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Electro-optical mechanically flexible coaxial microprobes for minimally invasive interfacing with intrinsic neural circuits

Central to advancing our understanding of neural circuits is developing minimally invasive, multi-modal interfaces capable of simultaneously recording and modulating neural activity. Recent devices have focused on matching the mechanical compliance of tissue to reduce inflammatory responses. However, reductions in the size of multi-modal interfaces are needed to further improve biocompatibility and long-term recording capabilities. Here a multi-modal coaxial microprobe design with a minimally invasive footprint (8–14 µm diameter over millimeter lengths) that enables efficient electrical and optical interrogation of neural networks is presented. In the brain, the probes allowed robust electrical measurement and optogenetic stimulation. Scalable fabrication strategies can be used with various electrical and optical materials, making the probes highly customizable to experimental requirements, including length, diameter, and mechanical properties. Given their negligible inflammatory response, these probes promise to enable a new generation of readily tunable multi-modal devices for long-term, minimally invasive interfacing with neural circuits.

and mechanical properties. Given their negligible inflammatory response, these probes promise to enable a new generation of readily tunable multi-modal devices for minimally invasive interfacing with neural circuits.
Microelectrode recordings are the gold standard for measuring individual neurons' activity at high temporal resolution in any nervous system region and central to defining the role of neural circuits in controlling behavior. Microelectrode arrays, such as the Utah 3 or Michigan 4 arrays, have allowed tracking of distributed neural activity with millisecond precision. However, their large footprint and rigidity lead to tissue damage and inflammation that hamper long-term recordings 1,5 . State of the art Neuropixel 6 and carbon fiber probes 7 have improved on these previous devices by increasing electrode density and reducing probe dimensions and rigidity.
Although these probes have advanced the field of recordings, next-generation devices should enable targeted stimulation in addition to colocalized electrical recordings 1,2 . Optogenetic techniques enable high-speed modulation of cellular activity through targeted expression and activation of light-sensitive opsins 8 . However, given the strong light scattering and high absorption properties of neural tissue 9 optogenetic interfacing with deep neural circuits typically requires the implantation of large-diameter rigid fibers, which can make this approach more invasive than its electrical counterpart 10,11 .
The ideal neural probe would combine optical and electrical modes while maintaining small cross-sectional dimensions and tunable lengths. The ability to bi-directionally interface with genetically defined neuron types and circuits is key to ultimately being able to understand how the nervous system computes and controls behavior. It is also fundamental for determining the mechanistic basis of sensorimotor disorders, defining how circuit activity is affected by injury, and how it might be restored or facilitated. Approaches to integrating optical and electrical modalities have ranged from adding fiber optics to existing Utah arrays 12 to the Optetrode 13,14 or of the substrate, and then butt-coupled to a cleaved SMF (Fig. 1a). We used active alignment to maximize mode overlap between the microfiber and SMF. The coupling was locked in using a UV-curable optical adhesive droplet on the end of the SMF (Fig. 1b). To create a robust detachable interface for in vivo testing, the SMF was inserted into a ceramic ferrule. The distal end of the ferrule assembly was machine polished to allow coupling to a patch cable ( Fig. 1g-h). Other interface designs for different applications are conceivable ( Supplementary Fig. 11).
To form a low noise conductive layer around the probe tip, a 379 ± 43 nm layer of iridium oxide (IrO x ) was sputtered on the microfibers 20 followed by a 362 ± 137 nm electrochemically deposited layer of poly(3,4-ethylene dioxythiophene) polystyrene sulfonate (PEDOT:PSS) (Fig.   1c) 20 . The porous nature of IrO x allowed better adhesion of the conductive PEDOT:PSS layer and enhances the overall electrical performance of the probe. The probe was passivated with 1.76 ± 0.16 µm of Parylene-C to electrically isolate the probe and provide a biocompatible surface ( Supplementary Fig. 2). In order to expose the electrical and optical surfaces, a focused ion beam was used to cleave off the tip ( Fig. 1e; see Methods). Fig. 1g displays the final probe design and Fig. 1h shows a photograph of a completed probe. Through the combination of IrO x and PEDOT:PSS, electrical impedances of < 1 MΩ at 1 kHz were achieved from electrode areas of < 15 µm 2 (Fig. 1f).
