Injectable Conductive Hydrogels for Use in Neuroprosthetic Intervention

Neuroprosthetic interventions are strategies aimed at treating a wide range of neurological disorders, long-term neuroprosthetic treatments traditionally require the implantation of hard metallic electrodes that must sustain their electrical connection with neural tissues for prolonged periods of time. However, surgical introduction of these electrodes and their mechanical mismatch with neural tissues results in inflammation, which disrupts their electrical interface. Our aim was to develop soft, injectable, and conducting hydrogel-based electrode materials and characterize their sustained mechanical and electrical properties before and after sterilization and injection. These gels were made from poly (3-4, ethylenedioxythiophene) (PEDOT) (a conductive polymer) and poly (acrylic acid) (PAAc) and were polymerized at subfreezing temperatures to generate soft 3D macroporous structures. These porous hydrogels exhibited enhances mechanical properties. When optimized, gels exhibited softness consistent with neural tissues (<100 kPa), excellent toughness (>2 kJ/m), and excellent strain-at-failure (survived >90% compression without failure). Additionally, these gels’ mechanical properties could be tuned by altering their compositions, though their conductivity remained almost constant and independent of gels composition at about 1 S/cm. This conductivity was much higher than neural tissues making them well-suited for stimulating the sensing in neuroprosthetic applications. Finally, because of their optimized mechanical properties, these gels were highly compressible, exhibited further enhanced electrical properties when compressed, and were capable of surviving injection through 16-gauge needles.


