Zwitterionic pedot:PSS hydrogel (ZIPH) as implantable conductors with suppressed foreign-body responses

Abstract

A mixed ion / electron conducting material that leads to suppressed fibrotic responses when implanted inside a living body is described. The mixed ion / electron conducting material includes a zwitterionic PEDOT:PSS hydrogel having a first network containing interconnected fibers of poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS); and a second network containing a zwitterionic polymer hydrogel, where the first network is formed in situ in the second network to form a double network hydrogel. A method of manufacturing the mixed ion / electron conducting material and an implantable medical device including the mixed ion / electron conducting material are also described.

Description

Attorney Docket No.507814.5000508 ZWITTERIONIC PEDOT:PSS HYDROGEL (ZIPH) AS IMPLANTABLE CONDUCTORS WITH SUPPRESSED FOREIGN-BODY RESPONSES CROSS REFERENCES

[0001] This application is based on and claims the benefit of and priority to U.S. Provisional Patent Application No.63 / 645,429, filed on May 10, 2024, which is herein incorporated by reference in its entirety. GOVERNMENT LICENSE RIGHTS

[0002] This invention was made with government support under EB034563 awarded by the National Institutes of Health and 2105367 awarded by the National Science Foundation. The government has certain rights in the invention. BACKGROUND

[0003] For long-term, continuous operation of implantable biosensors and electrophysiological devices, adverse foreign body response (FBR) to the implanted device is a major obstacle that needs to be overcome. As the FBR progresses, any implanted device will become damaged and isolated from its physiological environment due to, for example, encapsulation by fibrotic tissue and inflammatory immune cells. To achieve more compatible and low-impedance biointerfaces, conducting polymers, such as poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate (PEDOT:PSS), have been extensively explored as candidate implantable materials. However, FBR to such conducting polymers remains an unmet challenge. Therefore, design strategies for improving the intrinsic immunocompatibility of PEDOT:PSS are highly desirable. SUMMARY

[0004] This disclosure generally relates to a mixed ion / electron conducting material that leads to suppressed fibrotic responses when implanted inside a living body.

[0005] In some example embodiments, the mixed ion / electron conducting material includes a zwitterionic PEDOT:PSS hydrogel comprising a first network comprising interconnected fibers of PEDOT:PSS; and a second network comprising a zwitterionic polymer hydrogel, where the first network is formed in situ in the second network to form a double network hydrogel.Attorney Docket No.507814.5000508

[0006] In some example embodiments, a method for manufacturing a mixed ion / electron conducting material with suppressed foreign body responses is disclosed. The method includes forming a zwitterionic PEDOT:PSS hydrogel comprising a first network comprising interconnected fibers of PEDOT:PSS; and a second network comprising a zwitterionic polymer hydrogel, where the first network is formed in situ in the second network to form a double network hydrogel.

[0007] In other example embodiments, an implantable medical device for implanting inside a living body comprising a mixed ion / electron conducting material with suppressed foreign body responses is disclosed. BRIEF DESCRIPTION OF THE DRAWINGS

[0008] FIG. 1 is a schematic illustration of an example mixed ion / electron conducting material comprising a zwitterionic PEDOT:PSS double-network hydrogel (abbreviated as ZIPH) including a first network comprising interconnected fibers of PEDOT:PSS and a second network comprising a zwitterionic polymer hydrogel.

[0009] FIG.2 is a flow chart showing exemplary steps of a method of forming the ZIPH double network hydrogel.

[0010] FIG. 3A is a schematic illustration of morphological changes that occur to ZIPH after being subjected to ethanol-TFE annealing treatment cycles.

[0011] FIG.3B shows cryo-TEM (Tunneling Electron Microscopy) images of ZIPH subjected to various solvent annealing conditions.

[0012] FIG. 3C shows C-AFM (Atomic Force Microscopy) images of ZIPH before (ZIPH-E) and after 3 cycles of ethanol-TFE annealing treatment (ZIPH-3c).

[0013] FIG. 3D shows GIXD (Grazing Incidence X-ray Diffraction) patterns of poly(sulfobetaine methacrylate) (PSB) hydrogel and ZIPH subjected to various solvent annealing conditions.

[0014] FIG.3E shows the conductivity values of ZIPH prepared with various solvent annealing conditions. All hydrogels were pre-equilibrated in phosphate buffered saline (PBS) and autoclaved before measurement.

[0015] FIG.3F shows the tensile curves of ZIPH before (ZIPH-E) and after 3 cycles of ethanol-TFE annealing treatment (ZIPH-3c).

[0016] FIGS.4A and 4B each shows representative Masson’s Trichrome Staining (MTS)-stained tissue sections of explanted PEDOT:PSS and ZIPH-3c, respectively.Attorney Docket No.507814.5000508 The collagen layer is marked with arrows, and the locations of implants appear clear and are indicated by asterisks.

[0017] FIG. 5 shows the average normalized collagen density and conductivity value of ZIPH-3c compared to PEDOT:PSS.

[0018] FIG.6 shows representative MTS-stained tissues from 6 types of hydrogel 1-, 4-, and 12-weeks post implantation. Locations of implants appear clear and are indicated by asterisks.

[0019] FIGS.7A and 7B show collagen density (blue pixel density) as a function of a distance from the surface of the implant 4-weeks post-implantation, determined from MTS-stained tissue.

[0020] FIGS. 8A and 8B show measurements of collagen capsule density of morphology of the collagen capsule encapsulating ZIPH-3c and PEDOT:PSS.

[0021] FIGS. 8C and 8D show measurements of collagen capsule thickness of morphology of the collagen capsule encapsulating ZIPH-3c and PEDOT.

[0022] FIG. 8E shows the MTS-stained microscope images of morphology of the collagen capsule encapsulating ZIPH-3c and PEDOT:PSS of FIGS. 8A-D at various timepoints.

[0023] FIG.9A shows the total counts of macrophages associated with each type of implant, determined by flow cytometry and 4-week post-implantation. Macrophages were defined as F4 / 80+cells.

[0024] FIG.9B shows the M2 / M1 ratios associated with each implant, determined by flow cytometry at 4 weeks post implantation. M1 macrophages were defined as CD86+CD206+macrophages, and M2 macrophages were defined as CD86- CD206+macrophages. The M2 / M1 ratios are upregulated for PEDOT:PSS but are not for the rest.

[0025] FIG. 9C shows representative immunofluorescence microscopy images of tissues stained with F4 / 80 (macrophage), CD86 (M1), CD206 (M2). Locations of implants appear clear and are marked by asterisks.

[0026] FIGS. 9D and 9E show key cytokines associated with each implant, determined via LegendPlex assay at 1-week post-implantation.

[0027] FIGS. 9F and 9G show key cytokines associated with each implant, determined via LegendPlex assay at 4-week post-implantation. monocyteAttorney Docket No.507814.5000508 chemoattractant protein-1 (MCP-1) is consistently upregulated for both ZIPH-3c and ZIPH-u.

[0028] FIG. 9H shows total counts of neutrophils associated with each implant, determined by flow cytometry 1-week post-implantation. Neutrophils were defined as Ly-6G+cells.

[0029] FIG.9I shows complement C3a concentration associated with each implant, determined by enzyme-linked immunosorbent assay (ELISA) 1-week post- implantation. The high number of neutrophils is correlated with the generation of C3a. PSB has a distinctly higher neutrophil number and C3a concentration than the rest of the hydrogels that contain poly(styrene sulfonate (PSS).

[0030] FIG. 10A shows principal component analysis (PCA) of a 50-gene panel from 4-week post-implantation samples plotted against each other. The fibrotic and mildly fibrotic samples occupy distinct areas in the plot.

[0031] FIG. 10B shows genes associated with the negative direction of the third principal component (PC3). The expression of Tbx21 (T-bet) makes ZIPH-3c unique among the samples in the experiment. T helper cell-related genes are highlighted.

[0032] FIG.10C shows differentially expressed genes associated with the positive direction of the first principal component (PC1). All the genes have some connection to fibrotic diseases. Genes that are statistically significantly different from sham control and have a strong connection with fibrosis are highlighted.

[0033] FIG. 10D shows total counts of fibroblasts associated with each implant, determined by flow cytometry 4 weeks post-implantation. Fibroblasts are defined as CD45- Lin- CD140a+, in which Lin consists of CD31, epithelial cell adhesion molecule (Ep-CAM), Tie-2, and Ter-119.

[0034] FIG. 10E shows the representative immunofluorescence microscopy images of tissues stained with collagen I, SMA (myofibroblast), and F4 / 80 (macrophage). Locations of implants are marked by asterisks.

[0035] FIG.11A shows MCP-1, vascular endothelial growth factor (VEGF), and IL- 18 peri-implant concentrations of different implant types at 1 week.

[0036] FIG. 11B shows MCP-1, VEGF, and IL-18 peri-implant concentrations of different implant types at 2 weeks.Attorney Docket No.507814.5000508

[0037] FIG. 11C shows MCP-1, VEGF, and IL-18 peri-implant concentrations of different implant types at 4 weeks.

[0038] FIG. 11D shows MCP-1, VEGF, and IL-18 peri-implant concentrations of different implant types at 8 weeks.

[0039] FIG. 11E shows MCP-1, VEGF, and IL-18 peri-implant concentrations of different implant types at 12 weeks.

[0040] FIG. 11F shows MCP-1, VEGF, and IL-18 peri-implant concentrations of different implant types at various timepoints (FIGS. 11A -11E) and its evolution over time for ZIPH-3c.

[0041] FIG.12A is a schematic illustration of multilayer functionalization method for the introduction of methacrylate groups for stable hydrogel adhesion.

[0042] FIG.12B shows an example fabrication process for integrating ZIPH-3c into a device.

[0043] FIG.12C shows example electrode placement and electrical connection on mice for ECG measurements.

[0044] FIG.12D shows an example ECG trace recorded with ZIPH-3c electrodes 12-weeks post-implantation.

