ANALYTE SENSING

Field of Embodiments

[0001] The present embodiments relate generally to non-invasive or minimally invasive transdermal measurement systems. More specifically, the embodiments relate to micro fluidic transdermal glucose measurement systems in which a thin electrode is contained within a fluid- transmitted capillary, and processes for their production and use.

Background

[0002] Minimally invasive transdermal systems are described in, for example, co-owned U.S. Patents 6,887,202 and 7,931,592, both entitled "Systems and Methods for Monitoring Health and Delivering Drugs Transdermally," as well as co-owned U.S. Appn. No. 13/459,392, each of which is incorporated herein by reference in its entirety. These systems, like the embodiments described herein, provide for a minimally invasive sampling technique and device suitable for rapid, inexpensive, unobtrusive, and pain-free monitoring of important biomedical markers, such as glucose.

[0003] Existing systems remain open to improvement in various aspects, including consistency in sampling and measurement.

Summary

[0004] A sensing device, designed to be used in contact with the skin, is provided. The device contains a plurality of individually controllable sites for electrochemically monitoring an analyte, such as glucose, in interstitial fluid of a user. The device includes:

a hydrophobic layer, designed to contact the skin;

an overlaying first structural layer;

an overlaying metal electrode layer;

an overlaying second structural layer;

for each such detection site, a capillary channel traversing these layers, thus providing an opening adjacent the skin;

wherein said metal electrode layer is discontinuous at the circumference of said capillary channel, such that two non-contiguous edge portions of electrode are present within the circumference of said channel;

appli ed to one such edge portion of the metal electrode layer, such that it is exposed to the interior of said capillary channel, a sensing layer effective to measure said analyte; and surrounding the lower end of said capillary channel, adjacent said hydrophobic layer, an electronic element (microheater) effective to produce heat when a sufficient voltage is applied thereto.

Also provided are electrical conduits and contacts such that a voltage can be applied to the microheater, and an additional voltage can be applied between the two edge portions of the electrode layer, and an electrochemical response from the sensing material/electrode layer, indicative of the concentration of analyte in the sample fluid, can be detected.

[0005] In selected embodiments, the hydrophobic layer is hydrophobic silicone. The first structural layer may be a glass or ceramic-like material. The metal electrode layer is preferably gold or platinum, and the sensing layer, for use in detecting glucose, is preferably a conducting polymer, such as polypyrrole (PPy), modified with glucose oxidase (GOx), and preferably further containing an effective amount of a mediator such as ferricyanide. The second structural layer is preferably non-absorbent and/or hydrophobic, and may also be a layer of hydrophobic silicone.

[0006] The diameter of the capillary channel, in one embodiment, is about 50 μιη.

[0007] The thickness of the metal electrode layer is generally in the range of 100 nm to 1 micron range, e.g. 100-500 nm, 500-1000 nm, 500-800 nm, 250-750 nm, 300-500 nm, etc. An exemplary thickness is about 500 nm. The structural layers generally have thicknesses such that the overall thickness of the device is about 1 mm or less.

[0008] The thickness of the applied sensing layer, measured in a direction perpendicular to the capillary channel length, may be 200 nm or less, 100 nm or less, or 50 nm or less, in selected embodiments.

[0009] In use, a voltage is applied to the microheater sufficient to ablate the stratum corneum of the underlying skin, e.g. a voltage of about 3 V, typically for about 30 msec. This ablation allows interstitial fluid to enter the capillary channel, where it rises via both capillary action and the body's hydrostatic pressure and contacts the sensing material (e.g. PPy/GOx) within the capillary. A second voltage, typically 0.2-0.4 V, is applied to the electrode layer, i.e. between the two above-described edge portions of the electrode layer, and the level of analyte (e.g. glucose) contacting the sensing material is electrochemically detected, in accordance with known methods.

