Organic electro-scattering antenna

Abstract

Described are sensing devices including a substrate, an electrically conductive layer positioned on the substrate, and an electrochemical dopable material contacting the electrically conductive layer. The electrochemical dopable material may decrease its refractive index in the visible domain when subjected to a positive bias voltage and increase its refractive index in the visible domain when subjected to a negative bias voltage. The electrochemical dopable material may convert received electrical signals into visible light and scatter the visible light. Also described are sensing arrays including a plurality of the sensing devices and methods of monitoring electrical potentials in liquid media.

Description

[0001] Attorney Docket No. MML-081W001

[0002] ORGANIC ELECTRO-SCATTERING ANTENNA

[0003] GOVERNMENT INTEREST

[0004] This present disclosure was made with government support under grant

[0005] R01HL172065 from the National Institutes of Health. The government has certain rights in the present disclosure.

[0006] RELATED APPLICATIONS

[0007] This application claims the benefit under 35 U. S. C. § 119(e) of U. S. Provisional application serial number 63 / 726,791 filed December 2, 2024, the disclosure of which is incorporated by reference herein in it’s entirety.

[0008] BACKGROUND

[0009] The ability to probe electrical potentials with high recording site density and micrometer spatial resolution in a wet environment is a holy grail of biosensing. It can unlock unparalleled opportunities for enhanced multiplexing, throughput, and spatial resolution across a wide range of applications, including biomolecule detection [Bellin, D. L. et al. Nat. Commun. 5, 3256 (2014); Mensack, M. M. et al. Lab. Chip 13, 208-211 (2013)], DNA sequencing [Rothberg, J. M. et al. Nature 475, 348-352 (2011)], impedance sensing [Bounik, R. et al. BME Front. 2022, (2022)], functional imaging [Abbott, J. et al. Lab. Chip 22, 1286-1296 (2022)], and electrophysiology [Abbott, J. et al. Nat. Biomed. Eng. 4. 232-241 (2020); Abbott, J. et al. Nat. Nanotechnol. 12, 460-466 (2017); Müller, J. et al. Lab. Chip 15, 2767–2780 (2015)]. Over recent decades, organic electronic materials (OEMs) have led to remarkable advances toward this goal [Kaushal, J. B. et al. Biosensors 13, 1-48 (2023)]. Their unique electrochemical properties in aqueous electrolytes make them materials of choice for developing highly sensitive reporters of bioelectrical signals.

[0010] For instance, OEMs’ mixed ionic and electronic conductivities enabled neuroprosthetics with enhanced tissue-electrode interfaces, facilitating electrical recording and stimulation [Wilks, S. J. et al. Front. Neuroengineering 2, 591 (2009); Go, G. T. et al. Adv. Mater. 34, 2201864 (2022); Airaghi Leccardi, M. J. I. & Ghezzi, D. Healthc. Technol. Lett. 7, 52-57 (2020)]. In addition, electrical modulation of OEM doping levels led to the development of organic electrochemical transistors (OECTs) [White, H. S. et al. J. Am. Chem. Soc. 106, 5375-5377 (1984)], a class of highly sensitive biosensors extensively used in various applications, particularly electrophysiology [Jimbo. Y. et al. Proc. Natl. Acad. Sci. U. S. A. 118, 1-8 (2021); Khodagholy, D. et al. Nat. Commun. 4, (2013); Rivnay, J. et al. Sci. Adv. 1, 1-5 1

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[0012] (2015)], impedance sensing (Jimison, L. H. et al. Adv. Mater. 24, 5919-5923 (2012); Yao, C. et al. Adv. Mater. 25, 6575-6580 (2013); Faria, G. C. et al. MRS Commun. 4. 189-194 (2014); Rivnay, J. et al. Appl. Phys. Lett. 106, 43301 (2015); Bonafe, F. et al. Nat. Commun.

[0013] 13, 1-9 (2022)], and analyte detection [Zhu, Z. T. et al. Chem. Commun. 4, 1556-1557 (2004); Pappa, A. M. et al. Adv. Healthc. Mater. 5, 2295-2302 (2016); Liu, H. et al. Adv. Sci.

[0014] 2305347 (2024)]. Interestingly, the dependence of OEM optical absorbance properties on doping levels — and therefore on voltage — permitted wireless monitoring of cellular electrophysiological activities [Alfonso, F. S. et al. Proc. Natl. Acad. Sci. U. S. A. 117, 17260-17268 (2020); Zhou, Y. et al. J. Am. Chem. Soc. 2022, (2022)]. Based on this principle, the electrochromic optical recording device (ECORE) enabled high-sensitivity single-point measurements with single-cell resolution. More generally, OEM doping levels affect not only their optical absorbance but also their complex refractive index [Dingler, C. et al. Macromolecules 55, 1600–1608 (2022)]. Recently, electrically switchable optical diffraction gratings were developed, leveraging this principle [Doshi, S. et al. Nanophotonics 1-10 (2024)]. When coated onto plasmonic nanoantennas, OEMs were also demonstrated to enable electrical modulation of the plasmonic resonance [Habib, A. et al. Sci. Adv. 5, eaav9786 (2019); Jeon, J.-W. et al. Chem. Mater. 28, 2868-2881 (2016); Locarno, M. & Brinks, D. Am. J. Phys. 91, 538 (2023)]. Arrays of these electro-plasmonic nanoantennas enabled wireless probing of cardiomyocyte activity from a single recording site of a few hundred micrometers in diameter. Beyond electro-plasmonic antennas, electrochemically-modulated plasmonic nanoantennas fully composed of OEMs were successfully developed. Notably, poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOTPSS) displays a metallic behavior at high doping levels with a negative real dielectric permittivity, enabling plasmonic resonance in the infrared domain [Karst, J. et al. Science 374, 612-616 (2021)]. Upon de-doping, it exhibits the characteristics of a dielectric (i.e., positive real dielectric permittivity), quenching the plasmonic resonance of the antenna. Arrays of polymer plasmonic nanoantennas were recently implemented in the context of integrated photonics to modulate light beam intensities electrically [Chen, S. et al. Nat. Nanotechnol. 15, 35-40 (2020); Karki, A. et al. Commun. Mater. 3, 1-8 (2022); Karki, A. et al. Adv. Mater. 34, (2022)].

[0015] While OEMs have led to remarkable advances in biosensing, the spatial resolution and recording site density of organic biosensors remain problematic. For OECTs, and more generally for microelectrode-based sensors, the spatial resolution is typically limited to a few hundred micrometers and the number of recording sites to a few tens because of the

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[0018] conductive traces connecting each sensing unit to its electrical instrumentation. Nevertheless, studying complex biological systems requires enhanced spatial resolution to accurately recapitulate physiological processes occurring at subcellular scales. Specifically, singlemicrometer resolution across thousands of recording sites is necessary to enable functional readouts with spatial context and open new avenues in biosensing. Y et, none of the existing technologies have enabled such features.

[0019] SUMMARY

[0020] The present disclosure provides, among other things, a sensing device [sometimes referred to as an organic electro-scattering antenna (OCEAN)] configured to wirelessly probe small potential fluctuations using visible light in a physiological solution. In some embodiments, an array of sensing devices leverages the unique dependence of PEDOTPSS light scattering on voltage to enable wireless sensing of millivolt electrical signals, potentially from thousands of recording sites and with micrometer resolution. These features make sensing devices of the present disclosure unique compared to state-of-the-art technologies and position them as a central class of devices to perform multisite probing of bioelectrical signals with high spatial resolution, potentially opening new avenues in spatial biology.

[0021] A theoretical model was developed to describe how external potentials applied across sensing devices including PEDOT: PSS affect their doping level and complex permittivity in a physiological environment. This model was leveraged to study how changes in permittivity affect their scattering spectrum and dynamic signal in the visible domain. Furthermore, a robust and reliable nanofabrication process (combining convention microfabrication techniques with next-generation focused-ion beam lithography) was established to manufacture sensing arrays of different dimensions and their electro-optic modulation characteristics were studied in terms of sensitivity, noise, signal-to-noise ratio (SNR), limit of detection, time constant, and long-term stability. The present disclosure demonstrates a single sensing device can be used to monitor 100 mV voltage pulses wirelessly with SNRs up to 48 and a limit of detection approaching 2.5 mV at millisecond time scales and 5 μm spatial resolution. Furthermore, sensing devices exhibited milliseconds time constants and exceptional long-term stability, enabling continuous optical recordings for at least 10 hours.

[0022] By leveraging the dependence of organic electronic material scattering properties on their electrochemical doping levels, sensing devices of the present disclosure offer unprecedented sensing characteristics in terms of spatial resolution (5 μm) and recording density (4.106 cm-2). This advancement opens a new range of readout capabilities.

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[0025] Sensing devices of the present disclosure offer beneficial spatial resolution and recording site density’, enabling detailed functional mapping at the cellular and subcellular levels.

[0026] A first aspect of the present disclosure relates to a sensing device comprising: a substrate; an electrically conductive layer positioned on the substrate; and an electrochemical dopable material contacting the electrically conductive layer, wherein the electrochemical dopable material: has a refractive index in the visible domain that changes based on whether the electrochemical dopable material is subjected to a positive or negative bias voltage; converts received electrical signals into visible light; and scatters the visible light.

[0027] In some embodiments of the first aspect, the substrate is optically transparent.

[0028] In some embodiments of the first aspect, the substrate comprises glass.

[0029] In some embodiments of the first aspect, the substrate is opaque.

[0030] In some embodiments of the first aspect, the electrically conductive layer comprises at least one of metal and conductive polymer.

[0031] In some embodiments of the first aspect, the metal comprises metal oxide.

[0032] In some embodiments of the first aspect, the metal oxide comprises indium tin oxide. In some embodiments of the first aspect, the electrochemical dopable material comprises at least one organic semiconductor and at least one electrolyte.

[0033] In some embodiments of the first aspect, the at least one organic semiconductor comprises at least one conductive polymer.

[0034] In some embodiments of the first aspect, the at least one conductive polymer comprises poly(3,4-ethylenedioxythiophene).

[0035] In some embodiments of the first aspect, the at least one electrolyte comprises at least one polystyrene.

[0036] In some embodiments of the first aspect, the at least one polysty rene comprises polystyrene sulfonate.

[0037] In some embodiments of the first aspect, the at least one organic semiconductor and at least one electrolyte are present at a ratio in the range of 1: 1 to 1: 20 v / v.

[0038] In some embodiments of the first aspect, the sensing device further comprises an electrically insulating layer.

[0039] In some embodiments of the first aspect, the electrically conductive layer is positioned between the substrate and the electrically insulating layer.

[0040] In some embodiments of the first aspect, the electrochemical dopable material contacts both the electrically conductive layer and the electrically insulating layer.

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[0043] In some embodiments of the first aspect, the electrically insulating layer comprises at least one of silicon-containing material, electrically insulating polymer, and ceramic.

[0044] In some embodiments of the first aspect, the silicon-containing material comprises silicon nitride.

[0045] In some embodiments of the first aspect, the electrochemical dopable material comprises a cap portion with a stem portion extending therefrom; and the stem portion extends into and along an aperture in the electrically insulating layer.

[0046] In some embodiments of the first aspect, the cap portion has a circular surface extending along the electrically insulating layer.

[0047] In some embodiments of the first aspect, the circular surface has a diameter of 1 pm. In some embodiments of the first aspect, the electrically conductive layer contacts an end surface of the stem portion.

[0048] In some embodiments of the first aspect, the end surface is circular and has a diameter in the range of 10 nm to 1 mm.

[0049] In some embodiments of the first aspect, the end surface is circular and has a diameter of 250 nm.

[0050] A second aspect of the present disclosure relates to a method of monitoring electrical potential of a biological material in a liquid medium, comprising: contacting the liquid medium with the sensing device of the first aspect; and measuring the electrical potential of the biological material the liquid medium.

[0051] In some embodiments of the second aspect, the liquid medium is a physiological medium.

[0052] In some embodiments of the second aspect, the biological material comprises a cell. In some embodiments of the second aspect, the cell is a neuron.

[0053] In some embodiments of the second aspect, the cell is a cardiac cell.

[0054] In some embodiments of the second aspect, the biological material comprises a protein.

[0055] In some embodiments of the second aspect, the biological material comprises a small molecule.

[0056] In some embodiments of the second aspect, the biological material comprises a ribonucleic acid (RNA) molecule.

[0057] In some embodiments of the second aspect, the biological material comprises a deoxyribonucleic acid (DNA) molecule.

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[0060] In some embodiments of the second aspect, measuring the electrical potential of the biological material comprises directing light at the sensing device at an angle in the range of 61° to 90° with respect to vertical of the electrically conductive layer.

[0061] In some embodiments of the second aspect, the light has a wavelength in the range of 380 nm to 700 nm.

[0062] In some embodiments of the second aspect, the light has a wavelength of 637 nm. In some embodiments of the second aspect, measuring the electrical potential of the biological material comprises using dark-field microscopy.

[0063] In some embodiments of the second aspect, measuring the electrical potential of the biological material comprises measuring visible light scattered by the electrochemical dopable material of the sensing device using a microscope.

[0064] A third aspect of the present disclosure relates to a sensing array comprising a plurality of the sensing device of the first aspect.

[0065] A fourth aspect of the present disclosure relates to method of monitoring electrical potentials in a liquid medium, comprising: contacting the liquid medium with the sensing array of the third aspect; and measuring the electrical potential of the liquid medium.

[0066] In some embodiments of the fourth aspect, the liquid medium is a physiological medium.

[0067] In some embodiments of the fourth aspect, the physiological medium comprises a cell.

[0068] In some embodiments of the fourth aspect, the cell is a neuron.

[0069] In some embodiments of the fourth aspect, the cell is a cardiac cell.

[0070] In some embodiments of the fourth aspect, the physiological medium comprises a protein.

[0071] In some embodiments of the fourth aspect, the physiological medium comprises a small molecule.

[0072] In some embodiments of the fourth aspect, the physiological medium comprises a ribonucleic acid (RNA) molecule.

[0073] In some embodiments of the fourth aspect, the physiological medium comprises a deoxyribonucleic acid (DNA) molecule.

[0074] In some embodiments of the fourth aspect, measuring the electrical potential of the liquid medium comprises directing light at the sensing array at an angle in the range of 61° to 90° with respect to vertical of the electrically conductive layer.

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[0077] In some embodiments of the fourth aspect, the light has a wavelength in the range of 380 nm to 700 nm.

[0078] In some embodiments of the fourth aspect, the light has a wavelength of 637 nm. In some embodiments of the fourth aspect, measuring the electrical potential of the liquid medium comprises using dark-field microscopy.

[0079] In some embodiments of the fourth aspect, measuring the electrical potential of the liquid medium comprises measuring visible light scattered by the electrochemical dopable materials of the sensing array using a microscope.

[0080] BRIEF DESCRIPTION OF THE DRAWINGS

[0081] Fig. 1 A-B provides schematic diagrams illustrating an embodiment of the present disclosure to show the overall concept of organic electro-scattering antenna (OCEAN). Fig.

[0082] 1A is a schematic cross-section representation of a single OCEAN under positive voltage bias. Under this condition, holes are injected from the ITO electrode into the PEDOT chains, increasing its doping level and releasing the solvated cations that were previously compensating for the fixed sulfonate ion charges in the PSS strands (insert, top). The augmented charge density in PEDOT decreases its refractive index, therefore reducing its scattering properties. Fig. IB is a schematic image illustrating how under negative voltage bias, solvated cations are injected into the PEDOT: PSS from the electrolyte and electrostatically compensate for the fixed sulfonate anion charges, subsequently enabling hole extraction through the ITO electrode (insert, bottom). In its dedoped state, PEDOT shows an augmented refractive index and, consequently, enhanced scattering properties. In each configuration, the light scattered by OCEANs originates from an evanescent wave generated by total internal reflection (TIR). Schematic representations in the inserts were adapted with permission from Proctor et al. J. Polym. Sci. Part B Polym. Phys. 54, 1433-1436 (2016).

[0083] Fig. 2A-H provides graphs demonstrating modeling of single OCEAN electro-optic characteristics in one embodiment of the present disclosure. Fig. 2A shows real (ai) and Fig.

[0084] 2B shows imaginary (82) part of PEDOT: PSS permittivity for different bias voltages VB in PBS versus Ag / AgCl reference electrode. Fig. 2C and D show simulated scattering cross-section (σsc) spectra for a single PEDOT: PSS OCEAN with a cap diameter of 750 nm and 1500 nm, respectively, in response to voltage biases ranging from 0.2 V to -0.8 V in PBS. Fig. 2E shows scattering cross-section and Fig. 2F shows sensitivity against voltage biases simulated for single OCEANs with cap diameters of 750 nm (red) and 1500 nm (blue) at a wavelength of 640 nm. Fig. 2G is a graph showing dynamic voltages anticipated across 750

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[0087] nm and 1500 nm in diameter OCEANs (VPEDOT:PSS,750nm and VPEDOT:PSS,1500nm, respectively) in response to an applied -100 mV voltage pulse (V stimulation) with respect to a -0.45 V bias voltage. Fig. 2H is a graph showing dynamic electro-optic modulation simulated in response to the stimulation voltage pulse described in Fig. 2G for the two different OCEAN geometries tested.