The probes' optical properties were first assessed by imaging the output cone angle in a dye solution (Fig. 1d, inset), which showed a uniform output with a divergence angle of 10 -15°. Importantly, after the cladding layers are placed on the probe, no detectable scattering light is observed from the microfiber/SMF interface (Fig. 1d) as compared to the pre-cladding probe (Fig. 1b). The optical losses between a laser-coupled patch cable and the EO-Flex output were quantified using three different wavelengths (473 nm, 543 nm, 600 nm) with all devices showing < 7 dB (n=4). These values match up well with simulated results for an ~2 µm mode misalignment in the ferrule sleeve ( Supplementary Fig. 1). Electrochemical impedance spectroscopy (EIS) was performed on the probes while submersed in a 1x phosphate-buffered solution (PBS). All probes fabricated and tested showed an average electrical impedance of 844 ± 179 kΩ at 1 kHz (n=4; Fig. 1f and Supplementary Fig. 2a-d) after cladding deposition and milling the tip, compared to >10 MΩ before PEDOT deposition.
To confirm that EO-Flex probes allow high-sensitivity electrical measurements in vivo, we performed simultaneous extracellular recordings and two-photon imaging in the cortex of isoflurane-anesthetized mice 21,22 . Imaging of fluorescently labeled cells (see Methods) enabled the monitoring of insertion and targeted movement of the probe through the tissue ( Fig. 2a and Supplementary Fig. 8). Probes readily penetrated the exposed dura with minimal buckling when using water immersion and reached target regions in optically accessible cortical layer 2/3. When mounted to a three-axis micromanipulator, fine adjustment of the probe tip position, once inside the tissue, was feasible to optimize the signal-to-noise ratio and target individual neurons.
However, the lateral movement was typically limited to < 30 µm. Using this approach, we acquired endogenous multi-and single-unit activity (Fig. 2b). Principle component analysis (PCA) and Gaussian clustering of electrical recordings were used to determine the number of distinct units (Fig. 2c). Spiking rates were calculated using a Bayesian Adaptive Kernel Smoother (BAKS) algorithm 23 applied to the full duration of the recording (Fig. 2d). Fig. 2b-e shows a representative recording.  . We varied the stimulation frequency (10 Hz -60 Hz), pulse width (0.6 -9.8 ms), and output power (5 -208 µW) to determine the optimal settings to excite ChR2-expressing neurons. Using waveform analysis on simultaneously recorded electrical activity, we found that a minimum power of 29 µW (2,849 mW mm -2 ) was required for firing of the cells in sync with the optical pulse train (Fig. S4). In the example recording shown in Fig. 3a-b, PCA combined with a mixed Gaussian fit for the clustering of the data yielded two primary clusters ( Fig. 3c) with two different waveforms (Fig. 3d) occurring during the stimulation period ( Fig. 3e-g). We found minimal interference (e.g., Becquerel effect) between the proximal optical and electrical channels, as demonstrated by retracting the EO-Flex probe away from ChR2-expressing neurons, or placing it in a buffer solution, while optically pumping at maximum power (208 µW) using the same optogenetic pulse trains ( Supplementary Fig. 7). At this maximum power, neural circuits responded with minimal temporal lag (Fig. S4f) and could follow frequency stimuli up to 40 Hz before struggling to maintain synced firing ( Supplementary Fig. 6).
The ability of EO-Flex probes to optically evoke neural activity was further verified by twophoton calcium imaging in Vglut2-GCaMP6f mice. Four to five weeks after the cortical injection of an AAV2-CaMKII-C1V1-mCherry vector (see Methods), expressing the green light-activated ion channel C1V1 in neurons, probes were inserted into layer 2/3 regions with C1V1 expression ( Supplementary Fig. 8). A 556 nm DPSS laser was coupled to the EO-Flex probe, and stimulation parameters were swept while simultaneously monitoring neuronal calcium transients. Delivered optical pulses led to correlated calcium spiking in C1V1-positive neurons within the field of view ( Supplementary Fig. 8). The successful optical evocation of neural activity was also verified by simultaneous electrical recordings ( Supplementary Fig. 8e).