Brief Description of the Drawing

[0010] Figure 1 illustrates an embodiment of a sensing device as disclosed herein. Detailed Description

[0011] A section of an exemplary sensing device, designed to be used in contact with the skin, is shown in Figure 1. The device typically contains a plurality of individually controllable sites, of which one is illustrated in the Figure, for electrochemically monitoring an analyte, such as glucose, in interstitial fluid of a user. The device, in a preferred embodiment, includes:

a hydrophobic layer 12, designed to contact the skin;

an overlaying first structural layer 14;

an overlaying metal electrode layer 16;

an overlaying second structural layer 18;

a capillary channel 20 traversing these layers, thus providing an opening adjacent the skin; wherein said metal electrode layer is discontinuous at the circumference of said capillary channel, such that two non-contiguous edge portions of electrode are present within the circumference of said channel;

appli ed to one such edge portion of the metal electrode layer, such that it is exposed to the interior of said capillary channel, a sensing layer 22 effective to measure said analyte; and surrounding the lower end of said capillary channel, adjacent said hydrophobic layer, an electronic element (microheater) 24, effective to produce heat when a sufficient voltage is applied thereto.

[0012] The diameter of the capillary channel, in one embodiment, is about 50 μιη.

[0013] The thickness of the metal electrode layer is generally in the range of 100 nm to 1 micron range, e.g. 100-500 nm, 500-1000 nm, 500-800 nm, 250-750 nm, 300-500 nm, etc. An exemplary thickness is about 500 nm. The thickness of the structural layers is not generally critical (although layer 12 should be sufficiently thick to insulate sensing material 22 from heat produced by microheater 24), but these may also be in the general range of hundreds of nm, e.g. 100-500 nm, 500-1000 nm, 500-800 nm, 250-750 nm, 300-500 nm, etc. The overall thickness of the device is generally less than 1 mm.

[0014] The diameter of the capillary channel 20, in one embodiment, is about 50 μιη. Other diameter ranges, e.g. 10 - 100 μιη, or 25-75 μιη, could also be effective.

[0015] Also provided, though not shown in the Figure, are electrical conduits and contacts such that a voltage can be applied to the microheater, and an additional voltage can be applied to the electrode layer (i.e. between the two above-described edge portions of the electrode layer), and an electrochemical response from the sensing material/electrode layer, indicative of the concentration of analyte in the sample fluid, can be detected. The multiple detection sites within a device are preferably individually controllable; i.e. voltages can be selectively applied to a given detection site or sites by a user of the device.

[0016] In selected embodiments, the hydrophobic layer 12 is hydrophobic silicone, though any biocompatible/non-irritating hydrophobic material can be used. The structural layer 12 may be a glass or ceramic-like material, which provides thermal insulation between the microheater 24 and sensing material 22, or other structurally stable, nonabsorbent, preferably thermally insulating material. The metal electrode layer 16 is preferably gold or platinum.

[0017] A sensing layer 22 effective to measure the analyte is present on one of the above- described edge portions of the metal electrode layer, such that it the sensing material is exposed to the interior of the capillary channel. The sensing layer 22, for use in detecting glucose, is preferably a conducting polymer, such as polypyrrole (PPy), modified with the enzyme glucose oxidase (GOx).

[0018] Preferably, in fabrication, the PPy-GOx layer is electrodeposited, in accordance with known methods (see e.g. Liu et al., Matl. Sci. Eng. C 27(l):47-60 (Jan 2007); Yamada et al, Chem. Lett. 26(3):201-202 (1997); Fortier et al, Biosens. Bio electronics 5:473-490 (1990)) as an extremely thin layer on an exposed face of the metal electrode, as shown in the Figure. Measuring in the direction perpendicular to the capillary length, the thickness of the applied layer may be e.g. 200 nm or less, 100 nm or less, or 50 nm or less, in selected embodiments.

[0019] In one embodiment, a mediator such as ferricyanide, as known in the art, is co-deposited along with the PPy and GOx. This system allows electrochemical measurement of the analyte to be carried out at a voltage of about 0.2 -0.4V. A somewhat higher voltage (e.g. 0.7 V), which can lead to interference from other molecules in the interstitial fluid, would typically be required without the mediator. Other electron-accepting mediators known in the art for use with GOx include ferrocene derivatives, conducting organic salts such as tetrathiafulvalen- tetracycloquinodimethane (TTF-TCNQ), quinone compounds, phenothiazine compounds, and phenoxazine compounds.

[0020] The multilayer structure of the device containing the capillary channels can be fabricated by known deposition and etching methods. The sensing material 22 is, in one embodiment, applied by electrodeposition to one of the exposed edges of the gold electrode within the formed capillary channel 20, as noted above. As noted above, two noncontiguous edges of electrode are present within the microchannel, to allow for electrochemical detection. These edges could be visualized as two distinct semicircles within the inner surface of the channel, one of which is treated with the sensing material.