Introduction
Up to one billion people worldwide are affected by neurological disorders (e.g. Alzheimer's, Parkinson disease, and epilepsy). This number is estimated to increase considerably (Fig. 1A), particularly due to the rise of life expectancy. The degree to which people are debilitated by these neurological problems rivals and exceeds some of the worst ailments in existence (e.g. HIV, heart disease, and cancer) (Fig. 1B) (World Health Organization, 2006). Neuroprosthetic devices are developed in order to return the function or to reduce the symptoms exhibited by patients suffering from neurological diseases (Leache et al.). For instance, electrical therapies such as auditory implantation (Zeng et al., Wilson et al.), deep brain stimulation (Williams et al., Lyons), spinal cord functional electrical stimulation (FES) (Carmel et al., Minassian et al.), spinal cord stimulation (SCS) (Tilley et al.), and vision prosthesis (Fridman et al., Pezaris et al) are promising strategies in treating patients with severe neurological diseases. While promising, these strategies are not without limitations (Leach et al.). These strategies require a stable electrode-tissue interface which does not vary over time (Hassarati et al.). Electrodes must: (i) minimize the overpotential at the tissue-electrode interface and (ii) maintain an electrically conductive path between the electrode and the tissue being stimulated. Traditionally, metallic electrodes are used in these situations and have outstanding electrical properties over a short period of time. However, they do not maintain these properties over protracted periods of time (Green et al. (2008), Leach et al., Rousche et al.), which is of particular importance for treating neurological disorders that require long-term intervention (e.g. deafness, blindness, epilepsy, Parkinson's Disease, etc.). Implanting these electrodes into neural tissue ( Fig. 2A) results in an initial period of electrical conductivity; however, this conductivity is thought to be lost over time for number of reasons. First, electrode introduction requires invasive surgery (Zeng et al., Wilson et al) which results in blood vessel and neuron rupture. This traumatic electrode introduction results in an initial degree of conductivity, but starts the foreign body response (Fig 2B) (Leach et al.). Later on, loss of conductivity is likely due to the micromotion (Rousche et al., Leach et al.) which results from a mismatch in stiffness between the electrodes and the tissue. This mismatch causes shear stress at the interface, resulting in inflammation, glial tissue development (Fig. 2C), increased  Green et al. (2012), Rousche et al.). Thus, conductive materials are needed that are injectable and soft in order to reduce inflammation and to maintain electrical connectivity with neural tissues for longer periods of time. In this study, we therefore aimed to develop a novel hydrogel-based electrode whose mechanical properties mimic those of native tissue and that can be introduced through a minimally invasive injection. Normal neural tissue before electrode insertion (tissue is not in an inflamed state). That is, microglia is inactivated and astrocytes are not in a reactive state. B. After the insertion of the electrode inflammation begins, but neurons and the electrode are still in contact with neurons. C. As time goes on, this inflammation results in scare tissues, which isolates the electrodes from the neurons. Activated microglia and reactive astrocytes start blocking the electrode and from neurons D. Chronic inflammation due to mismatch between brain tissue and electrode will result shear stress and causes more migration and attachment of scarring and glial tissue at the surface of the electrode. This further isolates prevents electrical connectivity between the electrode and the neurons.
In order to help reduce inflammation and the associated loss in conductivity when implanted for prolonged periods of time, electrodes developed here were designed to have tissue-like mechanical properties by virtue of being made in a hydrogel format.  (Pan et al.), that have the potential to be loaded with and locally deliver therapeutic agents such as anti-inflammatories (Kearney and Mooney). Unlike traditional electrode materials, hydrogels have similar mechanical properties as physiological tissues, which make them useful in a wide variety of biomedical applications (Dee et al.). However, they normally have poor conductivity (Saracino et al.). Although poor conductivity is not an issue in many applications, for use as electrodes in neuroprosthetic interventions, conductivity is a critical priority (Cheong et al.; Leach et al.). Conductive polymer hydrogels are a unique class of hydrogels that not only have similar mechanical properties as neural tissue (Aregueta-Rables et al.), but also have electrical properties similar to metals and inorganic semiconductors (Zhao et al.). It has recently been demonstrated that hydrogels composed of Poly (3-4, ethylenedioxythiophene) (PEDOT) provide electrical conductivity (Fig. 3A) and stability in oxidized environments (Cheong et al.;Abidian et al. (2006); Abidian et al. (2008); Huange et al.; Naficy et al.; Sekine et al.). Moreover, PEDOT hydrogels are biocompatible (Cho et al.) and can have similar Young's moduli as neural tissues (Fig.   3B). For these reasons, PEDOT was employed as the conductive component of the hydrogel-based electrodes described here. A. Demonstration of PEDOT hydrogel conductivity showing that it can be used to complete an LED circuit (Naficy et al.). B. A table comparing the Young's moduli of a traditional hydrogels (poly acrylic acid) with conducting hydrogels (PEDOT) and neural tissues. C. Images showing that a gel retains its shape after injection (Bencherif, et. al) Finally, to further minimize inflammation, these hydrogel-based electrodes were designed to be injectable by being endowed with shape-memory properties. Shape memory polymers are a class of polymers which can change their shape temporarily due to the change in the temperature, pressure, or pH and regain its original shape (Pilate et al.). This property allows for minimally invasive implantation (Wang et al.). These hydrogel-based electrodes were fabricated using a cryogelation approach, which results in hydrogel structures with interconnected macropores (Kumar and Sirvastava). The sponge-like "cryogels" are capable of being injected by being passed through a standard 16-gauge needle (Fig. 3C, top). The cryogels retain their structural integrity after injection (Fig. 3C, bottom). Additionally, the cryogelated hydrogels have moduli (Kennedy et al.) similar to tissue (i.e., 10s-100s kPa) (Engler et al.), and have sufficient mechanical toughness to remain intact and mechanically consistent when exposed to cyclic compression (50% compression, 1000 cycles) (Cezar et al.).
We hypothesized that PEGDM-cross-linked poly(AAc) gels that encapsulate PEDOT polymers will exhibit softness, toughness, compressibility, be injectable, while also being electrically conductive. Our goal in this study was to create an injectable conductive hydrogel which we were expected to be extremely tough comparing to other hydrogels. Our approach involves the use of a PEGDM-cross-linked anionic poly(AAc) hydrogel network to entrap cationic PEDOT polymer. Beyond the covalent PEGDM cross-links holding the gel together (which are highly bio-compatible), these polymer choices provide a high density of electrostatic interactions between negatively charge poly(AAc) and positively charged PEDOT (i.e., ionic cross-links). These electrostatic interactions are broken, absorbing energy when the gel is compressed. However, after breaking and absorbing energy, these broken ionic cross-links can re-cross-link with other positive/negative charges in the network, as they are bountiful in interpenetrating poly(AAc)/PEDOT networks. We believe this strategy will produce gels with the enhanced toughness required for injectability and for long-term survival after injection, while retaining the electrical conductivity of PEDOT and while still maintaining the softness of a hydrogel material.

Materials
Acrylic acid (AAc), Sodium hydroxide ( from Sigma Alrdrich. Except Acrylic acid, the rest of materials were used without any changes. Acrylic acid was purified with aluminum oxide column to remove inhibitor.