[0088] Fig. 3A-D provides a schematic diagram and images illustrating a non-limiting example of nanofabrication of OCEANs. Fig. 3A is a schematic representation highlighting the two main steps of nanofabrication process flow developed to manufacture OCEANs. Fig.

[0089] 3B provides scanning electron microscope (SEM) images of nanostructures following FIB patterning (top) and PEDOT: PSS electrodeposition (bottom). Fig. 3C is an SEM micrograph of an array of OCEANs. Fig. 3D is a brightfield microscopy image of an ITO-SiNx pad decorated with an array of OCEANs on an otherwise SiNx-SU8 passivated substrate.

[0090] Fig. 4A-H provides graphs illustrating electro-optic characterization of an embodiment of OCEANs. Fig. 4A is an optical and electrical voltammogram. Relative variation of scattered light intensity (Z score) emitted by single OCEANs of different diameters in response to voltage sweeps (top, N = 16, mean ± standard deviation).

[0091] Corresponding electrical current measured at the array level for each geometry tested (bottom). The insert shows a schematic representation of the experimental setup used for electro-optic characterization. Fig. 4B shows Z score modulation in response to voltage pulses of amplitudes between -100 mV and 100 mV measured for single OCEANs (N = 16, mean ± standard deviation). Box charts summarizing the sensitivity (Fig. 4C), signal-to-noise ratio for 100 mV voltage pulses (Fig. 4D); optical noise (Fig. 4E), and limit of detection (Fig.

[0092] 4F) for single OCEANs of different geometries and operating at voltage biases ranging from 0 V to -0.6 V vs. Ag / AgCl reference electrode (N=16). Fig. 4G shows Z score temporal evolution over 10 hours of electrical stimulation with 500 ms-long and -100 mV voltage pulses at 1 Hz (VB = -0.3 V, N=16, IED = 45 s). Fig. 4H is a box chart summarizing the influence of OCEAN geometry and operating voltage on time constants (N=6).

[0093] Fig. 5 is a schematic diagram showing doping and dedoping of PEDOT: PSS in anon-limiting embodiment of the present disclosure. The image shows the PEDOT: PSS dedoping process when a negative voltage bias is applied across the organic electro-scattering antennaelectrolyte interface. Adapted with permission from Proctor et al. J. Polym. Sci. Part B Polym. Phys. 54, 1433-1436 (2016).

[0094] Fig. 6A-H provides graphs showing electrochemical modulation of PEDOTPSS complex permittivity in a non-limiting embodiment of the present disclosure. Fig. 6A shows 8

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[0096] real and Fig. 6B shows imaginary part of the PEDOT: PSS complex permittivity in its fully doped (red curve) and fully undoped (black curve) state, fitted from Dingler, et al., Macromolecules 55, 1600–1608 (2022). The fully doped state is fitted by a single Drude model, while the fully undoped state by a combination of two Lorentz and two Tauc-Lorentz models (dashed lines). Fig. 6C shows real and Fig. 6D shows imaginary parts of the PEDOT: PSS complex permittivity for intermediate doping levels a. Fig. 6E shows real and Fig. 6F shows imaginary parts of the PEDOT: PSS complex permittivity for different voltage bias levels VB in phosphate-buffered saline (PBS) solution. Fig. 6G shows variations of the PEDOT: PSS real permittivity with respect to wavelength for different voltage bias levels. Fig. 6H shows relative variation of scattered light intensity (Z score) collected from individual PEDOT: PSS organic electro-scattering antennas in response to a linear voltage sweep from 0.7 V to -0.8 V in PBS vs Ag / AgCl (N=16, mean ± standard deviation).

[0097] Assuming that the optical signal variations are solely caused by a variation of the PEDOT doping level, the link between voltage bias and doping level can be established by fitting the optical signal with a sigmoidal function.

[0098] Fig. 7A-B provides schematic diagrams and a photographic image of Nanofabrication and interfacing of OCEANs in a non-limiting embodiment of the present disclosure. Fig. 7A shows a nanofabrication process flow developed to manufacture OCEANs. Fig. 7B is an optical image of the device following electrical and mechanical interfacing with a custom-designed PCB. A glass ring is mounted on top to facilitate experiments in liquid.

[0099] Fig. 8A-J provides images and graphs illustrating OCEAN diameter analysis by optical microscopy in a non-limiting embodiment of the present disclosure. Optical microscopy image of an array of OCEANs following 30 s (Fig. 8A); 45 s (Fig. 8B); and 60 s (Fig. 8C) of PEDOTPSS electrodeposition. The images were made with a 60x water dipping objective. Fig. 8D-F provides corresponding thresholded images established from Fig. 8A-C and used to determine the cap diameters and variability. Fig. 8G-I shows distributions of cap diameters extracted from thresholded images in Fig. 8D-F for the three different electrodeposition times tested. Fig. 8F is a box chart summarizing the relationship between PEDOT: PSS electrodeposition time and cap diameter (N = 130). These OCEANs were grown through SiNx openings of 250 nm in diameter.

[0100] Fig. 9A-C shows graphs of electrochemical impedance spectroscopy on OCEANs in a non-limiting embodiment of the present disclosure. Impedance module (top row) and phase (bottom row) of PEDOT: PSS OCEAN arrays (2 x 16 x 16 = 512 OCEANs), versus frequency, for different voltage biases in PBS and cap diameters ranging from 0.7 pm (trn =

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[0103] 30s) in Fig. 9A; 1.4 gm (tED = 45s) in Fig. 9B: to 1.8 pm (tED = 60s) in Fig. 9C. Dots and solid curves represent the experimental data and theoretical fit based on the electrical equivalent circuit shown in Fig. 9 A, respectively.

[0104] Fig. 10A-F provides a schematic diagram and graphs illustrating electrical equivalent circuit of single OCEAN-electrolyte interfaces in a non-limiting embodiment of the present disclosure. Fig. 10A is a schematic representation of the electrical equivalent circuit of an OCEAN-electrolyte interface. Plots showing B) Qo (Fig. 10B); n (Fig. 10C), Rs + RPEDOT:PSS (Fig. 10D), and Cstray (Fig. 10E) with respect to voltage bias amplitudes for small (IED = 30 s, diameter 0.7 pm), intermediate (tED = 45 s, diameter 1.4 pm), and large (IED = 60 s, diameter 1.8 pm) cap diameters. Numerical values extracted from the EIS experimental dataset shown in Fig. 9. Fig. 10F shows time constants computed by multiplying RPEDOTPSS + Rs with CPEDOT: PSS for each condition investigated.

[0105] Fig. 11 is a graph of relative irradiance spectra of PEDOTPSS OCEAN arrays for different bias voltages in PBS in a non-limiting embodiment of the present disclosure.

[0106] Fig. 12A-C provides photographs illustrating total internal reflection dark-field microscopy in a non-limiting embodiment of the present disclosure. Fig. 12A is an optical image showing an experimental setup developed to characterize the electro-optic characteristics of OCEANs. Fig. 12B is a picture of the prism-based total internal reflection illumination module. Fig. 12C is an image of an OCEAN chip mounted on the experimental setup and electrically connected to a potentiostat. The diascopic green light is used to image the device using reflected brightfield microscopy.

[0107] Fig. 13A-F provides optical (top) and electrical (bottom) voltammograms measured in a non-limiting embodiment of the present disclosure from OCEANs of small (tED = 30 s, diameter 0.7 pm) in Fig. 13A, intermediate (tED = 45 s, diameter 1.4 pm) in Fig. 13B; and large (IED = 60 s, diameter 1.8 pm) in Fig. 13C dimensions. The optical traces were constructed by averaging the individual voltammogram of each OCEAN composing the array (N = 16, mean ± standard deviation). The electrical voltammogram displays the current going through the entire array (512 OCEANs). The black curve represents the first cycle, while the red represents the second one. Potentials are swept with respect to an Ag / AgCl reference electrode. Total internal reflection dark-field microscopy images of small in Fig. 13D; intermediate in Fig. 13E; and large in Fig. 13F OCEAN arrays under a voltage bias of -0.8V vs Ag / AgCl.

[0108] Fig. 14A-H provides graphs showing electro-optic modulation, in a non-limiting embodiment of the present disclosure, of single OCEANs (tED = 30 s, diameter 0.7 pm) in 10

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[0110] response to voltage pulses with amplitude ranging from -100 mV (black), -50 mV, -25 mV, 25 mV. 50 mV, to 100 mV (red) in PBS. Operating biases are swept from 0V in Fig. 14A; -0.1 V in Fig. 14B; -0.2 V in Fig. 14C; -0.3 V in Fig. 14D; -0.4 V in Fig. 14E; -0.5 V in Fig.

[0111] 14F; to -0.6 V in Fig. 14G with respect to Ag / AgCl reference electrode. Fig. 14H shows optical Z score with respect to voltage pulse amplitude for the different operating biases tested. (N=16, mean± standard deviation).

[0112] Fig. 15A-H provides graphs showing electro-optic modulation, in a non-limiting embodiment of the present disclosure, of single OCEANs (ten = 45 s, diameter 1.4 pm) in response to voltage pulses with amplitude ranging from -100 mV (black), -50 mV, -25 mV, 25 mV, 50 mV, to 100 mV (red) in PBS. Operating biases are swept from 0V in Fig. 15A; - 0.1 V in Fig. 15B; -0.2 V in Fig. 15C, -0.3 V in Fig. 15D; -0.4 V in Fig. 15E; -0.5 V in Fig.

[0113] 15F; to -0.6 V in Fig. 15G with respect to Ag / AgCl reference electrode. Fig. 15H shows optical Z score with respect to voltage pulse amplitude for the different operating biases tested. (N=16, mean± standard deviation).

[0114] Fig. 16A-H provides graphs showing electro-optic modulation, in anon-limiting embodiment of the present disclosure, of single OCEANs (tED = 60 s, diameter 1.8 pm) in response to voltage pulses with amplitude ranging from -100 mV (black), -50 mV, -25 mV, 25 mV, 50 mV, to 100 mV (red) in PBS. Operating biases are swept from 0V in Fig. 16A; -0.1 V in Fig. 16B; -0.2 V in Fig. 16C; -0.3 V in Fig. 16D; -0.4 V in Fig. 16E; -0.5 V in Fig.

[0115] 16F; to -0.6 V in Fig. 16G with respect to Ag / AgCl reference electrode. Fig. 16H shows optical Z score with respect to voltage pulse amplitude for the different operating biases tested. (N=16, mean± standard deviation).

[0116] Fig. 17A-C shows graphs illustrating dynamic characteristics of OCEANs in anon-limiting embodiment of the present disclosure. Optical traces and exponential fits collected from single OCEANs with diameters ranging from 0.7 pm (tED = 30 s) in Fig. 17A; 1.4 pm (IED = 45 s) in Fig. 17B; to 1.8 pm (tED = 60 s) in Fig. 17C for operating biases between 0 V and -0.6 V with respect to Ag / AgCl reference electrode (N=6, mean ± standard deviation).

[0117] Fig. 18A-F provides graphs showing long-term stability of OCEANs in a non-limiting embodiment of the present disclosure. Fig. 18A-F shows electro-optic modulation of single OCEANs (1.4 pm in diameter) following 10 min and up to 10 hours of electrical stimulation with 500 ms-long and - 100 mV voltage pulses at 1 Hz (VB = -0.3 V, N=16, mean ± standard deviation).

[0118] Fig. 19A-D provides schematic diagrams and graphs illustrating modeling of cardiomyocyte action potential recording with OCEANs in a non-limiting embodiment of the 11

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[0120] present disclosure. Fig. 19A is a schematic representation of the cardiomyocyte-OCEAN interface according to a Luo Rudy model [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)]. in a pseudo current clamp configuration, where an ideal current source of infinite input impedance is connected between the OCEAN and the bath. Fig. 19B shows simulated voltage traces across the PEDOTPSS OCEAN (VPEDOTPSS) and cell membrane (Vm) in response to a 20 ms long stimulation current pulse (Istim) of 0.5 nA (left panel) and 1 nA (right panel) amplitude. 1 nA is enough to trigger the generation of an action potential in the ventricular cardiomyocyte, but this electrophysiological potential is not transferred across the PEDOT PSS structure. Fig. 19C is a schematic illustration of the cardiomyocyte-OCEAN interface electrical equivalent circuit in a pseudo voltage clamp configuration, where an ideal voltage source is connected between the OCEAN and the bath to apply the optimal bias voltage. Fig. 19D shows computed voltage traces VPEDOTPSS and Vm in response to the applied voltage bias Vmas when a cardiomyocyte fires an action potential. In this case, the action potential is generated by an additional transmembrane current pulse (10 ms long, 0.4 nA amplitude), mimicking the contribution of the neighboring cells through gap junctions. For Fig. 19A and C, CPEDOT:PSS represents the capacitive charge transfer between the PEDOTPSS and the bath, RCT is the charge transfer resistance, Rj is the junctional resistance of the cell membrane, Rseai is the seal resistance, and Cm is the non-junctional capacitance of the cell membrane. As defined in the Luo Rudy model [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)] fast sodium currents, slow inward currents, time-dependent potassium currents, time-independent potassium currents, plateau potassium currents, and background currents are represented by their respective transconductance gNa(t, V), gsi(t, V). gK(t, V), gKi(V), gKp(V), and gB and reversal potentials E\a. Esi, EK, EKI, and EKP. IGJ represent the stimulation current originating from the neighboring cells through gap junctions and initiating the action potential. The numerical values Rseai = 200 Mil and Rj = 500 MQ were used in this model [Spira, M. E. & Hai, A. Nat. Nanotechnol. 8, 83-94 (2013)]. See Example 10 herein for the numerical values of the remaining components of the circuit.

[0121] Fig. 20A-E provides schematic diagrams and graphs illustrating modeling the electrooptic characteristics of a single OCEAN with cells in a non-limiting embodiment of the present disclosure. Fig. 20A is a. schematic illustration of the simulated OCEAN-cell interface under total internal reflection illumination. Fig. 20B is a simulated spatial distribution of the electric field enhancement at the cell-antenna interface (E: Electric field, Eo: electric field of the incident illumination). Fig. 20C is a graph showing dynamic voltage anticipated across a

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[0124] 1500 nm in diameter OCEAN (VPEDOT: PSS) in response to an applied -100 mV voltage pulse (V stimulation) with respect to a -0.4 V bias voltage. Fig. 20D shows absolute and Fig. 20E shows relative dynamic electro-optic modulation of the system scattering cross-section (σsc) simulated in response to the stimulation voltage pulse described in Fig. 20C with or without a cell covering the OCEAN. The relative variation of scattering cross section is defined as <7SC-CTSC,t=0,..,....

[0125] - o‘sc,t=0. ' wherein σsc,t=o is the scattering&cross-section at the time t = 0 s.

[0126] DETAILED DESCRIPTION

[0127] Sensing Devices

[0128] With reference to Fig. 1, a sensing device [sometimes referred to herein as an Organic Electro-Scattering Antenna (OCEAN)] of the present disclosure includes a substrate 102an electrically conductive layer 106, an electrochemical dopable material 108. and optionally an electrically insulating layer 104. When the electrically insulating layer 104 is present, the electrically conductive layer 106 may be positioned between the substrate 102 and the electrically insulating layer 104. Moreover, when the electrically insulating layer 104 is present, the electrochemical dopable material 108 may be positioned to contact both the electrically insulating layer 104 and electrically conductive layer 106. When the electrically insulating layer 104 is not present, the electrochemical dopable material 108 is positioned to at least contact the electrically conductive layer 106, which is positioned on the substrate 102.

[0129] In some embodiments, the electrochemical dopable material 108 may include a cap portion 110 with a stem portion 112 extending therefrom. In some embodiments, the stem portion 112 may extend into and along an aperture in the electrically insulating layer 104, when present. In some embodiments, an end surface of the stem portion 112 may contact the electrically conductive layer 106. For example, the electrically conductive layer 106 may be positioned between the electrically insulating layer 104 and the substrate 102, and electrochemical dopable material may undergo electrodeposition within the aperture of the electrically insulating layer 104 to form the electrochemical dopable material 108 having the cap portion 110 and stem portion 112.

[0130] While the cross-section of the stem portion 112 is not to be limited to any particular shape, in some embodiments the stem portion 112 may have a circular cross-section (e.g., corresponding to a circular cross-section of the aperture of the electrically insulating layer 104 within which it is positioned). The circular cross-section may have a diameter in the range of about 10 nm to about 1 mm. In some embodiments, the circular cross-section may

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[0133] have a diameter of about 250 nm. However, the present disclosure envisions the stem portion 112 being of any size and shape that permits the electrochemical dopable material 108 to scatter light.

[0134] In embodiments where the electrochemical dopable material 108 includes the cap portion 110 and stem portion 112, the stem portion 112 may extend from a surface of the cap portion 110 (e.g., that extends along a surface of the electrically insulating layer 104). In some embodiments, the surface of the cap portion 110 may be circular, although other shapes are envisioned. In some embodiments, the circular surface may have a diameter of about 1 pm. However, one skilled in the art will appreciate that other shapes and sizes of the surface of the cap portion 110 (e.g., that extends along the surface of the electrically insulating layer 104) are possible. The present disclosure envisions the cap portion 110 being of any size and shape that permits the electrochemical dopable material 108 to scatter light.

