AgCl (3M NaCl) reference electrode. The electrolyte was 1 XPBS and the potentiostat was Biologic SP 200. Electrochemical Impedance Spectroscopy (EIS) and Cyclic Voltammetry (CV) tests were performed. The basic operation of potentiostatic EIS involves keeping a reference voltage constant while injecting current into the system at a range of frequencies (1Hz to 100kHz) and measuring the voltage transient at the working electrode. EIS describes the charge transfer resistance and phase angle. CV involves sweeping a series of voltages (-0.5V to +0.5V) at a constant rate (10 mV/s). CV illustrates the charge storage capacity and capacitance of the measured configuration.

To analyze these samples, PEDOT was wrapped so that only a small area was exposed to the electrolyte, while enough material was exposed above the electrolyte level to allow the electrodes to be connected to a potentiostat. Patterned vinyl decals were made using a vinyl cutter, with the opening diameter of the electrode contact area being 0.5 or 1 mm. Patterned vinyl decals were then attached to both sides of the PEDOT sample to insulate all of the immersed areas of the sample (except for the contact points). The test was arranged as shown in figure 14.

CV and EIS were performed on PEDOT: PSS and VPP PEDOT with electrode contact diameters of 1mm and 0.5 mm. The comparative data are shown in fig. 15. Both systems exhibit capacitive and pseudocapacitive behavior. According to the equivalent circuit model, about 30% of the current response is contributed by the true capacitance, while the pseudocapacitive behavior contributes the remaining 70% of the current response. Assuming a nerve outputting 100mV action potential, PEDOT: PSS will yield 5mC/cm for a 50 μm diameter area of the nerve sensor²Signal or 0.1 μ C, whereas VPP PEDOT will yield 10mC/cm for a 50 μm diameter neural sensor area²Signal or 0.2 μ C. The increase in capacitance of VPP PEDOT compared to PEDOT: PSS may be due to a significant increase in exposed surface area due to the nanofibrous morphology of the substrate.

Sample patterning

The electrode structure requires micron-scale patterning of the conductive portion of the electrode to achieve an interface with individual nerve fibers. In this regard, various methods of imparting macroscopic patterns to electrospun nanofiber layers have been investigated, including template patterning, photo-patterning, and photolithography. Template patterning was found to be suitable for larger scale patterns (millimeters) but not for the micron-scale resolution required for neural electrodes. Photopatterning, while feasible, is not recommended because of the need for functionalization of the nanofiber chemistry, which can significantly affect the resulting conductivity. Therefore, photolithography is currently the preferred patterning method to obtain the electrodes of the present invention.

Preliminary patterning experiments were performed to assess whether the chemistry used was compatible with the solvents and processing conditions of the lithography shown in fig. 16A and 16B. Some blistering and sample burning is evident due to inefficient adhesion of the sample to the silicone wafer, solvent penetration into the nanofiber layer, and sub-optimal temperatures of the "burning" surface. However, in general, the basic compatibility of the photolithographic process with these electrospun samples has been demonstrated.

The nylon-6 nanofiber mat has been electrospun directly onto a metal coated silicon wafer for subsequent etching of the nylon-6 layer using the process outlined in figure 17. As shown in fig. 17, the nylon-6 pad was first coated with a 1-2 μm layer of parylene-C to enhance insulation, stability and improve the patterning process. The results of this process are shown in the photographs of FIGS. 18A-18C.

To facilitate implantation in vivo, the nerve electrode shown in fig. 1 was created from a Nylon-6 nanofiber substrate, a Parylene-C layer, a 400nm patterned gold conductive layer, and an insulating Parylene-C top layer. Such electrodes perform significantly better than conventional neural electrodes. Images of such electrodes are shown in fig. 19 and 20A-20C.