[0045] FIG. 12E shows the normalized voltage retention of the R-wave peak voltage at different timepoints. Voltages were normalized to day 0 R-wave peak voltage of the same electrode pair. DETAILED DESCRIPTION

[0046] Introduction

[0047] Implantable medical devices (IMDs), including cardiac pacemakers, cochlear implants, and continuous glucose monitors, are playing an increasingly important role due to an aging population and the associated increasing prevalence of chronic diseases. However, one major obstacle that limits IMDs from realizing their full potential is the foreign body response (FBR). The FBR is characterized by inflammatory and fibrotic processes that surround any implant that enters the body, which can lead to degradation, mechanical failure, and isolation via fibrotic encapsulation of the devices. Fibrotic encapsulation is particularly detrimental for biosensors and electrophysiological devices because it impedes the transport of analytes and ions.Attorney Docket No.507814.5000508

[0048] The FBR is initiated by the adsorption of proteins on the implant surface, a process commonly known as fouling. This triggers an inflammatory response that leads to the recruitment of macrophages, which become a continuous, long-term source of degradative enzymes and chemicals such as reactive oxygen species. In addition, macrophages and other immune cells recruit fibroblasts that facilitate the formation of fibrotic tissue. Even for IMDs that have been approved by the Food and Drug Administration (FDA) and in clinical use, e.g., pacemakers, the FBR is commonly still an issue that limits the longevity of the device function and causes additional side effects. A number of approaches have been and are being pursued to reduce the FBR, yielding FBR-suppressing, i.e., immunocompatible properties.

[0049] So far, common strategies to suppress the FBR to IMDs can be broadly divided into two types: (i) permanent surface modification; and (ii) controlled release of soluble immunomodulatory drugs. The latter has limitations because the FBR resumes once the drug reservoir is depleted. This leaves surface modification as a better option for a long-term solution. Among the surface modification and coating strategies, zwitterionic polymers / hydrogels are a promising class of materials due to their antifouling properties arising from their superhydrophilicity. Previous work has shown that zwitterionic hydrogel implants can nearly eliminate the FBR for at least three months in mice, due to their antifouling properties. However, electronically inactive zwitterionic coatings may suffer from two major drawbacks: the stability of such surface coatings during long-term implantation and the negative impacts on the transport of analytes and ions. For this reason, having electronic materials (i.e., conductors and semiconductors) directly exposed on device surfaces is more favorable for such applications.

[0050] As electronic materials for IMDs, conjugated polymers are promising candidates due to their low mechanical moduli, broad chemical design space, and lower interfacial impedance. Specifically, poly(3,4-ethylenedioxythiophene): poly(styrene sulfonate) (PEDOT:PSS) has attracted the most attention because it is chemically stable, highly conductive, highly processable, and non-cytotoxic. PEDOT:PSS has demonstrated its usefulness in applications ranging from biosensors to electrophysiological recording and stimulation.Attorney Docket No.507814.5000508

[0051] Although PEDOT:PSS has been shown to be biocompatible in terms of cytotoxicity and cell proliferation, there is still a lack of understanding of the FBR against PEDOT:PSS and PEDOT derivatives. For characterizing the FBR against PEDOT:PSS, most of the previous studies primarily used immunofluorescence (IF) to measure the infiltration of immune cells (e.g., macrophages). However, this approach has major limitations in giving complete and accurate information about the FBR for the following reasons. First, because fibrosis and collagen deposition are the primary outcome of the FBR, Masson’s trichrome staining (MTS) has been used as a more direct characterization of the FBR level, but has rarely been utilized in PEDOT:PSS implantation studies. Second, imaging with IF suffers from intrinsic limitations of artifacts, such as autofluorescence and non-specific binding of antibodies, which could lead to questionable conclusions. In addition, IF only can detect a limited number of markers, leading to an incomplete picture of the FBR. Third, using IF to compare the number of infiltrating immune cells without information about specific phenotypes (e.g., M1 and M2 for macrophage, etc.) does not necessarily reflect FBR severity. For example, it has been shown that tissue-resident macrophages, rather than recruited macrophages, are responsible for FBR-related fibrosis. Due to the above reasons, the conclusion of PEDOT:PSS being immunocompatible in some of the previous works needs reexamination.

[0052] In fact, negatively charged polymers have been reported to lead to poor FBR outcomes. As such, it is reasonable to speculate the negatively charged PSS in PEDOT:PSS to follow this observation. By combining multiple immunological assays, including histology, immunofluorescence, flow cytometry, cytokine analysis, and transcriptomics, PEDOT:PSS has been found to be associated with fibrosis.

[0053] In this disclosure, example design and synthesis of a zwitterionic hydrogel- based double-network for PEDOT:PSS is described. No studies for implantable biomaterials have explored double-network morphologies as a design strategy for suppressing FBR-mediated fibrosis. As such, the design strategy presented in this disclosure could provide unique insights for other types of biomaterials to achieve suppressed FBR.

[0054] The zwitterionic PEDOT:PSS double-network hydrogel (abbreviated as ZIPH) described herein is a mixed ion / electron conducting material that leads toAttorney Docket No.507814.5000508 suppressed foreign body responses against PEDOT:PSS when implanted inside a living body. By suppressing the foreign body response, electrical isolation of implantable devices can be inhibited. Hence, this double network hydrogel enables electrophysiological devices that can be implanted inside a living body for an extended period of time without deterioration of device performance. Combined with its mixed ion / electron conducting properties, substantial improvements in long-term performances of devices, such as brain-machine interfaces, deep brain stimulators, continuous biosensors, cochlear implants, and cardiac pacemakers, are anticipated.

[0055] In some implementations, electrically conducting materials were incapable of suppressing the fibrotic response, in some cases exacerbating it. The double- network hydrogel (ZIPH) not only suppresses the fibrotic response but also exhibits an electrical conductivity value that is one of the highest ever reported among similar material types. The double-network hydrogel has a unique combination of fibrosis- suppressing properties and excellent electrical properties.

[0056] Design of Immunocompatible Conductive Hydrogels

[0057] FIG. 1 shows a schematic illustration of a mixed ion / electron conducting material 100. The mixed ion / electron conducting material 100 comprises a zwitterionic PEDOT:PSS double-network hydrogel (ZIPH) 200 including a first network comprises interconnected fibers of poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) 300 and a second network comprises a zwitterionic polymer hydrogel 400. The interconnected fibers of PEDOT:PSS 300 may comprise aggregates of nanofibers having rod-shaped structures with diameters in the nanoscale, typically between about 1 nm to about 50 nm in size. The nanofibers may have lengths in excess of 100 nm. The PEDOT:PSS fibers 300 network may be embedded in a network of zwitterionic polymer hydrogel matrix 400, such that the first network is formed in situ in the second network to form the double network hydrogel ZIPH 200.

[0058] For combining PEDOT:PSS 300 and zwitterionic polymers 400 for implantable electronics, it is found that there are three major criteria that need to be satisfied. First, the double-network morphology should have the zwitterionic matrix dominate the overall immunocompatibility. For this, nanoscale phase separation is preferred so that PEDOT:PSS phases are smaller than the length scales (e.g., tens to hundreds of nm) that can be recognized by cells. Second, to achieve high conductivity,Attorney Docket No.507814.5000508 the PEDOT:PSS domain needs to be interconnected to facilitate charge transport. Third, to keep stable electrical performance in implantable environments, the design must not have severe swelling in water.

[0059] Based on the aforementioned criteria, the zwitterionic polymer in the zwitterionic hydrogel 400 may be poly(sulfobetaine methacrylate) (PSB). Sulfobetaine- based small molecules have been found to increase the conductivity of PEDOT:PSS films. Furthermore, due to the strong intra- and inter-chain electrostatic attraction, PSB hydrogels swell the least compared to other zwitterionic hydrogels.

[0060] In some examples, the ZIPH hydrogel may be in the form of a thin film or a bulk film. The thickness of the thin film may be in the range from about 100 nm to about 2000 m. The thickness of the bulk film may be in the range from about 100 m to about 3 mm.

[0061] A method for manufacturing a mixed ion / electron conducting material 100, which includes the above zwitterionic PEDOT:PSS hydrogel (ZIPH) 200 is described below with reference to FIG.2. Referring to FIG.2, the method of forming the ZIPH double network hydrogel 200 includes dissolving 202 a zwitterionic monomer and a crosslinking agent in an aqueous dispersion containing particles comprising poly(3,4- ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) to form a mixture. Then polymerizing 204 the mixture via thermal or UV polymerization to form the zwitterionic PEDOT:PSS hydrogel (ZIPH) 200, wherein the mixture includes an initiator. The particles of PEDOT:PSS may comprise nanoparticles having a diameter ranging from about 5 nm to about 15 nm in size. The nanoparticles may form aggregates.

[0062] In some examples, an additive may be added to the aqueous dispersion of PEDOT:PSS to facilitate the formation of rod-like structures for the PEDOT phase, through destabilization of the micellar structure of PETOD:PSS. The additive may be 4-ethylbenzenesulfonic acid (EBSA).

[0063] The zwitterionic monomer described herein may be selected from [2- (methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide (SBMA).

[0064] The crosslinking agent described herein may be selected from N, N’- methylenebisacrylamide (MBAA).Attorney Docket No.507814.5000508

[0065] The initiator described herein may be selected from ammonium persulfate and 2-hydroxy-2-methylpropiophenone (APS). In an example of preparing the thermally initiated hydrogel, N,N,N’,N’- tetramethylethylenediamine (TEMED) may be further added to the mixture. The time and temperature to achieve thermal polymerization may range from about 60 to about 900 °C for a period of 30 mins to 2 hours. In one example, the final mixture was placed under vacuum for 15 minutes, injected into a sandwich sheet mold, and placed in an 80 °C oven for 1 hour.

[0066] For the preparation of UV initiated hydrogels, 2-hydroxy-2- methylpropiophenone (HMPP) and 4-ethylbenzenesulfonic acid (EBSA) may be further added to the mixture. The final mixture may be exposed to UV irradiation from seconds to several minutes. In one example, the final mixture was sparged with nitrogen for 10 minutes, spin coated and was exposed to UV irradiation (365 nm wavelength, 4 mW / cm2) for about 5 minutes.

[0067] Morphology and Conductivity Evolution of ZIPH During Solvent Annealing

[0068] Following the polymerization step 204, the ZIPH hydrogel may be subjecting 206 to at least one solvent annealing treatment cycle whereby the particles transition to an interconnected fiber network. The solvent used in the solvent annealing treatment 206 may be selected from ethanol and 2,2,2-trifluoroethanol (TFE). An example solvent annealing method involves alternatively exposing ZIPH to ethanol and 2,2,2-trifluoroethanol (TFE) with a certain number of cycles (e.g., designated as ZIPH-nc, where nc denotes n cycles with n may be any number of cycles from 1-3) to further promote an aggregation and interconnectivity of the PEDOT phase in the ZIPH. In some examples, the ZIPH may be subjected to at least 2 to 3 solvent annealing treatment cycles. In other examples, the ZIPH may be subjected to 3 solvent annealing treatment cycles.