[0135] In some embodiments, a sensing device of the present disclosure may be configured such that the electrochemical dopable material 108 is supported by an optically transparent and electrically conductive material-coated substrate patterned on an electrically insulating layer. In some embodiments, a sensing device of the present disclosure may be configured such that the electrochemical dopable material 108 is supported by an optically transparent and electrically conductive indium tin oxide (ITO)-coated glass substrate patterned on an electrically insulating silicon nitride (SiNx) layer.

[0136] A sensing device may scatter the incident light from a microscope. As such, a sensing device of the present disclosure may be configured without any wires connected thereto (i.e., wires used to measure an output of the sensing device).

[0137] Sensing Arrays

[0138] Two or more sensing devices of the present disclosure may be configured into a sensing array. For example, the electrically conductive layer 106 may be positioned between the electrically insulating layer 104 and the substrate 102, the electrically insulating layer 104 may have two or more apertures therein, and separate instances of the electrochemical dopable material 108 may be positioned in and about the two or more apertures. For further example, when the electrically insulating layer 104 is not present, the electrically conductive layer 106 may be positioned on the substrate 102 and separate instances of the electrochemical dopable material 108 may be positioned on the electrically conductive layer 106. Electrodeposition may be used to form separate instances of the electrochemical dopable material 108.

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[0141] As noted above, the electrochemical dopable material 108 of a sensing device may be smaller in cross-section than a cell. As such, a sensing array may be used to measure various locations of electric activity of a single cell. A sensing array may be used to measure the electric activity' of specific portions of a cell, with each sensing device of the array measuring the electric activity of a different portion of the cell.

[0142] As noted previously, a sensing device may scatter the incident light from a microscope. As such, a sensing array of the present disclosure may be configured without any wires connected thereto (i.e., wires used to measure outputs of the sensing devices of the sensing array).

[0143] Substrates

[0144] In some embodiments, the substrate 102 may be optically transparent, meaning light having a visible spectrum wavelength is capable of passing through the substrate 102. In some embodiments, an optically transparent substrate may include or be made substantially or entirely of one or more of glass, quartz, sapphire, and one or more polymers (e.g., polystyrene).

[0145] In other embodiments, the substrate 102 may be optically opaque, meaning light having a visible spectrum wavelength is incapable of passing through the substrate 102. In such embodiments, illumination and imaging may be performed from the top. As described elsewhere herein, illumination may come from the bottom of the sample being analyzed, go through the sample, and excite the electrochemical dopable material 108. In such situations, the light scattered by the electrochemical dopable material 108 is collected from the top of the sample (e.g., the microscope objective is on top of the sample).

[0146] In some embodiments, an optically opaque substrate may include or be made substantially or entirely of one or more of ceramic material(s) (e.g., silicon) and polymer(s). If using an opaque substrate, illumination cannot go through the substrate to excite the electrochemical dopable material 108 but, instead, may come from the top of the sample. The incident light can come from the same microscope objective used for imaging or could also be sent through a different optical path.

[0147] Electrically Insulating Layers

[0148] When the electrically insulating layer 104 is present, the electrically insulating layer 104 may include one or more electrically non-conductive materials. Example electrically non-conductive materials include, but are not limited to, silicon-containing materials [e.g.,

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[0151] silicon nitride (SiNx)], electrically insulating polymers (e.g., polytetrafluoroethylene (PTFE), polyvinyl chloride (PVC), polycarbonate, nylon, and polypropylene), and ceramics (i.e., materials made by shaping and then firing an inorganic, nonmetallic material at a high temperature).

[0152] Electrically Conductive Layers

[0153] The electrically conductive layer 106 may include one or more electrically conductive materials. In some embodiments, the electrically conductive layer 106 may include or be made substantially or entirely of one or more metals. In some embodiments, the electrically conductive layer 106 may include or be made substantially or entirely of one or more metal oxides. In some embodiments, the electrically conductive layer 106 may include or be made substantially or entirely of indium tin oxide (ITO).

[0154] In some embodiments, the electrically conductive layer 106 may include or be made substantially or entirely of at least one conductive polymer.

[0155] A conductive polymer may include one or more aromatic cycles but no heteroatoms. Examples of such include, but are not limited to, poly(fluorene)s, polyphenylenes, polypyrenes, polyazulenes, and polybenzodifurandiones.

[0156] A conductive polymer may include one or more double bonds but no heteroatoms. An example of such includes, but is not limited to, poly(acetylene) (PAC).

[0157] A conductive polymer may include one or more aromatic cycles and one or more double bonds but no heteroatoms. An example of such includes, but is not limited to, poly(p-phenylene vinylene) (PPV).

[0158] A conductive polymer may include one or more aromatic cycles and contain one or more nitrogen heteroatoms. Examples of such include, but are not limited to, poly(pyrrole)s (PPY), poly carbazoles, polyindoles, polyazepines, and polyanilines (PANI).

[0159] A conductive polymer may include one or more aromatic cycles and one or more sulfur heteroatoms. Examples of such include, but are not limited to, poly(thiophene)s (PT), poly(3,4-ethylenedioxythiophene) (PEDOT), and poly(p-phenylene sulfide) (PPS).

[0160] In some embodiments, the electrically conductive layer 106 may be optically transparent, meaning light having a visible spectrum wavelength can pass through the electrically conductive layer 106. In other embodiments, the electrically conductive layer 106 may be optically opaque, meaning light having a visible spectrum wavelength is incapable of passing through the electrically conductive layer 106.

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[0163] Electrochemical Dopable Materials

[0164] The electrochemical dopable material 108 may be any material(s) whose electrochemical doping levels can be adjusted with voltage. The composition of the electrochemical dopable material 108 may be configured such that the refractive index of the electrochemical dopable material 108: (i) deceases in the visible domain when the electrochemical dopable material 108 is subjected to a positive bias voltage; and (ii) increases in the visible domain when the electrochemical dopable material 108 is subjected to a negative bias voltage. As used herein, the '‘visible domain” and “visible spectrum” refer to the 380 nm to 700 nm range of wavelengths that the human eye can see.

[0165] The electrochemical dopable material 108 may include an organic semiconductor and an electrolyte, where the organic semiconductor is doped / dedoped by the electrolyte depending on current flowing through the electrochemical dopable material 108.

[0166] The organic semiconductor portion of the electrochemical dopable material 108 may include at least one polymer one conductive polymer.

[0167] A conductive polymer may include one or more aromatic cycles but no heteroatoms. Examples of such include, but are not limited to, poly(fluorene)s, polyphenylenes, polypyrenes, polyazulenes, and polybenzodifurandiones.

[0168] A conductive polymer may include one or more double bonds but no heteroatoms. An example of such includes, but is not limited to, poly(acetylene) (PAC).

[0169] A conductive polymer may include one or more aromatic cycles and one or more double bonds but no heteroatoms. An example of such includes, but is not limited to, poly(p-phenylene vinylene) (PPV).

[0170] A conductive polymer may include one or more aromatic cycles and contain one or more nitrogen heteroatoms. Examples of such include, but are not limited to, poly(pyrrole)s (PPY), poly carbazoles, polyindoles, polyazepines, and polyanilines (PANI).

[0171] A conductive polymer may include one or more aromatic cycles and one or more sulfur heteroatoms. Examples of such include, but are not limited to, poly(thiophene)s (PT), poly(3,4-ethylenedioxythiophene) (PEDOT), and poly(p-phenylene sulfide) (PPS).

[0172] the electrolyte portion of the electrochemical dopable material 108 may include at least one polystyrene. An example of such includes, but are not limited to, polystyrene sulfonate.

[0173] The organic semiconductor and electrolyte portions of the electrochemical dopable material 108 may be present in different volume ratios. In some embodiments, when the electrochemical dopable material 108 starts as a solution, the organic semiconductor and 17

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[0175] electrolyte portions may be present at a ratio in the range of about 1: 1 to about 1:20 v / v. For example, the organic semiconductor and electrolyte portions may be present at a ratio of 1:1, 1:2, 1:3, 1:4, 1:5, 1:6, 1:7, 1:8, 1:9, 1:10, 1:11, 1:12, 1:13, 1:14, 1:15, 1:16, 1:17, 1:18, 1: 19, or 1:20 v / v.

[0176] The electrochemical dopable material 108 may be configured to have a doping level capable of being modulated by external voltage biases. Fig. 1 illustrates the concept of a sensing device of the present disclosure. Upon positive bias voltage in the electrically conductive layer 106, holes are injected in the electrochemical dopable material 108 (e.g., the organic semiconductor portion), increasing its doping level and decreasing its refractive index in the visible domain (Fig. 1A). Conversely, a negative electrical stimulus causes a reduction in the doping level of the electrochemical dopable material 108 (e.g., the organic semiconductor portion) and augments its refractive index within the same spectral window (Fig. IB). Furthermore, the scattering of a dielectric particle is directly related to its refractive index and that of the surrounding medium. When the refractive index of the electrochemical dopable material 108 is larger than the liquid medium contacted with the sensing device, its scattering is larger when its refractive index is higher. As a result, the light scattered by a sensing device of the present disclosure may be brighter at negative bias voltages — when the electrochemical dopable material 108 is dedoped — and vice versa. Practically, isolating light scattered by a sensing device of the present disclosure may require the geometric separation of illumination and detection paths, which can be achieved using a variety of dark-field imaging modalities [Priest, L. et al. Chem. Rev. 121, 11937-11970 (2021)].

[0177] In other embodiments (e.g., for one or more electrochemical dopable material 108 not made of PEDOTPSS), the refractive index of the electrochemical dopable material 108 in the visible domain may increase when subjected to a positive bias voltage and decrease when subjected to a negative bias voltage.

[0178] Methods of Use

[0179] A sensing device, or array of the present disclosure may be used to monitor electrical potential of a biological material in a liquid medium. The liquid medium may be contacted with a sensing device or array, and the electrical potential of the biological material in the liquid medium may then be measured.

[0180] In some embodiments, the electrical potential of the biological material may be measured by directing light at the sensing device or array using total internal reflection (TIR) illumination. The directed light may be an evanescent wave that decays once it passes 18

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[0182] through the electrochemical dopable material 108, such that the wave does not excite the biological material located within the liquid medium. Excitation of the biological material could result in unbeneficial light scattering from the biological material, which w ould add an optical background and decease the signal-to-noise ratio of the sensing device or array. By' utilizing an evanescent w ave, the contribution of the biological material to the optical signal of the sensing device or array can be minimized.

[0183] As part of TIR illumination, the light may be directed at the sensing device or array at an angle in the range of 61° to 90° wdth respect to vertical of the electrically conductive layer 106. For example, the light may be directed at an angle of 61°, 62°, 63°, 64°, 65°, 66°, 67°, 68°, 69°, 70°, 71°, 72°, 73°, 74°. 75°, 76°, 77°, 78°, 79°, 80°, 81°, 82°, 83°, 84°, 85°, 86° 87°, 88°, 89°, or 90° with respect to vertical.

[0184] In TIR illumination, the light directed at the sensing device or array may have a wavelength anywhere between the ultraviolet and infrared spectra, provided the electrochemical dopable material 108 is responsive and sensitivity' of the sensing device or array is specific to the chosen wavelength. For example, the light may have a wavelength in the range of 380 nm to 700 nm. In some embodiments, the light may have a wavelength of 637 nm.

[0185] Dark-field microscopy (sometimes referred to as dark-ground micrscopy), as known in the art, may be used as an alternative to TIR illumination. In dark-field microscopy, the unscattered beam is excluded from the captured image. As a result, the field around the biological material may be generally dark in a captured image of the present disclosure.

[0186] As discussed elsew here herein, a sensing device of the present disclosure is capable of scattering light in the visible domain. As such, measuring the electrical potential of the biological material may include using a microscope to observe (e.g., image or otherwise measure) visible light scattered by the electrochemical dopable material(s) 108 of the sensing device or array.

[0187] Liquid Media

[0188] Sensing devices of the present disclosure may be used to measure electrical potential of various biological materials.

[0189] In some embodiments, the liquid medium may be a physiological medium.

[0190] The biological material may include one or more cells. In some embodiments, the one or more cells may be one or more cells that exhibit electrical activity. Example cells include.

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[0193] but are not limited to, neurons and cardiac tissues. In some embodiments, the liquid medium may be that of a cell culture.

[0194] The biological material may include one or more proteins.

[0195] The biological material may include one or more small molecules. As used herein, a “small molecule” refers to an organic molecule having an upper molecular weight limit of about 900 daltons.

[0196] The biological material may include one or more peptides.

[0197] The biological material may include one or more ribonucleic acid (RNA) molecules. The biological material may include one or more deoxyribonucleic acid (DNA) molecules.

[0198] The liquid medium may be an aqueous-based solution. However, it need not be. The liquid medium could have a non-aqueous solvent, such as acetonitrile or isopropanol.

[0199] As noted above, the electrochemical dopable material 108 of a sensing device may be smaller in cross-section than a cell. As such, a sensing device or array may be used to measure the electric activity of a single cell. A sensing array may be used to measure the electric activity of specific portions of a cell, with each sensing device of the array measuring the electric activity of a different portion of the cell.

[0200] Applications

[0201] Sensing devices of the present disclosure are usable in various industries due to their ability to provide high-resolution and wireless monitoring of bioelectrical signals, and also their light modulation capabilities.

[0202] Sensing devices of the present disclosure have a wide range of applications.

[0203] Sensing devices may be used to conduct fundamental biomedical research to probe and understand biological processes at cellular and molecular levels. For example, sensing devices can be used to study neuronal activity with micrometer precision, enabling researchers to observe and map complex neural networks, and thereby better understand brain function and the development of neuroprosthetics. For further example, sensing devices can be used to monitor electrical signals in cardiac tissues, providing insights into heart function and aiding in the development of treatments for arrhythmias and other cardiac disorders.

[0204] Sensing devices may also be used in clinical diagnostics to provide detailed and realtime monitoring of biological signals for early diagnosis and treatment. For example, sensing devices can be used as part of non-invasive diagnostics that require high-resolution monitoring of bioelectrical signals without the need for invasive procedures. For further 20

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[0206] example, sensing devices can be used as part of personalized medicine (e.g., by providing detailed data on a patient's specific bioelectrical patterns (e.g. from biopsies), sensing devices can help tailor treatments to individual needs, enhancing the effectiveness of therapies and reducing side effects).

[0207] Sensing devices can also be used for drug discovery and pharmacology, to provide detailed and high-throughput data on cellular responses to potential treatments. For example, sensing arrays of the present disclosure can be used to (high throughput) screen many compounds simultaneously, assessing their effects on cellular electrical activity and identifying potential drug candidates quickly and efficiently. For further example, researchers can use sensing devices to study the mechanisms of action of drugs on a cellular level, providing insights into how treatments affect cellular function and helping to refine therapeutic strategies.

[0208] Sensing devices can also be implemented in the field of integrated photonics to modulate the intensity of scattered light with micrometer resolution. For example, sensing devices can be included in integrated photonics circuits to modulate light intensities in the visible domain.

[0209] EXAMPLES

[0210] Example 1

[0211] Modeling of OCEANs

[0212] The scattering cross-section of PEDOT: PSS OCEANs were simulated using the finite element modeling (FEM) method provided by CST Studio Suite software (Dassault Systemes, France). The OCEAN stem diameter (SiNxopening) was set to 250 nm. Two cap diameters, 750 nm and 1.5 pm, were investigated for which substrate lateral dimensions of 1000 nm and 1750 nm were defined, respectively. The SiNx, ITO, and glass substrate thicknesses were set to 50 nm, 70 nm, and 200 nm, respectively. Water was defined as the background material. Perfectly matched layers (PML) boundary conditions were used in all directions of the space, with extra space added in the upward vertical (+z) direction to ensure the reliability of the far field computation. The frequency-dependent permittivity of glass [Malitson, I. H. J. Opt. Soc. Am. 55, 1205 (1965)] and ITO [Konig, T. A. F. et al. ACS Nano 8, 6182-6192 (2014)] were extracted from publicly available datasets, while constant permittivity values of 4 and 1.77 were used for SiNx[Beliaev, L. Y. et al. Thin Solid Films 763, 139568 (2022)] and water [Lide, D. R. CRC Handbook of Chemistry and Physics, 84th Edition (2003)], respectively. PEDOT: PSS complex permittivities computed for different 21

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[0214] voltage biases (additional details in Example 6) were fed into the model to account for electrochemical modulation. A plane wave with an electric field pointing in the x direction and propagating downward (-z direction) was used as excitation. Spatial symmetries of the simulated structure were leveraged to shorten the simulation time, without compromising the accuracy of the simulation (yz plane defined as electric field symmetry plane, xz plane as magnetic field symmetry plane). Scattering cross-sections were simulated for all directions of space (angular resolution of 1°) and wavelength ranging from 400 nm to 1000 nm (with 10 nm increments) using the frequency domain solver. Angular integration of the 3D-resolved scattering radiation patterns was performed using MATLAB (Mathworks, USA), from which the total scattering cross-section of the OCEANs for all simulated wavelengths and voltage biases could be extracted.

