key: cord-0021310-36g0xdwy
authors: Tabet, Anthony; Antonini, Marc-Joseph; Sahasrabudhe, Atharva; Park, Jimin; Rosenfeld, Dekel; Koehler, Florian; Yuk, Hyunwoo; Hanson, Samuel; Stinson, Jordan; Stok, Melissa; Zhao, Xuanhe; Wang, Chun; Anikeeva, Polina
title: Modular Integration of Hydrogel Neural Interfaces
date: 2021-08-28
journal: ACS Cent Sci
DOI: 10.1021/acscentsci.1c00592
sha: fc15900f71727b6ea8bc1aacdfa823509e57a9dd
doc_id: 21310
cord_uid: 36g0xdwy

[Image: see text] Thermal drawing has been recently leveraged to yield multifunctional, fiber-based neural probes at near kilometer length scales. Despite its promise, the widespread adoption of this approach has been impeded by (1) material compatibility requirements and (2) labor-intensive interfacing of functional features to external hardware. Furthermore, in multifunctional fibers, significant volume is occupied by passive polymer cladding that so far has only served structural or electrical insulation purposes. In this article, we report a rapid, robust, and modular approach to creating multifunctional fiber-based neural interfaces using a solvent evaporation or entrapment-driven (SEED) integration process. This process brings together electrical, optical, and microfluidic modalities all encased within a copolymer comprised of water-soluble poly(ethylene glycol) tethered to water-insoluble poly(urethane) (PU-PEG). We employ these devices for simultaneous optogenetics and electrophysiology and demonstrate that multifunctional neural probes can be used to deliver cellular cargo with high viability. Upon exposure to water, PU-PEG cladding spontaneously forms a hydrogel, which in addition to enabling integration of modalities, can harbor small molecules and nanomaterials that can be released into local tissue following implantation. We also synthesized a custom nanodroplet forming block polymer and demonstrated that embedding such materials within the hydrogel cladding of our probes enables delivery of hydrophobic small molecules in vitro and in vivo. Our approach widens the chemical toolbox and expands the capabilities of multifunctional neural interfaces.

The fiber drawing process enables the fabrication of flexible neural probes that can simultaneously interrogate neuronal circuits via electrical, optical, and chemical modalities. During fiber drawing, a macroscopic model (the preform) of the desired probe is fabricated and drawn into hundreds of meters of fibers with microscale features. 1, 2 To date, these probes have enabled one-step optogenetics, 3 in vivo photopharmacology, 4 and in situ electrochemical synthesis of gaseous molecules for neuromodulation. 5 Despite these advancements, this approach has several limitations. To be codrawable, the constituent materials need to have similar glass transition temperatures (for polymers) and melting temperatures (for metals). The resulting melt viscosities must also be compatible to obtain stable draw conditions necessary to maintain the crosssectional geometry of the preform. Additionally, while thermal drawing yields hundreds of meters of fiber at once, each individual centimeter-long device must be be manually connected to back-end hardware, a laborious process that currently represents the fabrication bottleneck. Furthermore, the polymer cladding of these fibers has only served passive structural or electrical insulation purposes, significantly adding to the device footprint with little added functionality.

Hydrogels are an attractive class of materials for neural interfaces. 6 The mammalian brain itself is a weak hydrogel with a complex modulus G* on the order of 1 kPa. 7 While it has been shown that hydrogels alone can serve as neural interfaces, for example as optical waveguides 8 or electrodes, 9 their use in multifunctional neural probes has been more limited. 10, 11 Additionally, while hydrogels have been extensively used as depots for sustained release of bioactive molecules, 12 this drug delivery capability has not yet been extended to multifunctional neural interfaces.

In this article, we complement thermal drawing with a solvent evaporation or entrapment-driven (SEED) integration process to create multifunctional hydrogel-based neural interfaces capable of eluting drugs and nanomaterials. Fibers comprising of optical waveguides, microfluidic channels, or electrode arrays were first drawn individually ( Figure 1A −B).

The optical waveguide consisted of a poly(carbonate) (PC) core with a poly(methyl methacrylate) (PMMA) cladding (n PC = 1.586, n PMMA = 1.49), the recording electrode array fiber comprised four 25 μm tungsten (W) wires within a PC cladding, and the microfluidic channel was a hollow PC fiber. By using a copolymer that forms a physical gel upon introduction of an aqueous medium, we combined these individual modalities into a single physical hydrogel-based neural interface ( Figure 1C−D) . This SEED integration process did not require any free radicals or other toxic byproducts 11 and could accommodate a wide chemical toolbox previously inaccessible to multifunctional neural probes.

