Electrochemical Electrode Assembly

Abstract

An electrochemical electrode assembly having micro / nanoscale features, particularly at least one micro / nanoscale working electrode "active" surface for selective oxidation / reduction of a target analyte or its reaction product. The working electrode of the present invention has a surface area that is minimized relative to the surface area of ​​the surrounding insulating material surface, and target analytes in a biological fluid (e.g., blood, interstitial fluid, sweat, saliva, etc.) are required to diffuse over at least one region of the insulating surface toward the working electrode surface.

Description

[Technical Field]

[0001] The present inventive concept relates to electrochemical electrode assemblies, particularly, but not exclusively, to microelectrodes suitable for use as part of electrochemical sensors. [Background technology]

[0002] Electrochemical sensors implemented in implantable or skin-contacting devices have attracted considerable interest in the electroanalytical field due to their ability to quantitatively and qualitatively characterize target analytes in complex biological systems. Such micrometer-scale sensors, amenable to existing fabrication techniques, enable analysis with high time resolution and sensitivity.

[0003] The first electrochemical sensor for oxygen was reported in the 1960s. In particular, the concept of glucose sensing (in vitro) was proposed by Clark and Lyons (based on an oxygen sensor) in 1962, and the first in vitro commercial electrochemical glucose sensor was manufactured by YSI (based on a hydrogen peroxide sensor) in 1975. Subsequently, the first commercial in vivo glucose sensor became available in 1999. In recent years, with the development of micro- and nanoscale fabrication technologies, micro / nanoelectrode array sensors have emerged as attractive biofluid-compatible sensors offering multiplexing capabilities and robustness for bioanalysis in a variety of different environments, such as cells, tissues, and organs, as implantable and noninvasive wearable devices.

[0004] Micro- or nano-electrode array sensors configured for smart sensing, as invasive or wearable intelligent devices, typically comprise electrodes with micro- and nano-scale dimensions. Various types of micro- and nano-electrode arrays have been developed to address different measurement requirements. Typically, such electrode arrays are fabricated by layer deposition or layer growth, in which layers of metal, carbon, ceramic, etc. are deposited on a template substrate, such as silicon, glass, polymer, or ceramic. Photolithography, screen printing, film formation, laser ablation, and 3D printing techniques are used to fabricate complex microscale electrode patterns. One example of an electrode assembly is described in International Publication No. 2010 / 061229, in which a laminate structure comprises an insulating cap layer disposed on a conductive layer designed to provide exposed electrical contact areas at etched voids extending through the cap layer.

[0005] However, existing microelectrochemical electrode assemblies and sensors have disadvantages for several reasons. Existing devices typically do not provide the isolation between individual electrode elements to fully realize the benefits of micro / nanoscale electrodes. Available structures and sensors involve the use of thin-layer techniques, which require precise and complex surface patterning to expose regions of the conductive layer. Furthermore, existing devices fail to adequately control the diffusion of target analytes (in the carrier biological fluid) into the conductive region, resulting in excessive amounts of analyte oxidation or reduction, which can only be mitigated by unnecessarily long pauses in the applied potential (to allow for the continued diffusion of fresh analyte to the active conductive surface) or the inclusion of a limiting membrane, which reduces both the signal and the time response. Therefore, there is a need for electrochemical electrode and sensor configurations that address the above issues. Summary of the Invention

[0006] The objective of the present concept is to provide an electrode assembly that can be conveniently and efficiently manufactured, and further to provide a device that is suitable for use as an implantable or wearable (skin contact) electrochemical sensor, providing improved sensitivity, reliability, response speed, and reproducible results.

[0007] A further object is to provide an electrode assembly and electrochemical electrode-based sensor that provides significantly improved analyte diffusion control (non-planar diffusion control) during use. A more specific object is to provide an assembly that achieves a time-independent diffusion-controlled current during operation (typically when polarized to a potential sufficient to oxidize or reduce the target analyte). A more specific object is to provide an electrode assembly that controls the diffusion of a target analyte in a biological fluid to the working electrode surface such that the diffusion rate of the target analyte is greater than the rate of oxidation / reduction at the electrode surface. A more specific object is to avoid the problem of significant target analyte volumetric depletion near / in the working electrode surface, which would otherwise manifest as significantly reduced or zero faradaic current in the working electrode region (depending on the target analyte or its reaction products being excessively or completely oxidized / reduced by the conductive surface coating material).

[0008] Reference herein to a target analyte encompasses species of interest and / or species produced due to the presence of the target analyte, such species may be produced by enzymatic reactions as described herein.

[0009] A further specific object is to provide an electrochemical sensor, particularly an implantable or wearable device, capable of analyzing multiple target analytes in series or parallel.

[0010] These objectives are achieved by the electrochemical electrode assembly of the present invention, which has micro / nanoscale features, particularly a micro / nanoscale working electrode "active" surface for selective analysis of target analytes. The working electrode of the present invention has a surface area that is minimized relative to the surface area size of the surrounding insulating material surface. According to the concept of the present invention, the conductive surface is defined and surrounded by at least one insulating material surface, so that target analytes in a biological fluid (e.g., blood, sweat, saliva, etc.) must diffuse toward the working electrode surface over at least a region of the insulating surface. By configuring the relative size, shape, and / or relative spatial arrangement of the conductive surface to the surrounding insulating surface, the diffusion of the target analyte in the biological fluid is dominated by hemispherical diffusion, which exceeds the consumption rate (i.e., reaction, oxidation, or reduction) of the target analyte at the conductive surface.

[0011] The present concepts are configured to effectively determine the concentration or rate of change of a target analyte in a fluid sample, e.g., a biological fluid, in response to the presence of the target analyte at the active electrode surface. According to certain implementations, the active electrode surface comprises an analyte-responsive layer having a material configured to chemically and / or physically bind and interact with the target analyte. In particular, the analyte-responsive layer may include an enzymatic material that chemically interacts or reacts with the target analyte to generate an electrochemically active product of an enzymatic reaction with the target analyte. The target analyte may include an electrochemically active product of an enzymatic reaction with an initial species. That is, the target analyte may be a chemical product of an enzymatic reaction with an initial target species. For example, if the present concepts are configured to determine glucose levels / concentrations, the target analyte may include an enzymatic reaction product with glucose, i.e., hydrogen peroxide. The present concepts may be configured to measure a current between the working electrode and the auxiliary electrode resulting from the direct oxidation or reduction of the target analyte. Alternatively, or in addition, the present concepts can be configured to measure the current between the working electrode and the auxiliary electrode resulting from the oxidation or reduction of an electrochemically active reaction product of the target analyte and the analyte-responsive layer. For example, if the target analyte is glucose, the species subject to oxidation and reduction at the active electrode surface may include hydrogen peroxide (which is an electrochemically active reaction product of an enzymatic analyte-responsive layer material coated and / or positioned proximate to the active electrode surface). References herein to the oxidation / reduction of a target analyte include the oxidation / reduction of the analyte directly or indirectly via the oxidation / reduction of a reaction product of the analyte.

[0012] Thus, the electrochemical method of analyzing at least one target analyte in a biological fluid of the present invention includes applying an oxidation or reduction potential to a working electrode to oxidize or reduce the target analyte or its reaction product, and measuring a faradaic current at the working electrode. Optionally, the step of applying an oxidation or reduction potential includes performing cyclic voltammetry.

