On the basis of previous work applying multimaterial fibers to optical neuromodulation ( 9 , 10 ), polycarbonate (PC; refractive index n = 1.58; glass transition temperature T g = 145°C; Young’s modulus E = 2.38 GPa) and cyclic olefin copolymer (COC; n = 1.52, T g = 158°C, E = 3.0 GPa) were selected as the respective core and cladding of the optical fiber ( 11 – 13 ). To enable neural recording while minimizing the device footprint, we deposited uniform 1-μm-thick conductive layers of AgNWs (diameter d = 70 nm; length L = 40 μm) ( Fig. 1 , A and C) over the COC cladding via dip coating from isopropanol (IPA) solutions with different concentrations ( Fig. 1 , A, D to F). Because the hydrophobicity of COC limited the adhesion and deposition of AgNWs from IPA ( 14 ), oxygen plasma treatment of the fibers was essential to enhance the uniformity of AgNW mesh layers ( 14 , 15 ). An AgNW coating was chosen as the electrode material because of its high conductivity and compatibility with facile solution-based processing ( 16 ). It was hypothesized that the mesh formed by AgNWs would be more resilient with respect to bending and stretching deformation ( 17 , 18 ) than a continuous metallic film of comparable thickness, because the latter is anticipated to develop cracks under strains commonly experienced in spinal cords ( 19 ). The entire structure was then encapsulated within a layer of polydimethylsiloxane (PDMS; n = ~1.41 to 1.47; thickness, 5 μm) ( 20 ) to minimize direct contact of AgNW with tissue and prevent surface oxidation and mechanical degradation ( Fig. 1 , A and D). The final device diameter ranged from 105 to 135 μm and was constrained by the dimensions of the structural fiber core (100 to 130 μm).

( A ) Illustration of the fiber probe fabrication. ( B ) Spool of a fiber with PC core and COC cladding. ( C ) Transmission electron microscopy (TEM) image of the AgNWs. ( D ) Cross-sectional image of the fiber probe. ( E ) Scanning electron microscopy image shows a portion of the ring AgNW electrode cross section. ( F ) Scanning electron microscopy image of the AgNW mesh on top of the fiber surface.

To fabricate probes suitable for electrophysiological recording and optical neuromodulation in rodent spinal cords, we combined two techniques. First, we used thermal drawing to produce a flexible optical fiber that also served as a structural core for the probe ( Fig. 1A ). Being versatile and scalable, thermal drawing can be applied to macroscale templates (preforms) composed of multiple materials. It also allows us to reduce final device dimensions by up to 200 times (fig. S1) while producing hundreds of meters of fiber in a single draw ( Fig. 1B ) ( 7 , 8 ). By tuning the stress during the drawing process, a range of feature dimensions can be achieved without compromising the cross-sectional geometry defined within the preform (fig. S1).

Optical and electrical properties of the nanowire-coated fiber probes

To match the mechanical properties of neural tissues, flexible polymer-based optical waveguides have been recently introduced to replace conventional rigid silica fibers (21–27). Waveguides composed of SU-8 and poly(methyl methacrylate) (PMMA) fabricated via a lithographic process have been used in the context of optogenetic neuromodulation (23), whereas PDMS and hydrogel-based devices have been applied to fluorescence measurements and optical control of gene expression (20, 26, 27). In addition to their transparency across the visible spectrum (fig. S2), the polymers PC and COC are compatible with the thermal drawing process. Consequently, the geometry of the device can be easily altered to fit the application (28). The difference between the refractive indices of PC and PDMS is 0.18. Although this should, in principle, be sufficient to sustain multimode transmission through the fiber even in the absence of COC cladding (fig. S2), direct coating of AgNWs onto the PC surface resulted in significant losses due to scattering and evanescent coupling of light into the plasmon modes of these nanomaterials (Fig. 2A) (29). Addition of the COC cladding reduced the losses from 2.5 to 1.9 dB/cm (Fig. 2A). Because of their flexibility, these probes were able to maintain transmission under extreme deformations (Fig. 2B), including their use as sutures (Fig. 2C).

Fig. 2 Optical characterization of flexible neural probes. (A) Normalized transmission at a wavelength λ = 473 nm as a function of length for fiber probes with and without COC cladding to separate AgNW mesh from the PC optical core with a diameter of 120 μm. (B) Transmission at λ = 473 nm for PC/COC/AgNW/PDMS fiber probes (core diameter, 120 μm) bent at 90° or 180° as radii of curvature (0.5 to 10 mm) displayed relative to straight probes. All scale bars and shaded areas represent SEM. n = 5 samples for each data point. (C) Image of a PC/COC/AgNW/PDMS fiber probe connected to a laser source, threaded through a needle, and used to create several stitches on fabric.