Gel fabrication
To make non-macroporous gels, gels were fabricated at room temperature (RT).
Different amounts of AAc (4 wt % to 10 wt %), PEGDM (1, 2, and 5 wt %), NaOH network that physically entrapped PEDOT (Fig. 4B). Gels were soaked in PBS three days prior to running tests. To make gels with enhanced mechanical properties for softness and injectability, macroporous gels were made as described above, however, they were cryogelated in 10 x 10 mm cylindrical Teflon molds at -20 C and left to gel at -20 C overnight. This results in a concentrated PEGDM-cross-linked poly(AAc) network with entrapped PEDOT that exists between ice crystals. When the ice crystals are thawed, the result in a macro-porous structure with a concentration PEGDM-crosslinked network with entrapped PEDOT existing between macro-pores ( Fig 4C).

Gel mechanical and structural characterizations
Gel stiffness (modulus), strain at failure, and toughness were quantitatively measured using a compression test. Gels of various compositions were placed between the plates of an Instron (Model 3345) and compressed until they structurally failed ( Fig.   5A). After running the compression test on a given gel, a stress-strain curve (Fig. 5B  Pore interconnectivity was characterized using a water wicking test. Gel mass was measured before and after they were wicked with piece of tissue. The percent change in mass was recorded as the percent pore interconnectivity. The weight measurements were performed using an analytical balance (model AND SH-120). For detailed calculation please see appendix A.2.
The general microstructure of some of the macroporous cryogelated gels were assessed using scanning electron microscopy (SEM) (Zeiss SIGMA VP Field Emission). Gels were lyophilized immediately after cryogelation and placed on metal mounting post with adhesive and then sputter coated with platinum-palladium. Then, they were put in the SEM sample chamber and scanned to produce electron micrographic images.

Stress (kPa)
A B

Gel characterizations after sterilization
In order to sterilize the hydrogel-based electrodes, they were soaked for a day in PBS then placed in 70% ethanol for an hour and then were swollen for another two days in sterile PBS to remove excess ethanol prior to being passed through a 16 gauge needle.
As an alternative to ethanol sterilization, hydrogel-based electrodes were also sterilized using an autoclave. After removing gels from their molds, each gel was soaked in PBS in separate media bottles. Each media bottle was completely covered in aluminum foil and sealed with autoclave tape and autoclaved for 30 minutes (at 121 ℃). After

v A B
sterilization, in the manners described earlier (Sections 2.3 and 2.4), the electrical and mechanical properties of these injected gels were characterized and compared to those that were not sterilized. For detailed calculation please see appendix A.

Gel cyclic testing
In order to characterize gel robustness, gels were placed between the plates of an Instron (model 3345). Samples were cyclically loaded at 50% strain 10 times. Gels were also cycled 1 time at 90% strain in order to mimic the strain associated with injection. In the manners described in the sections 2.3 and 2.4 the mechanical and electrical properties of these gels were characterized and compared to those that were not cycled. For detailed calculation please see appendix A.

Gel injection
In order to investigate the injectability of gels, a subset of gels with desirable electrical and mechanical properties were made as described in section 2.2 at different sizes. These gels were cryogelated in 0.125 mm cylindrical Teflon molds with different (0.78, 1.08, 1.57, 2.35, and 3.14 mm) thicknesses. Then, cryogels were passed through 16 gauge needles and inspected afterwards for preservation of their structural integrity.

Data processing, representation, and statistical analyses
All data collected were represented as mean plus and minus standard deviations across repeat experiments. When comparing individual control single experimental conditions, student t-test was used to extract p-values. When comparing across multiple 13 conditions, Analysis of Variance (ANOVA) was employed with Tukey's post-hoc analysis. For all comparisons, a p-value of less than 0.05 was required as a benchmark for statistical significance.