Together, our in vivo data demonstrate the ability of EO-Flex probes to electrically record and optically modulate neural activity in the intact brain.
Lastly, we evaluated the brain's response to probe implantation. EO-Flex probes were implanted into the cortex of Cx3cr1 GFP/+ mice with labeled microglia for seven days. A 250 µmdiameter multi-mode fiber, suitable for optogenetic experiments, was inserted using the same stereotaxic coordinates but on the opposite hemisphere for comparison. Serial brain sections were prepared that included both implantation sites. Tissue slices were co-stained with anti-GFAP and anti-NeuN antibodies to quantify reactive astrogliosis and neuronal loss, respectively (Supplementary Fig. 9-10). We found that implantation of the multi-mode fiber was associated with significant neuronal loss ( Supplementary Fig. 10c)  Developing probes that can reach deeper brain regions is straightforward with the developed fabrication protocols as virtually any microfiber length can be generated ( Supplementary Fig.   11a). However, for a given set of cladding layers, the probe's stiffness decreases with length.
Therefore, longer probes might require additional tactics to overcome low buckling forces during the insertion process (e.g., dissolvable sugar coatings 30 , or rigid polymer layers). Alternatively, a surgical incision in the dura could facilitate probe insertion. Regardless of probe length, our implantation studies demonstrated that the small-footprint EO-Flex probes have a drastically reduced immune response in comparison to standard multi-mode fibers.
In summary, we report on the fabrication of novel multi-modal coaxial microprobes and demonstrate their ability to optically stimulate and electrically record from intrinsic neural circuits with minimal interference between the two modalities. The small footprint and high aspect ratio of the EO-Flex probes allow for minimally invasive interfacing with neural circuits. Further size reduction is possible with this coaxial design using smaller fiber optic cores, however, the tradeoff is an increase in optical losses and electrical impedance ( Supplementary Fig. 3).
Although the probes' capabilities were only tested in the brain, as a platform with excellent control over probe diameter and length, the choice of cladding materials with various chemical compositions, inherent mechanically flexibility, and a clear route to scaling up probe densities (e.g., translating the cladding deposition process to fiber bundles) ( Supplementary Fig. 11), this technology should find immediate applications as minimally invasive interfaces in diverse nervous system regions, including the spinal cord and peripheral nerves.

Acknowledgements
We would like to thank Drs. Anis Husain and Rob Saperstein (Ziva Corporation) for technical discussions and electromagnetic simulations of early EO-Flex prototypes and designs. We The SnO 2 nanofibers were synthesized using thermal evaporation of SnO powders at high temperatures according to published protocols 1 . Ceramic combustion boats were loaded with 1-5 grams of tin monoxide powder and placed in a tube furnace. The system was pumped down to < 1 mTorr as the furnace was turned on to 1000 °C. At operating temperature, system pressures were typically around 300 mTorr. The system was allowed to run for an hour, after which the furnace was turned off, and the system was allowed to cool while the vacuum pump remained on. The combustion boat was then removed, and nanowires found on the rim of the boat were transferred to a silicon substrate.
To enable efficient optical coupling of the waveguide to standard optogenetic hardware, a single-mode fiber (SMF) (Thorlabs S405-XP) with a mode field diameter (2.8-3.4 µm) slightly smaller than the microfiber core was chosen. This ensures high optical coupling efficiency. To create a robust detachable interface for in vivo testing, the SMF was inserted into a ceramic ferrule (Thorlabs CF126-10) and secured in place with quick cure epoxy (DevCon #20445). The ferrule assemblies were then machine polished until a smooth coupling interface was observed through a fiber inspection scope (Thorlabs, FS200), and the opposing fiber end (for coupling to the waveguide) was cleaved using a ruby scribe (Thorlabs S90R). The ferrule assembly was mounted on a three-axis stage, and the scribed end was maneuvered into a droplet of UV-cured Together, these procedures yielded ferrule assemblies for repeatable mounting to an in vivo imaging and optogenetics setup (Fig. 2). Alternative interface designs (e.g., for probe arrays) are shown in Supplementary Fig. 11. were made on each pulled needle to measure the virus volume injected into the brain.