Electrophysiological recording during optogenetic neuromodulation is commonly accomplished by integrating conductive electrodes within the probes (23, 30–34). Conductive polymer composites exhibit high flexibility and biocompatibility (35), but are limited by high impedances on the order of megohms (9, 10, 36). Metallic electrodes deposited on polymer substrates have low impedance but are subject to cracking (19). Fractal and serpentine metallic electrodes defined via contact printing address the flexibility challenge but offer limited spatial resolution (5, 6, 37–39). AgNW meshes were previously used as stretchable interconnects in flexible electronic applications (16), and their composites have been recently applied to monitoring of cardiac function (35). We found that to reproducibly achieve meshes with low resistivity, the concentration of AgNWs in a dip-coating solution needed to exceed 4 mg/ml. At concentrations >6 mg/ml, the resistivity of the AgNW mesh was proportional to the concentration of the dip-coating solution (Fig. 3A). To account for anticipated changes in mesh morphology during deformation, we chose the lowest resistance mesh (9.37 × 10−4 ohm∙cm) for our probes. PDMS was selected as a protective coating for the conductive layer because of its low modulus (tens of kPa for 30:1 polymer to curing agent by weight) and low refractive index, ensuring confinement of light to the PC/COC core (40). Following dip coating with PDMS, the probe tips were cut orthogonally to the fiber axis, exposing thin conductive AgNW ring electrodes. Similar to solid metallic electrodes, the impedance of the probes at 1 kHz had a significant dependence on the contact area but less on the length. Mesh electrodes within 1- and 10-cm fiber probes exhibited impedance values of similar orders of magnitude (|Z PC/COC, 1 cm | = 50 ± 26 kΩ, |Z PC/COC, 10 cm | = 58 ± 21 kΩ; mean ± SEM), which indicated that, following evaluation in small rodents, this fabrication approach may, in principle, be scaled to applications in larger animals (Fig. 3B).

Fig. 3 Electrical characterization of flexible neural probes. (A) Resistivity of the mesh as a function of AgNW solution concentration. Inset: TEM images of the AgNW meshes deposited from solutions with 2, 6, and 10 mg/ml concentrations. AgNW mesh deposited from the 10 mg/ml solution was used for further characterization and in vivo evaluation. (B) Impedance spectra of the AgNW mesh electrodes deposited on 1-, 5-, and 10-cm-long fibers with 120-μm PC/COC cores. All scale bars and shaded regions represent SEM. n = 5 samples for each data point.

In addition to bending, AgNW mesh concentric electrodes were also resilient to stretching deformation. Using a fabrication process identical to the one outlined for PC/COC fibers, we thermally drew stretchable fibers composed of COC elastomer (COCE; n = 1.51; melting temperature T m = 84°C; E = 34 MPa). We chose COCE because of its low modulus and compatibility with a range of drawing parameters. To establish stable processing conditions, we introduced sacrificial PMMA cladding into the preform and then removed it with acetone following drawing (Fig. 4). The resulting pillow-shaped COCE fibers (cross-section width × height in the range from 125 μm × 100 μm2 to 250 μm × 200 μm2) were similarly treated with oxygen plasma, dip-coated with AgNWs, and encapsulated with PDMS. Consistent with lower optical transmission of COCE as compared to PC and COC, higher optical losses of 3.98 dB/cm were measured for AgNW-coated COCE core fibers (fig. S3).

Fig. 4 Mechanical and electrical characterization of stretchable neural probes. (A) Tensile tests performed for a thermally drawn COCE fiber and a COCE fiber probe coated with three layers of AgNW mesh and a protective PDMS cladding (COCE/AgNW/PDMS) (n = 5 devices). (B) Impedance spectra of the AgNW mesh electrodes deposited onto 200 × 200 μm2 COCE core fiber with lengths of 1, 5, and 10 cm. (C) Impedance of a three-layer AgNW mesh within COCE/AgNW/PDMS probes with core dimensions of 200 × 200 μm2 as a function of tensile strain. (D) Scanning electron microscopy images of the three-layer AgNW mesh deposited onto COCE fiber at 0, 10, and 20% strain. (E) Impedance of fiber probes characterized in (C) measured over five extension and release cycles. All scale bars represent SEM. n = 5 samples for each data point.

Because COCE is a rubbery material, these fibers could sustain up to 230% strain, which was reduced to 200% following the coating with AgNWs and PDMS (Fig. 4A). AgNW mesh electrodes coated onto 1- and 10-cm COCE fibers exhibited somewhat greater difference in impedance (|Z COCE, 1 cm | = 34 ± 17 kΩ, |Z COCE, 10 cm | = 162 ± 50 kΩ; mean ± SEM; Fig. 4B), as compared to their PC/COC analogs. However, the absolute values of impedance were still well within the range suitable for extracellular recordings even for 10-cm-long fibers.

We found that a single-layer AgNW mesh coating could only withstand strains of ~30% before losing conductivity due to the disruption of the conductive network (fig. S4). In contrast, electrodes composed of a three-layer AgNW mesh maintained low impedance at strains up to ~100% (Fig. 4C). This is consistent with scanning electron microscopy images that do not reveal any structural differences between the AgNW mesh–coated fibers subjected to 0, 10, and 20% strain (Fig. 4D). Repeated extension of the COCE/AgNW/PDMS fibers resulted in negligible hysteresis of the electrode impedance, indicating resilience of these devices to deformation (Fig. 4E). Because the spinal cord and peripheral nerves only experience strains up to ~12% (41), the low-impedance AgNW mesh–coated fibers provide arbitrarily scalable and stretchable alternatives to polymer composite and metallic electrodes.