Characterizations of gels' mechanical properties
Mechanical compression experiments were performed to characterize how gel composition influenced mechanical properties such as stiffness, strain of failure, and toughness. Gels were made with different concentrations of polymer (poly(AAc)) and cross-linker (PEGDM) and a constant PEDOT concentration (2.6 wt %). This particular gel format was used to provide a negatively charged framework (poly (AAc) covalently cross-linked with PEGDM) for the conducting PEDOT to interpenetrate. 2.6 wt % PEDOT was the maximum concentration possible for all gel formulations and that in separate experiments, using higher PEDOT concentrations did not improve gel conductivity (data not shown). Additionally, it was thought that the electrostatic interactions between the positively charged PEDOT and the negatively charged AAc network would yield tough mechanical properties (Sun et al.). In these studies, it was found that increasing AAc and PEGDM concentrations increased gel stiffness ( Fig. 7A and 7B, trendlines increasing from left to right). On the other hand, increasing AAc and PEGDM concentrations decreased strain of failure ( Fig. 7C and 7D, trendlines decreasing from left to right), thus producing gels that, while stiffer, failed under lighter compression (i.e., were more brittle). Generally, cryogelation resulted in gels that were softer and could endure higher strains compared to their room temperature counterparts ( Fig. 7A and 7B, comparing blue and red bars, particularly at higher cross-linking concentrations). This enhanced softness and strain at failure likely was due to the sponge-like macro-porosity exhibited by the cryogels (Fig. 7C). Note that often the softness of the cryogels was not characterized by the Young's modulus values. This was likely due to the fact the highly macro-porous cryogels collapsed under their own weight and thus began compression tests in an already compressed (and therefore stiffer) state. Figure 7. The strain of failure and stiffness of these 2.6 wt% PEDOT hydrogel-based electrodes are a function of hydrogel composition. A. Young's modulus as a function of AAc concentration for (i) gels cast at room temperature (RT) and (ii) gels that were cryogelated at -20 ℃. Blue, red and green data points and exponential trend lines corresponds to gels cross-linked using 1, 2, and 5 wt % PEGDM, respectively. B. Young's modulus as a function of PEGDM concentration for (i) gels cast at room temperature and (ii) gels were cryogelated at -20 ℃. Blue, red, green, and black data points trend lines correspond to gels with AAc concentration of 4, 7, 9, and 10wt %, respectively. C. Strain at failure as function of AAc concentration for (i) RT and (ii) cryogels. Blue, red, and green data points and trend lines correspond to gels cross-linked using 1, 2, and 5 wt% PEGDM, respectively. D. Strain of failure as a function of PEGDM concentration for (i) RT and (ii) cryogels. Blue, red, green and black data points trend lines correspond to gels with AAc concentration of 4, 7, 9, and 10wt %, respectively. In parts A through D, N = 4. *, **, ***, and **** indicate statistically significant differences when comparing gels . Cryogels were generally softer and could withstand higher strains than gels cast at room temperature. A. Bar graph comparing Young's moduli for gels cast at room temperature (blue) to cryogels (red) with PEGDM concentrations of (i) 1 wt %, (ii) 2 wt %, and (iii) 5 wt %. B. Bar graph comparing strain at failure for gels cast at room temperature (blue) to cryogels (red) with PEGDM concentrations of (i) 1 wt %, (ii) 2 wt %, and (iii) 5 wt %. C. An SEM image of a cryogel, highlighting its macro-porosity (for cryogel with 9 wt % AAc, 1 wt % PEGDM, 2.6 wt % PEDOT). In parts A and B, N = 4. *, **, ***, and **** indicate p-values of < 0.05, 0.01, 0.001, and 0.0001, respectively. Gel toughness is a parameter related to both stiffness and strain of failure and is a mechanical parameter critical for injectability and long-term robustness. Generally, toughness increased with increasing Acrylic Acid and PEGDM concentration ( Fig. 9A and 9B, trendlines increasing from left to right). Though this trend generally holds true, some gel formulations either did not follow this trend or only did so slightly. This might be a consequence of how increasing Acrylic Acid and PEGDM concentrations both increases stiffness (which allows the gels to absorbed more energy and small strains) but also increasing brittleness (which results in gels failing and lower strains). In other words, the mechanical toughness can likely be optimized at Acrylic Acid and PEGDM concentrations that are neither too high nor too low. In this work (as will be examined more in Section 3.3), we will be optimizing our gels for enhanced softness and ability to be injected. Namely, we will be looking for gels that have both low modulus and high strain of failure.