To use other solvents, PEDOT: PSS films need to be coated with a barrier layer (e.g., silicon nitride) to avoid direct contact. Silicon nitride deposition leads to the formation of cracks and bubbles, and therefore alternative methods need to be developed. PSS films at 150 ℃ prior to deposition of silicon nitride have proven feasible and can then be removed by dry etching using fluorine plasma reactive ion etching. Upon inspection, some delamination of PEDOT was observed, and thus complete etching of 18 μm thickness was not completed. To alleviate this delamination, the solvent compatibility of cast PEDOT: PSS films was investigated and it was determined that isopropanol and acetone as compatible solvents did not delaminate the films. Solvents containing tetramethylammonium hydroxide (TMAH) in any proportion decompose PEDOT: PSS films and are therefore unusable. FIGS. 21A-21C are images of a PEDOT: PSS coated silicon wafer (FIG. 21A), a photolithographic and pattern transfer by reactive ion etching done using SU-8 resist (FIG. 21B), and an oxygen plasma etch of 6 μm PEDOT (FIG. 21C) after an annealing step and deposition of a 205nm SiN layer (FIG. 21A).

Description of the stability Properties

Preliminary soaking/stability tests performed by immersing the prototype interface material in saline have been demonstrated. The results of this experiment show that there is no significant change in the elastic modulus or peak tensile strength after one week. Qualitative analysis of the long term soaking of these preliminary samples showed no change in structure or stratification over a 6 month soak period.

Other attempts to characterize the stability of the samples included immersing three layers of nylon-6 wrapped PEDOT: PSS and VPP PEDOT samples in PBS for extended periods of time. In addition to a qualitative assessment of the stability of the samples, the electrical and mechanical properties of these samples were also characterized. Qualitative delamination of the thinner (-16 microns) samples and subsequent degradation of the conductive layer in these samples was observed. In these preliminary characterization tasks, thicker samples (-32 microns) appeared to retain higher stability.

Due to potential problems associated with electrode delamination after implantation, the use of several silane additives in PEDOT and nylon-6 compositions has been investigated. The sheet adhesion has been tested for shear and normal forces to evaluate adhesion of adjacent layers within the electrode (particularly between the nylon-6 layers encapsulating the patterned conductive layer). A modified version of ASTM D1876 and D3163 (T-peel and single lap test) was followed to evaluate the adhesion between two electrospun nylon-6 pads.

Adhesion testing was used to compare adhesion between nylon-6 samples with and without the (3-aminopropyl) triethoxysilane (APTES) silane additive (150 ℃ for 1 hour; used in the PEDOT lithography development process). These results are shown in FIGS. 22A-22B. No significant difference was observed in the results of the T-peel (T-peel) test. The results show that the APTES addition with heat treatment studied did not have a significant effect on the normal adhesion of the nylon-6 nanofiber self-adhesive. Interestingly, it appears that the post-production heat treatment did have a significant effect on the adhesive strength exhibited by these samples (p < 0.05). It is hypothesized that heating the sample above the glass transition temperature of nylon-6 (47 ℃) causes entanglement of the nylon-6 chains at the interface between the fibers of each layer; the magnitude of this increased adhesion is very significant when normalized across the overlap region. The test further indicates that APTES currently does not contribute to the adhesion that these samples have. However, the addition of a heat treatment protocol step appears to significantly increase the shear stress required to separate the two layers, a valuable property for in vivo constructs. Formulation of a "single layer" system with the PEDOT and silicone all impregnated into the nylon-6 nanofiber layer would eliminate all concerns about electrode delamination upon implantation/prolonged hydration.