[0013] According to a first aspect of the inventive concept, there is provided an electrochemical electrode assembly comprising: an electrically insulating material having at least one insulating surface; and an electrically conductive material having at least one conductive surface whose periphery is defined by and surrounded by the insulating surface, wherein the surface area of ​​the conductive surface, or each conductive surface if the assembly comprises multiple conductive surfaces, is less than 5% of the surface area of ​​the insulating surfaces surrounding the periphery of the conductive surface, or each of the insulating surfaces surrounding each periphery of each conductive surface, and wherein each surface area of ​​the insulating surfaces is defined between an inner boundary at the periphery of the conductive surface, or each inner boundary at the periphery of each conductive surface, and a respective outer boundary separated from the respective inner boundary by a respective distance of at least 5 times the diameter or smallest width of the conductive surface, or each conductive surface if the assembly comprises multiple conductive surfaces.

[0014] To control target analyte consumption at the working electrode surface, the inventive assemblies are configured such that the active oxidation / reduction electrode surface forms a discontinuity in the insulating material surface, presenting a significantly reduced "active" surface area relative to the surrounding insulating material surface. The inventive concepts are further configured with physical properties that promote the flow and diffusion of target analytes to the active working electrode surface. That is, the inventive concepts are configured to minimize or avoid locating the active surface within an obscured, restricted, or closed region or zone in the micro / nanoscale structure, where such restricted region would prevent the free flow of target analytes to the active surface. In one preferred implementation, both the active surface and the surrounding insulating surface are substantially flat and free of grooves, channels, pits, peaks, or troughs in the nano / microscale structure. However, depending on the particular manufacturing technique, such contouring may be required, and the active electrode surface may be provided on the sidewalls, edges, troughs, or valleys of a contoured / profiled nano / micro assembly structure, particularly a laminate structure in which the conductive layer is capped or defined with one or more insulating layers, and portions of the laminate structure are ablated or cut (by forming channels or cavities within the structure) to provide areas of exposed conductive layer surface that provide areas of active oxidation / reduction surface. Such configurations may provide active electrode surfaces within cavities that form part of a 3D array or structure. The electrochemical electrode assemblies of the present invention may be fabricated by a variety of techniques, including but not limited to atomic layer deposition (ALD), chemical vapor deposition (CVD), plasma-enhanced chemical vapor deposition (PECVD), slot-die coating, screen printing, inkjet printing, spray coating, 3D printing, and other techniques that are not limited to existing microprocessor or chip manufacturing techniques in the silicon chip manufacturing field.

[0015] Optionally, the surface area of ​​an individual conductive surface or each of the plurality of conductive surfaces is less than 300 μm 2 , 200 μm 2 , 150 μm 2 , or 100 μm 2 Optionally, the surface area of ​​the conductive surface or each of the plurality of conductive surfaces is less than 300 μm 2 , 200 μm 2 , 150 μm 2 , 100 μm 2 , 80 μm 2 , 60 μm 2 , 40 μm 2 , 20 μm 2 , 10 μm 2 , 5 μm 2 , 3 μm 2 , or 1 μm 2 is less than.

[0016] According to certain embodiments having a working electrode with minimized active surface area, the total conductive surface of the plurality of active conductive surfaces is less than 100 μm 2 Optionally, according to some implementations of the electrode configuration and / or electrochemical sensor, the total surface area of ​​the plurality of conductive surfaces may be less than 100 μm 2 However, in such a configuration, the surface area of ​​each conductive surface must not exceed 100 μm 2 It can be less than.

[0017] Optionally, the conductive surface, or if the assembly comprises multiple conductive surfaces, the total surface area of ​​the multiple conductive surfaces is between 0.001 and 500,000 μm 2 , 0.001~250,000μm 2 , 0.001~100,000μm 2 , 0.001 to 10,000 μm 2 , 0.01 to 8,000 μm 2 , 0.001 to 6,000 μm 2 , 0.001 to 4,000 μm 2 , 0.001 to 2,000 μm 2 , 0.001 to 1,000 μm 2 The range is.

[0018] Optionally, the assembly is configured for use in a minimally invasive device, such as an implantable or skin-contacting device, and for use in electroanalytical applications for quantitative and qualitative characterization of target analytes in complex biological systems. Optionally, in such configurations and applications, when the assembly comprises multiple conductive surfaces, the total surface area of ​​the conductive surfaces is between 0.001 and 1,000 μm 2 , 0.001~800μm 2 , 0.001~600μm 2 , 10 to 600 μm 2 , 100~600μm 2 , 100~400μm 2 , 100-200 μm 2 , 0.001~500μm 2 , 0.001 to 300 μm 2 , 0.01 to 300 μm 2 , 0.01 to 300 μm 2 , 0.01 to 200 μm 2 , 1 to 300 μm 2 , 1 to 200 μm 2 , 50~300μm 2 , 50~250μm 2 , 50~200μm 2 , 5 to 200 μm 2 , or 10 to 200 μm 2 Such a configuration contributes to controlled diffusion of the target analyte to the conductive surface.

[0019] Preferably, the surface area of ​​a single conductive surface, or each individual conductive surface when the electrochemical electrode assembly comprises multiple individual conductive surfaces, is at least an order of magnitude smaller than the surface area of ​​the surrounding insulation surrounding the single conductive surface or individual conductive surfaces. In particular, the conductive surface or each conductive surface is less than 300 μm 2 , 200 μm 2 , 150 μm 2 , or 100 μm 2 and the surrounding insulating material may be at least 1,000 μm 2 Optionally, the surface area of ​​the conductive surface or each insulating surface surrounding an individual conductive surface may be on the order of 5,000 μm 2, 10,000 μm 2 , 15,000 μm 2 , 20,000 μm 2 , 50,000 μm 2 , 100,000 μm 2 , 500,000 μm 2 , or 1 mm 2 may exceed.

[0020] Optionally, the width, diameter, or minimum distance of the single or individual conductive surfaces may be in the range of 1 to 1,000 nm, 1 to 800 nm, 1 to 600 nm, 1 to 400 nm, 1 to 800 nm, 1 to 600 nm, 1 to 400 nm, 1 to 200 nm, 1 to 100 nm, or 10 to 100 nm. Optionally, the width, diameter, or minimum distance of the single or individual conductive surfaces may be in the range of 10 to 1,000 nm, 50 to 800 nm, 50 to 600 nm, 50 to 400 nm, 50 to 200 nm, 100 to 1,000 nm, 100 to 800 nm, 100 to 600 nm, 100 to 400 nm, or 100 to 200 nm. Optionally, the width, diameter, or minimum distance of each conductive surface is less than 100, 80, 60, 40, 20, or 10 nm.

[0021] Optionally, the assembly may comprise between 10 and 100,000, between 10 and 50,000, between 10 and 10,000, between 10 and 8,000, between 10 and 3,000, between 10 and 1,000, between 10 and 800, between 10 and 600, between 10 and 400, between 10 and 200, between 10 and 100, between 10 and 80, or between 10 and 60 active surfaces.

[0022] Optionally, the conductive and / or insulating surfaces are any one or combination of substantially flat, convex, concave, and profiled to include peaks or valleys. Preferably, the conductive surfaces are located in areas that do not unnecessarily impede diffusion of target analytes to the conductive surface.

[0023] Optionally, the electrically conductive material is a layer disposed on the substrate, and the electrically insulating material is a layer disposed on a layer of electrically conductive material, the layer of insulating material being discontinuous to expose areas of the layer of electrically conductive material that provide at least one electrically conductive surface. Such exposed areas may be provided by ablating or cutting portions of the laminate structure.

[0024] Optionally, the assembly may include a single conductive surface. Alternatively, the assembly of the present invention may include multiple conductive surfaces, with the peripheries of each conductive surface spaced apart by a distance at least five times the diameter or minimum width of any one of the conductive surfaces. That is, the conductive surfaces are spaced apart by a distance greater than the width, diameter, or minimum distance of each conductive surface. References herein to the distance of a conductive surface refer to a line bisecting the conductive surface, extending from a first point on the periphery of at least one conductive surface to a second point on each periphery. Preferably, this distance is on the order of nanometers. Preferably, the corresponding distance between the inner boundary of an insulating surface and the outer boundary of the insulating surface is on the order of micrometers.