Characterization of gels' electrical properties
Electrical tests were performed to determine how gel composition influenced electrical conductivity. It was found that gel conductivity remained relatively constant for varying concentrations of AAc and PEGDM ( Fig.10A and B, 0.8-1.2 S/cm for all gel formulations with no striking trends). This was likely due to the same amount of PEDOT concentration (2.6 wt %, the maximum possible concentration for the adopted fabrication process) being used in all gels. Higher PEGDM cross-linking did seem to slightly reduce conductivity (Fig. 9B, a slight downward trendline slope from left to right). We attributed this to several phenomena. First, gels with higher PEGDM crosslinking had lower pore interconnectivities (Fig. 10C, downward trend moving left to right, with this being more prominent in the macro-porous cryogels), therefore took on fewer conducting ions, and therefore did had slightly lower conductivities. Also, these highly cross-linked gels often bulged and did not make flush contact with the electrodes, again slightly reducing conductivity. In general, there were no strong trends between gel conductivity and gel composition, which suggests that the mechanical properties of these gels can be tuned (Figs. 7, 8, and 9) independently from their conductivities.
Independence of conductivity from gel composition is beneficial, since it gives one the opportunity to choose gel composition based on the desirable mechanical properties determined by application, while retaining electrical conductivity. In any case, the conductivities exhibited by these PEDOT hydrogels were sufficient to light an LED with a battery (Fig. 10D), demonstrating relatively high conductivity. Figure 10. The electrical conductivity of these 2.6 wt % PEDOT gels does not dramatically vary based on AAc and PEGDM concentrations. A (i) RT and (ii) cryogelated gel conductivities at the indicated AAc concentrations for gels cross-linked using 1 (blue), 2 (red), and 5 (green) wt % PEGDM. B (i) RT and (ii) cryogelated gel conductivities at 1, 2, and 5 wt % PEGDM. 4 (black), 7 (green), 9 (red), and 10 (blue) wt% AAc. C (i) RT and (ii) cryogelated gel pore interconnectivity at 1, 2, and 5 wt % PEGDM. 4 (black), 7 (green), 9 (red), and 10 (blue) wt % AAc. D The conductivities provided by these PEDOT gels were sufficient to complete an LED circuit. Cryogel shown was made with 9 wt % AAc, 1 wt % PEGDM, and 2.6 wt% PEDOT. In part A, B, and C, N = 4. *, **, ***, and **** indicate statistically significant differences when comparing gels at constant AAc concentrations ( While the electrical conductivities of these gels remained unaffected by AAc and PEGDM concentrations, certain Acrylic Acid and PEGDM formulations resulted in gels that were capable of providing enhanced electrical properties by virtue of their mechanical properties. Specifically, formulations that yielded softer gels could be compressed more readily and this compression resulted in enhanced electrical conductance. For example, under moderate strain (50%), softer, more compressible gels exhibited enhanced conductances compared to stiffer gels (Fig. 11, higher measured conductance for gels with lower Young's moduli in both RT and cryogels). This was likely due to compression reducing the electrical path. Interestingly, gels made at room temperature exhibited slightly higher conductances under moderate strain (50%) than macro-porous cryogels (comparing parts (i) and (ii) in Fig. 11A). This might be due to the differences in how non-macro-porous and macro-porous gels compress. Non-macroporous room temperature gel volumes remained constant under compression: with reduction of height, gel surface area increased (Fig. 11B). This reduction in height and increase in cross-sectional surface area both enhanced electrical conductance. On the other hand, macro-porous cryogels tended to collapse in volume under compression: with reduction of height, surface area did not increase (the porous structure of the cryogels enabled this) (Fig. 11C). This type of collapse does enhance electrical conductance, but only due to reduction in height. Thus, macro-porous cryogels did not experience as much increase in conductivity when compressed. However, because this difference in conductance is only slight, and because the macro-porous cryogels exhibited excellent mechanics (softness and high strain of failure, Fig. 7), we believe they are more well-suited than PEDOT/EDOT gels (Naficy et al.;Dai et al. (2009);Dai et al. (2010); Cho et al.; Saski et al.) as injectable soft electrode materials for neuroprosthetic interventions. Furthermore, the electrical properties of these macroporous cryogels could be further enhanced using gel formations that resulted in higher degrees of pore interconnectivity. Specifically, cryogels with higher pore interconnectivity exhibited both higher conductivity at 0% strain (Fig. 12A, red data) and higher conductance at 50% strain (Fig. 12B, red data). Generally, cryogels could be made with higher pore interconnectivities by using formulations with modest amounts of AAc and PEGDM (i.e., formulations that utilized larger concentrations of water in lieu of polymer and cross-linker). This likely led to larger ice crystal development which promoted higher concentrations of PEDOT between ice crystals, leading to slightly higher electrical conductivity. Non-macro-porous room temperature gels, did not exhibit significant trends in this regard, likely due to their generally low pore interconnectivities (Fig. 12, blue data). Note that, again, while room temperature gels exhibited slighter higher conductances at 50% strain (Fig. 12A, blue data), these gels did not exhibit the high strain of failures required for injectability. Figure 11. When compressed, softer gels exhibited enhanced electrical conductances. A. The measured conductance of (i) RT gels and (ii) cryogels as a function of their moduli at 50% strain. B. Photographs of room temperature made gels before and after compression, highlighting how the gel crosssectional area increases during compression (thus enhancing conductance). C. Photographs of cryogels before and during compression, highlighting how gel cross-sectional area does not greatly increasing during compression.
In another study conducted by Dai et al. (2010) using different hydrogel networks ((EDOT concentrations ranging from 0.48 to 1.2 mol/L) and polystyrene sulfonate (PSS) concentrations ranging from 0.1 to 1 mol/L) conductivities from 5.7 x 10 -3 to 6.5 x 10 -2 S/cm were reported (lower than here). In a separate study, Dai et al. (2010) went on to measure EDOT hydrogel conductivity with other formulations (0.03 to 0.48 mol/L EDOT and 0.1 to 0.5 mol/L PSS). Conductivity of these gels were reported to be 6.7 x 10 -4 to 2.2 x 10 -3 S/cm. Cho et al. reported conductivities ranging between 10 -5 and 10 -3 S/cm (when stimulated at 0.1 and 10 6 kHz respectively at 25 ℃). One study (Saski et al.) was able to achieve enhanced conductivity between 40-80 S/cm by adopting a novel EDOT electropolymerized technique. However, this electropolymerization technique is more appropriate for thin film hydrogels or hydrogels as coatings. The electrical conductivities reported here and in these other studies (Naficy et al.;Dai et al. 2009;Dai et al. 2010;Cho et al.;Saski et al.) are likely appropriate for electrical stimulation and sensing of neural tissues. For instance, because brain tissue is about 0.0015 to 0.003 S/cm and cerebrospinal fluid is about 0.015 S/cm (Lorenzo et al.), neuroprosthetic electrode materials must exhibit higher conductivities for proper sensing/stimulation.
Our hydrogel-based electrodes (and some of those cited here) provide electrical conductivities well above these tissues and fluids (Figs. 10 and 12). Notably, however, our PEDOT cryogels provide the mechanical properties required for minimally invasive injection in addition to adequate electrical conductivities. While electrically conducting injectable hydrogels have been reported previously, they were fabricated with completely different materials (gelatin-based) using different fabrication methods, and only yielded conductivities around 7.25 x 10 -3 S/cm (Li et al.)