Description of cell compatibility characteristics

Cell compatibility screening using nylon-6/PEDOT samples with and without parylene-C and APTES additives has been accomplished. Samples were cultured in direct contact with S42 rat Schwann cells, and LDH and MTT assays were used to characterize the resulting cell compatibility. Overall, it was observed that the cells appeared healthy in appearance, with typical morphology, when cultured in contact with nylon-6 samples containing parylene-C, APTES or VPP PEDOT. However, it was observed that those cultures incubated with nylon-6 + PEDOT: PSS (with or without DMSO) contained low to moderately healthy adherent cells, which were significantly less robust and had an atypical morphology that might indicate a toxic effect. Cell viability (MTT) measurements indicate that these cell compatibility problems are due to undefined leachate during manufacturing. The evaluation of samples with and without DMSO or release agent was tested and the results of these experiments indicated that the leachant may be from the PEDOT: PSS component. Although the impact was only moderate, PEDOT: PSS appears to be a major contributor to toxicity in these studies, possibly due to contaminants, stock purity, or some aspect of preparation or handling. Luna found that applying a negative charge (-0.5V) to a conductive sample in PBS resulted in increased cell attachment of rat Schwann cells (S42; ATCC CRL-2942) to the sample. These results were confirmed by MTT analysis and LIVE/DEAD staining. To assess whether these changes were due to charge application or wash steps, the cellular compatibility of various sample treatments were assessed, including PEDOT with and without wash and charge steps, PSS and VPP PEDOT.

Samples were evaluated for cellular compatibility in both direct and indirect culture formats. Controls consisted of low density polyethylene (LDPE; negative toxicity), natural rubber latex (positive toxicity) and cells only (baseline). In the direct format, 8 mm material samples were incubated directly with S42 cell cultures for 24 hours. After incubation, cell viability was visually observed and confirmed using the MTT assay. Indirect methods include extraction of the material into a solution, followed by incubation of the extract with the culture. Based on the recommendations of ISO 10993 for the biological evaluation of medical devices, this protocol was used for in vitro cytotoxicity tests with 8 mm dressings in a medium (1.3 cm) containing 10% Fetal Bovine Serum (FBS)²/mL) for about 24 hours at 37 ℃ and 5.0% CO₂。

In the direct exposure study, it was observed that cells incubated with the rinsed dressing material were generally healthy in appearance, with a typical morphology, consistent in quality with the pure cell baseline and the LDPE negative toxicity control. In the direct format test, only very slight signs of toxic phenotype on the surface using the unwashed nylon-6 PEDOT: PSS/DMSO material were observed. This effect is clearly less pronounced than the strong response to the latex positive control, with almost all cells rounded and shed, which is a clear sign of poor health and low viability. Cultures exposed to nylon-6 VPP PEDOT and nylon-6 PEDOT: PSS/DMSO rinse variants showed no toxic effects, with no significant difference between the culture groups. Finally, nylon-6 PEDOT: PSS/DMSO material after washing for 2 hours with and without the-0.5V charge, the cultures were consistent in appearance, indicating that the charge may be a neutral variable. These results were confirmed using MTT assay results, as shown in fig. 23A. High survival was observed for both nylon-6/VPP PEDOT variants, and with increasing number of rinses, the survival of the PEDOT: PSS/DMSO test samples increased. For nylon-6 PEDOT: PSS/DMSO, there was no statistical difference between the 2 hour washes with and without a-0.5V charge, indicating that the charge may be a neutral variable. PSS/DMSO were statistically less viable than the LDPE negative control (p 0.01) only, and suggested that the unlushed nylon-6 VPP PEDOT was less viable (p 0.08) by ANOVA and Tukey analysis. This supports qualitative observations and indicates that the toxic components present in each chemical have been substantially reduced by inclusion of a washing process.