[0025] Optionally, each conductive surface is spaced apart from the other by a distance in the range of 10 to 100,000, 10 to 10,000, 10 to 1,000, or 10 to 100 times the width, diameter, or smallest distance of each conductive surface.

[0026] Preferably, when the assembly comprises multiple conductive surfaces, the surfaces are spaced apart by a distance at least 5 times the diameter or minimum width of each conductive surface. Optionally, this distance may be at least 10, 50, 100, or 1,000 times the diameter or minimum width of each conductive surface. This configuration prevents the conductive surfaces from interfering with each other and, in particular, functions to extract target analytes from the electrolyte medium in a controlled manner, thereby providing an electrode assembly with controlled diffusion, particularly oxidation / reduction characteristics and time-independent response.

[0027] Optionally, the assembly may include an analyte-responsive layer on or disposed on the conductive surface. Optionally, the analyte-responsive layer includes one of a material configured to bind to a target analyte, an enzymatic material that chemically interacts or reacts with the target analyte, or at least one aptamer, an antibody, or a molecularly imprinted polymer, or a combination thereof. Optionally, the assembly may include multiple analyte-responsive layers, each comprising a different material and binding to a different respective target analyte.

[0028] Optionally, the assembly may comprise multiple conductive surfaces each comprising a different material that attracts, captures, adsorbs, chemically reacts with, and / or binds a different respective target analyte. Thus, a single sensor may be configured to detect multiple different target analytes (either directly or indirectly via oxidation or reduction) in parallel and within a single or multiple data collection periods.

[0029] According to a further aspect of the inventive concept, there is provided an electrochemical sensor comprising at least one working electrode comprising an electrode assembly as claimed and described herein, and at least one auxiliary electrode. As will be appreciated, in use, the working electrode is electrically coupled to the auxiliary electrode through a sample fluid (biological fluid) in which a target analyte is present. That is, the working electrode and the auxiliary electrode are electrically isolated when not in use (e.g., "in open air"). Electrical coupling between the working electrode and the auxiliary electrode occurs only through the sample fluid (biological fluid), completing the electrical circuit.

[0030] Optionally, the sensor may include an additional reference electrode (to the working electrode and / or auxiliary electrode), so that the sensor may include two, three, or more electrodes.

[0031] Optionally, the sensor may further comprise a potentiostat electrically connected to the electrodes and applying a potential to at least the working electrode and / or measuring the potential between the working electrode and one of the auxiliary electrodes, and a control utility receiving charge, current, and / or charge or current data generated from the oxidation or reduction of the analyte or a reaction product of the target analyte at the working electrode and / or measuring the potential at the working electrode. Optionally, the sensor may further comprise an electrochemical potential module that generates and applies a potential to the working electrode. Such a module may be implemented as software and / or firmware stored and executed on a suitable operating system / platform (e.g., on the potentiostat device), as is well known to those skilled in the art. Optionally, the electrochemical potential module is configured to apply a pulsed potential to the working electrode.

[0032] According to further aspects of the inventive concept, there is provided a sensor as claimed and described herein. According to further aspects of the inventive concept, there is provided a wearable device attachable to a human or animal body, comprising a sensor as claimed and described herein. According to further aspects of the inventive concept, there is provided a method of manufacturing an electrode assembly as claimed and described herein. According to further aspects of the inventive concept, there is provided a method of manufacturing an electrochemical sensor as claimed and described herein.

[0033] According to a further aspect of the inventive concept, there is provided an electrochemical method for analyzing at least one target analyte in a biological fluid, the method comprising providing an electrochemical sensor as claimed and described herein; contacting the sensor with a body having a biological fluid containing at least one target analyte; applying an oxidation or reduction potential to a working electrode to oxidize or reduce the target analyte or a reaction product of the target analyte; and measuring a current between the working electrode and an auxiliary electrode resulting from the oxidation or reduction of the target analyte or a reaction product of the target analyte.

[0034] Optionally, the step of applying an oxidation or reduction potential includes performing at least one electroanalytical technique, such as voltammetry (e.g., cyclic voltammetry), potentiometry, or amperometry. Optionally, the electroanalytical technique includes electrochemical impedance spectroscopy (EIS). It will be understood that a wide range of electroanalytical techniques are compatible with the concepts of the present invention. Optionally, the method includes applying the potential for a predetermined period of time. Optionally, the method includes applying a pulsed potential to the working electrode, having a period of a first potential and a period of a second potential that is zero or lower than the first potential.

[0035] In accordance with a further aspect of the inventive concept, there is provided an electrochemical method for analyzing at least one target analyte in a biological fluid, the method comprising: providing an electrochemical sensor having an electrode assembly comprising an electrically insulating material having at least one insulating surface and an electrically conductive material having at least one conductive surface defined and surrounded by the insulating surface; contacting the sensor with a body having a biological fluid containing at least one target analyte; applying an oxidation or reduction potential to the working electrode to oxidize or reduce the target analyte or a reaction product of the target analyte; and measuring a current between the working electrode and an auxiliary electrode resulting from the oxidation or reduction of the target analyte or a reaction product of the target analyte, wherein the surface area of ​​the conductive surface relative to the surface area of ​​the surrounding insulating surfaces, or the surface area of ​​each conductive surface if the assembly comprises multiple conductive surfaces, is configured to control the rate at which the target analyte in the biological fluid diffuses to the conductive surface, such that the rate of consumption (i.e., oxidation or reduction) of a volume of the target analyte at the conductive surface is less than the rate at which the same volume of target analyte in the biological fluid diffuses to the conductive surface, thereby maintaining a supply of target analyte by diffusion at the conductive surface.

[0036] The electrochemical electrode structure of the present invention comprises at least one active electrode surface, which is sufficiently small, optionally formed as a discontinuity within a relatively large surface area of ​​insulating material, such that transport of electroactive species (target analytes) occurs under nonlinear diffusion control. Preferably, the inner boundary perimeter of the surrounding insulating surface is at least one-tenth of the outer boundary perimeter of the insulating surface. In effect, the outer boundary determines the minimum distance to the nearest or adjacent conductive surface or edge within the electrode assembly. Preferably, the conductive surface has a width, diameter, or other cross-sectional distance of less than 10 μm, more preferably less than 100 nm, in at least one dimension (2D or 3D). For example, for a conductive surface having an area a and a diameter or minimum width x, the peripheries of the conductive surface should be separated by at least 5x in all directions and in the major plane of the insulating surface, resulting in a minimum area of ​​25a. The relative dimensions of the conductive surface (oxidation / reduction) and the surrounding insulating material surface provide the desired diffusion control of the target analyte. The present concepts provide a sensor configuration that responds under spherical or hemispherical diffusion control, providing time- and biofluid volume-independent electrochemical sensing configurations and methods.

[0037] In accordance with a further aspect of the inventive concept, there is provided an electrochemical electrode assembly comprising an electrically insulating material having at least one insulating surface and an electrically conductive material having at least one conductive surface defined and surrounded by the insulating surface, wherein the surface area of ​​the conductive surface is less than 5% of the surface area of ​​the insulating surface defined between an inner boundary at the periphery of the conductive surface and an outer boundary spaced from the inner boundary by a distance at least 5 times the diameter or minimum width of the conductive surface.