Optimization of gel mechanical and electrical properties
Because of the complex interplay between gel composition and the many parameters of interest here (Young's modulus, strain of failure, toughness, and electrical conductivity), we have directly measured each of these parameters over a wide range of gel formulations. This has allowed us to identify select gel formulations that provide both desirable mechanical and electrical properties for use as injectable, conducting, and soft neuroprosthetic electrodes. In order to obtain relatively soft gels, only gels with stiffness less than 2 kPa were chosen (Fig 13A, bars identified with diamonds). Also for injectability, gels must be passed through 16 gauge needles and therefore should tolerate high strains. Thus, only gels which did not show any sign of failure at 90% strain were selected (Fig. 13B, bars identified with daggers). For long-term survival robustness and mechanical toughness are critical parameters. Thus, only gels which could absorb energy more than 2 kJ/m 3 were selected (Fig. 13C, bars identified with double daggers).
Finally, in order to receive and deliver electrical signals in neuroprosthetic applications, only the most conductive gels were desirable. Thus, only gels which had at least 1 S/cm conductivities were selected (Fig. 13D, bars identified with lightning bolts). The only gels that satisfied all these requirements were the 7 wt % AAc, 1 wt % PEGDM cryogels and the 9 wt % AAc, 1 wt % PEGDM cryogels (both containing 2.6 wt % PEDOT).
These optimized gels could both survive high strains ( Fig. 14A and B) and exhibited excellent electrical conductance under this high strain (Fig. 14C). In fact, the electrical conductance of gels when compressed from 0 to 90 % strain increased by an order of magnitude (Fig. 14C). This property may be very useful when these cryogels are compressed within tissue after injection and the electrode conductance may increase.