Mice were anesthetized with isoflurane (4% for induction; 1%-1.5% for maintenance) and positioned in a computer-assisted stereotactic system with digital coordinate readout and atlas targeting (Angle Two, Leica). Body temperature was maintained at 36°C-37°C with a DC temperature controller, and ophthalmic ointment was used to prevent eyes from drying. A small amount of depilator cream (Nair) was used to remove hair at the designated skin incision site.
The skin was cleaned and sterilized with a two-stage scrub of betadine and 70% ethanol. For injection, a drop of the virus was carefully pipetted onto parafilm (1-2 μl) for filling the pulled injection needle with the desired volume. Once loaded with sufficient volume, the injection needle was slowly lowered into the brain until the target depth (DV 0.2 mm) was reached. Manual pressure was applied using a 30-ml syringe connected by shrink tubing, and 0.4 μl of the AAV2-CaMKII-C1V1-mCherry vector (6.1E+12 VP/ml; undiluted; UNC Vector Core) was slowly injected over 5-10 min. Once the virus was injected, the syringe's pressure valve was locked. The position was maintained for approximately 10 min to allow the virus to spread and to avoid backflow upon needle retraction. Following the injection, the skin was sutured along the incision. Mice were given subcutaneous Buprenex SR (0.5 mg per kg) and allowed to recover before placement in their home cage. The vector was allowed to express for 4-5 weeks before in vivo recordings.
Animal preparation for in vivo recordings. Surgical procedures closely followed established protocols 5,6 . Briefly, mice were anesthetized with isoflurane (4-5% for induction; 1%-1.5% for maintenance) and implanted with a head plate on a custom surgical bed (Thorlabs). Body temperature was maintained at 36°C-37°C with a DC temperature control system, and ophthalmic ointment was used to prevent the eyes from drying. Depilator cream (Nair) was used to remove hair at the designated skin incision site. The skin was thoroughly cleansed and disinfected with a two-stage scrub of betadine and 70% ethanol. A scalp portion was surgically removed to expose frontal, parietal, and interparietal skull segments. Scalp edges were attached to the lateral sides of the skull using a tissue-compatible adhesive (Vetbond; 3M). A custom-machined metal plate was affixed to the skull with dental cement (cat. #H00335; Coltene Whaledent), allowing the head to be stabilized with a custom holder. An approximately 2 mm x 4 mm diameter craniotomy was made over the target area (e.g., AAV injection site). The dura mater overlying the cortex was kept intact. A 1% agarose solution and coverslip were applied to the exposed cortical tissue. To facilitate probe entry into the tissue, the agarose and coverslip were cut on one side to be flush with the craniotomy, allowing direct cortical access through the agarose. The coverslip was affixed to the skull with dental cement to control tissue motion. Recordings commenced immediately after optical window preparation. The depth of anesthesia was monitored throughout the experiment and adjusted as needed to maintain a breath rate of approximately 55-65 breaths per minute. Saline was supplemented as needed to compensate for fluid loss.

In vivo electrophysiology.
To characterize the electrophysiological properties of the EO-Flex probes, we performed extracellular single-and multi-unit recordings in the cortex of isofluraneanesthetized mice, as previously described 5,7 . The EO-Flex probes' electrical channel was  Fig. 4b).
Peristimulus plots correlating optical stimuli with spiking events were calculated using kernel bandwidth optimization, which has been shown to accurately estimate the underlying spiking rate 4 (Fig. 3b). For stimulation frequencies of 10, 20, and 30 Hz, we saw estimated firing rates closely following input frequencies when accounting for the echo after stimulation. In comparison, frequencies of 40 and 50 Hz initially tracked the input frequency for the first couple of light pulses, but eventually, the firing rate decreased as the optical train progressed ( Supplementary Fig. 6g).
Analysis of the two-photon calcium imaging data was performed using Suite2p 10 ( Supplementary Fig. 8). Optically evoked calcium spiking was observed in optical planes near the probe tip. Our analysis focused on the optical planes in which at least three cellular-size regions of interest (ROIs) consistently responded throughout the stimulation period ( Supplementary Fig. 8).
Immunostaining data were processed, analyzed, and plotted using ImageJ, Imaris, and