[0069] Because TFE is a good solvent for uncrosslinked PSB while having poor chemical compatibility with PEDOT:PSS, TFE could penetrate into the PSB hydrogel matrix to further destabilize the PEDOT:PSS phase, thereby promoting the formation of larger PEDOT:PSS aggregates in the PSB hydrogel matrix. In one example, before each TFE treatment, ZIPH hydrogel is treated with ethanol in each cycle to deswell the hydrogel, thereby preemptively counteracting the excessive swelling of ZIPH byAttorney Docket No.507814.5000508 TFE. In addition, the deswelling by ethanol causes the PEDOT:PSS phases to be in closer proximity to each other, which increases the likelihood of adjacent PEDOT:PSS phases to connect. FIG. 3A shows a schematic illustration of the morphological changes that occur to ZIPH after being treated with ethanol-TFE annealing cycles.

[0070] The effect of the solvent annealing process in facilitating the formation of PEDOT nanofibrous morphology is validated by cryogenic transmission electron microscopy (cryo-TEM) on ZIPH samples that went through different solvent annealing processes. FIG. 3B shows cryo-TEM images of ZIPH subjected to various solvent annealing conditions. With each ethanol-TFE cycle, the PEDOT:PSS particles transition to a more interconnected fibrillar morphology. For ZIPH blended with the additive EBSA for promoting aggregation but without any solvent annealing steps (referred to as ZIPH-E), dispersed PEDOT:PSS nanoparticles of 5-15 nm in diameter were observed. This shows that the favorable chemical compatibility between PSB and PEDOT:PSS does not provide enough driving force for PEDOT:PSS to grow nanoparticles into larger aggregates. After the solvent treatments in only ethanol or TFE, or 1 cycle of ethanol-TFE, there is minimal change in the nanoparticle morphology for PEDOT:PSS. After 2 to 3 solvent annealing cycles, the interconnected nanofibers started to emerge. The formation of the PEDOT:PSS fibrous structures from the solvent treatment processes can also be observed using conductive atomic force microscopy (C-AFM), as shown by the comparison of ZIPH-E and ZIPH-3c. FIG. 3C shows C-AFM images of ZIPH before (ZIPH-E) and after 3 cycles of ethanol-TFE annealing cycles (referred to as ZIPH-3c).

[0071] Grazing-incidence X-ray diffraction (GIXD) reveals that in contrast to the crystalline structure of pristine PEDOT:PSS, ZIPH suppresses the long-range crystallization of PEDOT. FIG. 3D shows GIXD diffraction patterns of PSB hydrogel and ZIPH subjected to various solvent annealing conditions. The solvent treatment processes have little effect in changing such amorphous structure. The diffraction pattern was nearly identical to that of PSB hydrogel.

[0072] As a result of the solvent annealing-induced formation of the interconnected nanofibrous morphology of PEDOT:PSS, the conductivity of ZIPH improved substantially from ~10-4S / cm for ZIPH-E to ~2.5 S / cm for ZIPH-2c (after 2 cycles of ethanol-TFE annealing treatment) and ZIPH-3c (after 3 cycles of ethanol-TEFAttorney Docket No.507814.5000508 annealing treatment) as shown in FIG. 3E. For TFE-only or 1 cycle of ethanol-TFE annealing, even though little changes were observed in the morphology, increases in conductivity have already started to take effect. The solvent treatment process also led to the increase in Young’s modulus, from 1.4 kPa for ZIPH-E to 9.9 kPa for ZIPH- 3c as shown in FIG.3F. This is also due to the formation of an interconnected network by the more rigid phase of PEDOT:PSS.

[0073] ZIPH-3c is Associated with the Lowest Collagen Density

[0074] The FBR behaviors against ZIPH in comparison with a single-network, unmodified PEDOT:PSS hydrogel was studied. The zwitterionic PEDOT:PSS double- network hydrogel (ZIPH) may significantly suppress the FBR, in addition to improving conductivity when compared to unmodified PEDOT:PSS hydrogels. For example, FIG. 4A shows representative MTS-stained tissue sections of unmodified PEDOT:PSS. The collagen layer is marked with arrows, and the locations of implants appear clear and indicated by asterisks. FIG.4B shows representative MTS-stained tissue sections of ZIPH-3c (after 3 cycles of ethanol-TEF annealing treatment), which shows a much less dense collagen layer. When compared to unmodified PEDOT:PSS hydrogels, ZIPH exhibits a reduction of FBR-associated fibrotic severity by 64% (as measured by the collagen density) as shown in FIG. 5, in addition to an increase in electrical conductivity by more than one order of magnitude (FIG. 5). Surprisingly, the FBR associated fibrotic severity is even markedly lower than that of the parent zwitterionic hydrogel, PSB, by 53%. (See, Example Section discussed below).

[0075] An implantable medical device for implanting inside a living body which includes a mixed ion / electron conducting material comprising the zwitterionic PEDOT:PSS hydrogel (ZIPH) described above is also provided. The implantable medical device may include cardiac pacemakers, implantable cardioverter defibrillators, cochlear implants, and continuous glucose monitors.

[0076] The examples below illustrate the disclosed zwitterionic PEDOT:PSS double-network hydrogel (ZIPH) and the manufacturing methods, but are not intended to limit the scope of any claims thereto. The ZIPH is evaluated according to a variety of experimental tests as described below in order to characterize its fibrosis- suppressing properties and electrical properties.

[0077] ExamplesAttorney Docket No.507814.5000508

[0078] Systematic immunological experiments combining histology, immunofluorescence, flow cytometry, RNA profiling, cytokine assays, and in vitro protein adsorption measurement are performed to provide deeper insights. First, the poor immunocompatibility of PEDOT:PSS was found to be mainly attributed to the high protein adsorptive properties of negatively charged PSS. Second, the higher immunocompatibility of ZIPH-3c compared to PSB may arise from four possible underlying reasons: minimal distortion of the macrophage phenotype ratio, inhibition of complement activation by PSS, suppressed fibrotic gene expression levels, and upregulation of TBX21 (a master regulator of type 1 T helper cells). These effects could come from the interplay between a unique combination of polymers and the resulting morphological structure. Overall, the chemical heterogeneity, which has been rarely explored, could be another design parameter for enabling facile “mix-and- match” approaches for imparting immunocompatibility to functional materials without changing the chemical structure. Also, for deepening the immunological understanding of the FBR mechanism, creating model systems with controlled chemical heterogeneity could be a useful tool. For the future uses of PEDOT:PSS in electrically biointerfaced devices, this disclosure provides the first thoroughly tested solution for addressing the most challenging problem of the foreign-body response.

[0079] Methods and Materials

[0080] Materials: PEDOT:PSS (PH1000) was purchased from Clevios and was filtered through a 0.7 m glass fiber filter (Tish Scientific) before use. The following chemicals were purchased from Sigma-Aldrich: [2-(Methacryloyloxy)ethyl]dimethyl-(3- sulfopropyl)ammonium hydroxide (SBMA, 95%), N,N’-Methylenebisacrylamide (MBA, 98%), ammonium persulfate (APS, 98%), N,N,N’,N’- tetramethylethylenediamine (TEMED, 99%), 2-hydroxy-2-methylpropiophenone (HMPP, 97%), 4- ethylbenzenesulfonic acid (EBSA, 95%), sodium styrene sulfonate (NaSS, Sigma- Aldrich, 90%), divinylbenzene (DVB, 80%), 4,4’-azobis(4-cyanovaleric acid) (ACVA, 98%), (vinylbenzyl)trimethyl ammonium chloride (VBTMAC, 99%), sodium dodecylbenzene sulfonate (SDS, 98.5%), and 4-cyano-4-(phenylcarbonothioylthio) pentanoic acid (CPhPA). 2,2,2-trifluoroethanol (TFE, 99%) was purchased from Sigma-Aldrich and Tokyo Chemical Industries. 2,2’-azobis[2-(2-imidazolin-2- yl)propane]dihydrochloride (VA-044) was purchased from Wako Chemicals (USA).Attorney Docket No.507814.5000508 The acetate buffer was prepared with 0.1 M acetic acid (glacial, Sigma-Aldrich, 99.85%), 0.1 M sodium acetate trihydrate (Sigma-Aldrich, >99%), and 0.5 M NaCl at pH 5.2.

[0081] Preparation of hydrogels

[0082] For preparing ZIPH-E hydrogels, 250 mg of [2-(Methacryloyloxy)ethyl] dimethyl-(3-sulfopropyl)ammonium hydroxide (SBMA, Sigma-Aldrich, 95%) and 1.4 mg (1 mol%) of N,N’-Methylenebisacrylamide (MBAA, Sigma-Aldrich, 98%) were dissolved in 1.0 mL of PEDOT:PSS dispersion. Then, for the preparation of thermally initiated hydrogels, 10.21 L (0.1 mol%) of APS (20 mg / mL) and 10.41 L (0.2 mol%) of TEMED (20 mg / mL) aqueous solutions were added dropwise to the PEDOT:PSS mixture under continuous agitation. For the preparation of UV initiated hydrogels, 7.5 L (0.1 mol%) of HMPP (20 mg / mL) aqueous solution was added. This was followed by the dropwise addition of 62.5 L EBSA aqueous solution (20 wt%) while the mixture is continuously agitated and vortexed for an additional minute. For thermal polymerization, the final mixture was placed under vacuum for 15 minutes, injected into a sandwich sheet mold, and placed in an 80 °C oven for 1 hour. For UV polymerization, the final mixture was sparged with nitrogen for 10 minutes, spin coated and exposed to UV irradiation (365 nm wavelength, 4 mW / cm2) for 5 minutes.

[0083] To prepare ZIPH-u hydrogels, the above procedures were carried out without the addition of EBSA. To prepare annealing cycle treated bulk ZIPH hydrogels, ZIPH-E hydrogels were briefly rinsed in ethanol. Subsequently, the hydrogels were immersed in a volume of ethanol three times the volume of the ZIPH sample. Thereafter, the ethanol was discarded, and the hydrogels were immersed in a volume of TFE three times the volume of the ZIPH sample. The TFE was discarded, and the hydrogels were briefly rinsed in water. The whole annealing process was repeated two more times to obtain ZIPH-3c. To prepare annealing cycle treated thin film ZIPH hydrogels, the films were immersed in ethanol for 1 minute and immersed in TFE for 1 minute.