Gel properties when cyclically compressed, sterilized, and injected
For use in neuroprosthetic applications, these hydrogel materials must be robust enough to retain their mechanical and electrical properties after cyclic compression and sterilization. As injectable electrode materials, they must also remain structurally intact when subjected to the high strains associated with injection. We therefore conducted experiments examining how cyclic compression and sterilization influenced gel electrical and mechanical properties and directly tested these gels for injectability. After selecting optimum cryogel formulations (7 and 9 wt % AAc, 1 wt % PEGDM, and 2.6 wt % PEDOT), they were sterilized with ethanol or in an autoclave. Results indicated that ethanol-sterilization had less effect on gel mechanical properties than autoclavesterilization. Specifically, Young's modulus moderately changed after ethanol treatment ( Fig. 15A, "ethanol" column) whereas autoclave treatment significantly altered the modulus (Fig. 15A, "autoclave" column). This was most prominent in the 9 wt % AAc cryogels (Fig. 15A, blue bars). Additionally, autoclaving gels increased stiffness, which is not desirable for this application. Most critical to injection, however, is that ethanol treated gels fully retained their ability to undergo large strains (Fig. 15B, "ethanol" column) while autoclaved gels did not. Furthermore, ethanol-sterilized gels retained their mechanical properties better after cyclic compression ( Fig. 15A and B, "ethanol + 10 cycle" and "autoclave + 10 cycle" columns). Taken altogether, these data indicate that autoclaving might change the structure of these gels, increasing their stiffness and eliminating their shape memory properties. Thus, ethanol treatment appears to be a more desirable method for sterilizing these gels. Testing the electrical conductivity after sterilization and cyclic testing also revealed some interesting insights. Surprisingly, sterilization moderately improved conductivity, except for the ethanol-treated 9 wt % cryogels (Fig. 15C, bars indicating positive percent changes). Although 9 wt % cryogels sterilized with ethanol electrical conductivity slightly increased, this result did not occur after ten cycles (Fig. 15C, "ethanol + 10 cycles" column, blue bar). These increases in electrical conductivity likely are due to both ethanol-and autoclave-treatments changing the macro-porous gel structure. For instance, ethanol-treatment causes these gels to collapse greatly, forcing macro-pores to close. Autoclave-treatment likely damages the gels, also forcing macropores to collapse. In both cases, this pore closure likely increased electrical conductivity (i.e., better interconnectivity of the conducting hydrogel when pore walls touch). While these increases in electrical conductivities may be desirable, they are likely temporaryespecially for the ethanol-treated gels, since the de-swelling effects of ethanol treatment may not persist after being subjected to physiological media for prolonged periods of time.
Success of injection (Fig. 16A) was dependent on the speed of injection, size of cryogel, and how well hydrated the gels were. As expected, cryogels with smaller diameters survived injection better than larger gels (Fig. 16B). Notably, gels made with 9 wt % Acrylic Acid, 1 wt % PEGDM, and 2.6 wt % PEDOT survived injection better than the 7 wt % Acrylic Acid cryogels (Fig. 16B, red bars only surviving at appreciable rates for gels < 1.28 mm thick, whereas blue bar indicate 100% survival of 1.58-mm-thick gels). Figure 16. These gels are injectable with some gel sizes working better than the others. A Different stages of injection before injection, inside the syringe, and after injection. B Gel survival percentage as a function of height for diameter = 0.125 mm for cryogels 7 wt% AAc, 1 wt % PEGDM, 2.6 wt % PEDOT (red) and 9 wt% AAc, 1 wt % PEGDM, 2.6 wt % PEDOT (blue).
There are other efforts to create injectable conducting hydrogels, though using different approaches to fabricate the hydrogels. Li et al. (2014) adopted an in situ crosslinking approach to form gelatin-graft-polyaniline (GP) gels at body temperature.
The solution was injected to animal hosts, and introduction to body temperature initiated gel polymerization. Conductivity of these gels was reported to be several orders of magnitude lower than those reported in the present study (between 1.21 x 10 -4 -4.45 x 10 -4 S/cm). Also, Li et al. (2014) reported swelling ratios that significantly increased over time after in situ injection (up to 500 %), which could be highly problematic in that this swelling could damage sensitive neural tissues. The optimized cryogels reported in our study swelled only 33.25 % and 36.75 % (7 wt % AAc, 1 wt % PEGDM, 2.6 wt % PEDOT and 9 wt % AAc, 1 wt % PEGDM, 2.6 wt % PDOT, respectively). The GP hydrogels reported by Li et al. (2014) also degraded and lost over 95% of their weight in 22 days, making them useful in many applications, but not well-suited for long-term electrode-neural interfaces. The optimized cryogels reported here were not designed to degrade and remained intact for several months in vitro. In another study (Li et al. (2015)), conductive hydrogels were made of gelatin-graft-polyanaline (GP) and carboxymethyl-chitosan, and in situ crosslinked with oxidized dextran via Schiff base at physiological conditions. These gels also gelated after injection and exhibited improvements in both electrical conductivity and swelling ratio from Li et al. (2014).
Increasing the amount of polyanaline reduced swelling ratio down to 120%.
Conductivity of these gels were improved (increased from 6.2 x 10 -3 S/cm to 7.3 x 10 -3 with increasing polyaniline), but were still orders of magnitude lower than what are reported in the present study. Elsewhere, Wu et al. adapted host-guest interactions to fabricated hydrogels. Hydrogels were used γ-cyclodextrin dimers as the host molecule and tetraaniline and poly (ethylene glycol) as the guest copolymer, enabling gelation post injection. While utilizing a highly novel gelation approach, the conductivity of these gels was in the range of 10 -4 S/cm. Our gels had comparatively higher conductivities (about 1 S/cm).