To avoid sophisticated cleaning steps associated with potentially toxic radical initiators, we employed a copolymer of poly(ethylene glycol) tethered to water-insoluble poly-(urethane) (PU-PEG), polymers with known biocompatibility Chronic recording of optogenetically invoked action potentials in the nucleus accumbens (NAc) in a Thy1-ChR2 mouse. (C) Demonstrating release profile of drugs injected through the microfluidic channel using Evans blue dye. Brain atlas image with NAc highlighted in blue (left). Crosssection of a Thy1-ChR2 brain injected with 3 μL of Evans blue dye at 33 nL/s followed by fixation with 4% paraformaldehyde, showing clearly the location of the NAc bolus (middle). Cross-section of the same brain approximately 0.7 mm away, showing the periphery of the depot. (D) Flow cytometry data demonstrating the capability of these neural interfaces to deliver cells with high viability: a mixture of approximately 50% live and 50% dead or dying RAW-Blue murine macrophages (left) and cells injected through the microfluidic channel with the backfill method at 1 μL/min (middle). We obtained similar viability when comparing to injections with 26G NanoFil syringes and live cells left on ice (right). The histogram was normalized to the mode. routinely used in clinical implants and pharmaceuticals. 13,14 Upon exposure to water, the PEG blocks facilitate hydration of the material while the hydrophobic forces between PU blocks prevent dissolution, resulting in a physical hydrogel. Since both blocks are soluble in ethanol, we first dissolved the copolymer in a 95% ethanol solution to form a PU-PEG bath. We then brought the individual fiber components together in this bath and used a heat source to evaporate the solvent, resulting in an integrated assembly ( Figure 1E ). This integration creates hydrogel fibers that maintain structural integrity upon insertion in a phantom brain model, and after implantation in vivo ( Figure S1 ). We then characterized the electrical, optical, and fluid delivery properties of these hydrogel-integrated probes ( Figure S2 −S4). The recording electrodes, 2 25 μm W wires, had impedance of 80 kOhm at 1 kHz which is well within the range suitable for extracellular recordings of neuronal potentials. 1 We chose W over nickel chromium (NiCr) used in tetrodes to avoid gold-plating, which is necessary to achieve sub-MOhm impedance, as that step would expose our hydrogel to an organic solvent. 15 The 25 μm W electrodes were selected over the 12.5 μm alternative because of a > 8-times lower impedance ( Figure S2 ). Optical losses in the PC/PMMA waveguide were measured as 0.76 dB/cm loss at a 473 nm wavelength, which was consistent with previously observed losses in PC-core fibers and sufficient for optical neural excitation mediated by channelrhodopsin-2 (ChR2) ( Figure  S3 ). We observed an injection efficiency of >90% for injection rates above 10 nL/s, confirming efficient fluid delivery through the microfluidic channels ( Figure S4 ). Finally, using dynamic mechanical analysis (DMA), we assessed the bending stiffness of the hydrogel neural interfaces ( Figure S5 ) and observed that these devices were flexible, in particular compared to other commonly used devices in neuroscience. 3, 16, 17 We then evaluated the functionality of our hydrogel-based probes in transgenic mice broadly expressing ChR2 fused to a yellow fluorescent protein under the Thy1 promoter. 18 ChR2 is a light-gated cation channel which, upon irradiation with blue light, causes neuronal depolarization and firing of action potentials. 19 We chronically implanted our neural probes into the nucleaus accumbens (NAc) of Thy1-ChR2 mice ( Figure  2A ). The NAc plays an important role in the cognitive processing of reward and motivation, and its aberrant function has been implicated in a wide range of mental disorders, including schizophrenia, substance addiction, and posttraumatic stress disorder. 20 The polymer-based optical waveguide and W recording electrodes allowed us to optically stimulate and simultaneously record evoked ( Figure 2B ; (= 473 nm, 10 Hz, 5 ms pulse width, 20 mW/mm 2 ) and spontaneous activity ( Figure S8 ) of NAc neurons in anesthetized mice. This activity was correlated with laser onset (jitter = 0.84 ms, mean peak latency = 5.86 ms, Figure  S6 ). To confirm that the observed action potentials were not due to the photoelectrochemical (Becquerel) effect, 21, 22 we also performed electrophysiological recordings during optical stimulation in a chronically implanted Thy1-ChR2 mouse following euthanasia and observed no evoked activity ( Figure  S7 ). We then validated our microfluidic capabilities in vivo by delivering Evans blue dye (2% in sterile saline) into the NAc of Thy1-ChR2 mice. Ten minutes following injection, the animals were transcardially perfused with 4% paraformaldehyde, and widefield microscopy of brain slices revealed a dye depot formation in the NAc (Figure 2C−D) .