[0084] To prepare PSB hydrogels, 1.5 g of SBMA (55 wt%), 4.2 mg of MBAA (0.5 mol%), and 0.882 mg of HMPP (0.1 mol%) were dissolved in PBS. The solution was vortexed for 1 minute and sparged with nitrogen for 10 minutes. The sparged solutionAttorney Docket No.507814.5000508 was injected into a sandwich sheet mold and irradiated with UV light (365 nm wavelength, 4 mW / cm2) for 30 minutes per side.

[0085] To prepare PSS hydrogels, 700 mg of NaSS (28 wt%) was dissolved in 1694 L of DMSO by stirring overnight. To the NaSS solution, 18.09 L of DVB (4 mol%), 43.4 L (0.1 mol%) of APS DMSO solution (20 mg / mL), and 44.2 L (0.2 mol%) of TEMED DMSO solution were added and vortexed for 1 minute. Thereafter, the solution was placed under vacuum for 15 minutes, injected into a sandwich sheet mold, and placed in an 80 °C oven for 1 hour. The resulting PSS gel was rinsed twice in PBS (phosphate buffered saline).

[0086] For the preparation of PEDOT:PSS bulk hydrogels, 16 v / v % of DMSO was added to PEDOT:PSS. Then, 10 mg / mL of 1-ethyl-3-methyllimidazolium trifluoromethaneulfonate (EMIM OTf) was added dropwise while the solution was continuously agitated. The resultant mixture was vortexed for 1 minute, placed under vacuum for 15 minutes, and injected into a sandwich sheet mold with silicone lining on top. The mold was placed on a hotplate, heated to 90 °C, for 90 minutes with the silicone lining side up. Then, the top piece with the silicone lining was removed, and the PEDOT:PSS gel was exposed to air. The PEDOT:PSS gel was dried at 40 °C for 16 hours, followed by another round of drying at 50 °C for 16 hours, at which point the thickness and lateral dimensions of the gel was approximately a third and a half of the original dimensions, respectively. The resulting PEDOT:PSS gel was rinsed twice in PBS.

[0087] All gels were washed in PBS for three days with changes of fresh PBS every 12 hours. All hydrogels used in in vivo experiments were autoclaved at 121 °C for 30 minutes.

[0088] Synthesis of polymers

[0089] Poly(styrene sulfonate) (PSS) was synthesized by dissolving 150 g of NaSS (727 mmol, 150 g) and 70 mg of ACVA (0.25 mmol, 70 mg) in 1.5 L of deionized(DI)H2O. The solution was sparged with argon gas at a high flow rate for 45 minutesbefore beingheated to 65 °C for 12 hours. A portion of the resulting solution was dialyzed using 10k dialysis tubes from (Thermo Scientific, SnakeSkin, 3.5k molecular weight cut-off) for 5 days against an excess of pure DI water which was changedAttorney Docket No.507814.5000508 every 12 hours. After purification, the solution containing polymer was collected, frozen with liquid nitrogen and lyophilized to give PSS as a dry white powder.

[0090] Poly((vinylbenzyl)trimethyl ammonium chloride) (PVBTMA) was synthesized by dissolving VBTMAC (945 mmol, 200 g) and ACVA (0.33 mmol, 93 mg) in 1.5 L of DI H2O. The solution was sparged with argon at a high flow rate for 45 minutes before being heated to 65 °C for 12 hours. The reaction solution was purified the same way as PSS, resulting in pure PVBTMA as a white powder. Poly(sulfobetaine methacrylate)-block-poly((vinybenzyl)trimethylammonium chloride) (PSB100-b- PVBT20) was synthesized according to the synthesis route depicted below.First, poly(sulfobetaine methacrylate) (PSB100) was synthesized by dissolving 20.0 g of SBMA (71.6 mmol, 20.0 g), VA-044 (0.0500 mmol, 16.2 mg), and 139.88 mg of CPhPA (0.5000 mmol, 139.9 mg) in 100 ml of 4:1 (v / v) acetate buffer / ethanol cosolvent. The solution was sparged with nitrogen gas for 30 minutes, heated to 50 °C for 6.5 hours, and quenched in liquid nitrogen. The polymer was precipitated in 20Attorney Docket No.507814.5000508 times the volume of methanol and redissolved in 0.5 M NaCl a total of three times. The polymer was dried under vacuum. Conversion: 70%.

[0091] To synthesize PSB100-b-PVBT20, PSB100 (5.00 g), VBTMAC (4.46 mmol, 945.2 mg), and 5.78 mg of VA-044 (0.0179 mmol, 5.78 mg) were dissolved in 11.3 mL of 4:1 (v / v) acetate buffer / ethanol cosolvent. The solution was sparged with nitrogen gas for 20 minutes and heated to 50 °C for 18 hours. The polymer was precipitated in 20 times the volume of methanol and redissolved in 30 mL of 7:1 (v / v) acetate buffer / ethanol cosolvent. The polymer was re- precipitated in 10 times the volume of methanol, and the supernatant was centrifuged at 4000 g's for 30 minutes. The precipitates were collected and washed with excess acetone. The polymer was lyophilized for 24 hours.

[0092] Zwitterionic self-assembled monolayer

[0093] For assembling zwitterionic polymer brush on hydrogels, hydrogel discs were transferred to 48-well plates. 150 µL of PVBTMA solution (30 mg / mL in PBS) was added, and the hydrogel in solution was shaken for 1 hour. The PVBTMA solution was discarded, and the hydrogels were washed with 1 mL of PBS three times. The previous steps were repeated except PSS solution (30 mg / mL in PBS) was used instead of PVBTMA solution. Then, 150 µL of PSB100-b-PVBT20 solution (13 wt% in PBS with 0.5 M total NaCl) was added and shaken for 1 hour. The solution was discarded, and the hydrogels were washed three times in PBS with 0.5 M total NaCl. Then, the hydrogels were washed once with regular PBS and soaked in PBS overnight.

[0094] Measurement of electrical properties

[0095] Conductivity measurements were conducted using a standard four-point geometry. Bulk hydrogels were prepared into discs (10 mm in diameter and 1 mm in thickness) using a biopsy punch, and solvent treatments were applied to the discs. Any changes in physical dimensions that resulted from the solvent treatments were accounted for when determining the conductivity values. Probes were put in direct contact with the hydrogels.

[0096] Measurement of electrochemical properties

[0097] The electrochemical measurements were made using the PalmSens electrochemical workstation. Hydrogels were prepared into discs (10 mm in diameterAttorney Docket No.507814.5000508 and 1 mm in thickness), pre-equilibrated in PBS, and inserted into a Swagelok cell. Electrochemical impedance spectroscopy (EIS) spectra were obtained over a range of 100 kHz to 0.1 Hz with an AC 20 mV sine wave and with no DC offset. The analysis of the EIS data was carried out using the Multitrace 4.5 software.

[0098] Measurement of mechanical properties

[0099] The tensile properties of hydrogels were tested on a universal testing machine (zwickiLine Z0.5, Zwick Roell Group, Ulm, Germany) with 500-N load cell at room temperature. Hydrogels were cut into JIS K6251-8 industrial standard dumbbells. For ZIPH-3c, ZIPH-u, PSB, and PSS hydrogels, the crosshead speed was set to 30 mm / s. For PEDOT:PSS hydrogel, the crosshead speed was set to 5 mm / s.

[0100] Cryo-TEM characterization

[0101] UV polymerized ZIPH films were prepared following a modified version of the procedures described earlier. The precursor solutions were spin-coated at 2000 RPM for 10 seconds on oxygen plasma-treated copper Quantifoil grids (200 mesh, R1.2 / 1.3; Electron Microscopy Sciences) supported on PDMS. A portion of the gel was scraped away to expose the mesh pattern of the grid. The hydrogel films were kept in a humidified chamber until further processing. The gels were then back blotted (force = 1, 20s) and plunge frozen into liquid ethane on a Thermo Scientific Vitrobot Mark IV. The Aquilos 2 cryo-FIB / SEM (Thermo Scientific) was used to prepare the lamella. Using the in-column sputter coater the grid was coated with platinum at 1kV and 10 mA current for 15 seconds. The grid was SEM imaged and then atlased using Maps 3.26 (Thermo Scientific) at 10kV accelerating voltage and 13 pA current to find suitable areas for milling. Prior to milling the grids were coated with a layer of organometallic platinum using a gas injection system for 10 sec. The milling process was automated into 4 steps (all at 30kV accelerating voltage): Rough Milling, Medium Milling, Fine Milling, and Thinning using 500, 300, 100, 30 pA current, respectively. The final lamella thickness was targeted between 100 -150 nm. Lamella were imaged on the Thermo Scientific Titan Krios 3Gi at 300kV at cryogenic temperature equipped with a Gatan K3 direct electron detector and BioContinuum in counting CDS mode.

[0102] Conductive AFM

[0103] The conductive AFM (C-AFM) measurements were performed by using the C-AFM mode of Cypher ES Environmental AFM with ASTELEC-01 probe and the dualAttorney Docket No.507814.5000508 gain C- AFM probe holder (gain= 1µA / 1nA / V). UV polymerized ZIPH-E and ZIPH- 3c films were prepared following the procedures described earlier. The precursor solutions were spin-coated at 6000 RPM for 10 seconds on oxygen plasma-treatedAu films on SiO2substrates and dried on ahotplate. Magnets were placed separatelyon Au and sample film and were sealed with Ag paste to realize ideal contact. Before measurement, the sample bias was zeroed by applying an offset voltage from the AFM hardware. During the measurement, 1 mV bias was applied across samples and the current was converted to a voltage and then amplified and recorded by the AFM hardware. Scans were acquired using scan rates of 2 µm / s.

[0104] Small-angle X-ray scattering

[0105] Small-angle X-ray scattering (SAXS) was performed at the Advanced Photon Source at Argonne National Laboratory on beamline 12-ID-B. Pristine PEDOT:PSS and PEDOT:PSS with 10 mg / mL of EBSA were loaded into capillary tubes (Charles Supper) and sealed. The sample-to-detector distance was 2 m, corresponding to a q-range of 0.003-0.5 Å-1.