[0038] In accordance with a further aspect of the inventive concept, there is provided an electrochemical electrode assembly comprising: an electrically insulating material having an insulating surface; and an electrically conductive material having a conductive surface whose periphery is defined by and surrounded by the insulating surface, wherein the surface area of ​​the conductive surface is less than 5% of the surface area of ​​the insulating surface surrounding the periphery of the conductive surface, and wherein the insulating surface is defined between an inner boundary at the periphery of the conductive surface and an outer boundary spaced from the inner boundary by a distance at least 5 times the diameter or width of the conductive surface.

[0039] In accordance with a further aspect of the inventive concept, there is provided an electrochemical electrode assembly comprising an electrically insulating material having an insulating surface and an electrically conductive material having a conductive surface whose periphery is defined by and surrounded by the insulating surface, wherein the surface area of ​​the conductive surface is an order of magnitude less than the surface area of ​​the insulating surface surrounding the periphery of the conductive surface.

[0040] Optionally, the surface area of ​​the insulating surface is defined between an inner boundary at the periphery of the conductive surface and an outer boundary that is spaced from the inner boundary by a distance multiple of the diameter or width of the conductive surface. Optionally, the distance is in the range of at least 5, 10, 50, 100, 500, 1,000, 10,000, or 100,000 times the diameter or width of the conductive surface. Optionally, the distance is in the range of at least 10-100,000, 10-10,000, 10-1,000, or 10-100 times the diameter or width of the conductive surface.

[0041] Preferably, the width of the conductive surface is the minimum width, which refers to the shortest surface distance in the case where the surface width or thickness is variable / non-uniform.

[0042] In accordance with a further aspect of the inventive concept, there is provided an electrochemical electrode assembly comprising: an electrically insulating material having at least one insulating surface; and an electrically conductive material having a plurality of conductive surfaces, each having a periphery defined by and surrounded by the insulating surface, wherein the surface area of ​​each conductive surface is less than 5% of the surface area of ​​a respective insulating surface surrounding the periphery of each of the conductive surfaces, and wherein each surface area of ​​the insulating surfaces is defined between an inner boundary at the periphery of each conductive surface and a respective outer boundary spaced from the respective inner boundary by a distance at least 5 times the diameter or minimum width of the respective conductive surface. [Brief explanation of the drawings]

[0043] Particular implementations of the present invention will now be described, by way of example only, and with reference to the accompanying drawings, in which: [Figure 1A] FIG. 1 is a plan view of a first stage manufacture of an electrode suitable for use as an electrochemical sensor according to a particular implementation. [Figure 1B] FIG. 1B is a plan view of the second stage of manufacture of the electrode of FIG. 1A. [Figure 1C] FIG. 2 is an enlarged plan view of the distal end of the electrode of FIGS. 1A and 1B. [Figure 2A] FIG. 2 is a plan view of a third stage of manufacture of the electrode of FIGS. 1A to 1C. [Figure 2B] FIG. 2B is a plan view of the final product of the electrode of FIGS. 1A-2A. [Figure 2C] FIG. 2C is an enlarged side view of the distal end of the electrode of FIG. 2B. [Figure 2D] FIG. 2D is an enlarged side view of the distal end of the electrode of FIG. 2C. [Figure 3] Cyclic voltammograms using the electrodes of Figures 2B and 2C with a silver chloride reference electrode and a Pt wire counter electrode in PBS containing 1 mmol / dm3 FCA at a scan rate of 50 mV / s. [Figure 4] 2B and 2C in PBS containing 1 mmol / dm 3 FCA after polarization to 0.35 V. [Figure 5]Graph of current versus time for the electrodes of FIGS. 2B and 2C in PBS containing 1 mmol / dm3 FCA for five repeated 100 ms, 0.35 V pulses with 10 s rest at open circuit potential between each repetition. [Figure 6] Graph of current versus time for a nanoband array electrode in PBS containing 1 mmol / dm3 FCA, with five 100 ms, 0.35 V pulses, with a 10 s rest at open circuit potential between each pulse. [Figure 7] Graph of current versus time for a 100 μm diameter platinum disk electrode in PBS containing 1 mmol / dm 3 FCA, with five 100 ms, 0.35 V pulses, with a 10 s rest at open circuit potential between each pulse. [Figure 8] Graphs of responses in a widest coherent window at open circuit potential for 0.35 V pulses, 0.1 ms sampling interval, 20 ms window, 1 s between each repetition, in PBS and PBS containing 1 mmol / dm3 FCA for the electrodes of Figures 2B and 2C. [Figure 9] Graphs of responses including FCA repetitions and earliest data points for the electrodes of Figures 2B and 2C in PBS and PBS containing 1 mmol / dm3 FCA, with 0.35 V pulses, 0.1 ms sampling intervals, 20 ms windows, and 1 s open circuit potential between each repetition. [Figure 10A] 2B and 2C, using a commercial ASIC (ADuCM355) in PBS containing 1 mmol / dm3 FCA, with 10 repeated 5 ms, 0.35 V pulses, with a 50 ms rest at open circuit potential between each repetition. [Figure 10B] This graph shows the results of ten repeated 5 ms, 0.35 V pulses with 50 ms rest at open circuit potential between repetitions in PBS containing 1 mmol / dm3 FCA for an electrode with a nanoband array, using a commercial ASIC (ADuCM355). [Figure 11]This graph shows ten repeated 5 ms, 0.35 V pulses with 50 ms rest at open circuit potential between pulses in PBS containing 1 mmol / dm3 FCA for an electrode with a 100 μm platinum disk, using a commercial ASIC (ADuCM355). [Figure 12A] 2B and 2C having an active electrode surface formed as a convex protrusion or ridge extending from the substrate. FIG. [Figure 12B] 2B and 2C having an active electrode surface formed as a recessed or indented / depressed cavity in the substrate surface. FIG. [Figure 13] 2D is a plan view of a dual electrode system having a pair of electrodes on the upper and lower surfaces of an intermediate substrate according to the embodiment of FIGS. 2B and 2C. FIG. [Figure 14] 14 is a side view of the distal ends of the electrodes of FIG. 13 disposed on the upper and lower surfaces of a substrate. [Figure 15] 2B and 2C with dual conductive layers, and FIG. 2B is a final product of the electrode of FIG. A with a triple conductive layer configuration, each layer comprising multiple active electrode surfaces according to further embodiments. [Figure 16] 2B and 2C, in plan view, an assembly with a plurality of laterally defined, individually addressable electrode regions; [Figure 17A] FIG. 10 is a first stage of manufacture view of a further embodiment of an electrode assembly. [Figure 17B] FIG. 17B is a diagram of the second stage manufacture of the electrode of FIG. 17A. [Figure 17C] FIG. 17B is a diagram of the third stage manufacture of the electrode of FIG. 17A. [Figure 17D] FIG. 17B is a diagram of the fourth stage manufacture of the electrode of FIG. 17A. [Figure 18A] FIG. 17E is a further plan view of the electrode of FIG. 17D. [Figure 18B]FIG. 18B is a side cross-sectional view taken along the CC cross section of FIG. 18A. [Figure 19] FIG. 10 is a plan view of a further embodiment of an electrode assembly comprising multiple active electrode surfaces defined in open recesses, openings, or open cavities in an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20A] FIG. 10 is a diagram of a first stage manufacture of a further embodiment of an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20B] FIG. 10 is a diagram of a second stage manufacture of a further embodiment of an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20C] FIG. 10 is a diagram of a third stage of manufacture of a further embodiment of an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20D] FIG. 10 is a diagram of a fourth stage of manufacture of a further embodiment of an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20E] FIG. 10 is a diagram of a fifth stage of manufacture of a further embodiment of an electrode assembly having a grid or mesh-like conductive layer structure. [Figure 20F] 10A and 10B are diagrams of final-stage manufacture of further embodiments of electrode assemblies having grid or mesh-like conductive layer structures. [Figure 21A] 10A-10C are diagrams of further embodiments of electrode assemblies configured with relatively large active electrode surfaces formed within a conductive layer grid configuration. [Figure 21B] 10A-10C are diagrams of further embodiments of electrode assemblies configured with relatively large active electrode surfaces formed within a conductive layer grid configuration. [Figure 21C] 10A-10C are diagrams of further embodiments of electrode assemblies configured with relatively large active electrode surfaces formed within a conductive layer grid configuration. [Figure 22A] FIG. 10 is a plan view of a hybrid electrode assembly with a nanoband array according to a further implementation. [Figure 22B] FIG. 10 is a plan view of a hybrid electrode assembly with a nanoband array according to a further implementation. [Figure 23] FIG. 10 is a plan view of a further embodiment of an electrode assembly comprising multiple active electrode surfaces, each surface disposed on a sidewall of a respective channel or groove formed in the multi-layer structure. [Figure 24] 10A-10C are diagrams of further embodiments of electrode assemblies comprising multiple cavities or openings formed in a multi-layer structure exposing multiple active electrode surfaces in the open cavities, openings, and / or open recesses. DETAILED DESCRIPTION OF THE INVENTION