Cell-based therapies remain an active area of research for fundamental understanding and future treatments of neurological diseases. 23, 24 We assessed whether our devices are compatible with these emerging therapeutic approaches by evaluating the viability of cells following delivery through the integrated microfluidic channel. Depending on the application, the therapeutic cargo can be front-filled into the tip of the microfluidic channel and delivered during device implantation surgery 3 or can be backfilled days or weeks following chronic implantation 1 ( Figure 2C ). We evaluated the viability of RAW-Blue macrophages (RBMs) using both delivery strategies via flow cytometric analysis with 4′,6-diamidino-2-phenylindole (DAPI) and Annexin V (AnnV) conjugated to Alexa Fluor 647 (AnnV-AF647). As a DNA-binding dye, DAPI was used to probe the viability of cells as fluorescence is only observed when cell membrane integrity is lost during cell death. Annexin V was used to identify exposed aminophospholipid phosphatidylserine (PS). PS is normally maintained on the inner leaflet of the cell membrane under physiological conditions but becomes exposed during the early stages of regulated cell death and serves as a phagocytic signal. Together, these markers enable quantification of apoptotic and necrotic processes in response to cell stresses or treatments. We applied these markers to compare viability of cells delivered through the microfluidic channel within the hydrogel neural probe to those kept on ice, injected with a 26G NanoFil syringe, or killed via heat shock ( Figure 2D ). With either the front-or backfill approach, our devices retained RBM cell viability >90% at a 1 μL/min injection rate ( Figures 2E and S9) , suggesting that these probes can be used to deliver live cells directly into the central nervous system. We next demonstrated that the hydrogel cladding enables delivery of small molecules along the entire length of the probe. We used fluorescein as a model drug and coloaded it into the hydrogel precursor PU-PEG ethanol solution. Since fluorescein is water-soluble, introduction into an agarose phantom brain results in diffusion of this molecule away from the hydrogel (Figures 3A and S10 ). Quantitative analysis of release into PBS shows a bolus release that peaks 30 min postinsertion ( Figure 3B ). We then demonstrated this capability in vivo by coloading Evans blue into the hydrogel and implanting the resulting neural probe into the NAc of Thy1-ChR2 mice ( Figure 3C−D) . Confocal microscopy images of brain slices revealed that hydrogel-loaded Evans blue has a different in vivo release profile compared to delivery through the microfluidic channel. Instead of convection-driven transport (first term on the right-hand side in eq 1) at the tip of the implant, Evans blue delivery is dominated by diffusiondriven transport (second term on the right-hand side in eq 1) ( Figure 3E −F) and happens along the whole length of the implant. This additional drug delivery modality enabled by the hydrogel may be more advantageous for certain applications, such as the modulation of the foreign body response using antifibrotic drugs eluted along the length of implants. 25