[0215] To estimate the dynamic response of an OCEAN at a specific wavelength (Fig. 2H), the voltage trace across PEDOT: PSS in response to a 100 mV voltage pulse was computed using the electrical equivalent circuit presented in Fig. 10A (assuming ideal RC-type behavior). Impedance characteristics of OCEANs with diameters of 0.75 µm and 1.5 µm were estimated from experimental measurements of OCEANs of 0.7 pm and 1.4 pm, respectively. Scattering cross-sections corresponding to the voltage levels anticipated at different time points were interpolated from the simulated scattering cross-section using a sigmoid characteristic to determine the OCEAN optical response.

[0216] The influence of the cell presence on OCEAN scattering properties was evaluated under total internal reflection illumination conditions (incidence angle = 70° with respect to the vertical axis), collecting the scattered light from the top. The cell membrane thickness was set to 5 nm, and 20 nm was placed above the top of the antenna. The cytosol thickness was set to 100 nm with perfectly matched layer boundary conditions in all directions of the space. Refractive indexes of 1.54 and 1.38 were used to simulate the cell membrane [Meyer, R. A. Appl. Opt. 18, 585-588 (1979)] and the cytoplasm [Zhang, Q. et al. Sci. Rep. 7, 2532 (2017)], respectively.

[0217] Results

[0218] The scattering cross-section of a particle - defined as the ratio between the scattered light power and the incident light intensity - is highly dependent on its complex refractive index - or complex permittivity - as well as that of the embedding medium (see Example 6 for details about the relationship between complex refractive index and complex permittivity) [Bohren, C. F. & Huffman, D. R. Absorption and Scattering of Light by Small Particles 22

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[0220] (1983)]. While analytical expressions relating permittivity to the scattering cross section can be derived in isolated settings (small-size approximation for Rayleigh scattering [Rayleigh, Lord. Philos. Mag. J. Sci. 47, 375-384 (1899)], spherical particle geometry for Mie theory [Mie, G. Ann. Phys. 330, 377-445 (1908)]), more complex geometries require numerical treatment. In the following section, the three-step modeling efforts are highlighted. First, studies were performed to assess how external voltage biases modulate the complex permittivity of PEDOTPSS via electrochemical doping and dedoping processes in phosphate-buffered saline (PBS) solution. Next, numerical simulations of a proposed OCEAN structure were conducted to investigate how changes in OCEAN permittivity related to changes in scattering cross-sections. Finally, an equivalent circuit model was used to extract dynamic scattering responses to applied voltages, providing insights into OCEAN’S operation bandwidth.

[0221] At equilibrium, positive charge carriers in the PEDOT backbone (i.e., polarons and bipolarons) were stabilized via electrostatic interactions with negatively charged PSS strands. As a result, the PEDOTPSS was naturally in a highly conductive (i.e., doped) state. Under negative voltage, cations from the electrolyte infused into the PEDOTPSS, neutralized the negatively charged PSS strands, and locally reduced PEDOT into its dedoped state, making it less conductive and changing its dielectric properties (see Example 6). In its fully doped state, the complex permittivity of PEDOT PSS can be described by a Drude model, which is consistent with a metallic behavior (Fig. 6A-B). On the other hand, when PEDOTPSS was fully dedoped, its permittivity was best fitted by a combination of Lorentz and Tauc-Lorentz functions, consistent with a purely dielectric behavior. Note that these permittivities vary linearly with their respective volumetric density of charge carriers. Consequently, computing the complex permittivities for intermediate doping levels comes down to a linear interpolation between the fully doped and dedoped permittivities. Complex permittivities of PEDOTPSS for different doping levels are presented and discussed in detail below herein. The relationship between the PEDOTPSS doping levels and the corresponding voltage amplitudes applied across it in PBS was experimentally established and is also reviewed in Example 6 herein.

[0222] Fig. 2A-B show the PEDOTPSS permittivity’s real and imaginary parts under different voltage bias conditions in PBS. The resulting scattering cross-section simulated for a single OCEAN with a diameter of 750 nm and 1500 nm are presented in Fig. 2C-D, respectively. Predictably, larger OCEANs yielded larger scattering magnitudes. At negative voltages, the real part of PEDOTPSS permittivity increased, especially for wavelengths near 23

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[0224] 700 nm, thus enhancing its scattering properties and leading to brighter OCEANs within this spectral window. Fig. 2E highlights how static voltage biases applied across OCEANs influenced their scattering cross-section at a wavelength of 640 nm. Interestingly, because the variation of PEDOT: PSS permittivity is not linear with voltage, the OCEAN sensitivity -defined here as the variation of scattering cross-section with respect to voltage - is expected to be maximal at operating biases near -0.5 V (Fig. 2F).

[0225] Based on the electrical equivalent circuit of an OCEAN detailed in Example 8 herein, the dynamic voltage trace transmitted across the PEDOT: PSS structure in response to a -100 mV in amplitude voltage pulse applied between the ITO and the bath was computed at optimal operating bias of -0.5 V (Fig. 2G). Due to their larger volumetric capacitance, larger OCEANs exhibited slower dynamics than their smaller counterparts (Fig. 10). Fig. 2H presents how these electrical stimuli were simulated to modulate the optical signal of single OCEANs. Interestingly, a negative voltage pulse leads to a positive variation of the optical signal, highlighting the central role of the real part of PEDOT: PSS permittivity in light scattering. Results of the simulations demonstrated that the intensity of the light scattered by single OCEANs could be electrochemically modulated by applying voltage stimuli with millivolt amplitudes and millisecond kinetics.

[0226] Example 2

[0227] Nanofabrication and interfacing of OCEANs

[0228] ITO-coated glass substrates (dimensions 22 x 22 mm2, thickness #1.5 -170 pm, resistivity 70-100 Q / n from SPI Supplies, USA) were subsequently sonicated for 5 minutes in deionized water and isopropyl alcohol. A multimeter was used to identify the side of the chips presenting the ITO layer, and their orientation was carefully maintained throughout the fabrication process (step 1, Fig. 7).

[0229] ITO patterning and passivation (steps 2 and 3, Fig. 7)

[0230] The substrates were treated with a 300W, 0.8 mBar, and 2 min-long oxygen plasma step (Nano-QL-PCCE7, Diener Electronic, Germany) and functionalization with hexamethyldisilazane for 10 s at 150°C in a YES-3 IOTA vacuum oven (Yield Engineering Systems, USA) in preparation for photolithography. A 1.8 pm-thick AZ3310 positive photoresist layer (MicroChemicals, Germany) was spin-coated at 1’500 rpm for 60 s using a CEE 200X-F spin coater (Brewer Science, USA) and baked at 90°C for 60 s. The photoresist was exposed using an MLA-150 maskless aligner (Heidelberg, Germany) with a dose of 600

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[0233] mJ / cm2, baked at 110°C for 60 s, and developed for 30 s in AZMIF300 (MicroChemicals, Germany) to define the traces and pads of the devices. A 15 min-long hard-bake step at 95°C was performed to ensure maximal stability of the photoresist mask during the subsequent etching step. The ITO layer was patterned by 5 min-long chlorine-based reactive-ion etching step using a 230i inductively-coupled plasma reactive ion etcher (Samco, Japan). The photoresist was removed by subsequent steps of oxygen plasma asher (300W, 0.8 mBar, 2 min, Nano-QL-PCCE7, Diener Electronic, Germany) and sonication in N-Methyl-2-pyrrolidone for 1 h 30 min. The resulting ITO traces and pads were passivated by a 50 nm-thick layer of SiNx deposited by plasma-enhanced chemical vapor deposition performed in a PD220NL deposition tool (Samco, Japan) using silane and ammonia and validated with an F20 thin film measurement system (Filmetrics, USA).

[0234] SU8 patterning (step 4, Fig. 7)

[0235] The SiNx-coated patterned substrates were treated with a 300W, 0.8 mBar, and 2 minlong oxygen plasma step (Nano-QL-PCCE7, Diener Electronic, Germany) and functionalization with hexamethyldisilazane for 10 s at 150°C in a YES-310TA vacuum oven (Yield Engineering Systems, USA) in preparation for photolithography. An Apogee spin coater (Cost Effective Equipment, USA) was used to spin coat a 3 pm-thick layer of GMPI1040 SU8 (Gersteltec, Switzerland) at 1’100 rpm for 62 s (11 s of acceleration and deceleration at 100 rpm / s and 40 s of steady rotation at 1 ’ 100 rpm). After 5 min of relaxation, the substrates were baked for 5 min at 65°C and 10 min at 95°C. An MLA 150 maskless aligner (Heidelberg, Germany) was used to expose the negative SU8 photoresist at a dose of 1’200 mJ / cm2Following exposure, the post-exposure bake was performed (5 min at 65°C and 10 min at 95°C) and followed by a 10 min-long rehydration delay. The SU8 was manually developed by immersion in propylene glycol methyl ether acetate for 40 s and hard-baked at 135°C for 2 hours.

[0236] SiNx patterning (step 5, Fig. 7)

[0237] A Velion FIB (Raith Nanofabrication, Germany), was used to pattern 250 nm in diameters openings in the exposed SiNx layer. High-sputtering yield gold double-plus ions (Au++) were used to maximize the patterning throughput, and a 10 pm aperture was set to maintain a beam current of ~9 pA, enabling fine spatial resolution. The optimal patterning dose was found to be 0.2 nC / pm2distributed across ten outward loops (0.02 nC / pm2 / loop).

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[0240] Interfacing

[0241] H20E conductive epoxy (Epotek, USA) was manually dispensed on the pads of a custom-designed PCB using a precision fluid dispenser (APPLSRA105-10CC, SRA Soldering Products, USA). Subsequently, the chip was accurately positioned and pressed onto the PCB (setpoint of 10 g) using an M-3 pick and place tool and cured at 80°C for 3 hours. Then, insulating epoxy 302-3M (Epotek, USA) was manually applied at the interface between the chip and the PCB before being baked at 65 °C for 3 hours to render the interface mechanically stable and waterproof. Finally, a 6 mm-high glass ring (36 mm outer and 32 mm inner diameter, Friedrich and Dimmock, USA) was assembled using the 302-3M epoxy and baked for 3 hours at 65°C to finalize the interfacing.

[0242] PEDOT. PSS electrodeposition (step 6. Fig. 7)

[0243] To make the chip hydrophilic and prevent air bubbles from being trapped in the chip cavities, a 30 W, 30 s, and 0.6 mBar oxygen plasma was performed using an Atto plasma asher (Diener Electronic, Germany). Subsequently, an electrochemical cleaning step was performed by sweeping the potentials between the ITO layer and a 50 mM solution of potassium hydroxide from -0.2 V to -0.8 V versus Ag / AgCl for at least 10 cycles, using a three-electrode setup and a Compactstat potentiostat (Ivium, Netherlands). A platinum wire was used as the counter electrode, and a RE-1B Ag / AgCl reference electrode (ALS-japan, Japan) was used as the reference. The solution used to electrodeposit PEDOT: PSS was composed of 10 mM of 3,4-Ethylenedioxythiophene (EDOT, 483028-10G, Sigma Aldrich, USA) and 0.1 M of Poly(sodium 4-styrene sulfonate) (NaPSS, 243051-100G, Sigma Aldrich, USA). Note that the 0.1 M of NaPSS corresponds to the concentration of sodium 4-styrene sulfonate (NaSS) monomers (206 mg of NaPSS in 10 ml of deionized water). PEDOT PSS OCEANs were electrodeposited by applying 0.9V versus Ag / AgCl for 30 s, 45 s, or 60 s using the same three-electrode setup.

[0244] Results

[0245] Nanofabrication of OCEANs

[0246] Arrays of OCEANs were manufactured following the nanofabrication process flow presented in Fig. 7. First, ITO traces and pads were patterned on an otherwise insulating glass substrate by photolithography and reactive ion etching (Fig. 7, steps 1 and 2). To passivate the ITO structures, a 50 nm thick SiNx layer was deposited by plasma-enhanced chemical vapor deposition and coated with a 3 pm thick film of SU8. 80 pm square openings were 26

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[0248] paterned by photolithography in the SU8 to access the ITO-SiNx pads and defined the location of the OCEAN arrays (Fig. 7. steps 3 and 4). As highlighted in Fig. 3A, focused-ion beam (FIB) lithography was then performed to patern arrays of nanoholes in the exposed SiNx layer (diameter ranging from 100 nm to 500 nm, pitch of 5 pm), locally revealing the ITO layer underneath. Finally, protruding PEDOT: PSS structures were grown through each cavity by electrodeposition and formed arrays of OCEANs (Fig. 7, steps 5 and 6). The influence of the electrodeposition time on the OCEAN diameter is described in Example 7 herein. Fig. 3B-D show scanning electron microscopy (SEM) and brightfield microscopy images of the resulting structures at different steps of the fabrication process.

[0249] Each chip comprised 25 arrays, each composed of 16 x 16 OCEANs. Each individual array was electrically addressed independently, and light scattered by individual OCEANs within the arrays was collected at the single-unit level. Chips were mounted on custom-designed printed circuit boards (PCBs) and assembled with glass rings to facilitate electrical interfacing and experiments in liquid (see Fig. 7B and additional details in Example 7 herein).

[0250] Example 3

[0251] Electrochemical impedance spectroscopy on OCEANs

[0252] Electrochemical impedance spectroscopy of OCEANs w as performed in PBS (806552-500ML, Sigma Aldrich, USA) using a Compactstat potentiostat (Ivium, Netherlands) in a three-electrode configuration. A 10 mV in amplitude sinusoidal voltage signal with frequencies ranging between 1 Hz and 100’000 Hz was applied between the ITO and the reference electrode of individual arrays. The fitting of the experimental data was directly performed using the Ivium software.

[0253] Example 4

[0254] Electro-optic characterization of OCEANs

[0255] Scattering spectra acquisition

[0256] Scattering spectra of PEDOT:PSS OCEAN arrays (50 x 50 OCEANs, ~1.8 pm in diameter) were acquired in PBS using a Ti2E inverted microscope (Nikon, Japan) configured with a dark-field condenser to enable dark-field microscopy. An SLS302 broadband light source (Thorlabs, USA) was connected to the microscope via an LLG3-4Z liquid guide (Thorlabs, USA) and used for diascopic illumination. A 40x objective was used to collect the light scattered by the OCEANs, which was subsequently transferred to an Ocean-HDX-XR spectrometer (Ocean Insight, USA) through a 1000 pm optical fiber (QP1000-2 -VIS-NIR,

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[0259] Ocean Insight, USA). The spectrometer slit was set to 100 pm. The OCEAN spectra were subtracted by the one acquired in dark conditions and normalized by the light source spectrum in the relative irradiance method of the spectrometer Ocean View software (Ocean Insight, USA).

[0260] Total internal reflection darkfield microscope setup

[0261] A custom-designed upright wide-field fluorescence microscope body was built from the DIY Cema components available at Thorlabs, USA. A Chrolis-C1 (Thorlabs, USA) multiwavelength light engine was used as the episcopic light source for fluorescence and brightfield microscopy. An Orca Fusion-BT (Hamamatsu, USA) camera was selected as the readout sensor for fluorescence, brightfield, and dark-field microscopy. A water dipping 60x objective (N60X-NIR, Thorlabs, USA) was used for the entire electro-optic characterization of OCEANs. A prism-based total internal reflection illumination module was built using a 637 nm pig-tailed laser diode (70 mW, LP637-SF70, Thorlabs, USA) mounted on a CLD1010LP controller (Thorlabs, USA), and its output collimated with an F230FC-B collimator (Thorlabs, USA). The collimator was mounted at 30° with respect to the horizontal plane on an M30XY / M motorized stage (Thorlabs, USA) and aligned with an ADB-10 prism (Thorlabs, USA). The entire illumination module was assembled on an L490 / M lab jack (Thorlabs, USA) for coarse Z adjustments. A custom 3D-printed sample holder was assembled on an MPRC / M recording chamber holder (Thorlabs, USA) and fixed on an MP20 rigid stand mounted on a 2D motorized translation stage (PLS-XY, Thorlabs, USA) to enable accurate motion of the sample.

[0262] Electro-optic characterization of OCEANs

[0263] A compactstat potentiostat (Ivium, Netherlands) was used in a three-electrode configuration to apply electrical stimuli (e.g., cyclic voltammetry and chrono-amperometry) to OCEANs in PBS. A platinum wire was used as the counter electrode, and a RE-1B Ag / AgCl reference electrode (ALS-japan, Japan) was used as the reference. The camera exposure time and frame rate were set to 1 ms and 200 fps for chronoamperometry and 1 ms and 20 fps for cyclic voltammetry. For time-constant measurements, the exposure time was set to 200 ps and the frame rate to 2’000 fps. The sensitivity of each OCEAN was estimated by performing a linear fit on their optical Z score response to voltage pulses of varying amplitudes (Figs. 14H. 15H. and 16H). The sensitivity, represented by the slope of this fit.