[0106] Grazing-incidence X-ray diffraction

[0107] Grazing-incidence X-ray diffraction (GIXD) was performed at the Advanced Photon Source at Argonne National Laboratory on beamline 8-ID-E. UV-initiated ZIPH samples were spin-coated at 1500 RPM for 10 seconds. Thermally initiated ZIPH samples were prepared by sandwiching 10 µL of the precursor solution between apiece ofsilicon (O2plasma treated) and polyethylene terephthalate (PET) laminatedcoverslip (10 mm x 17 mm) and heating to 80 °Cfor 15 minutes. PSB hydrogelsamples were prepared similarly by sandwiching 3 µL of the precursor solution between a piece of silicon (O2 plasma treated) and PET laminated coverslip (10 mm x 17 mm) and exposing to UV light for 10 minutes. The coverslip was gently removed while submerged in water. All samples were dried on a hotplate after polymerization of solvent annealing.

[0108] Scanning electron microscopy - energy dispersive X-ray spectroscopy

[0109] Scanning electron microscopy - energy dispersive X-ray spectroscopy (SEM-EDX) was conducted using the TESCAN LYRA3 field-emission SEM with Oxford X-Max detectors for EDX. UV-initiated ZIPH samples were spin-coated at 2000Attorney Docket No.507814.5000508 RPM for 10 seconds. An accelerating voltage of 4 kV was used to locate and analyze the samples.

[0110] Confocal Raman spectroscopy

[0111] Raman microscopy was conducted using the HORIBA LabRAM HR Evolution Confocal Raman Microscope. Thermally initiated ZIPH and PEDOT:PSS hydrogels were sealed between a glass slide and coverslip to prevent dehydration during measurement. An objective magnification of 10x was used to focus the 532 nm laser beam on the sample. The scattered light was analyzed using a grating of 600 lines / mm.

[0112] In vitro plasma adsorption test

[0113] Human plasma adsorption on the hydrogels was evaluated using the 3-(4- carboxybenzoyl)quinoline-2-carboxaldehyde (CBQCA) assay and bicinchoninic acid (BCA) assay. Hydrogel samples (5 mm in diameter and 1 mm in thickness) were incubated in 150 µL of human plasma (Innovative Research, Inc.) in 48-well plates at 37 °C for 2 hours. The hydrogels were washed with 1 mL of PBS 5 times and transferred to 96-well U-bottom plates.150 µL of 1% SDS-PBS solution was added and shaken for 1 hours.75 µL of the protein-SDS solution was transferred to 96-well flat bottom plates and diluted with 60 µL of PBS. Then, CBQCA assay was directly carried out to determine the relative amount of proteins adsorbed on the hydrogels. The fluorescence intensities at 550 nm (excitation / emission of 465 / 550 nm) were recorded using the TECAN Inifinite M200 Pro plate reader. For the BCA assay, 30 µL of the protein-SDS solution was transferred to 96-well flat-bottom plates and standard BCA procedures were applied. The absorbance at 562 nm was recorded using the TECAN Inifinite M200 Pro plate reader.

[0114] Cytotoxicity assay

[0115] NIH-3T3 fibroblasts were cultured in Dulbecco’s modified eagle medium (DMEM, from Gibco) supplemented with 10% fetal bovine serum (FBS, Gibco) and 1% penicillin / streptomycin (Gibco). Cytotoxicity of the hydrogels was determined by directly culturing fibroblasts on the hydrogels and using the LIVE / DEAD Viability Cytotoxicity Kit (Invitrogen). Hydrogel discs (5 mm in diameter and 1 mm in thickness) in 96-well flat bottom plates were pre-conditioned in 200 µL DMEM, and cells were directly seeded at a density of 1300 cells per well. After 96 hours, 5 µL of SDS solutionAttorney Docket No.507814.5000508 (2 wt%) per 100 µL media were added to half of the wells and incubated at room temperature for 10 minutes to establish dead controls. 100 µL of working solution (4 µM calcein AM and 4 µM ethidium homodimer) was added to the experimental wells, and 100 µL of calcein AM solution (4 µM) or ethidium homodimer solution (4 µM) was added to the dead control wells. The cells were incubated at room temperature for 45 minutes and immediately evaluated (LIVE: 494 / 517 nm, DEAD: 528 / 617 nm) using the TECAN Infinite M200 Pro plate reader.

[0116] Implantation surgeries

[0117] 6- to 8-week-old male C57BL / 6 mice were purchased from Jackson Laboratories. Preoperatively, all mice received subcutaneous injections of 5 mg / kg of Meloxicam for pre-surgery analgesia. All mice were anesthetized using 3% isoflurane in oxygen, and the backs of all mice were shaved. All mice received eye lubricants to prevent dehydration-induced blindness. Then, the shaved backs of all mice were sterilized by scrubbing with betadine and alcohol pads in an alternating fashion, repeated two times. Two transverse incisions, 1 cm in length, were made in the upper and lower backs. The blade tips of a pair of scissors were inserted into the subcutaneous space through the incisions, and two subcutaneous pockets, flanking each incision, were made via blunt tissue dissection using the blunt sides of a pair of scissors. Two implants (5 mm in diameter and 1 mm in thickness) were inserted in the subcutaneous pockets on both sides of each incision, towards the limbs of the mice, for a total of four implants per mouse. For sham control, subcutaneous pockets were made but no implants were placed. The incisions were closed via wound clips. Following surgery, all mice were given subcutaneous injections of 5 mg / kg of Meloxicam once every 24 hours for two days. Wound clips were removed after 10 days.

[0118] Histological processing for hematoxylin and eosin (H&E) and Masson's trichrome (MTS) staining

[0119] After 1-, 4-, or 12-weekspost-implantation, all mice were euthanized by CO2 asphyxiation, followed by cervical dislocation. Afterward, the backs of all the mice were shaved. The implants and the surrounding tissues were explanted and fixed in 10% neutral buffered formalin for 24 hours. After fixation, the implants and tissues were transferred to 70% ethanol. The materials were then paraffin-embedded, sectioned (5 µm), and stained for H&E using standard procedures by the HumanAttorney Docket No.507814.5000508 Tissue Resource Center at the University of Chicago. Masson's trichrome staining (MTS) was carried out in-house using standard procedures. Whole-section scans of H&E and MTS-stained tissue sections were conducted using the Olympus VS200 Slideview Slide Scanner.

[0120] Immunofluorescence

[0121] Materials were fixed, embedded, and sectioned as described above. Tissue sections were de-paraffinized in xylene and rehydrated using an ethanol-to-water gradient. Antigen retrieval was conducted in citrate buffer (pH 5) for 45 minutes at 55 °C. Samples were then rinsed in water, permeabilized by incubating in 10% DMSO in PBS for 5 minutes, washed in phosphate buffered saline with Tween 20 or PBS-tween 20 (PBS-T), and rinsed in water again. Thereafter, the samples were blocked for 1 hour using a 1% bovine serum albumin (BSA) solution. The blocking buffer was removed, and the sections were incubated in either the macrophage panel or fibrosis panel antibodies overnight at 4 °C. The macrophage panel antibodies included: rat anti-mouse F4 / 80 (1:100 dilution, Abcam, Cat. #ab6640), rabbit anti-mouse CD86 (1:200 dilution, Biorbyt, Cat. #orb49101), and goat anti-mouse CD206 (1:100 dilution, R&D Systems, Cat. #AF2535). The fibrosis panel antibodies included: rat anti-mouse F4 / 80 (1:100 dilution, Abcam, Cat. #ab6640), rabbit anti-mouse collagen I (1:100 dilution, Bio-Rad, Cat. #2150-1410), and goat anti-mouse SMA (1:100 dilution, Novus Biologicals, Cat. #NB300-978). The next day, the sections were washed twice in PBS-T and once in PBS. The washed sections were incubated for 2 hours at room temperature in a secondary antibody cocktail, consisting of donkey anti-rat IgG AlexaFluor 647 (1:200 dilution, Southern Biotech, Cat. #OB643031), donkey anti-rabbit IgG AlexaFluor 488 (1:500 dilution, Fisher Scientific, Cat. #A21206), and donkey anti-goat IgG AlexaFluor 555 (1:500 dilution, Fisher Scientific, Cat. #A21432). Afterward, the sections were washed twice with PBS-T and once with PBS. The washed sections were counterstained with 4',6-diamidino-2-phenylindole (DAPI, 1:1000 dilution, Thermo Scientific) for 2 minutes and washed with PBS. Prolong Antifade Gold mounting medium (Invitrogen) was added, and the sections were sealed with coverslips and nail polish. Images were acquired using either Zeiss Axiovert 200m inverted epifluorescence microscope with a Hamamatsu Flash 4.0 camera run by SlideBook 6.0 software (Intelligent Imaging Innovations).Attorney Docket No.507814.5000508

[0122] Cytokine analysis

[0123] Cytokine quantifications were performed using a 14-cytokine LegendPlex custom panel (BioLegend). After 1- or 4-weeks post-implantation, all mice were euthanized by CO2 asphyxiation, followed by cervical dislocation. Afterward, the backs of all the mice were shaved. The explanted implant and associated tissues were frozen immediately on dry ice and stored in a -80 °C freezer for later use. The frozen tissues were thawed at room temperature and minced with scissors. The minced implants and tissues were transferred to lysis matrix tubes (MP Biomedicals) with 300 µL of T-PER solution (Thermo Scientific), consisting of 5 mM EDTA and protease inhibitor (Thermo Scientific). Three cycles of homogenization were conducted on the FastPrep-24 5G homogenizer (MP Biomedicals) with three minutes on ice between cycles. Subsequently, the tubes were centrifuged for 20 minutes at 10,000 g's at 4 °C, after which a fat layer, a protein layer, and a debris layer formed. Avoiding the fat layer, the protein layer was transferred to low-bind tubes (Fisher Scientific) and centrifuged again for 10 minutes at 10,000 g's at 4 °C. Similar to the first centrifugation step, the protein layer was transferred to fresh low- bind tubes. The centrifugation and protein layer transfer step were conducted once more. The protein solutions were stored in a -80 °C freezer for future analysis.