Controlled drug delivery of hydrophobic small molecule drugs remains a formidable obstacle to their translational utility. Despite recent setbacks, 26 emergent clinical applications of hydrophobic molecules, such as cannabinoids, have garnered renewed interest in their effective delivery. 27 It was recently demonstrated that rationally designed polymers can overcome the delivery challenges of hydrophobic small molecules by, for example, forming nanodroplets that can carry these molecules into the cytosol. 28 These custom polymers are melts at room temperature, with glass transition temperatures >150°C, lower than that of PU-PEG, and are not codrawable with structural polymers typically used in fiber drawing of neural probes. SEED integration allows us to overcome these challenges and thus expands the drug delivery capabilities of neural interfaces. To enable delivery of hydrophobic compounds, a custom block copolymer (Scheme 1) of PEG and poly(caprolactone) (PCL) was synthesized with an acid-labile ether linkage. 28 This poly(acetal), named PA11 (1:1 ratio of PEG:PCL), was blended with the PU-PEG precursor solution at a 3:17 ratio, and the resulting blend was applied during integration of a fiber assembly. We found that PA11 releases from the PU-PEG matrix and forms nanodroplets approximately 25 nm in diameter in saline solution under mild perturbations ( Figure  4A−C) . As a positive control, we repeated the experiment with a thick film of the PU-PEG/PA11 blend and vigorously shook it overnight, which resulted in PA11 nanodroplets of similar size. Transmission electron microscopy (TEM) images of the dehydrated samples further corroborated formation and dimensions of these nanodroplets ( Figure 4B ). We then assessed the ability of PA11 to deliver hydrophobic compounds into aqueous media. We found that when the hydrophobic small molecule Nile Red was mixed in with the PA11:PU-PEG composite, PA11 could escape and carry the hydrophobic dye with it ( Figure 4C ). By immersing Nile Redloaded hydrogel fibers with and without PA11 into PBS and measuring the fluorescence after gentle shaking overnight, we observed release of the hydrophobic small molecule only when PA11 was coloaded into the hydrogel ( Figure 4D−E) . To further illustrate this functionality in the context of neurobiology, we incubated primary rat dorsal root ganglion (DRGs; sensory neuronal structures) with Nile Red in the presence or absence of PA11 ( Figure 4F and S11). We found that, after a 24 h incubation with PA11, 96% of neurons were Nile Redpositive ( Figure 4F ). No dye was found in neuronal cytoplasm or nuclei in the absence of PA11 ( Figures 4G and S11) , suggesting that this polymer is sufficient for effective intraneuronal delivery of hydrophobic small molecule drugs. Our lab recently reported hydrogel-based neural probes utilizing a UV-cured poly(acrylamide)−alginate based hydrogel. 11 Although these devices enabled long-term electrophysiological and optogenetic studies, they had some limitations. Namely, assembly of these devices was labor intensive, required device fabrication in an oxygen-free environment due to oxygen-sensitive chemistry, and utilized neurotoxic radicals that had to be thoroughly eliminated through an extensive washing process prior to deployment in neuroscience applications. By contrast, SEED integration utilizing PU-PEG hydrogels is a robust and translatable method that overcomes all of these limitations and requires only a few minutes of active fabrication time. Additionally, unlike previous neural probes reported from our group that utilized passive cladding, 1,3,5,11,17 devices fabricated with SEED integration feature active cladding which can be coloaded with both small molecule drugs and drug nanocarriers for delivery in vivo. These properties render these modular hydrogel neural interfaces well posed for fundamental and translational biological research.

By leveraging a SEED integration approach that employs amphiphilic copolymers, we created modular hydrogel neural interfaces capable of optogenetics, electrophysiology, and microfluidic delivery. We demonstrated that these devices are capable of delivering a variety of cargo, including cellular therapies with a high viability at fast injection rates. By loading model drugs or nanomaterials into the hydrogel itself, we also demonstrated a separate drug delivery modality with a unique driving force and release profile. We further showed the potential to extend this concept to new directions, including The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acscentsci.1c00592.

Experimental details and supporting Figures S1−S11 (PDF)

Multifunctional fibers for simultaneous optical, electrical and chemical interrogation of neural circuits in vivo

Customizing MRI-Compatible Multifunctional Neural Interfaces through Fiber Drawing

One-step optogenetics with multifunctional flexible polymer fibers

Vivo Photopharmacology Enabled by Multifunctional Fibers

In situ electrochemical generation of nitric oxide for neuronal modulation

Characterization of Mechanically Matched Hydrogel Coatings to Improve the Biocompatibility of Neural Implants. Sci. Rep. 2017

Mechanical Characterization of Human Brain Tissue and Soft Dynamic Gels Exhibiting Electromechanical Neuro-Mimicry

Light-guiding hydrogels for cell-based sensing and optogenetic synthesis in vivo

Supramolecular Peptide Hydrogel-Based Soft Neural Interface Augments Brain Signals through a Three-Dimensional Electrical Network

Multimaterial and multifunctional neural interfaces: From surface-type and implantable electrodes to fiber-based devices

Multifunctional Hydrogel Hybrid Probes for Long-Term Sensing and Modulation of Neural Activity

Efficacy and Safety of the mRNA-1273 SARS-CoV-2 Vaccine

The Design and Manufacture of Medical Devices

Creating lowimpedance tetrodes by electroplating with additives

Neural recording and modulation technologies

Next-generation interfaces for studying neural function

In Vivo Light-Induced Activation of Neural Circuitry in Transgenic Mice Expressing Channelrhodopsin-2

Millisecond-timescale, genetically targeted optical control of neural activity

The role of the nucleus accumbens in psychiatric disorders

Functional recovery following traumatic spinal cord injury mediated by a unique polymer scaffold seeded with neural stem cells

2020, 21, 103. (25) Farah, S. Long-term implant fibrosis prevention in rodents and non-human primates using crystallized drug formulations

Off-target toxicity is a common mechanism of action of cancer drugs undergoing clinical trials

Optical Control of Cannabinoid Receptor 2-Mediated Ca2+Release Enabled by Synthesis of Photoswitchable Probes

Polymer-Based Dual-Responsive Self-Emulsifying Nanodroplets as Potential Carriers for Poorly Soluble Drugs