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[0266] quantifies the change in the Z score per unit of applied voltage (Sensitivity=

[0267]

[0268] — where AZ is the variation in the optical Z score in response to a voltage change AZ based on the linear interpolation of experimental data). The noise was defined as the standard deviation of the Z score over a one-second period in the absence of stimuli Noise = ~ Z)2where N

[0269]

[0270] is the noise, Zi represents the Z score at each time point, Z is the mean Z score during the one-second interval, and N is the total number of time points recorded). The SNR was calculated by dividing the Z score response of each OCEAN to a -100 mV voltage pulse by

[0271]

[0272] its corresponding noise level (SNR =Z

[0273]

[0274] , where Zioomv is the Z score in response to the -100 mV voltage pulse). Finally, the limit of detection was defined as the voltage for which the SNR was equal to one. Practically, it was computed by finding the intersection between the linear relationship describing the Z score versus voltage stimulus amplitudes and the noise level for each OCEAN (SNR = 1 =y =Noise y^

[0275]

[0276] is the limit of detection).

[0277] Data analysis:

[0278] Image sequences were first processed with a vibration compensation algorithm [Evangelidis, G. & Psarakis, E. IEEE Trans. Pattern Anal. Mach. Intell. 30, 1858-1865 (2008)] to prevent motion artifacts from biasing the electro-optic characterization. Circular regions of interest (ROIs) with a radius of 20 pixels were defined around individual OCEANs. For each pixel of the ROIs, the average (ytpix) and standard deviation (<jpix) of the optical signal in the absence of electrical stimulation were computed and used to calculate a Z score per pixel (Zscoreptx (t) =Xpix

[0279] a^t-> tipix. with Xpix(t) the raw optical signal from a single Pix

[0280] pixel) for each time point. Pixels with a Z score variation larger than 3 were considered part of the OCEAN and averaged to define its optical response. Electro-optic characterization results were summarized in box charts displaying the median, lower, and upper quartiles, outliers, as well as the minimum and maximum values that are not outliers (Matlab, USA).

[0281] Results

[0282] Electro-optic characterization of OCEANs

[0283] Electro-optic properties of OCEANs grown through 250 nm in diameter openings and for electrodeposition time (tᴇᴅ) of 30 s, 45 s, and 60 s were investigated in this section. The

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[0286] corresponding cap diameters were 0.7, 1.4, and 1.8 pm, respectively (see details in Example 7 herein).

[0287] First, relative irradiance spectra of OCEAN arrays were characterized in PBS for bias voltages ranging from 0 V to -0.7 V using dark-field spectroscopy (details provided in Example 9 herein). As presented in Fig. 11, larger negative biases lead to brighter OCEANs. The scattered light intensity peaked at wavelengths neighboring 700 nm. These experimental results confirm the influence of PEDOT doping and dedoping on the scattering properties of OCEANs and demonstrate the theoretical predictions.

[0288] For the rest of the electro-optic characterization, a total internal reflection-based illumination (Fig. 4A insert, and additional details in Example 9 herein) was preferred over a dark-field condenser to confine the incident light to the OCEANs plane and minimize potential scattering background coming from the cells during biological experiments. An illumination wavelength of 637 nm was chosen as a trade-off between the OCEAN’S relative irradiance spectrum and the camera's quantum efficiency. Additionally, the variation of scattered light intensity for each OCEAN was expressed in terms of Z score. It represented how much the optical signal deviates from its baseline in terms of standard deviations and enabled systematic comparison between OCEANs. Additional information about Z score computation can be found in the Materials and Methods section in Examples 1-3.

[0289] Fig. 4A shows optical (top) and electrical (bottom) voltammograms acquired from OCEANs of different diameters in PBS. The optical signal scattered by each individual OCEAN was collected at the single-unit level. The OCEANs brightness follows a sigmoidal relationship with respect to bias voltages with a maximum slope between -0.3 V and -0.4 V. In addition, larger structures showed stronger electro-optic responses accompanied by larger currents (see additional details below herein, Fig. 13D-F for dark-field images of arrays comprising OCEANs of different diameters). Voltammograms comprising two cycles are shown below herein. As expected from the theoretical model, both voltage biases and cap diameters played a central role in the electro-optic characteristics of OCEANs. These contributions were systematically characterized by applying 1 s long voltage pulses with amplitudes ranging from -100 mV to 100 mV across OCEANs with diameters of 0.7, 1.4, and 1.8 pm and operating at voltage biases between 0 V and -0.6 V while monitoring their optical responses in PBS. Fig. 4B shows representative optical traces measured from 1.4 pm in diameter OCEANs operating at a bias voltage VB = -0.3 V. As anticipated by the theoretical model, a negative voltage pulse increased the OCEAN brightness, confirming enhanced PEDOT:PSS scattering properties in dedoped states. The overall dataset for the three

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[0292] different dimensions and the seven biases tested can be found in Example 9 herein. Fig. 4C-F and Table 2 summarize the influence of geometries and operating biases on single OCEAN sensitivity, optical noise, signal-to-noise ratio, and limit of detection. Predictably, the lowest limits of detection were consistently achieved within the bias voltage range of -0.3 V and -0.4 V. The optimal operating voltage was -0.4 V for cap diameters of 0.7 pm and 1.8 pm and -0.3V for 1.4 pm in diameter structures (see details below herein, Table 2 for limit of detection numerical values). Additionally, the smallest OCEANs exhibited the most limited electro-optic responses. Even though higher sensitivity was observed for larger structures, the associated optical noise increase made the OCEANs of intermediate dimensions (tED = 45 s) perform similarly compared to the large ones (tED = 60 s) with a limit of detection between 2 mV and 3 mV at optimal bias.

[0293] The dynamic characteristics of OCEANs were experimentally investigated and are presented in Fig. 4-H and in Example 9 herein. Due to their larger capacitance, large OCEANs showed longer time constants than their smaller counterparts (Fig. 4H and Fig. 17). At optimal bias, time constants were measured to be 6.0 ms, 34.7 ms. and 233.9 ms (mean values) for cap diameters of 0.7 pm, 1.4 pm, and 1.8 pm, respectively. Furthermore, negative voltage biases were associated with longer time constants, particularly for structures of intermediate dimensions. The increase of the serial resistance RPEDOTPSS caused by the dedoping of PEDOT at negative voltages (Fig. 10) is likely responsible for this observation as the time constant is defined as T = (RPEDOT-. PSS + RS) ■ CPEDOT-. PSS (see additional details in Example 8 herein).

[0294] Finally, the long-term stability of the OCEAN electro-optic properties was studied and discussed in detail in below herein in Example 9 herein. As demonstrated in Fig. 4G, single OCEANs enabled wireless detection potentials of -100 mV for a continuous duration of 10 hours. While the Z score median value remained quite steady over the course of the experiment, its variation between OCEANs, as well as the number of outliers, became larger. Furthermore, the dynamic characteristics of several OCEANs seem to have significantly- slowed down, as if PEDOT was steadily getting dedoped. The contribution of these affected OCEANs led to an overall signal-to-noise ratio that decreased overtime (Fig. 18).

[0295] Interestingly, the electro-optic characteristics of PEDOT: PSS OCEANs were not altered by three days of passive immersion in PBS at 37°C, without electrical stimulation.

[0296] Example 5

[0297] Modeling of cardiomyocyte action potential recording with OCEANs

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[0300] The cell-OCEAN interface model was adapted from a Matlab implementation of the Luo Rudy model (GitHub permalink: see:

[0301] / / github.com / meeravarshneya1234 / ventricular_myocyte_models / tree / a57aaf7a632b3808fe0fe93355c2a8d9eed16dd0 / Luo_Rudy91_Model). The updated code is available in the dedicated data repository. Equations, numerical values of the key components used in the model, and additional discussions related to its implementation are presented in certain Examples below herein.

[0302] Example 6

[0303] Electrochemical doping and dedoping of PEDOT: PSS

[0304] In this section, the doping and dedoping processes taking place within the PEDOT backbone are discussed. Subsequently, the PEDOT: PSS complex permittivity is presented for different doping levels and voltages are discussed.

[0305] Doped and dedoped states of PEDOT: PSS

[0306] Fig. 5 illustrates how the PEDOT doping level was modulated by external voltage biases across the PEDOT: PSS-electrolyte interface. At equilibrium, holes (i.e., polarons and bipolarons) in the PEDOT backbone were stabilized by the fixed sulfonate anions in the PSS strand. Under negative voltage, solvated cations from the electrolyte neutralized the negatively charged PSS strands and locally reduced the PEDOT into its dedoped state. The resulting hole was extracted through the indium tin oxide (ITO) electrode. Conversely, injecting holes under positive voltage biases across the interface could oxidize the PEDOT backbone into its doped state [Proctor, C. M. et al. J. Polym. Sci. Part B Polym. Phys. 54, 1433-1436 (2016)]. In response to electrochemical doping / dedoping processes, structural changes in the PEDOT backbone directly modulated the optical absorbance and scattering properties of the film, which were naturally captured in the complex (frequency-dependent) refractive index denoted m = n + IK.

[0307] The real part of the refractive index n indicated the phase velocity of light in the material and the imaginary' part K, which represents the material’s extinction coefficient. In light scattering theory, the real refractive index n is directly associated with a particle’s scattering cross-section - defined as the ratio between the scattered light's power and the incident light's intensity - and modulates its scattering characteristics [Mie, G. Ann. Phys. 330, 377-445 (1908)]. A material's complex refractive index can be expressed in terms of its

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[0310] complex permittivity ε = ε1+ iε2, where ε1= n2− κ2and ε2= 2nκ [Meyzonnette, J. L. et al. Springer Handbook of Glass 997-1045 (2019)]. For low K values, ε1tends to n2and ε2to zero.

[0311] Relationship between PEDOT:PSS doping level and permittivity

[0312] The complex permittivities of PEDOTPSS in its fully doped and fully dedoped states were fitted from a recent in-situ spectroscopic ellipsometry study [Dingier, C. et al.

[0313] Macromolecules 55, 1600-1608 (2022)] (Fig. 6A and B). It was assumed these extreme-state permittivities are a universal property of the material and invariant with respect to the electrolyte in which they were measured. The transient increase in absorption occurring near 1000 nm for intermediate biases was neglected as its exact voltage onset is electrolytedependent and difficult to estimate accurately from the data provided [Dingler, C. et al. Macromolecules 55, 1600–1608 (2022)]. The doped state complex permittivity was best described by a Drude model (red curves in Fig. 6A and B), which was consistent with a metallic behavior. The dedoped state complex permittivity was instead best fitted by a combination of two Lorentz and two Tauc- Lorentz functions, consistent with a purely dielectric behavior of the PEDOT: PSS in this regime (black curves in Fig. 6A and B, resulting from the combination of the four models depicted in dashed lines).

[0314] The corresponding complex permittivity for the Drude [Drude, P. Ann. Phys. 306, 566–613 (1900)] (εD(ω)), Lorentz [Lorentz, H. A. The Theory of Electrons and Its Applications to the Phenomena of Light and Radiant Heat (1916]) (εL(ω)), and Tauc-Lorentz [Jellison, G. E. & Modine, F. A. Appl. Phys. Lett. 69, 371-373 (1996)] (εTL(ω)) models are recapitulated below:

[0315] εD(ω) = ε∞− ωP2 / ω(ω+iγ) Equation 1

[0316] εL(ω) = ε∞+ ωP2 / (ω02− ω2− iΓω) Equation 2

[0317] εTL(ω) = ε∞+ χTL(E) Equation 3

[0318] fl AE0C(E-Egf

[0319] With E — —. 5{ZT- / ,(E)} = \E(E2-E02f +C2E2’9. and

[0320]

[0321] 271( 0, if E < Eg

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[0324] ℜ{χTL(E)} = (2 / π) ∫Eg∞ξℑ{χTL(E)} / (ξ2−E2) dξ

[0325] Where:

[0326] • to is the angular frequency

[0327] • ECO is the static dielectric constant

[0328] • top is the plasma frequency

[0329] • y is the characteristic collision frequency

[0330] • too is the resonance frequency of the oscillator

[0331] • T is the relaxation time

[0332] • T = - is the inverse of the relaxation time

[0333] • E is the photon energy

[0334] • h = 6.626. 10-34J. Hz-1is the Planck constant

[0335] • ZTL(E) is related to the electric susceptibility

[0336] • A is a fitting parameter related to the strength of the Lorentzian oscillator

[0337] • ED is a fitting parameter related to the resonant frequency of the Lorentzian oscillator • C is a fitting parameter related to the broadening of the Lorentzian oscillator

[0338] • Eg is a fitting parameter related to the bandgap material

[0339] The complex permittivity for the fully doped and fully dedoped states of PEDOT: PSS can be expressed as follows:

[0340] εDoped(ω) = εD(ω) Equation 4 εDedoped(ω) = εL,1(ω) + εL,2(ω) + εTL,1(ω) + εTL,2(ω) Equation 5

[0341] The values of the parameters fitted from the experimental data presented in reference are summarized in Table 1.

[0342] Table 1 Values of the parameters of the Drude, Lorentz, and Tauc-Lorentz models contributing to the fully doped and fully dedoped complex permittivity of PEDOT: PSS, fitted from experimental data presented in reference [Dingler, C. et al. Macromolecules 55, 1600–1608 (2022)]. _ _ _ _

[0343] Model Parameter Value Unit

[0344] £co, Doped 2.02 -

[0345]

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[0348] ωP' 1.58 · 1015rad · s-1εD γ 0.630 · 1015rad · s-1Sco, Dedoped 1.9 - ωP' 0.512 · 1015rad · s-1εL,1 ω( 3.24 · 1015rad · s-1Γ 0.539 · 1015rad · s-1ωP' 0.273 · 1015rad · s-1εL,2 ω0 3.57 · 1015rad · s-1Γ 0.419 · 1015rad · s-1A 3.90 • IO’19J £TL, I Eo2.83 • IO’19J C 2.34 · 10-19J Eg2.48 · 10-19J A 1.05 • IO’19J £TL,2 Eo 3.09 • IO’19J C 2.84 · 10-19J Es2.48 • IO’19J

[0349]

[0350] All the complex permittivities presented above exhibit a linear relationship with their respective number density (z.e., the volumetric density of the charge carriers contributing to the permittivity) through the plasma frequency parameter ωP2in their numerators. Note that A is related to the strength of the Lorentzian oscillator inherent to the Tauc-Lorentz model, which is proportional to ωP2[Lorentz, H. A. The Theory of Electrons and Its Applications to the Phenomena of Light and Radiant Heat (1916)].

[0351] Indeed, ωP2is defined as:

[0352] e2N

[0353] ωP2= Equation 6

[0354]

[0355] ε0m*

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[0358] Where e is the elementary charge, ε0is the vacuum permittivity, m* is the effective mass of the charge carrier considered, and / V is the volumetric density of charge carriers contributing to the permittivity. For εDoped, the PEDOT: PSS complex permittivity in its doped state, N = NDopedand represents the number density of polarons and bipolarons in the material.

[0359] Conversely, for εDedoped, the dedoped-state permittivity of PEDOT: PSS, N = NDedopedand represents the number density of electrons (absence of polarons and bipolarons) in the material. These number densities can be expressed as a function of the total density of carriers participating in the electrochemical doping and undoping process, NTot, as follows:

[0360] N Tot Doped T N Dedoped (X ' N Tot 3“ ( 1 tt) ' N Tot Equation 7

[0361] Where the doping level, a =NDoped

[0362] N, isdefined as the ratio between the doped state and the Tot

[0363] total carrier density. From there, computing the PEDOT: PSS complex permittivity for intermediate doping levels (0 < a < 1) comes down to a linear interpolation between the extreme permittivity εDoped(ω) and εDedoped(ω) as per the following relation:

[0364] εPEDOT:PSS(α,ω) = α · εDoped(ω) + (1 − α) · εDedoped(ω) Equation 8

[0365] The resulting complex permittivities are presented in Fig. 6C-D.

[0366] Relationship between voltage biases and PEDOT: PSS complex permittivity

[0367] PEDOT doping levels can be adjusted by applying an external potential across the PEDOT: PSS material in an electrolyte (see Example 5). The link between doping levels and voltage amplitudes was experimentally established by performing cyclic voltammetry while monitoring the relative variation of scattered light intensity (Z score) from PEDOT: PSS organic electro-scattering antennas (OCEANs, Fig. 6H). At positive voltages, the PEDOT is quasi-fully doped and Z score is minimum. At negative voltages, the PEDOT is quasi-fully dedoped and Z score is maximal. The linear voltage sweep across these two extreme voltage levels resulted in a sigmoidal relationship between Z score and the applied voltage, which is reminiscent of the PEDOT: PSS doping level variations with voltage biases [Rivnay, J. et al. Nat. Rev. Mater. 3, 1-14 (2018)]. Assuming that the optical signal is only modulated by a

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[0370] change in PEDOT: PSS doping levels, fitting it with a sigmoid function enables the expression of the relationship between doping levels and voltage amplitudes as follows:

[0371] α(V) = 1 / (1 + e(V−V₀) / β) Equation 9

[0372] Where V0= −445 mV is the sigmoid centering voltage where α = 0.5, and β = 135 mV is a fitting parameter that defines the steepness of the transition.

[0373] The PEDOT: PSS complex permittivity can therefore be expressed as a function of the applied voltage across the PEDOT: PSS-phosphate-buffered saline (PBS) interface by the following expression:

[0374] εPEDOT:PSS(V,ω) = α(V) · εDoped(ω) + (1 − α(V)) · εDedoped(ω) Equation 10

[0375] Fig. 6E-F shows the real and imaginary parts of the PEDOT: PSS permittivity under different voltage bias conditions. The variation of the real part of the permittivity with respect to wavelength for different bias levels is shown in Fig. 6G. At negative bias voltages and large wavelengths, the real permittivity of PEDOT: PSS exhibits maximum sensitivity to voltage.