[0124] The protein solutions were thawed at room temperature, and LegendPlex assay was conducted using the un-diluted protein solutions, following manufacturer's protocol on the Agilent Penteon 5-30 flow cytometer. C3a was quantified using a C3a ELISA kit (Novus Biologicals), following manufacturer's instructions. The protein solution was diluted 20-fold in PBS before conducting ELISA. The absorbance was recorded at a wavelength of 450 nm using the TECAN Inifinte M200 Pro plate reader.

[0125] Flow cytometry

[0126] After 1- or 4-weeks post-implantation, all mice were euthanized by CO2 asphyxiation, followed by cervical dislocation. The backs of all the mice were shaved, and subsequently, a depilatory cream was applied for one minute. Afterward, the cream and the hair were washed off with water and 70% ethanol. Approximately 2 cm2of skin, including the implant, was extracted from the back and placed on ice. The implants and associated tissues were finely minced with scissors and digested by adding 200 µL of RPMT-1640 (Sigma), supplemented withLiberase TL (0.5 mg / mL,Attorney Docket No.507814.5000508 Roche), DNase I (0.5 mg / mL, Grade II, Roche), and CaCl2 (2 mM). The samples were shaken for 2 hours at 37 °C. Immediately, 200 µL of ice-cold dissociation buffer (1x PBS, 10 mM EDTA, 2% FBS) were added to halt digestion. The digested tissues were pestled through 70-µm cell strainers (Fisher Scientific) and rinsed with 2 mL of ice-cold dissociation buffer. Then, the washed digested tissues were centrifuged for 8 minutes at 400 g's at 4 °C and resuspended in 200 µL of FACS buffer (eBiosciences). The enriched single-cell suspension was transferred to 96-well U- bottom plates, centrifuged, and resuspended in 200 µL of FACS buffer another time.

[0127] Aliquots of the single-cell suspension were then stained with propidium iodide (1:200 dilution, BD Bioscences) and counted on an Agilent Penteon 5-30 flow cytometer. Subsequently, 106cells were incubated in Zombie Aqua Fixable Viability Dye (1:500 dilution, BioLegend) for 20 minutes on ice followed by washing with FACS buffer. The cells were treated with Fc block (1:200 dilution, TruStain FcX Plus, BioLegend) for 10 minutes on ice followed by washing with FACS buffer. Then, the cells were stained for 20 minutes on ice with the following antibody panel: CD11b Pacific Blue (Cat. #101224), CD45 BrilliantViolet 605 (Cat. #103139), CD86 BrilliantViolet 650 (Cat. #105035), Ly-6C PerCP-Cy5.5 (Cat. #128011), CD31 PE (Cat. #102407), Tie-2 PE (Cat. #124007), Ter-119 PE (Cat. #116207), Ep-CAM PE (Cat. #118205), F4 / 80 PE-Cy7 (Cat. #123113), CD3 PE-Dazzle594 (Cat. #100245), CD140a PE-Cy5 (Cat. #135919), CD206 APC (Cat. #141708), and Ly-6G APC-Cy7 (Cat. #127623). All flow cytometry antibodies were purchased from BioLegend and were used at a 1:200 dilution. The cells were fixed with Cytofix / Cytoperm buffer (BD Bioscences) and run on the Agilent Penteon 5-30 flow cytometer the next day. All analyses were performed in FlowJo Flow Cytometry Analysis Software (Treestar) using the gating strategy. PE was used as the exclusionary channel for CD31, Tie-2, Ter-119, and Ep-CAM, which are collectively abbreviated as Lin.

[0128] NanoString analysis

[0129] Gene expression was evaluated using a custom-built multiplexed 50-gene FBR panel (NanoString Technologies). After 4-weeks post-implantation, all mice wereeuthanizedby CO2 asphyxiation, followed by cervical dislocation. Afterward, the backsof all the mice wereshaved. The explanted material and associated tissues were soaked in RNAlater solution (Invitrogen) overnight at 4 °C. The next day, the excessAttorney Docket No.507814.5000508 RNAlater solution was discarded, and the tissue samples were stored at -80 °C. For RNA isolation, the tissues were thawed, finely diced, and 600 µL of Trizol reagent (Invitrogen) with 10 µL / mL of 2-mercaptoethanol was added to approximately 100 mg of tissue. The tissues were mechanically ground down and incubated for 5 minutes at room temperature. After the addition of 120 µL of chloroform, the tube was shaken vigorously for 15 seconds and incubated for 2 minutes at room temperature. The tubes were centrifuged at 12,000 xg for 15 minutes at 4°C. 350 µL of the aqueous phase was transferred to columns from the RNeasy mini kit (Qiagen), and the RNA isolation was completed using the manufacturer's instructions. All samples were checked for a 260 / 280 ratio of 2 before proceeding to further steps. 100 ng of RNA was processed according to NanoString manufacturer protocols, and RNA levels (absolute copy numbers) were obtained via nCounter (NanoString Technologies). Group samples were analyzed using nSolver analysis software (NanoString Technologies). GUSB, ACTB, B2M, and TBP were used as reference genes.

[0130] ZIPH-3c consistently exhibit the lowest collagen density across all timepoints.

[0131] In addition to the collagen density discussion above, the following are additional studies on the FBR behaviors against ZIPH in comparison with a single- network PEDOT:PSS hydrogel and PSB hydrogel. Also, comparisons were made with the surface self-assembly of zwitterionic polymer brushes, referred to as ZI-SAM-PH, which only changes the surface chemistry of PEDOT:PSS films.

[0132] Disc-shaped samples were implanted subcutaneously into mice for periods of 1, 4, and 12 weeks (See FIG. 6), after which the mice were sacrificed and histological analysis of explanted tissues around the implants was conducted. Using MTS (Masson’s trichrome stain), the density of the collagen layer formed as a result of the FBR was quantified (FIGS.7A and 7B). In FIGS.7A and 7B, the scale bars = 50 m. Error bars = mean ± s.e.m. (standard error of the mean). The number of replicates per condition (N)=5-6 mice per treatment. Analysis of variance (ANOVA) was used to compare collagen densities between treatments. **** = P <0.0001. P is the probability and is a measure of how much statistical difference between two conditions. The lower the P value, the bigger the difference between the two conditions and the higher the likelihood that the difference is real.Attorney Docket No.507814.5000508

[0133] Overall, fibrosis severity was found to increase with time for all materials (FIG.6). For PEDOT:PSS hydrogel, the high FBR-induced fibrosis could largely come from the negatively charged PSS. Tests are performed on PSS hydrogels, which, indeed, gave the highest collagen density at the surface of the implant. The 4-month timepoint is used to directly compare collagen deposition on different implants. As shown by the collagen density as a function of the distance away from the implant surface (FIGS. 7A and 7B), ZIPH-3c indeed had substantially decreased collagen density level than PEDOT:PSS hydrogel, which validated that the incorporation of the zwitterionic hydrogel PSB effectively suppresses the FBR. Surprisingly, ZIPH-3c had an even much lower collagen density than the PSB hydrogel (FIG. 6 and 7A). This indicates that the double-network morphology might be having some unique effect in helping to further suppress the FBR. Along this line, the higher collagen density of ZIPH-u (i.e., ZIPH without the addition of EBSA or solvent treatment) compared to ZIPH-3c implies that the interconnected nanofibrous morphology of PEDOT:PSS is also favorable for suppressing the FBR.

[0134] Along another line, the comparison with the ZI-SAM-PH sample indicates that the surface modification strategy was less effective in suppressing the FBR than the double-network design of ZIPH. Although the immunocompatible properties of zwitterionic polymers mostly come from the antifouling properties, protein adsorption assays indicate that this may not be the case for the ZIPH designs. Among all the tested samples, both ZIPH-3c and ZIPH-u have some of the highest levels of protein adsorption. This suggests that the embedment of PEDOT:PSS in the PSB matrix may not be enough to negate the high protein adsorptive behavior of PEDOT:PSS. The result also shows that PSS is indeed responsible for the high protein adsorptive behavior of PEDOT:PSS, which is consistent with the in vivo collagen density results. Another finding is that the surface assembly of zwitterionic polymer brushes (e.g., ZI- SAM-PH) had one of the lowest levels of protein adsorption. Overall, the trend from these protein adsorption results does not fully agree with the in vivo histology results discussed above, which indicates that in the above designs, protein adsorption only has a loose correlation with the fibrosis associated with each implant type.

[0135] In a further experiment, to monitor the evolution of the collagen capsule over time, both collagen density and thickness at 1, 2, 4, 8, 12, and 24-weeks post-Attorney Docket No.507814.5000508 implantation of various implants were measured. Although the collagen density fluctuated considerably over time, ZIPH-3c consistently exhibited the lowest collagen density (FIGS. 8A and 8B). However, ZIPH-3c consistently was associated with one of the thickest collagen capsules over time, though it remained at a lower level than PEDOT:PSS (FIGS. 8C and 8D). FIG. 8E shows the MTS-stained images used to quantify the relative collagen densities and thicknesses for FIGS. 8A – 8D. The low collagen density and thickness seem to be due to the abundant presence of cells, most of them presumably macrophages, at the surface (FIG.8E).

[0136] Immune response for ZIPH-3c diverges from its parent materials

[0137] In the FBR, macrophages are known to play a central role in the final outcome. To gain a deeper understanding of the different immune responses induced by different hydrogel designs, both flow cytometry (FC) and immunofluorescence (IF) are used to analyze the population size and phenotypes of macrophages surrounding each implant. FC data (See, FIG.9A) indicated that ZIPH-3c was associated with the greatest number of F4 / 80+macrophages compared to other samples including PEDOT:PSS, PSB, PSS, and sham. Furthermore, the polarization states of F4 / 80+macrophages were analyzed using CD86 and CD206, which correspond to classical M1 and M2 phenotypes, respectively. M1 and M2 macrophages are thought to be pro- and anti-inflammatory, respectively, but studies have shown that an imbalance in M2 / M1 macrophage ratio in either direction compared to sham can be a primary driver for poor FBR and fibrosis outcomes. After 4 weeks, the M2 / M1 ratio for all implants except for PEDOT:PSS had no statistically significant differences compared to sham control (FIG. 9B), indicating the dysregulation of the M2 / M1 balance (specifically increased M2 / M1 ratio) for PEDOT:PSS. This suggests that, over time, ZIPH-3c and ZIPH-u are somehow damping the progression to skewed M2 / M1 ratios, and PEDOT:PSS is exacerbating the increase in M2 / M1 ratio. It should be noted that besides M1 and M2 macrophages, a hybrid CD86+CD206+macrophage population was also observed, indicating the existence of other macrophage phenotypes. Immunofluorescence (IF) microscopy confirmed that the F4 / 80+macrophages analyzed using FC were localized at the surface of the implants (FIG. 9C). These results suggest that despite attracting the greatest number of macrophages, ZIPH-3cAttorney Docket No.507814.5000508 better maintains the balance between M2 and M1 phenotypes, which could be part of the reason for the low collagen density.