[0021] In use, a voltage is applied to the microheater sufficient to ablate the stratum corneum of the underlying skin, e.g. a voltage of about 3 V, typically for about 30 msec. This ablation (which typically produces a temperature of about 130°C) allows interstitial fluid to enter the capillary channel, where it rises via capillary action and hydrostatic pressure and contacts the sensing material (e.g. PPy/GOx) within the capillary. A second voltage, typically 0.2-0.4 V, is then applied between the two above-described edge portions of the electrode layer), and the level of analyte (e.g. glucose) contacting the sensing material is electronically detected, preferably amperometrically detected, in accordance with known methods.

[0022] The device design presents various advantages, including the following. The sensor electrode pair, including the metal (e.g. gold) electrode and PPy/GOx-treated electrode (i.e. the two edge regions described above, one treated with sensor material), are located within the microcapillary channel and thus separated spatially from the microheater. This configuration avoids possible heat degradation of the enzyme. Further to this aspect, structural layer 14 is preferably formed of a heat-insulting material, such as a glass or ceramic material.

[0023] The detection of glucose is typically realized using chronoamperometry

(measurement of current generated versus time for a voltage step). Ideally, every glucose molecule reaching the GOx sensing electrode should immediately release its electrons to produce the measured current. To achieve this condition, the electrode should have a low surface area to sample volume ratio, to ensure that glucose i s not depleted in the vicinity of the sensing electrode during analysis. Accordingly, the sensing electrode is fabricated to be extremely small; i.e. essentially the width of the metal electrode layer 16, as shown in the Figure. Preferred thicknesses (measured in the direction perpendicular to the channel length) of the applied sensor layer 22 may be e.g. 200 nm or less, 100 nm or less, or 50 nm or less, in selected embodiments, in the 100 nm to 1 micron range, e.g. 500 nm. Diffusion times (i.e. the time for glucose molecules to reach the GOx enzyme) are reduced for similar reasons.

[0024] In general, the thickness of a metal layer applied via conventional metal deposition methods, e.g. electrodeposition or vapor deposition, can be precisely controlled, as compared to control of lateral dimensions of the planar surface area. Accordingly, high consistency in the effective sensor area (which is, again, the width dimension of the metal electrode layer 16), as well as roughness of the electrode layer, is achieved, giving high consistency between one sensor element and another, within a single device or between different devices. In fabrication of the multilayer device, the gold thickness can be easily reproduced with very little sidewall imperfections/roughness, and the exposed region (at the capillary wall) becomes the sensor electrode area.

[0025] Although glass/ceramic/polymeric substrate layers are exemplified, other materials, such as paper or other cellulose substrates, electrospun fibers, or other polymers, could also be used for the non-metal layers (12, 14, 18) in the device. However, surfaces contacting the skin, such as the lower surface of layer 12, should be non-absorbent and preferably hydrophobic in nature, in order to direct fluid flow from the skin into and though the capillary channel 20 to the sensing material 22. Methods of treating materials such as paper to render selected portions hydrophobic and/or non- absorbent are known in the art; see e.g. Martinez et al., Anal. Chem. 2010, 82, 3-10. The surface of structural layer 12 contacting the interior of the microchannel should be non- absorbent but should not repel water, so that sample fluid travels efficiently to the sensing area without volume loss.

[0026] Integrated circuitry (IC), including radio frequency (RF) communication capability, may be included peripheral to the device in order to transmit data readings to a remote location. By way of example, this transmission may employ Bluetooth™ devices, or it may be facilitated as part of a home area network (HAN) in a first instance, e.g., using protocols such as those described as part of the Zigbee standards. Further still, the data readings may be further transmitted outside of the HAN in accordance with a home health or telehealth communications system using existing wide area networks (WANs) such as the Internet. [0027] One skilled in the art recognizes the other areas of application for the devices described herein. While the examples specifically described herein are directed to glucose monitoring, adaptations could be made to ascertain other information from the biomolecules and biomarkers in the interstitial fluid. For example, the individual sites could monitor for infectious disease (microbial, fungal, viral); hazardous compounds; heart or stroke indicators (troponin, C-reactive protein); chemical or biological toxins; cancer markers (PSA, estrogen); drug efficacy and dosing (metabolites): and the like. Such applications of the device as described are considered to be within the scope of the present invention.