[0376] Example 7

[0377] Nanofabrication of OCEANs

[0378] The complete process flow developed to manufacture and interface OCEANs, as well as the method used to systematically characterize the OCEAN geometries following each fabrication batch, are presented and discussed in detail in this section.

[0379] Extended nanofabrication of OCEANs

[0380] Fig. 7A illustrates the complete nanofabrication process developed to manufacture OCEANs. The process started with a 170 pm-thick indium tin oxide (ITO)-coated glass substrate. First, the 70 nm thick ITO layer was patterned by chlorine-based reactive ion etching to define conductive traces and pads, eventually enabling electrical access to OCEANs. These ITO structures were subsequently passivated with a 50 nm-thick layer of silicon nitride (SiNx) deposited by plasma-enhanced chemical vapor deposition. A 3 pm-thick layer of SU8 was patterned with 80 pm square openings by photolithography to maximize the electrical trace stray impedances while permitting direct access to the ITO-SiNxpads for 37

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[0382] OCEAN patterning. Focused-ion beam (FIB) lithography was used to pattern arrays of 250 nm in diameter nanoholes in the exposed SiNx layer. Electrodeposition of PEDOT: PSS was then performed to locally grow OCEANs from each cavity and form OCEAN arrays.

[0383] The OCEAN chip was electrically interfaced with a custom-designed printed circuit board (PCB) - before PEDOT: PSS electrodeposition - by dispensing conductive epoxy on each PCB interfacing pad and employing a die bonder to position the chip on the PCB precisely (Fig. 7B). The PCB comprised four conductive layers. The two outer layers formed ground planes that shielded the two inner layers patterned with electrical traces connecting the PCB standard connectors to the different electrical sections of the chip. Each chip comprised 25 arrays of OCEANs - each composed of 16 x 16 OCEANs - electrically interfaced by 12 conductive traces (2 arrays per conductive trace for the first 11 traces and 3 arrays for the last one). Any electrical instrument could independently address each trace using the 12-position DIP switch connected to a standard U. FL connector marked WE for working electrode. The two other U. FL connectors designated as CE and REF for counter and reference electrodes, respectively, were directly connected to standard pin connectors where platinum and silver / silver chloride (Ag / AgCl) reference electrodes could easily be mounted and positioned to contact the bath. The custom-designed PCB permitted robust and low-noise interfacing of OCEANs with a three-electrode potentiostat and strongly facilitated their fabrication and characterization. The Omnetics connector on the right side of the PCB will be used in subsequent work to interface an array of microelectrodes seamlessly integrated into OCEANs and used as control. The U. FL connector labeled STIM for stimulation will be used to interface an integrated stimulation dipole, potentially useful to pace the culture.

[0384] Geometrical characterization of OCEANs

[0385] To account for potential variability in the process, the PEDOT: PSS electrodeposition step was systematically monitored under an optical microscope in real-time. At the end of the process, a custom algorithm was used to estimate the diameter of the freshly developed OCEANs and their variability. Fig. 8 presents the algorithm's outcome for the three electrodeposition times (IED) studied in this work. Assuming an isotropic electrodeposition process, the heights of the antenna cap (HOCEAN) can be estimated from the cap and stem diameter (0capand 0stem, respectively) using the following equation.

[0386] JJ _ capNst

[0387] H OCEAN = — ~2- Equation 11

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[0390] Consequently, the height of the small, intermediate, and large antennas (i.e., 0cap = 0.7 pm, 1.4 gm, and 1.8 gm) was estimated to be 225 nm, 575 nm, and 775 nm, respectively (0stem = 250 nm).

[0391] Example 8

[0392] Electrochemical characterization of OCEANs

[0393] The electrochemical properties of OCEANs were systematically characterized in PBS. The outcome of this procedure is presented in detail in the following section.

[0394] Experimental data acquired by electrochemical impedance spectroscopy (EIS) are first introduced and discussed, and then followed by the implementation of a simple equivalent electrical circuit describing the PEDOT: PSS OCEAN- electrolyte interface.

[0395] Electrochemical impedance spectroscopy on OCEANs

[0396] Fig. 9 presents the EIS experimental data acquired for OCEAN arrays composed of OCEANs with diameters ranging from 0.7 pm to 1.8 pm and voltage biases between 0 V and -0.6 V. At fixed bias, the OCEAN-electrolyte interfaces behave similarly to the conventional microelectrode-electrolyte interfaces, suggesting a double layer constant phase element in serial with a spreading resistance, and in parallel with a stray capacitance as an electrical equivalent circuit [Franks, W. et al. IEEE Trans. Biomed. Eng. 52, 1295-1302 (2005)]. However, an additional resistance, whose value is modulated by the bias voltage amplitude, is necessary to explain the EIS plot's dependence on voltage (Fig. 10A). This resistance, RPEDOTPSS, represents the electrical resistance of the PEDOT strands and becomes particularly significant when PEDOT: PSS becomes dedoped.

[0397] Modeling of the OCEAN-electrolyte interface

[0398] Fig. 10A displays the electrical equivalent circuit implemented to fit the EIS experimental data presented in Fig. 9. It comprises a constant phase element CPE,PEDOT:PSSrepresenting the transfer of charge at the PEDOT:PSS-electrolyte interface (of impedance ZCPE= 1 / (Q0(jω)n), where Q0is a measure of ZCPEmagnitude, n is a constant between 0 and 1,

[0399]

[0400] and a> is the pulsation), a spreading resistance Rs, a stray capacitance Cstray, and a voltagedependent resistance RPEDOTPSS accounting for the loss of conductivity in PEDOT during dedoping [Proctor, C. M. et al. J. Polym. Sci. Part B Polym. Phys. 54, 1433-1436 (2016)].

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[0403] Fig. 10B-E summarize the numerical value of each component included in the electrical equivalent circuit and how they are modulated by the OCEAN geometry and operating voltage. The computed respective time constants τ = (RPEDOT:PSS+ Rs) · CPEDOT:PSSare presented in Fig. 10F. The interface was assumed to be purely capacitive for simplification purposes, and consequently, CPEDOT SS was used instead of CPEPEDOTPSS ZCPEDOT: PSS= 7. CPE, for n = 1). At negative biases, the OCEAN kinetics were anticipated to become slower due to the increased RPEDOT:PSS. Additionally, larger OCEANs were expected to show slower time constants due to larger capacitances. Nevertheless, increases in RPEDOT:PSS + Rs partially compensated for the increase in CPEDOTPSS in the case of the largest OCEANs (tED = 60 s), and, therefore, their time constants remained comparable to OCEANs of intermediate size.

[0404] Example 9

[0405] Electro-optic characterization of OCEANs

[0406] This section provides a comprehensive description of the experimental setup used to characterize the electro-optic characteristics of OCEANs and presents the complete dataset collected during this process.

[0407] Relative irradiance spectra of OCEAN arrays for different voltage biases

[0408] Relative irradiance spectra of OCEAN arrays (50 x 50 OCEAN per array, ~1.8 pm cap diameter) were acquired in PBS while applying different voltage biases between the ITO and the bath. The OCEAN array s were illuminated by a broadband light source through the diascopic dark-field condenser of an inverted microscope. The light scattered by the array was collected by a 40x objective and analyzed by an optical spectrometer connected to the output port of the microscope via an optical fiber. The acquired spectra were normalized by the light source spectrum to compensate for its non-uniformity. Voltage biases were applied across the OCEANs using a potentiostat in a three-electrode configuration, where the ITO acted as the working electrode, an Ag / AgCl reference electrode as the reference electrode, and a platinum wire as the counter electrode. The resulting spectra are presented in Fig. 11. At negative voltages, the PEDOT composing the OCEANs becomes dedoped, which increases the real part of its permittivity (Fig. 6E). As a result, the OCEANs are more effective at scattering light.

[0409] Total internal reflection darkfield microscopy

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[0412] A custom total internal reflection darkfield microscopy setup was developed to confine the incident illumination to the plane defined by the array of OCEANs and collect only the light they scattered. The optical setup comprises an upright microscope body coupled to a complementary metal-oxide semiconductor (CMOS) camera, a 60 x water dipping objective, a multiwavelength light source to image the device using reflected brightfield microscopy, and a custom-designed holder mounted on a motorized XY stage (Fig. 12A and C). To confine the illumination to the plane defined by the OCEANs, a prism-based total internal reflection illumination module was integrated into the setup (Fig. 12B). The output of a 637 nm pig-tailed laser diode was collimated, mounted at an angle of 30° with respect to the horizontal plane, and aligned with an N-BK7 prism to illuminate OCEANs above the critical angle and achieve total internal reflection. Immersion oil was placed between the prism and the sample to ensure a continuous high-refractive index medium. Under these conditions, an evanescent wave propagated vertically towards the electrolyte with an amplitude decreasing exponentially in the direction of propagation, thus confining the excitation light to the first hundreds of nanometers following the SiNx-electrolyte interface. Such illumination will permit the minimization of background scattering from cells and, therefore, will play a central role in maintaining a low limit of detection when translating to biological experiments. Furthermore, because the laser is collimated, it enables the imaging of entire arrays of OCEANs with submicrometer spatial resolution, facilitating data collections during characterization workflows and, potentially, enabling wireless electrophysiological studies in cell networks with subcellular resolution. Note that the incident illumination should come from the direction opposite to the microscope objective to collect the forward scattering from the OCEANs, whose intensity is brighter than the backward scattering. A three-electrode potentiostat was used to apply voltage stimuli across OCEANs using an external platinum wire and an Ag / AgCl reference electrode as counter and reference electrodes, respectively.

[0413] Optical and electrical cyclic voltammetry with OCEANs

[0414] Fig. 13A-C shows optical and electrical voltammograms measured from OCEANs of different dimensions. The two different cycles composing each optical and electrical plot show, overall, similar behaviors. However, electrical traces displayed redox currents at voltages around -0.1 V and 0.1 V. Interestingly, their amplitude decreased with cycles, and their contribution to the optical signal variation seemed minimal, especially for larger

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[0417] OCEANs. Fig. 13D-F shows total internal reflection darkfield microscopy images of OCEAN arrays of different dimensions under a voltage bias of -0.8V vs Ag / AgCl.

[0418] Electro-optic responses of OCEAN to voltage pulses

[0419] The influence of cap diameters and operating biases on OCEAN electro-optic performance was systematically characterized by applying voltage pulses of amplitudes ranging from -100 mV to 100 mV across OCEANs in PBS while monitoring their optical Z score individually. Figs. 14-16 present the overall resulting dataset.

[0420] Dynamic characteristics of OCEANs

[0421] The time constants of single OCEANs were experimentally determined for different cap diameters and operating biases. 200 ms-long and 100 mV amplitude voltage pulses were applied across the OCEANs in PBS while monitoring their optical signal at 2’000 frames per second. The optical traces from each OCEAN were fitted with an exponential to estimate their time constant. Fig. 17A-C present the average experimental and fitted optical traces used to study the influence of cap geometries and operating biases on the dynamic characteristics of OCEANs. Fig. 4H summarizes the experimentally measured time constants for the different conditions tested.

[0422] Long-term stability of OCEANs

[0423] The long-term stability of OCEANs was investigated by applying a 1 Hz and 100 mV in amplitude voltage square wave with respect to an operating bias of -0.3 V for 10 hours. The optical responses of single OCEANs were intermittently monitored throughout the experiment and are presented in Fig. 18. Interestingly, the average Z score only decreased from 6 to 5 following 36’000 electrochemical modulation cycles, demonstrating the remarkable long-term stability of most OCEANs (Fig. 4G). Nevertheless, both the standard deviations and the number of outliers also increased throughout the duration of the experiment, suggesting that some OCEANs showed deteriorated electro-optic properties. No morphological differences were observed between the OCEANs showing long-term stability versus the other ones. This heterogeneity could originate from a deterioration of the ITO-PEDOT: PSS interface with time, potentially through delamination, hindering electrochemical doping and dedoping in some OCEANs. Alternatively, variability in the structural reorganization of PEDOT: PSS structure in response to cyclic stimulation could also explain the

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[0426] deteriorated electro-optic properties. For instance, the repeated injection of solvated cations in the PEDOT PSS structure could wash away excess of the hydrophilic poly(sodium 4-styrene sulfonate) and affect the electro-optic properties of the OCEAN. In such a case, posttreatment with strong sulfuric acid could be investigated to remove the excess of poly(sodium 4-styrene sulfonate) and re-organize PEDOTPSS before experiments [Kim, N. et al. Adv. Mater. 26, 2268-2272 (2014)]. The slower OCEAN kinetics observed after hours of stimulation suggests a steady dedoping of PEDOTPSS. which could potentially be explained by similar phenomena. Nevertheless, further investigations should be performed to accurately identify the underlying mechanisms limiting the long-term stability of some OCEANs.

[0427] Performance summary and comparison to state-of-the-art voltage-to-light

[0428] transducers

[0429] Table 2 summarizes OCEAN electro-optic characteristics with respect to electrodeposition time and bias voltage.

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[0432] Table 2 Summary of OCEAN experimental electro-optic characteristics with respect to voltage bias and antenna diameter (mean ± standard deviation)

[0433] Voltage

[0434] Limit of detection [mV| Time constant bias Sensitivity |mV '] Noise SNR

[0435] [ms]

[0436] |V|

[0437] tED [s| 30 45 60 30 45 60 30 45 60 30 45 60 30 45 60 0 -0.0131 -0.017 ± -0.0211 0.234 ± 0.1301 0.1931 6.191 18.461 14.901 22581 7671 9.101 6.11 29.91 175.01 0004 0003 0006 0064 0039 0065 251 413 332 21 18 211 162 1 8 427 2603 -0.1 -0.019 ± -0.031 1 -0.036 ± 0.237 ± 0.123 ± 0.1741 10.581 32.521 28.081 15351 4061 4.801 5.61 19.31 141.51 0.006 0.004 0.008 0.085 0017 0.065 4.36 5.10 5.16 13.61 0.84 0.85 1.9 15.7 171.2 -0.2 -0.025 + -0.044 ± -0.059 ± 0.246 ± 0.1271 0.191 + 14.081 42.401 40.301 10761 2901 3.20 + 6.1 ± 21.81 228.01 0.007 0.003 0.013 0.085 0021 0.061 5.92 7.33 7.09 6.25 0.51 0.56 3.0 19.2 346.5 -0.3 -0031 -0.0541 -0.082 ± 0.246 ± 0.1321 0.1961 16.081 48.181 53.471 9.231 2471 2.341 4.81 34.71 235.41 0.008 0.004 0.017 0.082 0026 0.065 6.82 1093 11.05 5.72 0.49 0.51 0.6 39.1 284.2 -0.4 -0031 ± -0053 ± -0084 + 02421 0 1371 0 192 + 14941 4252 + 51 661 867 + 259 + 2231 60 + 3271 23391 0.007 0.003 0.016 0.078 0021 0.071 5.66 5.75 11.76 4.34 0.42 0.50 1.1 20.4 299.4 -0.5 -0.0281 -0.0501 -0.0681 0.246 ± 0.1461 0.2171 15.131 38.001 35.881 9.481 2961 3.101 7.01 36.61 120.71 0006 0003 0014 0094 0030 0074 647 854 667 525 067 058 08 194 540

[0438] 23.1563

[0439] -0.6 -0.0281 -0.043 ± -0.045 ± 0.249 ± 0.1501 0.2131 17.91 1 31.821 9.121 3491 4.661 81.41 50.51 119.1 1

[0440] +

[0441] 0.003 0.003 0.006 0.082 0037 0.085 6.58 8.23 3.44 0.80 1.49 99.6 16.2 70.9

[0442]

[0443] 9.2596

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[0446] Table 3 summarizes the electro-optic performance characteristics of OCEANs compared to state-of the- art transducers. To ensure a meaningful comparison, studies specifically focused on wireless technologies that operate under conditions analogous to those of OCEANs, allowing us to extract and directly compare relevant performance characteristics.

[0447] Table 3: Performance comparison between OCEANs and state-of-the-art voltage-to-light transducers.

[0448] Number of Limit of Time Technology Spatial resolution

[0449] recording sites detection constant Ecore [Zhou, Y. et Single Probe diameter: 25 pm 5 pV 0.2 ms al. J. Am. Chem.

[0450] Soc. 2022, (2022)]

[0451] Electroplasmonic

[0452] nanoanteruias Single Probe diameter: 100’s pm* —10’s mV* 0.191 ms [Habib, A. et al.

[0453] Sci. Adv. 5,

[0454] eaav9786 (2019)]

[0455] Probe diameter: 1.4 pm

[0456] This work Up to 12’000 2.5 mV 26.8 ms Pitch: 5 pm

[0457]

[0458] Estimated from the 20x objective field of view diameter used in Fig. 4 of [Habib. A. et al. Sci. Adv. 5, eaav9786 (2019)].

[0459] * Estimated from Fig. 4-B of [Habib, A. et al. Sci. Adv. 5. eaav9786 (2019)].