[0138] Cytokines secreted during the FBR also provide in-depth information on the different immune responses. A 14-cytokine LegendPlex panel is used to quantify cytokine concentrations (FIGS. 9D-9G). Overall, similar to the macrophage phenotypes, cytokine expression profiles of ZIPH-3c and ZIPH-u are found to significantly diverge from their parent materials, PEDOT:PSS and PSB. In particular, ZIPH-3c and ZIPH-u were found to be associated with significantly higher concentrations of MCP-1 and VEGF compared to the rest at both the 1-week and 4- week timepoints (FIGS. 9D and 9F). The significant upregulation of MCP-1 is a potential explanation for why ZIPH-3c is associated with the greatest number of macrophages. Interestingly, even though MCP-1 and VEGF are known to be powerful arteriogenic and angiogenic cytokines, respectively, vascularization in the surrounding areas was not observed for either of the implants (Supplementary Fig. 16). Anotherimportant cytokine is IL-1 is one of the first cytokines released during tissue injury.All the implants except for PSB had suppressed IL-1 levels, and ZIPH-3c had one ofthe lowest IL-1 levels for both time points (FIGS.9D and 9F). This difference betweenZIPH-3c and PSB may explain the lower fibrotic tendency of ZIPH-3c. ForPEDOT:PSS and PSS hydrogels, significant upregulation of TNF- was found after 1week, suggesting an inflammatory response (FIG.9E). Such upregulation persisted at4 weeks for PSS (FIG. 9G). Because TNF- is a powerful pro-inflammatory cytokineleading to fibrosis, this partially explains the fibrotic response against PEDOT:PSS andPSS. No statistically significant differences in the concentrations of IFN- , IL-4, andIL-13 between all implants were observed despite being the major cytokinesassociated with M1 (IFN- ) and M2 (IL-4, IL-13) macrophages. However, this may notbe so surprising, considering that macrophage polarization states often exist on a spectrum between M1 and M2.

[0139] Another important type of cell in the FBR is neutrophils, which act as the first responders to the implants. Fow cytometry is used to quantify Ly-6G+neutrophils. At 1-week post implantation, the number of Ly-6G+neutrophils associated with PSB was found to be significantly greater compared to rest (FIG.9H). It is known that the recruitment of neutrophils can be facilitated by the activation of the complementAttorney Docket No.507814.5000508 system. Therefore, complement activation of different samples is further compared by quantifying protein C3a, which is generated as a result of complement activation. At the 1-week timepoint, PSB had substantially higher C3a concentrations in the peri- implant tissue than the rest (FIG.9I), suggesting that PSB is activating the complement cascade. In FIG 9I, the error bars = mean ± s.e.m., N = 5-6 mice per treatment. One- way ANOVA was used for statistical comparison of multiple means. * = P < 0.05; ** = P < 0.01; **** = P < 0.0001; ## = P < 0.01 compared to rest; #### = P < 0.0001 compared to rest. This finding is consistent with a previous study showing that a zwitterionic polymer, poly(2-metharyloyloxyethyl phosphorylcholine), facilitates complement activation. For the ZIPH samples, despite having PSB as a significant fraction, the lower trafficking of neutrophils and complement activation (FIGS.9H and 9I) could come from PSS, which has been shown to suppress complement activation. This difference between ZIPH and PSB could also be part of the reason for the lower collagen density from ZIPH.

[0140] ZIPH-3c is associated with a unique gene expression pattern that suppresses fibrosis

[0141] To gain further insight into the mechanisms of the FBR against ZIPH-3c, the tissues associated with the implant materials were analyzed with a 50-gene NanoString panel 4 weeks post implantation. Principal component analysis (PCA) was used to analyze the data. Gene expression analysis coupled with characterization of fibroblast and myofibroblast populations show that ZIPH-3c is associated with a unique, mildly fibrotic gene expression pattern.

[0142] When the data are plotted on a vector space defined by the first (PC1) and third (PC3) principal components, there is a clear separation between the mildly fibrotic implants, consisting of ZIPH-3c and ZIPH-u, and the more severe fibrotic implants, consisting of PEDOT:PSS and PSB (See, FIG.10A).

[0143] Compared to the other implants, the gene expression pattern for ZIPH-3c is unique. ZIPH-3c forms a tight, unique cluster on the PC1-PC3 vector space and is one of the main drivers of variance for PC3 (See, FIG.10A). Among all the representative PC3 genes, the gene with the most distinct expression level for ZIPH-3c is Tbx21, a gene for T-bet. Although not statistically significant (P =0.054), Tbx21 was upregulated for ZIPH-3c. T-bet is the master regulator for Type 1 T helper cells (Th1), which, underAttorney Docket No.507814.5000508 certain circumstances, is implicated in the suppression of fibrosis. But for Th2 and Th17, which are implicated in the induction, their master regulators, Gata3 (Th2) and Rorc (Th17), from ZIPH-3c were found to be indistinguishable from the baseline sham control (FIG.10B). On the other hand, PC1 is composed of genes that have the biggest correlation with fibrosis, including Ccr2, Sphk1, Retnla, Acta2, Fap, and Plod2. Strong upregulations of these fibrotic genes are consistently observed for PEDOT:PSS, PSB, and ZISAM-PH, but not for ZIPH-3c or ZIPH-u (FIG. 10C). Not surprisingly, collagen genes also contribute strongly to PC1. Their expression patterns overall follow the trend of collagen deposition on different implants (FIG.10C). However, given that the differences between implants are not statistically significant, there should be other genes and pathways important for the FBR outcome.

[0144] Fibroblasts and myofibroblasts are responsible for collagen deposition, with the increasing presence of myofibroblasts associated with increasing FBR severity. Acta2, a gene expressed by myofibroblasts, was found to be upregulated with statistical significance for PEDOT:PSS and PSB implants, while ZIPH-3c maintained a similar level with sham (FIG.10C). This agrees with IF imaging using SMA (Acta2), which shows concentrated myofibroblasts at the surface of PEDOT:PSS but not for ZIPH-3c (FIG. 10E). In addition, CD45- Lin- CD140a+cells, which encompass both fibroblasts and myofibroblasts, were quantified via flow cytometry. ZIPH-3c, PEDOT:PSS, and PSB were associated with the greatest number of CD45- Lin- CD140a+cells, without statistically significant differences (FIG.10D). IF imaging also confirmed that collagen I, SMA, and F4 / 80 were all co-localized at the implant surface, confirming that the fibrotic response was directed against the implants and both F4 / 80+macrophages and CD45- Lin- CD140a+cells are part of the same immunological milieu (FIG.10E). In FIG.10E, error bars = mean ± s.e.m. N=5-6 mice per treatment. One-way ANOVA was used for statistical comparison of multiple means. * = P < 0.05; ** = P < 0.01; *** = P < 0.001; # = P < 0.00001.

[0145] In contrast to the FC result for CD45- Lin- CD140a+cells, the SMA expression on ZIPH-3c indicates a low number of myofibroblasts, which is consistent with the current understanding of the FBR.

[0146] ZIPH-3c is associated with unique cytokine upregulation patterns.Attorney Docket No.507814.5000508

[0147] Cytokines were quantified at 1, 2, 4, 8, and 12-weeks post-implantation to gain some mechanistic insights into the FBR suppression by ZIPH-3c. ZIPH-3c was consistently associated with the highest level of MCP-1 concentrations across timepoints (FIGS. 11A-11E). ZIPH-3c was also associated with the highest level of VEGF at earlier timepoints (FIG. 11A and 11B). IL-18 concentrations were similar across implant types until 8-week post-implantation when a significant spike in concentration was seen for ZIPH-3c (FIG. 11D). Over time, trends in MCP-1 and VEGF concentrations were observed to be synchronized, while IL-18 showed the opposite trend (FIG.11F). The upregulation of MCP-1, VEGF, and IL-18 may suggest a Th1-driven immune response for ZIPH-3c, which could explain the low collagen density. IL-18 may be upregulated at 8-weeks post-implantation to compensate for the significant decrease in MCP-1 at the timepoint. It is likely that MCP-1 is the reason, if not, one of the reasons, for the high infiltration of macrophages in the peri-implant region of ZIPH-3c. Further investigation is needed to understand the relationships between these cytokines and how they relate to the FBR-induced fibrotic response.

[0148] ZIPH-3c electrocardiographic electrodes demonstrate similar performance to dexamethasone-eluting PEDOT:PSS electrodes.

[0149] To demonstrate that the FBR-suppressing properties of ZIPH-3c are applicable at the device level, electrocardiographic (ECG) electrodes were fabricated with ZIPH-3c and implanted in mice. Due to its super hydrophilic nature, ZIPH-3c had poor adhesion to substrates under aqueous conditions, making device integration difficult. A novel surface multilayer functionalization method that is stable under aqueous conditions was conceived to overcome this problem. The method involves the assembly of poly(glycidyl methacrylate) (PGMA) and polyallylamine (PAH) monolayers to eliminate hydrolysis prone covalent bonds and functionalizing the surface with methacrylate groups (FIG. 12A). The method was further integrated, along with ZIPH-3c, on a parylene C-based device (FIG.12B).