[0044] Referring to FIGS. 1A-1C, the electrode assembly 10 includes a substrate 18 formed from an electrically insulating material. All dimensions shown in FIGS. 1A-2C are in microns. A conductive layer 11 is formed as an elongated strip disposed on a surface 19 (FIG. 2C) of the substrate 18, with a proximal end connected to an electrically conductive tab 12 and a distal end 13 including a plurality of laterally extending conductive fingers 14. Each finger 14 includes a respective terminal end 15. According to a specific embodiment, the conductive layer 11 has a thickness of approximately 0.05 microns and is disposed on the substrate 18, which itself has an approximate thickness of approximately 250 microns. The approximate width of the electrode at the conductive fingers 14 (between the respective terminal ends 15) is approximately 600 microns. The width of the conductive elongated strip 11 is approximately 60 microns, and the approximate total length of the elongated electrode (including the strip 11 and fingers 14) is approximately 40,000 microns. Figure 1A shows the first stage of manufacture, which involves the placement of conductive strips 11 and fingers 14 on substrate 18. Figure 1B shows the second stage of manufacture, which involves the placement of a layer of insulating material 16 over both substrate 18 and conductive layer 11. Insulating layer 16 is approximately 600 μm wide and approximately 40,000 μm long, extending nearly the entire length of conductive strips 11.

[0045] FIG. 2A shows a third-stage manufacturing process in which selected regions of the assembly of FIG. 1B are identified for ablation using a laser cutter or similar technique. After ablation cutting, the completed electrode assembly is shown in FIG. 2B, with the structure of FIG. 2A cut / ablated at ablation lines 17 (shown schematically). Thus, referring to the side view of FIG. 2C, an active electrode surface 15 is provided at the end of each conductive finger 14 and exposed on each sidewall surface of the laminated electrode assembly (comprising substrate 18, conductive layer 11, and insulating material 16). Importantly, each electrode active surface 15 is spatially separated from the others, providing a distinct conductive surface area for oxidation / reduction of target analytes contained in biological fluids (e.g., sweat, saliva, blood, etc.). In accordance with the presently described embodiment, insulating layer 16 is of a material configured and selected to suit the end use application and, when used as a minimally invasive device, is biocompatible. Such layer materials provide compatibility, adhesion, etc. for additional functionality required for each active surface 15 region to chemically interact with the target analyte, i.e., to bind, react, adsorb, etc., the target analyte, for oxidation / reduction, reaction, or interaction with the active surface resulting in a potential change.

[0046] According to the present embodiment of the electrode assembly described herein, an active electrode coating can be provided on each exposed active electrode surface 15. Such coatings, by existing configuration, are configured to chemically interact with the target analyte, binding, reacting, adsorbing, etc., and retaining the analyte on the active surface 15 due to oxidation and / or reduction or other changes in effect at the active surface that result in a change in potential. In some cases, the desired properties of the active electrode coating can be incorporated into the insulating layer 16, thereby eliminating steps. Such active electrode coatings are well known to those skilled in the art of sensor fabrication, particularly biosensor fabrication. Additional coatings can be applied to advantageously modify the overall sensor performance, including, but not limited to, the rejection of potential interferents. This adds the possibility of utilizing the structure for potentiometry and, if the insulating layer 16 is passive, the inclusion of an "exclusion membrane" to reject interferents.

[0047] All of the embodiments described herein are suitable for use in minimally invasive diagnostic sensors, implantable or wearable electrochemical sensors.

[0048] According to the particular electrode assembly of FIGS. 2B and 2D , the material of substrate 18 comprises polyethylene terephthalate (PET) and provides a first insulating layer. Conductive layer 11 optionally comprises platinum deposited using a suitable technique. The final electrode structure of FIGS. 2B and 2D can be achieved, for example, using laser ablation or photolithography to create an electrically conductive pattern with distinct, spatially separated active electrode surface areas for oxidation / reduction or surface interaction of the target analyte or other species generated by the presence of the target analyte, such as a reaction product of the analyte reacting with an active coating, such as an enzymatically active coating.

[0049] According to the configuration of Figures 2B and 2D, the electrode assembly has fingers 14 sized and spaced to ensure the electrode structure is electrically addressable, twelve active conductive regions 15 with appropriate dimensions (each active region 30 µm long and 50 nm wide), and sufficient separation from adjacent active regions 15 (90 µm) to minimize interactions between them while ensuring efficient diffusion characteristics (50 nm width). At the proximal end, second and third contact pads 12 allow for optional incorporation of auxiliary electrodes (not shown) as needed. The dimensions of the pads 12 can be varied to accommodate the selected electrical connection method (pad size, pitch, etc.). These auxiliary electrodes can be incorporated on the backside of the structure, requiring through-holes for connection, or adjacent to the working electrode structure on the same side as the contact pads 12.

[0050] 2D, the insulating surface surrounding each of the active electrode surfaces 15 is formed by a substrate 18 and a layer of insulating material 16. The substrate 18 and the insulating layer 16 collectively provide a peripheral insulating surface that confines and bounds each of the active electrode surfaces 15. As shown, the relative dimensions of each active electrode surface 15 are significantly smaller than the surrounding surface area of ​​the insulating material (16, 18). In particular, the surface area of ​​each active electrode surface 15 is approximately 1.2 μm 2 , or 1 μm 2 , 0.5 μm 2 , or 0.1 μm 2 It may be less than.

[0051] Each active electrode surface 15 is bounded and surrounded by an insulating material region 24. Each region 24 has a surface area on the order of square microns, while each active electrode surface area 15 is on the order of square nanometers. Specifically, the insulating material region 24 is defined between an inner boundary 21 corresponding to the periphery 20 of each active electrode surface 15 and a respective outer boundary 23. The outer boundaries 23 are separated by a distance d that is five times the distance e, which is the diameter or minimum width of each active electrode surface 15. This configuration is advantageous for controlling the diffusion of analytes (target species) to the active conductive surface 15, particularly by avoiding unnecessarily high analyte diffusion rates and achieving controlled oxidation / reduction of the target species.