[0460] Example 10

[0461] Modeling intracellular cardiomyocyte action potential recording with OCEANs

[0462] A study was performed to investigate the feasibility of using OCEANs to record cardiomyocyte action potentials wirelessly following intracellular access by electroporation. The protruding geometry of the OCEAN, together with its ~1 pm dimension, promotes an enhanced coupling at the cell-OCEAN interface and, therefore, a large seal resistance [Abbott, J. et al. Acc. Chem. Res. 51, 600-608 (2018)]. The electrically insulated substrate also significantly contributes to minimizing the current leaks. A large seal resistance is essential to minimize the electrophysiological signal attenuation at the cell-OCEAN interface and guarantees recordings of quality [Desbiolles, B. X. E. et al. Nano Lett. 19, 6173-6181 (2019)]. It also plays a central role in enabling intracellular access to the OCEAN following electroporation, minimizing the required electroporation voltage amplitude, and confining pore formations to the junctional cell membrane [Desbiolles, B. X. E. et al. Microsyst.

[0463] Nanoeng. 6, 1-12 (2020)]. Under these conditions, the cell transmembrane potential is

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[0466] anticipated to be efficiently transferred across the PEDOT:PSS structure with minimal attenuation and to modulate its scattering properties.

[0467] Modeling the cardiomyocyte-OCEAN interface

[0468] An electrical equivalent circuit of the cell-OCEAN interface was developed to predict how much of the electrophysiological signal can be transferred across the OCEAN under two different electrical configurations: pseudo-current clamp and voltage clamp. A pseudocurrent clamp configuration represents the electrical arrangement used in multi-electrode arrays (MEAs), where the ideally infinite input impedance of the amplifier separates the electrode from the ground. However, under this condition, no voltage bias can be applied across the OCEAN, therefore minimizing their electro-optic sensitivity. On the other hand, a pseudo-voltage clamp configuration would permit the application of an operating bias voltage across OCEANs. Nevertheless, the bias voltage hypothetical clamping effects on the transmembrane potentials of cells should be investigated. In this section, we analyzed these two different electrical arrangements to better understand their respective implications on OCEAN recording capabilities and their potential adverse effects on cell electrophysiology. A Luo Rudy model was used to simulate the electrogenic characteristics of a ventricular cardiomyocyte [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)].

[0469] Fig. 19A shows a schematic representation of the electrical equivalent circuit developed to study the feasibility of electrophysiological recordings using OCEANs in a pseudo-current clamp configuration. Due to the infinite input impedance of the current source connected between the ITO and the bath, the current passing through the PEDOTPSS structure is purely dictated by the current source. When a small stimulation current pulse Istim is applied (Fig. 19B, left), the transmembrane potential Vm becomes slightly depolarized, while VPEDOTPSS only reflects the charge of the capacitance CPEDOTPSS in response to Istim, independently from Vm. At the end of the current pulse, Vm returns to its resting value, and CPEDOT: PSS discharges through the equivalent circuit; however, with a slower kinetic as no charges reach the ground through the current source (Istim = 0). In response to larger currents (Fig. 19B, right), the transmembrane potential Vm overcomes its threshold value and generates an action potential. Nevertheless, the potential across the OCEAN remains fully dictated by the stimulation current source and does not reflect the cell electrophysiological activity. When Istim is null, no current can flow from the cell to the OCEAN as Istim = IPEDOT: PSS. Consequently, VPEDOTPSS, and therefore the optical signal of OCEANs, cannot be modulated by the cell potential in a pseudo-current clamp configuration.

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[0472] In an ideal pseudo-voltage clamp configuration (Fig. 19C), the voltage source input impedance is null. As a result, current originating from the electrophysiological activity of a cell can flow through the PEDOT: PSS OCEAN, modulate the potential across its terminals, and, therefore, regulate its optical properties. In Fig. 19D, a bias voltage VBias of -0.1 V is applied between the ITO and the bath to simulate the operating voltage necessary to enhance the electro-optic sensitivity’ of the OCEAN. Interestingly, due to the large impedance of the PEDOT: PSS structure with respect to the rest of the circuit, most of VBias is transferred across the OCEAN and does not interfere with the transmembrane potential of the cell. Importantly, the cell transmembrane potential is not clamped by the external voltage source, compared to whole-cell voltage clamp experiments performed using the patch-clamp technique. In the event of an action potential, VPEDOTPSS is directly modulated by the cell electrophysiological activity, demonstrating the feasibility of using OCEANs to perform electrophysiological studies. In consequence, OCEANs should be used in a pseudo- voltage clamp configuration to record the transmembrane potentials of cells without interfering with their physiology. Note that the seal resistance of a single OCEAN can theoretically be estimated by measuring its optical time constant in response to a voltage pulse, as the PEDOT: PSS capacitance loading kinetic is directly related to the quality of the cell-OCEAN interface.

[0473] In summary, with MEAs, the amplifier's input impedance (i.e., between the electrode and the bath / ground) must be much higher than the electrode impedance to minimize the voltage drop across the electrode-electrolyte interface. Ideally, it should be infinite so that the full amplitude of the electrophysiological signal originating from the cell and reaching the electrode would be transferred to the amplifier input without attenuation. With arrays of OCEANs, the sensing element is the OCEAN itself. Consequently, in an ideal scenario, the voltage drop across the PEDOT: PSS structure should not be minimized like with MEAs but maximized so that the entirety of the electrophysiological signal amplitude will contribute to modulating the scattering properties of the OCEAN. This condition is achieved by connecting the ITO to the ground or a voltage source (i.e., null input impedance) when operating biases are needed.

[0474] The equations describing the pseudo-current and voltage clamp models presented in Fig. 19 are introduced below herein. The last part discusses the assumptions made during the definition of the main components of the cell-OCEAN interfaces as well as the numerical values that were used in the models.

[0475] Equations defining the cardiomyocyte-OCEAN interface model

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[0478] Pseudo current-clamp configuration (1)

[0479] From the electrical equivalent circuit presented in Fig. 19A. the following equations can be defined:

[0480] Vseal - Vm= Vj Equation 12 Where VSeal is the voltage drop across the seal resistance RSeal, Vj across the junctional cell membrane, and Vm the transmembrane potential.

[0481] Vseai = / Leak ’? Seal Equation 13 Where Leak is the leaking current passing through Rseai.

[0482] V j — Rj ' 12 ~ Rj ’ IJ ~ ( stim ■ Iiea ) Equation 14 Where Rj is the resistance of the junctional cell membrane, 12 the current passing through the junctional cell membrane, and IStim the applied stimulation current.

[0483] From Equations 12 to 14, the following relationship could be determined:

[0484] ILeak = (Rj·IStim+Vm) / (RSeal+Rj) Equation 15

[0485]

[0486] KSeal+tij

[0487] The current L passing through the cell membrane capacitance Cm can be expressed as:

[0488] Ic = Istim − (INa + Isi + IK + IK1 + IKp + IB) − ILeak Equation 16

[0489] Where INa, Isi, IK, IK1, IKP, and IB are the fast sodium currents, slow inward currents, timedependent potassium currents, time-independent potassium currents, plateau potassium currents, and background currents going through the cell membrane, as defined in the Luo Rudy model [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)].

[0490] As a result, the variation of transmembrane potential with time can be defined as:

[0491] dVm_ IC

[0492] dt CmEquation 17

[0493] Furthermore, because the current passing through the OCEAN-electrolyte interface IPEDOTPSS is equal to the stimulation current Istim, the variation voltage across the OCEAN VPEDOTIPSS with time can be expressed as follows:

[0494] dVPEDOT:PSS / dt = (1 / CPEDOT:PSS) · (IStim − VPEDOT:PSS / RCT) Equation 18

[0495]

[0496] dtr\ ^Stlm

[0497] LPEDOT-PSS V KD

[0498] CT

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[0501] This system of differential equations was solved numerically using Matlab to find VPEDOT:PSS and Vm. The code implementing the pseudo-current clamp model is available in the dedicated data repository (see above herein).

[0502] Pseudo-voltage clamp configuration (2)

[0503] The potential drop across the seal resistance Vseaican be expressed as follows:

[0504] VSeal = VBias − VPEDOT:PSS Equation 19

[0505] Where Visias is the applied voltage bias and VPEDOTPSS the voltage drop across the OCEAN-electrolyte interface. The junctional potential Vj across the junctional cell membrane is therefore defined:

[0506] Vj = VSeal − Vm Equation 20

[0507] With Vm, the transmembrane potential of the cell of interest. Following Ohm’s law, the resulting current b passing through the junctional resistance of the cell membrane Rj is described as:

[0508] 12= — Equation 21

[0509]

[0510] Ri

[0511] And Ic, the current passing through the cell membrane capacitance as:

[0512] Ic = I2 − (INa + Isi + IK + IK1 + IKp + IB − IGJ) Equation 22

[0513] Where INa, Isi, IK, IK1, IKP, and IB are the fast sodium currents, slow inward currents, timedependent potassium currents, time-independent potassium currents, plateau potassium currents, and background currents going through the cell membrane, as defined in the Luo Rudy model [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)]. IGJ is the stimulating current coming from the neighboring cells through gap junctions and triggering the action potential.

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[0516] The leaking current keak passing through Rseai is defined as:

[0517] eak = 7^ Equation 23

[0518] Rseai

[0519] And the resulting current IPEDOT: PSS passing through the OCEAN-electrolyte interface as:

[0520] IPEDOT:PSS = I2 + ILeak Equation 24

[0521] The variation of Vm and VPEDOT: PSS in time followed the following relationships:

[0522] dVm_ Ic

[0523] dt CmEquation 25

[0524] Where Cm is the cell membrane capacitance and:

[0525] dVPEDOT:PSS / dt = (1 / CPEDOT:PSS) · (IPEDOT:PSS − VPEDOT:PSS / RCT) Equation 26

[0526]

[0527] dtr’ I ‘PEDOT. PSS

[0528] CDI PEDOT: PSS VHCT >

[0529] Where Rcr the charge transfer resistance at the OCEAN-electrolyte interface.

[0530] This system of differential equations was solved numerically using Matlab to find VPEDOT PSS and Vm. The code implementing the pseudo-voltage clamp model is available in the dedicated data repository (see above herein).

[0531] Numerical values of the main components defining the cardiomyocyte-OCEAN interface model

[0532] The numerical values used for each component of the cell-OCEAN electrical equivalent circuit (Fig. 19) are summarized in Table 4. A 1.4 pm in diameter OCEAN operating at VBias = -0.1 V was evaluated. Note that non-faradaic coupling at the PEDOT: PSS-electrolyte interface was assumed purely capacitive and modeled by CPEDOT:PSS for simplification purposes. Additionally, RPEDOT: PSS and Rs (Fig. 10A) were considered negligible and a charge transfer resistance RCT was added to account for the faradaic charge transfer contributing to the discharging of CPEDOT:PSS at the OCEAN-electrolyte interface. Its value was experimentally measured to be 0.75 TQ by cyclic voltammetry. In the pseudocurrent clamp configuration, a value of 1 GQ was used instead to keep discharge time

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[0535] constant in a reasonable window. Ion channel reversible potentials and transconductances were taken from the Luo Rudy model [Luo, C. H. & Rudy, Y. Circ. Res. 68, 1501-1526 (1991)].

[0536] Table 4: Table summarizing the numerical value of each component comprised in the cell-OCEAN interface electrical equivalent circuit.

[0537] Component Numerical value Unit

[0538] CPEDOT: PSS 26.6 pF

[0539] RCT 0.75 TΩ

[0540] RJ 500 MΩ

[0541] RSeal 200 MΩ

[0542] Cm 90.8 pF

[0543]

[0544] Modeling the electro-optic characteristics of OCEAN with cells

[0545] Under total internal reflection illumination, an evanescent wave propagates vertically with an intensity that decays exponentially along the vertical axis. This wave is scattered by the antenna in an electrochemically dependent manner, enabling the wireless probing of local potential fluctuations. Due to the evanescent nature of the incident wave, only the portion of the cell in close proximity to the substrate is expected to interact with the light and potentially contribute to background scattering. For this reason, we focused on modeling the junctional cell membrane - 5 nm thick and located 20 nm above the antenna - and the cytoplasm (Fig.

[0546] 20 A). The spatial distribution of the electric field enhancement (Fig. 20B) demonstrates that the incident light contribution is minimal when it reaches the junctional membrane, validating our assumption. Fig. 20C to E further confirm that the electro-optic modulation of single OCEANs can be measured with minimal interference from the cell under total internal reflection. When an electrical stimulus from the cell modulates the scattering properties of the antenna (simulated as VStimulationin Fig. 20C), the intensity of the scattered light is modulated, regardless of the presence of the cell on the antenna (Fig. 20D, red and blue curves).

[0547] Although the absolute scattering cross-section is slightly reduced in the presence of the cell, the relative change in the optical signal remains nearly identical (Fig. 20E). Given the high brightness of the antenna, this small attenuation in absolute scattering cross-section is not expected to impact the electrooptic performance of the recordings. Overall, this model demonstrates the feasibility of using OCEANs in the presence of cells.

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[0550] Discussion (Examples 1-10)

[0551] OCEANs stand out by their fundamental mechanism that converts voltage fluctuations in an aqueous electrolyte into light variations, setting them apart from current state-of-the-art methods. For instance, OEM-based plasmonic antennas rely on their high charge density when highly doped (metallic state) to emit light through plasmonic resonance. By removing charges from the organic polymer through electrochemical dedoping processes, the brightness of the plasmonic antennas is quenched. The plasmonic resonance wavelength of PEDOT: PSS nanoantennas - defined by its geometry, complex permittivity, and the permittivity of the surrounding medium - is typically limited to the infrared window, which hinders experiments on regular biomicroscopy setups. Conversely, OCEANs leverage the dielectric state (low charge density ) of dedoped PEDOT: PSS to emit light through scattering in the visible domain. The maxima in PEDOT: PSS real dielectric permittivity near the 700 nm wavelengths leads to bright scattering characteristics. By injecting charges in the PEDOT: PSS through electrochemical doping processes, the OCEAN scattering is quenched. This fundamental principle difference enabled OCEANs to perform efficiently in the visible domain, where current biomicroscopy setups are optimal.

[0552] To achieve satisfactory SNRs, optical signals originating from multiple polymer plasmonic and electroplasmonic antennas are typically pooled together, hindering their inherent spatial resolution and number of recording sites in the context of biosensing. These features are also limited in electrode and OECT-based sensors due to the conductive traces connecting each sensing unit to an electrical instrument. Conversely, OCEANs enable wireless probing of potentials with high SNRs at the single-antenna level, eliminating the need to pool contributions from multiple sites and for individual conductive traces.

[0553] Consequently, they permit unprecedented spatial resolution and an outstanding number of recording sites compared to existing biosensing technologies. Notably, the pitch between two OCEANs in an array was set to 5 pm throughout this work (details below herein and Fig. 13D-F), theoretically leading to 4 million OCEANs on a 1 x 1 cm2area. When using a 60x or a 40x objective on a 25 mm in diameter field of view microscope (417 pm and 625 pm in diameter on the sample, respectively), it was estimated that -5’000 or -12’000 OCEANs, respectively, could be imaged simultaneously. Prospectively, this number could be enhanced by decreasing the pitch between OCEANs. Simultaneous multisite imaging is a significant advantage of OCEANs compared to absorbance-based technologies, such as ECORE, which can only monitor a single recording site. Table 3 below summarizes the electro-optic

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[0556] performance characteristics of OCEANs compared to state-of-the-art voltage-to-light transducers.

[0557] The theoretical model presented in this work describes how intermediate voltage biases drive PEDOTPSS intermediate doping levels in PBS and how its complex permittivity is affected. These findings enabled accurate prediction of the electrochemical modulation of OCEAN scattering spectra and anticipate the central role of cap diameters and operating bias voltages on the antenna brightness and sensitivity. While larger structures were shown to be brighter and more sensitive than their smaller counterparts, their time constant was also predicted to be slower following the systematic electrochemical characterization of the OCEAN-electrolyte interface by electrochemical impedance spectroscopy. Due to this tradeoff between electro-optic modulation and dynamic performances, antennas of small, intermediate, and large diameters were fabricated to determine an optimal dimension experimentally.

[0558] The nanofabrication process developed to manufacture OCEANs was reliable and led to homogeneous structures across fabrication batches. The FIB lithography step was performed using a Velion FIB (Raith Nanofabrication, Germany). Compared to traditional FIBs, the Velion FIB was designed for large-scale nanofabrication. It comprises a laser-interferometric stage enabling stitching with a resolution of 1 nm and, therefore, permitting nanoscale patterning across large areas. Additionally, its vertical ionic column uses gold double-plus ions (Au++). whose sputtering yield is higher than gallium ions, therefore minimizing the patterning time. These characteristics enhanced the scalability of the process. Importantly, while using the Velion FIB proved convenient for rapid design iterations, it was not essential to successfully fabricate OCEANs. Any conventional (e g., electron-beam or deep ultraviolet lithography) or alternative (e.g., scanning probe lithography) nanofabrication technique capable of patterning 250 nm hole arrays in a 50 nm-thick layer of silicon nitride would be suitable for manufacturing OCEANs, ensuring broad accessibility to the proposed technology. For example, using a stepper, 250 nm diameter openings could be patterned in a photoresist layer at the wafer scale and subsequently transferred into the silicon nitride by reactive ion etching. This approach could offer a high-throughput manufacturing solution for OCEANs, significantly reducing the cost per chip. The optical monitoring of PEDOT: PSS electrodeposition in real time enabled compensating for potential vari ability in the process and was pivotal in establishing consistency between batches.