For indirect studies, the unwashed nylon-6 PEDOT: PSS/DMSO extract cultures again had only slight visible signs of toxicity. The mild effect here is more easily observed as round adherent cells; this is clearly less pronounced than the latex positive toxicity control extract, where almost all cells are round and isolated. Cells cultured with 2 and 48 hour washed nylon-6 PEDOT: PSS/DMSO extract material appeared healthier and similar to the negative control culture, indicating that both washing protocols substantially reduced the presence of toxic components in the dressing. Again, these results were confirmed with MTT assay (fig. 23B). Cell viability was statistically lower for the unwashed sample extract alone than for the LDPE negative control (p) by analysis of variance (ANOVA) and Tukey post hoc (post-post hoc)<0.001), the two wash samples were statistically comparable. This supports qualitative observations and indicates that the presence of toxic components in nylon-6 PEDOT: PSS/DMSO chemistry has been substantially reduced by the rinsing process. All tests and negative controls were smaller than the latex positive control (p.ltoreq.1X 10)^-8). There was no statistically significant difference between the 2 hour and 48 hour rinse treatments, indicating that the short and extended rinse process had no significant benefit for toxicant extraction.

Overall, these findings generally indicate cellular compatibility and confirm the benefits of immersion rinse in DPBS after preparation. The observations and measurements herein show that the use of undushed nylon-6 PEDOT: PSS/DMSO and nylon-6 VPP PEDOT effectively reduced mild cytotoxicity in both chemistries by including a post-preparation DPBS rinse. This rinsing step successfully leaches toxic residues from the material, thereby reducing the cytotoxicity of the rinsed dressing. Thus, both the rinsed VPP PEDOT and PEDOT: PSS based materials have been identified as cell-compatible candidates suitable for the conductive properties in the final dressing.

Description of sterility

Sterilization and cellular compatibility of materials and systems were investigated. Sterilization was performed using gamma radiation (Steris Isomedix) and ethylene oxide treatment. Sterility was confirmed using direct inoculation (soaking) techniques in liquid Thioglycolate (Fluid Thioglycolate) and Trypticase Soy Broth (TSBs) without adverse effects on electrical, mechanical or biological properties.

Neural devices have been studied in vivo models for hindlimb electrode implantation of neural interfaces. To prepare for implantation, rat sciatic nerves were exposed near their trifurcations; the TIME electrode implant is for placement in the tibial branch. An 80 μm tungsten needle with a Kevlar fiber pigtail loop was used to penetrate the nerve bundle and pass the loop through. After the needle is passed, the ring is pulled to a position that intersects the introducer tip of the TIME interface to be implanted. This loop is closed by pulling on the introducer tip of the electrode and pulls the electrode into the nerve.

To increase the stability of the electrode during implantation, the TIME interface was connected to a cuff electrode placed on the upper part of the nerve. To prevent the introducer ring from sliding off the electrode tip during implantation, Kevlar (Kevlar) "support fibers" are attached. The nanofiber-based electrode is free to conform during implantation, with no critical bend radius observed. In fact, this flexibility has been identified as a possible reason for the introducer tip slipping out of the Kevlar introducer loop during implantation.

As mentioned above, a method has also been developed to connect the TIME interface 14 to a cuff electrode 30 having electrode contacts 30a, 30b for stability during implantation. Such an embodiment is depicted in fig. 24. In addition to attaching the TIME interface 14 to the cuff electrode 30 for stability, Kevlar fibres 35 are adhered over the length of the electrode 10 to reduce tensile stress on the electrode 10. Point 35a represents a potential connection area of Kevlar fibres 35 to TIME interface 14. By adding Kevlar support fibers 35 to the TIME interface 14, the process of securing the electrodes is simplified. After successful electrode implantation, the Kevlar fibers 35 within the tract may be severed and pulled out of the nerve. The different ends of the fibre 35 will now be on either side of the nerve and act as connection points where sutures can be tied to secure the electrodes in place. Advanced packaging systems are intended to facilitate implantation and eliminate subsequent movement of the device that would affect the recording. These systems have been demonstrated in vivo using Qualia control electrodes.

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While the invention has been described in connection with what is presently considered to be the most practical and preferred embodiment, it is to be understood that the invention is not to be limited to the disclosed embodiment, but on the contrary, is intended to cover various modifications and equivalent arrangements included within the spirit and scope thereof.

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