[0052] The auxiliary electrodes are omitted from Figures 1A-1C for clarity. In this embodiment, the auxiliary electrodes were fabricated using commercially available silver / silver chloride screen printing ink and patterned on the backside of the substrate 18. Alternative auxiliary electrodes and their fabrication materials are definable and known to those skilled in the art. The structure was then screen printed with a commercially available dielectric ink to form a dielectric layer 16 on the front side, as shown in Figure 1B, covering the entire structure except for the electrical contact pads 12. The thickness of the dielectric layer was approximately 30 μm. This was more than sufficient to ensure that the active areas exposed in the final step (detailed below) were sufficiently encapsulated with an electrochemically inert material to achieve optimal diffusion (sometimes referred to as hemispherical). Alternatively, a second screen printing step could be incorporated to minimize areas of exposed conductor (commonly referred to as pinholes). The backside was also screen printed with dielectric ink up to 700 μm into the distal end of the structure, leaving the silver / silver chloride layer exposed at the distal end (not shown for clarity). In Figure 2A, the singulation ablation lines 17 of the structure 10 are shown, which cut through the dielectric / conductor / PET substrate at the distal end, exposing six conductor end faces 15, each 30 μm long and 50 nm wide, on each side. Each surface 15 of the singulated device is 90 μm (120 μm center-to-center) from the adjacent surface, as shown in Figures 2B and 2C. The total working electrode area in this embodiment is 18 μm. 2, which is about 8,000 μm 2 This is compared to a typical structure in this scenario (i.e., approximately 400 times smaller surface area). This is advantageous in reducing the specific capacitance of the electrode. The singulation process was effected by a femtosecond infrared laser, and the width of the final structure was approximately 400 μm. Note that the performance of the device is largely independent of its width; performance is determined by the number, dimensions, spacing, and environment of the active areas (facets 15). This is not the case when considering conventional electrode structures. This becomes important when the embodiment is used in applications such as minimally invasive implants (e.g., when applied to continuous glucose monitoring applications), because the overall size of the structure has a significant impact on the foreign body response (FBR) elicited in situ. The larger the implant, the faster the FBR and the shorter the implant's operational lifetime. Therefore, the flexibility to adopt an optimal width without compromising performance is an important attribute.

[0053] The electrode embodiments in Figures 2B and 2C were tested with ferrocene carboxylic acid (FCA) and hydrogen peroxide (HP). These two analytes were chosen as alternatives to first-generation biosensors, which utilize enzymes to generate HP, which is subsequently measured at an indicator electrode, and second-generation biosensors, which utilize artificial mediators (FCA). First-generation biosensors require the platinum electrode surface to be catalytically active for HP electrochemistry. Unless otherwise noted, measurements were performed in phosphate-buffered saline (PBS) solution at room temperature. The potential ranges for cyclic voltammetry were 0 to 0.45 V versus the silver / silver chloride standard for FCA and -0.2 to 0.75 V for HP (scan rate 50 mV / s). Chronoamperometry measurements were performed at 0.35 V for FCA and 0.6 V for HP. For comparison, measurements were performed using a 100 μm platinum disk electrode as a model of current implantable devices, and a conventional 50 nm platinum nanoband array electrode (30 μm apertures on a 60 μm pitch) as a current teaching example for nanoband electrode structures.

[0054] As is clear from FIG. 3, the electrodes of FIGS. 2B and 2C have a 1 mmol / dm 3 When examined using cyclic voltammetry in the presence of FCA, they exhibit typical nanoband performance in that there is no evidence of mass transport limitation (the oxidation wave plateaus above 0.35 V and does not pass through a maximum). Similar waveforms are observed when conventional nanoband arrays are examined.

[0055] When this embodiment was investigated using a potential step, here to 0.35 V vs. a silver / silver chloride electrode, 2This deviation from the behavior of either a conventional nanoband array or a macroscale platinum electrode with an area of ​​100 μm (Figure 4) is evident. For the nanoband array and macroelectrode, the continuous decrease in current over the 60-second measurement window indicates that these systems are mass-transport limited and unable to replenish the FCA consumed at the electrode surface. In this embodiment, a steady state is reached within approximately 6 seconds, at which point the current appears to have stopped decreasing. This time-independent response is characteristic of the establishment of optimal (hemispherical) diffusion. At this point, the mass transport rate of the analyte to the electrode surface is sufficient to replenish the FCA consumed at the electrode. This condition is typically observed only when the flow of the electrode surface area is very carefully controlled, but in this example, the solution was stationary.

[0056] In this embodiment, 1 mmol / dm 3 The electrode was tested in PBS containing FCA and in PBS ("blank") with 100 ms 0.35 V pulses and 10 s pauses between pulses (the electrode was held at open-circuit potential). This behavior was compared to that of a 100 μm Pt disk electrode under the same conditions. The results are shown in Figures 5-7. In this embodiment, the five repeated pulses are randomly ordered, whereas the response of the 100 μm Pt disk is ordered, decreasing from the first response to the fifth. This indicates that even with a short consumption period (100 ms) and a long recovery time (10 s), the 100 μm disk consumes more FCA than mass transport can replenish under quiescent conditions, whereas this is not the case in this embodiment. This improves measurement accuracy in applications where controlling solution flow is not possible or desirable.

[0057] This embodiment was tested at the shortest time interval available for the instrument used (0.1 ms), with a 20 ms sampling window (with a 1 s open-circuit potential between repetitions), demonstrating that the response of this embodiment is primarily Faradaic (i.e., redox processes dominate over charging phenomena). Five pulse repetitions were measured. The potentiostat used exhibits instrument artifacts at these high sampling rates, resulting in complex waveforms. Synchronizing the waveforms for each response reveals a clear Faradaic response (Figures 8 and 9). With a 0.1 ms sample interval, starting from the first data point (t = "0"), FCA repetition 4 shows no evidence of charging at the initial point t = "0" or the first point t = 0.1 ms (Figure 9). Therefore, this embodiment exhibits very low non-Faradaic (capacitive charging) components, demonstrating that poling at kilohertz frequencies is possible. t="0" is some time after the embodiment has been polarized, and the duration is <<0.1 ms, so it is clear that with appropriate hardware higher frequency / shorter sampling interval investigations are possible.

[0058] Additionally, this embodiment was tested using a commercially available Application Specific Integrated Circuit (ASIC) designed for mobile / wearable scenarios (using a PalmSens Sensit BT potentiostat utilizing the Analog Devices ADuCM355 ASIC). The shortest software-allowed sample interval of 1 ms was used with a minimum number of samples (5). Note that the native capability of the ADuCM355 is 400 ksamples / s. Under software limitations, a 5 ms 0.35 V pulse with a 50 ms rest period (electrodes held at open circuit potential) was used to measure the concentration of 1 mmol / dm 3 PBS containing FCA and PBS alone ("blank") were used. The responses under these conditions are shown in Figures 10A, 10B, and 11. Results are shown for the electrodes of Figures 2B and 2C, 100 μm diameter disk electrodes, and nanoband arrays.

[0059] The electrode assembly of the present invention can be embodied in a variety of forms to suit the application, analyte of interest, and its diffusion characteristics in the matrix under consideration. Because device sensitivity scales linearly with length, if the fingers 14 are defined over a length of 1440 μm (i.e., 12 fingers 14 instead of 6 fingers 14 on each side, with the same arm width and spacing), the signal current doubles. Similarly, if the number of fingers 14 over a length of 720 μm is doubled, a similar increase in signal current is observed. Thus, the structure can vary from a single, monolithic finger 14 (720 μm long in this example embodiment) to a highly segmented finger structure (e.g., 72 5 μm-wide fingers separated by 5 μm spacing) to a sparsely packed finger structure (as in this embodiment). Appropriate structures can be fabricated (designed) to provide optimal performance depending on the application under consideration. Furthermore, the device of the present invention can be made significantly narrower (400 μm in this embodiment) without compromising performance. The width limitations of the device of the present invention depend on the length and resistivity of the conductor tracks 11 (FIGS. 1A-2C) and the manufacturing tolerances of the manufacturing method and tools used. For this embodiment and the manufacturing tools used, embodiments with a width of 60 μm are possible. A further practical consideration is whether the desired mechanical properties are maintained when the width of the embodiment is reduced. However, it is important to note that the electrochemical performance will not be compromised.