[0150] To record the ECG signal, two electrodes were implanted subcutaneously in the dorsal flanks of the murine rib cage and were connected to a transcutaneous button with external connectors for connection with peripheral electronics (FIG.12C). Chronic ECG measurements were successfully obtained for 12 weeks (FIG.12D). As controls, PEDOT:PSS electrodes and PEDOT:PSS electrodes with dexamethasoneAttorney Docket No.507814.5000508 (Dex)-eluting silicone backings (PEDOT:PSS-Dex) were also implanted. Dex is a steroidal anti-inflammatory drug that is used in various implantable medical devices for extending the lifetime. At early timepoints, all electrodes experienced large variations and fluctuations in signal between timepoints. However, by week 8, all electrode types started to stabilize and variations between electrodes of the same type decreased. Out of all the electrode types, ZIPH-3c was the most consistent across timepoints. By week 12, the average normalized voltage retentions of ZIPH-3c and PEDOT:PSS-Dex were twice that of PEDOT:PSS. The average normalized voltage retention of ZIPH-3c was nearly identical to that of PEDOT:PSS-Dex (FIG. 12E). These results demonstrate the effectiveness of ZIPH-3c at suppressing the FBR and serving as biointerfacing electrodes. Furthermore, ZIPH-3c did not lead to side effects in mice, such as delayed hair growth and the muscular atrophy of panniculus carnosus, which were observed for dexamethasone-releasing implants.

[0151] While various embodiments have been described, it will be apparent to those with ordinary skill in the art that many more embodiments and implementations are possible. Accordingly, the embodiments described herein are examples, not the only possible embodiments and implementations.

[0152] The subject-matter of the disclosure may also relate to the following aspects:

[0153] A first aspect relates to a mixed ion / electron conducting material with suppressed foreign body responses. The mixed ion / electron conducting material includes a zwitterionic PEDOT:PSS hydrogel comprising a first network comprising interconnected fibers of poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) and a second network comprising a zwitterionic polymer hydrogel, whereby the first network is formed in situ in the second network to form a double network hydrogel.

[0154] A second aspect relates to the mixed ion / electron conducting material of the first aspect, wherein the zwitterionic PEDOT:PSS polymer hydrogel comprises poly(sulfobetaine methacrylate) (PSB).

[0155] A third aspect relates to the mixed ion / electron conducting material of the first or second aspect, wherein the zwitterionic PEDOT:PSS hydrogel has a form of a thin film or a bulk film.Attorney Docket No.507814.5000508

[0156] A fourth aspect relates to the mixed ion / electron conducting material of any preceding aspect, wherein the fibers of PEDOT:PSS comprise nanofibers.

[0157] A fifth aspect relates to a method for manufacturing a mixed ion / electron conducting material with suppressed foreign body responses. The method includes forming a zwitterionic PEDOT:PSS hydrogel comprising a first network comprising interconnected fibers of poly(3,4- ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) and a second network comprising a zwitterionic polymer hydrogel, whereby the first network is formed in situ in the second network to form a double network hydrogel.

[0158] A sixth aspect relates to the process of the fifth aspect, wherein the zwitterionic PEDOT:PSS hydrogel has a form of a thin film or a bulk film.

[0159] A seventh aspect relates to the process of the fifth or sixth aspect, wherein the fibers of PEDOT:PSS comprise nanofibers.

[0160] An eighth aspect relates to the process of any preceding aspect, wherein forming the zwitterionic PEDOT:PSS hydrogel includes dissolving a zwitterionic monomer and a crosslinking agent in an aqueous dispersion containing particles comprising poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) to form a mixture; polymerizing the mixture via thermal or UV polymerization to form the zwitterionic PEDOT:PSS hydrogel, wherein the mixture includes an initiator; and subjecting the zwitterionic PEDOT:PSS hydrogel to at least one solvent annealing treatment cycle, whereby the particles transition to an interconnected fiber network.

[0161] A ninth aspect relates to the process of any preceding aspect, wherein the particles of PEDOT:PSS comprise nano-particles.

[0162] A tenth aspect relates to the process of any preceding aspect, wherein the solvent used in the solvent annealing is ethanol or 2,2,2-trifluoroethanol (TFE).

[0163] An eleventh aspect relates to the process of any preceding aspect, wherein each solvent annealing treatment cycle comprises exposing the zwitterionic PEDOT:PSS hydrogel to ethanol first and then to 2,2,2-trifluoroethanol (TFE).

[0164] A twelfth aspect relates to the process of any preceding aspect, wherein the zwitterionic PEDOT:PSS hydrogel is subjected to at least 2 to 3 solvent annealing treatment cycles.Attorney Docket No.507814.5000508

[0165] A thirteenth aspect relates to the process of any preceding aspect, wherein the zwitterionic monomer is [2-(methacryloyloxy)ethyl]dimethyl-(3- sulfopropyl)ammonium hydroxide (SBMA).

[0166] A fourteenth aspect relates to the process of any preceding aspect, wherein the crosslinking agent is N, N’-methylenebisacrylamide (MBAA).

[0167] A fifteenth aspect relates to the process of any preceding aspect, wherein the initiator is ammonium persulfate or 2-hydroxy-2-methylpropiophenone.

[0168] A sixteenth aspect relates to the process of any preceding aspect, the process further includes adding an additive to the mixture prior to the polymerizing step to promote PEDOT:PSS aggregation in the hydrogel.

[0169] A seventeenth aspect relates to the process of any preceding aspect, wherein the additive is 4-ethylbenzenesulfonic acid (EBSA).

[0170] An eighteenth aspect relates to an implantable medical device for implanting inside a living body. The implantable medical device includes a mixed ion / electron conducting material according to any preceding aspect.

[0171] A nineteenth aspect relates to the implantable medical device of the eighteenth aspect, wherein the zwitterionic PEDOT:PSS hydrogel suppresses foreign body response-mediated fibrosis compared to unmodified PEDOT:PSS hydrogels.

[0172] A twentieth aspect relates to the implantable medical device of the eighteenth or nineteenth aspect, wherein the zwitterionic PEDOT:PSS hydrogel exhibits a reduction in foreign body response associated fibrotic severity by at least 64% compared to unmodified PEDOT:PSS hydrogels.

[0173] A twenty-first aspect relates to the implantable medical device of any preceding aspect, wherein the zwitterionic PEDOT:PSS hydrogel increases the conductivity of the mixed ion / electron conducting material by more than one order of magnitude compared to unmodified PEDOT:PSS hydrogels.

[0174] A twenty-two aspect relates to the implantable medical device of any preceding aspect, wherein the implantable medical device is selected from the group consisting of cardiac pacemakers, implantable cardioverter defibrillators, cochlear implants, and continuous glucose monitors.

[0175] In addition to the features mentioned in each of the independent aspects enumerated above, some examples may show, alone or in combination, the optionalAttorney Docket No.507814.5000508 features mentioned in the dependent aspects and / or as disclosed in the description above and shown in the figures.

Claims

Attorney Docket No.507814.5000508 CLAIMS 1. A mixed ion / electron conducting material with suppressed foreign body responses, the mixed ion / electron conducting material comprising: a zwitterionic PEDOT:PSS hydrogel comprising: a first network comprising interconnected fibers of poly(3,4- ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS); and a second network comprising a zwitterionic polymer hydrogel, wherein the first network is formed in situ in the second network to form a double network hydrogel.

2. The mixed ion / electron conducting material of claim 1, wherein the zwitterionic PEDOT:PSS polymer hydrogel comprises poly(sulfobetaine methacrylate) (PSB).

3. The mixed ion / electron conducting material of claim 1, wherein the zwitterionic PEDOT:PSS hydrogel has a form of a thin film or a bulk film.

4. The mixed ion / electron conducting material of claim 1, wherein the fibers of PEDOT:PSS comprise nanofibers.

5. A method for manufacturing a mixed ion / electron conducting material with suppressed foreign body responses, the method comprising: forming a zwitterionic PEDOT:PSS hydrogel comprising a first network comprising interconnected fibers of poly(3,4- ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) and a second network comprising a zwitterionic polymer hydrogel, wherein the first network is formed in situ in the second network to form a double network hydrogel.

6. The method of claim 5, wherein the zwitterionic PEDOT:PSS hydrogel has a form of a thin film or a bulk film.Attorney Docket No.507814.5000508 7. The method of claim 5, wherein the fibers of PEDOT:PSS comprise nanofibers.

8. The method of claim 5, wherein forming the zwitterionic PEDOT:PSS hydrogel comprises: dissolving a zwitterionic monomer and a crosslinking agent in an aqueous dispersion containing particles comprising poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) to form a mixture; polymerizing the mixture via thermal or UV polymerization to form the zwitterionic PEDOT:PSS hydrogel, wherein the mixture includes an initiator; and subjecting the zwitterionic PEDOT:PSS hydrogel to at least one solvent annealing treatment cycle, whereby the particles transition to an interconnected fiber network.

9. The method of claim 8, wherein the particles of PEDOT:PSS comprise nano-particles.

10. The method of claim 8, wherein the solvent used in the solvent annealing is ethanol or 2,2,2-trifluoroethanol (TFE).

11. The method of claim 8, wherein each solvent annealing treatment cycle comprises exposing the zwitterionic PEDOT:PSS hydrogel to ethanol first and then to 2,2,2-trifluoroethanol (TFE).

12. The method of claim 8, wherein the zwitterionic PEDOT:PSS hydrogel is subjected to at least 2 to 3 solvent annealing treatment cycles.

13. The method of claim 8, wherein the zwitterionic monomer is [2- (methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide (SBMA).

14. The method of claim 8, wherein the crosslinking agent is N, N’- methylenebisacrylamide (MBAA).Attorney Docket No.507814.5000508 15. The method of claim 8, wherein the initiator is ammonium persulfate or 2-hydroxy-2-methylpropiophenone.

16. The method of claim 8, further comprising adding an additive to the mixture prior to the polymerizing step to promote PEDOT:PSS aggregation in the hydrogel.

17. The method of claim 16, wherein the additive is 4-ethylbenzenesulfonic acid (EBSA).

18. An implantable medical device for implanting inside a living body comprising a mixed ion / electron conducting material according to claim 1.

19. The implantable medical device of claim 18, wherein the zwitterionic PEDOT:PSS hydrogel suppresses foreign body response-mediated fibrosis compared to unmodified PEDOT:PSS hydrogels.

20. The implantable medical device of claim 18, wherein the zwitterionic PEDOT:PSS hydrogel exhibits a reduction in foreign body response associated fibrotic severity by at least 64% compared to unmodified PEDOT:PSS hydrogels.

21. The implantable medical device of claim 18, wherein the zwitterionic PEDOT:PSS hydrogel increases the conductivity of the mixed ion / electron conducting material by more than one order of magnitude compared to unmodified PEDOT:PSS hydrogels.

22. The implantable medical device of claim 18, wherein the implantable medical device is selected from the group consisting of cardiac pacemakers, implantable cardioverter defibrillators, cochlear implants, and continuous glucose monitors.

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