Conclusion
We were able to generate electrically conducting hydrogels with unique mechanical properties by interpenetrating a PEDOT network with a poly (acrylic acid) network. The gels' mechanical properties could be tuned (1 to 70 kPa) by altering the amount of both poly (acrylic acid) polymer and PEGDM cross-linker contained in the gel. Notably, altering gel composition in order to achieve tuned mechanics did not overly affect gel conductivity (maintained around 1 S/cm). Cryogelation of these conducting gels produced gels that were very soft (< 10 kPa) and exhibited excellent strain of failures (> 90%) and mechanical toughness (2 -8 kJ/m 3 ). Owing to these enhanced mechanical properties, specific gel formulations (7 and 9 wt % AAc with 1 wt % PEGDM and 2.6 wt % PEDOT) were identified to be highly compressible and proved to be capable of sterilization (ethanol-treatment), cyclical compression, and injection through a 16-gauge needle. Also, owing to this compressibility, when compressed, these gels exhibited enhanced electrical conductances. We believe the combination of softness, toughness, compressibility, and electrical conductivity exhibited by these PEDOT-based hydrogels make them very well suited for use in neuroprosthetic applications where injectability and softness are desirable material attributes for minimizing inflammation, reducing scar tissue, and prolonging electrical contact with neural tissues.
Will be plotted vs strain, which is degree of deformation. The force unit in SI system is Newton (N). The surface area unit in SI system is square meter (m 2 ). Therefore, the stress unit in SI system is Newton per square meter (N/m 2 ) or Pascal (Pa). Engineering strain is a ratio of total deformation to the initial dimension. Eq. 4 = ∫ 0 Toughness unit in SI system is J/m 3 .

A.2 Pore interconnectivity Equations:
For pore interconnectivity, the mass of the sample will be measured with scale.
Eq.5 % = − Where: Eq.6 Wb = Wg + Wm + Wp (Wp are just the pours which are connected to the surface) Eq.7 = + A.3 Electrical Equations: Figure A2. Electrical properties of each gel will be measured. A. Schematic of the electrical circuit used for recording the current. 8 volts is applied across a cylindrical gel (blue cylinder) and the current is measured. B. The set-up for recording the current under various strains with a zoomed-in image of a PEDOT cryogel under compression.
Samples are placed in the circuit similar to Figure A2. Different voltages (V) are applied, currents (I) are recorded and the resistances (R), resistivity (ρ) and electrical conductivity (σ) of the samples are calculated with ohm's law: Eq.8 = I is current and its unit is Amperes (A), V is voltage and its unit is volts (V) and resistance (R) unit is Ohms (Ω).

Eq. 9 =
Where ρ is electrical resistivity, L is length (its unit is m) and S is surface area (its unit is m 2 ) and R is resistance (and its unit is Ω) and the resistivity unit is (Ω.m) Eq. 10 = 1 Where σ is electrical conductivity. The SI unit of conductivity is Siemens per meter (S/m).
Eq. 11 = Where G is electrical conductance. The Si unit of conductance is Siemens (S)