[0559] The total internal reflection dark -field microscopy setup that was developed with these studies enabled the selective collection of the light scattered by OCEANs over the 53

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[0561] incident light. Compared to conventional dark-field microscopy, this approach minimizes the background scattering - which could come from the cell culture - and ensures a consistent translation of OCEAN electro-optic characteristics to biological experiments. Additionally, compared to methods using single photodetectors, it enables the formation of an image of the array with submicrometer resolution, providing spatial information. The illumination wavelength of 637 nm was not only selected to maximize OCEAN electro-optic characteristics but also to minimize potential adverse effects on cells. In addition to phototoxicity, the photobleaching of pharmacological compounds at shorter wavelengths is a well-known issue that could have hindered the applications of this technology7. For instance, blebbistatin - a cardiomyocy te contraction inhibitor widely used to prevent mechanical artifacts with optical electrophysiology techniques - is known to photo-bleach at wavelengths below 500 nm [Kolega, J. Biochem. Biophys. Res. Commun. 320, 1020-1025 (2004)]. Hence, using a broadband light source or a blue-shifted illumination w ould inhibit the blebbistatin effect and render electrophysiological studies with cardiomyocytes irrelevant.

[0562] The systematic experimental electro-optic characterization of OCEANs demonstrated that single units could transduce potential fluctuations into light variations with SNRs of 48 for 100 mV voltage pulses and limits of detection of 2.5 mV. The theoretically predicted relationships linking operating voltage biases and cap diameters to the antennas’ brightness and sensitivity are consistent with the experimental observations. In particular, the fact that positive voltage pulses led to negative variations of the OCEAN optical signal demonstrates that the fundamental mechanism governing OCEAN is not based on absorbance- where a positive voltage pulse w ould induce a positive fluctuation of the optical signal - but instead relies on the modulation of the scattering properties of the device. The relationship between the scattering signal and voltage is linear near the optimal operating voltage bias. A systematic comparison between the electro-optic performance of OCEANs and state-of-the-art voltage-to-light transducers is difficult as experimental conditions and outcomes are rarely comparable. For instance, ECORE show ed an enhanced limit of detection compared to individual OCEANs but with an illumination intensity orders of magnitude higher [Alfonso, F. S. et al. Proc. Natl. Acad. Sci. U. S. A. 117, 17260-17268 (2020)]. Interestingly, single antennas exhibited an improved limit of detection compared to arrays of electroplasmonic nanoantennas in physiological solution under reasonably equivalent experimental conditions (estimated from Fig. 4B of Habib, A. in Sci. Adv. 5, 2019). Nevertheless, potential strategies to improve the limit of detection of OCEANs should be explored, including optimizing the optical setup and exploring materials with enhanced sensitivity. The optical setup-OCEAN 54

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[0564] system as a whole defines the detection limit. In particular, the optical setup’s mechanical stability should be prioritized, as methods relying on scattering with coherent light sources are naturally prone to speckle formation resulting from partial reflections at optical elements [Peters, A. et al. HardwareX 14, e00424 (2023)]. This background interference pattern can be quickly altered by changes in the optical path due to temperature fluctuations, laser wavelength drift, or mechanical instability and was identified as the primary source of noise in the experiments. Enhancing the optical setup mechanical stability through passive or active mechanisms42 can potentially improve OCEAN’S detection limit. Other polymers and treatments optimized in the context of OECTs should be investigated for material optimization [Rivnay, J. et al. Nat. Rev. Mater. 3, 1-14 (2018)]. Alternatively, where bandwidth is not a limiting factor, increasing the antenna diameter beyond 1.8 pm could potentially enhance OCEANs sensitivity and limit of detection, as long as the noise could remain controlled.

[0565] Furthermore, experimental measurements of OCEAN dynamic characteristics agree with the theoretical predictions of time constants ranging from a few milliseconds to more than a hundred milliseconds at optimal biases for small and large structures, respectively. Currently, OCEAN time responses are mostly limited by large values of RPEDOT PSS, mainly imposed by the small stem diameter of the PEDOT: PSS structure. Prospectively, faster dynamic characteristics could be achieved by modifying the OCEAN geometry from a mushroom-like to a disk-like structure of the same volume but with an augmented contact area with the ITO electrode. This way, the volumetric PEDOT: PSS capacitance would remain constant and RPEDOTPSS be reduced, thus enabling faster time constants.

[0566] Finally, OCEANs demonstrated exceptional long-term characteristics with continuous electro-optic modulation capabilities for at least ten hours. This long-term stability is a distinct advantage over fluorescent-based indicators (e.g., voltage-sensitive dyes or genetically encoded voltage indicators), which photobleach within minutes. It demonstrates the potential of OCEANs to perform biosensing studies over time.

[0567] As a case study, a Luo Rudy-derived analytical model of the cell-OCEAN interface was developed to investigate the feasibility of recording intracellular cardiomyocyte action potentials using OCEANs (see Example 10 for further details). The model's outcome suggests that even for a conservative estimate of the seal and junctional resistance Rseai = 200 MQ and Rj = 500 MQ [Spira, M. E. & Hai, A. Nat. Nanotechnol. 8, 83-94 (2013); Abbott, J. et al. Acc. Chem. Res. 51, 600-608 (2018)]. respectively, a significant voltage drop of 30 mV - far above the detection limit of OCEANs - is expected across the PEDOT: PSS structure.

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[0570] demonstrating the feasibility of using OCEANs for intracellular electrophysiological studies. Additionally, theoretical modeling of OCEAN electro-optic scattering properties in the presence of cells revealed minimal scattering contribution from cells under total internal reflection illumination, indicating that similar electro-optic performance can be expected during biosensing applications (see Example 10 for further details). Compared to state-of-the-art plasmonic antenna arrays and absorbance-based technologies, OCEANs protrude from an otherwise insulating substrate. This ensures an enhanced seal resistance at the cell-OCEAN interface and enables intracellular access following electroporation, maximizing the electrophysiological signal amplitude transferred to the sensor. Interestingly, the next iteration of OCEANs could have their stem exposed - similar to mushroom-like microelectrodes - to leverage cell engulfment and reinforce their interface with the cell [Hai, A. et al. J. Neural Eng. 6, 66009 (2009); Fendyur, A. & Spira, M. E. Front. Neuroengineering 5, 1-10 (2012); Shmoel, N. et al. Sci. Rep. 6, 1-11 (2016)]. Besides demonstrating feasibility, the presented model illustrates fundamental aspects to consider when using OCEANs for biosensing compared to conventional electrode-based sensors (see Example 10 for further details). In this context, OCEANs have the potential to enhance both the spatial resolution compared to multi-electrode arrays and the long-term stability compared to fluorescent reporters, thus potentially opening new opportunities in electrophysiology7. As a next step, proof-of-concept electrophysiological recordings from electrogenic cell networks is performed to demonstrate the feasibility of OCEANs experimentally. Cardiomyocyte monolayers are used due to their cell-electrode interface reliability and their electrical activity synchronicity. Hence, integrating planar microelectrodes with OCEAN arrays could provide simultaneous ground truth signals. Importantly, OCEANs of various geometries are explored to enhance the cell-OCEAN interface and improve the quality of the recordings.

[0571] Conclusion (Examples 1-10)

[0572] The studies presented herein introduce the concept of organic electro-scattering antennas and demonstrate their feasibility to wirelessly probe electrical activity with high spatial resolution for biosensing applications. OCEAN is the first technology leveraging the intrinsic dependence of PEDOT: PSS scattering properties on its doping levels to transduce small potential fluctuations into variations of scattered light intensity in the visible spectral window7. A theoretical model describing the relationship between voltage, doping level, and complex permittivity in PEDOT: PSS was developed to predict the scattering properties of single OCEANs with varying dimensions and operating at different voltage biases in

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[0575] response to electrical stimulation. A reliable nanofabrication process combining nextgeneration FIB lithography with conventional microfabrication techniques was established to precisely manufacture arrays of OCEANs with diameters as low as 0.7 pm. The electro-optic properties of OCEANs were systematically characterized by applying various electrical stimuli in PBS while monitoring scattered light intensity using a custom-designed total internal reflection dark-field microscope. Single OCEANs showed SNRs up to 48 in response to a 100 mV voltage pulse with a limit of detection of 2.5 mV at optimal operating biases. Time constants ranged between 6.0 ms to 233.9 ms for structures of 0.7 pm and 1.8 pm, respectively. OCEANs also demonstrated exceptional long-term stability, enabling continuous electro-optic modulation for ten hours. Finally, an analytical cell-OCEAN model derived from Luo Rudy was implemented to demonstrate the feasibility of recording cardiomyocyte intracellular action potentials using OCEANs.

[0576] OCEANs can potentially enable functional readout studies from thousands of recording sites simultaneously, with micrometer spatial resolution, and over extended periods of time. Such recording characteristics make OCEANs a great candidate to overcome the technical limitations of current biosensing approaches, opening new opportunities in bioelectronics and, prospectively, accelerating progress in both fundamental and clinical research.

[0577] Equivalents

[0578] It is to be understood that the methods, compositions, and apparatus which have been described above are merely illustrative applications of the principles of the present disclosure. Numerous modifications may be made by those skilled in the art without departing from the scope of the present disclosure. Although the present disclosure has been described in detail for the purpose of illustration, it is understood that such detail is solely for that purpose and variations can be made by those skilled in the art without departing from the spirit and scope of the present disclosure.

[0579] Although several embodiments of the present disclosure have been described and illustrated herein, those of ordinary skill in the art will readily envision a variety of other means and / or structures for performing the functions and / or obtaining the results and / or one or more of the advantages described herein, and each of such variations and / or modifications is deemed to be within the scope of the present disclosure. More generally, those skilled in the art will readily appreciate that all parameters, dimensions, materials, and configurations described herein are meant to be illustrative and that the actual parameters, dimensions.

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[0582] materials, and / or configurations will depend upon the specific application or applications for which the teaching of the present disclosure is / are used. Those skilled in the art will recognize or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments of the present disclosure described herein. It is, therefore, to be understood that the foregoing embodiments are presented by way of example only and that, within the scope of the appended claims and equivalents thereto, the present disclosure may be practiced otherwise than as specifically described and claimed. The present disclosure is directed to each individual feature, system, article, material, and / or method described herein. In addition, any combination of two or more such features, systems, articles, materials, and / or methods, if such features, systems, articles, materials, and / or methods are not mutually inconsistent, is included within the scope of the present disclosure.

[0583] All definitions, as defined and used herein, should be understood to control over dictionary definitions, definitions in documents incorporated by reference, and / or ordinary meanings of the defined terms.

[0584] The indefinite articles “a” and “an.” as used herein in the specification and in the claims, unless clearly indicated to the contrary, should be understood to mean “at least one.” The phrase “and / or,” as used herein in the specification and in the claims, should be understood to mean “either or both” of the elements so conjoined, i.e., elements that are conjunctively present in some cases and disjunctively present in other cases. Other elements may optionally be present other than the elements specifically identified by the “and / or” clause, whether related or unrelated to those elements specifically identified, unless clearly indicated to the contrary.

[0585] The contents of all references, patents, and published patent applications cited throughout this application are incorporated herein by reference in their entirety.

[0586] What is claimed is:

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Claims

Attorney Docket No. MML-081W001CLAIMS1. A sensing device comprising:a substrate;an electrically conductive layer positioned on the substrate; andan electrochemical dopable material contacting the electrically conductive layer, wherein the electrochemical dopable material:has a refractive index in the visible domain that changes based on whether the electrochemical dopable material is subjected to a positive or negative bias voltage;converts received electrical signals into visible light; andscatters the visible light.

2. The sensing device of claim 1, wherein the substrate is optically transparent.

3. The sensing device of claim 2, wherein the substrate comprises glass.

4. The sensing device of claim 1, wherein the substrate is opaque.

5. The sensing device of claim 1, wherein the electrically conductive layer comprises at least one of metal and conductive polymer.

6. The sensing device of claim 5, wherein the metal comprises metal oxide.

7. The sensing device of claim 6, wherein the metal oxide comprises indium tin oxide.

8. The sensing device of claim 1, wherein the electrochemical dopable material comprises at least one organic semiconductor and at least one electrolyte.

9. The sensing device of claim 8, wherein the at least one organic semiconductor comprises at least one conductive polymer.

10. The sensing device of claim 9, wherein the at least one conductive polymer comprises poly(3.4-ethylenedioxythiophene).59#18667993vlAttorney Docket No. MML-081W00111. The sensing device of claim 8, wherein the at least one electrolyte comprises at least one polystyrene.

12. The sensing device of claim 11, wherein the at least one polystyrene comprises polystyrene sulfonate.

13. The sensing device of claim 8, wherein the at least one organic semiconductor and at least one electrolyte are present at a ratio in the range of 1: 1 to 1:20 v / v.

14. The sensing device of claim 1, further comprising an electrically insulating layer.

15. The sensing device of claim 14, wherein the electrically conductive layer is positioned between the substrate and the electrically insulating layer.

16. The sensing device of claim 14, wherein the electrochemical dopable material contacts both the electrically conductive layer and the electrically insulating layer.

17. The sensing device of claim 14, wherein the electrically insulating layer comprises at least one of silicon-containing material, electrically insulating polymer, and ceramic.

18. The sensing device of claim 17, wherein the silicon-containing material comprises silicon nitride.

19. The sensing device of claim 14, wherein:the electrochemical dopable material comprises a cap portion with a stem portion extending therefrom; andthe stem portion extends into and along an aperture in the electrically insulating layer.

20. The sensing device of claim 19, wherein the cap portion has a circular surface extending along the electrically insulating layer.

21. The sensing device of claim 20, wherein the circular surface has a diameter of 1 pm.60#18667993vlAttorney Docket No. MML-081W00122. The sensing device of claim 19, wherein the electrically conductive layer contacts an end surface of the stem portion.

23. The sensing device of claim 22, wherein the end surface is circular and has a diameter in the range of 10 nm to 1 mm.

24. The sensing device of claim 23, wherein the end surface is circular and has a diameter of 250 nm.

25. A method of monitoring electrical potential of a biological material in a liquid medium, comprising:contacting the liquid medium with the sensing device of claim 1; and measuring the electrical potential of the biological material the liquid medium.

26. The method of claim 25, wherein the liquid medium is a physiological medium.

27. The method of claim 25, wherein the biological material comprises a cell.

28. The method of claim 27, wherein the cell is a neuron.

29. The method of claim 27, wherein the cell is a cardiac cell.

30. The method of claim 25, wherein the biological material comprises a protein.

31. The method of claim 25, wherein the biological material comprises a small molecule.

32. The method of claim 25, wherein the biological material comprises a ribonucleic acid (RNA) molecule.

33. The method of claim 25, wherein the biological material comprises a deoxyribonucleic acid (DNA) molecule.61#18667993vlAttorney Docket No. MML-081W00134. The method of claim 25, wherein measuring the electrical potential of the biological material comprises directing light at the sensing device at an angle in the range of 61° to 90° with respect to vertical of the electrically conductive layer.

35. The method of claim 34, wherein the light has a wavelength in the range of 380 nm to 700 nm.

36. The method of claim 35, wherein the light has a wavelength of 637 nm.

37. The method of claim 25, wherein measuring the electrical potential of the biological material comprises using dark-field microscopy.

38. The method of claim 25, wherein measuring the electrical potential of the biological material comprises measuring visible light scattered by the electrochemical dopable material of the sensing device using a microscope.

39. A sensing array comprising a plurality of the sensing device of claim 1.

40. A method of monitoring electrical potentials in a liquid medium, comprising:contacting the liquid medium with the sensing array of claim 39; and measuring the electrical potential of the liquid medium.

41. The method of claim 40, wherein the liquid medium is a physiological medium.

42. The method of claim 41, wherein the physiological medium comprises a cell.

43. The method of claim 42, wherein the cell is a neuron.

44. The method of claim 42, wherein the cell is a cardiac cell.

45. The method of claim 41, wherein the physiological medium comprises a protein.

46. The method of claim 41, wherein the physiological medium comprises a small molecule.62#18667993vlAttorney Docket No. MML-081W00147. The method of claim 41, wherein the physiological medium comprises a ribonucleic acid (RNA) molecule.

48. The method of claim 41, wherein the physiological medium comprises a deoxyribonucleic acid (DNA) molecule.

49. The method of claim 40, wherein measuring the electrical potential of the liquid medium comprises directing light at the sensing array at an angle in the range of 61° to 90° with respect to vertical of the electrically conductive layer.

50. The method of claim 49, wherein the light has a wavelength in the range of 380 nm to 700 nm.

51. The method of claim 50, wherein the light has a wavelength of 637 nm.

52. The method of claim 40, wherein measuring the electrical potential of the liquid medium comprises using dark-field microscopy.

53. The method of claim 40, wherein measuring the electrical potential of the liquid medium comprises measuring visible light scattered by the electrochemical dopable materials of the sensing array using a microscope.63#18667993vl

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