[0060] In the described embodiment, the active area surface 15 is flat. According to further embodiments, such active surface 15 may be concave or convex in two or three dimensions, as shown in Figures 12A and 12B.

[0061] In a further alternative embodiment, a second electrode is defined on the backside of the insulating substrate 18 by depositing and patterning an additional conductor layer, followed by an insulating layer 16 as described above, as shown in Figures 13 and 14. In this embodiment, the "top" and "bottom" electrodes are coincident. In a further embodiment, the active areas 15 of the two electrodes are offset or staggered. In yet a further embodiment, the active areas 15 and the spacing between the two electrode active areas 15 are not identical.

[0062] In alternative embodiments, additional conductive and insulating layers and patterning are added to create two or more stacked, independently addressable active areas 15 (collectively referred to as electrodes) on one side of a single substrate (FIG. 15A and B). In the illustrated embodiment, the active areas 15 of the two electrode structures are identical and coincident. In some embodiments, the active area structures may not be identical (e.g., if configured to oxidize / reduce different target analytes or their reaction products), and in further embodiments, the active area structures may not be coincident.

[0063] Multiple individually addressable active areas 15 can also be defined laterally, some examples are shown in Figure 16. A basic insulating substrate 18 is covered by a conductive layer 11 and a dielectric insulating top layer 16. The active areas 15 can be of the same size, spacing, and structure, although in further embodiments they may differ depending on the sensing application. A combination of any or all of the above approaches can also be employed.

[0064] In other embodiments of the assembly of the present invention, the PET-based substrate 18 can be replaced with any suitable material (e.g., polyester, polyalkane, etc.) having the required mechanical, biological, biocompatibility, and electrical properties. Similarly, the Pt conductive layer 11 can be replaced with any suitable material (e.g., platinum group metals, metals, organic conductors, etc.) having the required mechanical, biological, biocompatibility, electrical, and electrochemical properties. The thickness of the conductive layer 11 can also be varied as needed (typically in the range of 0.1 to 100 nm). Additionally, the insulating top layer 16 can be replaced with any suitable material having the required mechanical, biological, biocompatibility, and electrical properties.

[0065] In some cases, it may be necessary to separate the process of fabricating the active region 15 from the singulation process, an example of which is shown in Figures 17A-17D and 18A and 18B. Singulation can occur anywhere outside or within the ablation region 17, as long as the lines are outside the inner lines of ablation.

[0066] In this example, the same basic electrode pattern (as in Figures 2B and 2C) is used, with six biaxial fingers 14 projecting laterally at their distal ends from a central conductor strip. The patterning and layer structure are also identical to the previous example. Specifically, the starting point for fabrication is a PET substrate 18 with a 50 nm Pt layer 11 into which a suitable pattern has been ablated. In this further example, the fingers 14 are 30 μm wide and spaced 120 μm on center. The pattern and most of the remaining Pt / PET area are screen printed with a suitable dielectric ink layer 16, leaving small remaining areas that electrically insulate the Pt layer 11 and define the electrical contact pads 12, as shown in Figure 20B. As with the previous example of Figures 2B and 2C, the contact pad dimensions can be optimized for the electrical connection system being used. FIG. 20C shows the marking out of the ablation track 17 (which could be a single track area if desired), and FIG. 20D shows the assembly after forming the ablation track (through the dielectric layer 16, the conductor layer 11, and onto the PET substrate 18, here acting as an effective "etch stop"). Ablation can be performed using laser machining or die cutting, or similar methods, including methods such as reactive ion etching or chemical etching. FIG. 20E shows a cross section of this example extending slightly into the insulating substrate 18. The device can be singulated to any size and shape suitable for the desired application. Materials and structures can be varied as with the first-described embodiment.

[0067] The two approaches can be combined to increase the signal generated. An example is shown in Figure 19, where eight additional active areas 15 are defined on the surface when the structure is singulated, increasing the number of defined active areas 15 to ten. The number, spacing, and size of the active areas defined within the surface can be varied using different patterning of the conductor layer and / or the shape of the openings themselves to optimize the performance of the structure for a selected application.

[0068] The advantages of the present invention are demonstrated by utilizing an extreme embodiment suitable for minimally invasive in vivo applications such as continuous glucose monitoring. The inventors recognized that the size of the electrode assembly of the present invention is important and determines the longevity and performance of the device in use. For applications where this is not a critical issue, i.e., where sample flow management is not possible / desirable and / or the sample volume to be addressed is unknown or small, the basic approach can be extended to assemblies with active surfaces 15 over larger surface areas. An example of this approach is shown in Figures 20A-20F. In some applications, it may be desirable to repeat the surface ablation pattern over a larger area. An example of this approach is provided below.

[0069] In such a case, the starting point for fabrication is a PET substrate 18 with a 50 nm Pt layer, with the appropriate pattern ablated into the Pt layer 11. The pattern and most of the remaining Pt / PET layers 11, 18 are screen-printed with a suitable dielectric ink 16, leaving small remaining areas to electrically isolate the Pt and define the electrical contact pads 12. Figure 20C shows the ablation tracks 17, which may alternatively be a series of five individual tracks 17. Figure 20D shows the assembly after the ablation tracks 17 have been formed (through the laminate and into or onto the PET substrate 18). Figure 20E shows the markout 17 for singulation cutting / laser processing 22 of the device, with the finished structure shown in Figure 20F. Materials and structures can be varied as in the first-described embodiment. In this example, the active area 15 is defined by the sidewalls of the laminate material structure (formed by ablation cutting) during the singulation process, in addition to those created in the surface ablation stage. This is optional and, if not required for a given application, singulation may be performed so that no additional active areas 15 are created during this process.

[0070] The concepts of the present invention can be implemented with other electrode structures, for example, with nanoband arrays, as shown in Figures 22A and 22B. In the embodiment of Figure 22A, the nanoband arrays are defined and structured as independently addressable electrodes, in line with the previous embodiment. In this example, contact pads 12 are on both sides of the structure, but alternatively, they can be located on the same side of the structure, providing a convenient means for connecting to a commonly preferred multi-way electrical connector. The embodiment of Figure 22B focuses on the present invention and a structure in which the nanoband array electrodes are addressed as a single operating / indicating electrode through a single connection pad.

[0071] For example, the above embodiments can be combined with one or more conventional large scale electrodes using screen printing techniques if desired.

[0072] In a further embodiment, as shown in FIG. 23 , a series of interdigitated fingers 14 are defined in the conductor layer 11. The active surfaces 15 defined by ablation / cutting appear on the sidewalls of the formed channels or grooves (formed in the laminate structure). In this example, the surfaces 15 on opposite walls (within each channel or groove) are not directly opposite each other, but are offset or staggered along the length of each channel or groove. This is advantageous in ensuring that each active surface 15 faces a passive region within the structure (i.e., is not directly opposite another active surface 15), further reducing the ability of an active region 15 in close proximity to an oxidizable / reducible analyte to impair a competing active region 15. In this example, no additional edges or active surfaces 15 are created during singulation, although of course, these could be created if additional fingers 14 were defined extending to the periphery of the structure.

[0073] The surface patterning approach can be limited to discrete areas as desired, an example being shown in Figure 24. The number of discrete ablation apertures (open cavities, recesses, or pits) 22 can be varied as desired, and the number, size, and spacing of the active areas 15 defined within each ablation aperture 22 can also be varied to provide the desired performance in a selected application.

Claims

1. 1. An electrochemical electrode assembly comprising: an electrically insulating material having at least one insulating surface; an electrically conductive material having at least one conductive surface whose periphery is defined by and surrounded by said insulating surface; Equipped with 1. An assembly wherein the surface area of ​​the conductive surface, or of each conductive surface if the assembly comprises multiple conductive surfaces, is less than 5% of the surface area of ​​the insulating surface surrounding the periphery of the conductive surface, or of each of the insulating surfaces surrounding each periphery of each conductive surface, and each surface area of ​​the insulating surfaces is defined between an inner boundary at the periphery of the conductive surface, or a respective inner boundary at the periphery of each conductive surface, and a respective outer boundary of the conductive surface, or of each conductive surface if the assembly comprises multiple conductive surfaces, separated from the respective inner boundary by a respective distance of at least 5 times the diameter or smallest width of the respective conductive surface.

2. The surface area of ​​each of the conductive surfaces or each of the plurality of conductive surfaces is 300 μm 2 , 200 μm 2 , 150 μm 2 , 100 μm 2 , 80 μm 2 , 60 μm 2 , 40 μm 2 , 20 μm 2 , 10 μm 2 , 5 μm 2 , 3 μm 2 , or 1 μm 2 The assembly of claim 1 , wherein the axial length is less than 1 / 2 mm.

3. When the assembly comprises multiple conductive surfaces, the total surface area of ​​said conductive surfaces is between 0.001 and 500,000 μm 2 ,0.001~250,000μm 2 ,0.001~100,000μm 2 ,0.001~10,000μm 2 ,0.01~8,000μm 2 ,0.001~6,000μm 2 ,0.001~4,000μm 2 ,0.001~2,000μm 2 ,0.001~1,000μm 2 10. The assembly of claim 1, wherein the axial length is in the range of 100 .mu.m.

4. When the assembly comprises multiple conductive surfaces, the total surface area of ​​said conductive surfaces is between 0.001 and 1,000 μm 2 ,0.001~800μm 2 ,0.001~600μm 2 , 10 to 600 μm 2 , 100 to 600 μm 2 , 100 to 400 μm 2 , 100 to 200 μm 2 ,0.001~500μm 2 ,0.001~300μm 2 ,0.01~300μm 2 ,0.01~300μm 2 ,0.01~200μm 2 , 1 to 300 μm 2 , 1 to 200 μm 2 , 50 to 300 μm 2 , 50 to 250 μm 2 , 50 to 200 μm 2 , 5 to 200 μm 2 , or 10 to 200 μm 2 10. The assembly of claim 1, wherein the axial length is in the range of 100 .mu.m.

5. 10. The assembly of claim 1, wherein the width, diameter, or smallest distance of a single or individual conductive surface is in the range of 10 to 1,000 nm, 50 to 800 nm, 50 to 600 nm, 50 to 400 nm, 50 to 200 nm, 100 to 1,000 nm, 100 to 800 nm, 100 to 600 nm, 100 to 400 nm, or 100 to 200 nm.

6. 6. The assembly of any one of claims 1 to 5, comprising 10 to 100,000, 10 to 50,000, 10 to 10,000, 10 to 10,000, 10 to 8,000, 10 to 3,000, 10 to 1,000, 10 to 800, 10 to 600, 10 to 400, 10 to 200, 10 to 100, 10 to 80, or 10 to 60 conductive surfaces.

7. 7. The assembly of claim 1, wherein the conductive surface and / or the insulating surface is any one or combination of a substantially flat surface, a convex surface, a concave surface, and a surface profiled to include peaks or valleys.

8. 8. The assembly of claim 1, wherein the electrically conductive material is a layer disposed on a substrate and the electrically insulating material is a layer disposed on a conductive material layer, the insulating material layer being discontinuous and exposing areas of the conductive material layer that provide the at least one conductive surface.

9. 9. An assembly according to any one of claims 1 to 8, comprising a plurality of conductive surfaces, the peripheral edges of each of said conductive surfaces being spaced apart from one another by a distance at least five times the diameter or smallest width of any one of said conductive surfaces.

10. The assembly of any one of claims 1 to 9, further comprising an analyte responsive layer disposed on the conductive surface.

11. The analyte responsive layer comprises: a material configured to bind to a target analyte - an enzymatic material that chemically interacts or reacts with the target analyte 11. The assembly of claim 10, comprising:

12. 12. The assembly of claim 10 or 11, comprising a plurality of analyte responsive layers each comprising a different material and binding to a different respective target analyte.

13. The assembly of claim 12 comprising a plurality of conductive surfaces each comprising one of the different materials.

14. 1. An electrochemical sensor comprising: At least one working electrode comprising an electrode assembly according to any one of claims 1 to 13; at least one auxiliary electrode; A sensor comprising:

15. a potentiostat electrically connected to the working electrode and the auxiliary electrode for applying a potential to the working electrode and / or measuring a potential between the working electrode and one of the auxiliary electrodes; a control utility that receives charge, current, and / or charge or current data resulting from the oxidation or reduction of an analyte or a reaction product of the target analyte at the working electrode and / or measures the potential at the working electrode; The sensor of claim 14 further comprising:

16. 16. The sensor of claim 14 or 15, further comprising an electrochemical potential module that generates and applies a potential to the working electrode.

17. 17. The sensor of claim 16, wherein the electrochemical potential module is configured to apply a pulsed potential to the working electrode.

18. An implantable device comprising a sensor according to any one of claims 14 to 17.

19. A wearable device comprising the sensor according to any one of claims 14 to 17 and attachable to the body of a human or animal.

20. A method for manufacturing an electrode assembly according to any one of claims 1 to 13.

21. A method for manufacturing the electrochemical sensor according to any one of claims 14 to 17.

22. 1. An electrochemical method for analyzing at least one target analyte in a biological fluid, comprising: Providing an electrochemical sensor according to any one of claims 14 to 17; contacting the sensor with a body having a biological fluid containing at least one target analyte; applying an oxidation or reduction potential to the working electrode to oxidize or reduce the target analyte or a reaction product of the target analyte; measuring a current between the working electrode and the auxiliary electrode resulting from the oxidation or reduction of the target analyte or a reaction product of the target analyte; A method comprising:

23. 23. The method of claim 22, wherein the step of applying an oxidation or reduction potential comprises performing any one of voltammetry, potentiometry, or amperometry.

24. 24. A method according to claim 22 or 23, comprising applying a potential for a predetermined period of time.

25. 24. The method of claim 22 or 23, comprising applying a pulsed potential to the working electrode having a period of a first potential and a period of a second potential that is zero or lower than the first potential.

26. 1. An electrochemical method for analyzing at least one target analyte in a biological fluid, comprising: an electrically insulating material having at least one insulating surface; an electrically conductive material having at least one conductive surface defined by and surrounded by said insulating surface; providing an electrochemical sensor having an electrode assembly comprising: contacting the sensor with a body having a biological fluid containing at least one target analyte; applying an oxidation or reduction potential to a working electrode to oxidize or reduce the target analyte or a reaction product of the target analyte; measuring a current between the working electrode and an auxiliary electrode resulting from the oxidation or reduction of the target analyte or a reaction product of the target analyte; Including, The surface area of ​​the conductive surface relative to the surface area of ​​the surrounding insulating surfaces, or the surface area of ​​each conductive surface if the assembly comprises multiple conductive surfaces, is configured to control the rate at which the target analyte in the biological fluid diffuses to the conductive surface, so that the consumption rate of a volume of the target analyte on the conductive surface is less than the rate at which the same volume of target analyte in the biological fluid diffuses to the conductive surface, thereby maintaining the supply of target analyte by diffusion on the conductive surface.

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