JP2003527599A - Embedded analyte sensor - Google Patents

Abstract

  (57) [Summary] The embedded analyte sensor comprises a substrate, an electrode on the substrate, and a membrane on the electrode. The membrane is composed of elemental silicon. The sensor shows less than 20% signal drift per day in vivo.

Description

Detailed Description of the Invention

[0001] BACKGROUND The present invention relates to an implantable analyte sensor.

Various implantable glucose sensors have been developed so far. As an example,
U.S. Pat. Nos. 5,387,327; 5,411,647; and 5,476,776; and PCT International Publication WO 91/15993; WO 94/20602; WO 96/06947; and
Some are described in WO 97/19344. These embedded glucose sensors usually include a polymer substrate, and a metal electrode is printed on the surface of the substrate. The electrode is covered with a biocompatible membrane, allowing glucose to reach the electrode, while excluding other molecules such as proteins. The amount of glucose present is measured by using electrochemical techniques, often with the help of enzymes attached to the electrodes. When implanting the glucose sensor in the patient's body, the electrodes can be attached via wires. This wire exits the patient's body and connects to an external circuit that controls the electrodes and also measures and records glucose levels. Alternatively, all or part of this external circuit can be miniaturized and housed in the implantable glucose sensor. Even a transmitter, such as that described in WO 97/19344, can be housed within the implantable glucose sensor, thus completely eliminating the need for leads from the patient's body.

A problem with amperometric glucose sensors is the instability of the signal. This is due to enzyme degradation, enzyme leakage, and / or enzyme interactions resulting from protein interactions.
Alternatively, it is considered to be a result of the contamination of the electrodes. A common remedy to overcome this is to use the biocompatible membranes or coatings described above. However, these membranes also have some problems. For example, Nafion-based biosensor membranes show cracks, peeling, protein attachment, calcium deposition, and the like. In biological environments, polymer-based membranes undergo calcification, cracking, and changes in permeability. The poor porosity associated with polymer membranes has also been shown to be highly involved in membrane stability and calcification in vivo. Biological components that penetrate pores and voids in the material give rise to metabolically negative sites where ions and calcium accumulate. This situation poses a potentially serious problem for implantable glucose sensors, coupled with the fact that mineral deposition is known to spread surface disruption of polymeric membranes.

In polymer membranes, their pore size distribution usually follows some kind of probability distribution (eg Gaussian distribution), which leaves a finite possibility that large proteins will eventually move through the membrane. Become. This leakage or inadequate diffusion properties may also lead to body-imaging such as biofouling and protein adsorption, encapsulation by fibrous proliferating tissue, and deterioration of device material over time.
Drift can be caused by events at the sensor interface.

Membranes with nominal pore sizes as small as 20 nm are currently available. That said, filtration at these pore sizes is far from perfect. The most common filter is a polymer membrane formed by a solution casting method, but the one obtained by this method has a large pore size distribution with a deviation of 30%. Membrane (eg
Ion-track etchin to form MILLPORE ISOPORE)
With g), a narrower pore size distribution (± 10%) can be made. However, this membrane has a low porosity (<10 ⁹ holes / cm ² ) and a limited pore size,
And the pores are randomly distributed on the surface. Porous alumina (eg WHATMAN) has also been used to obtain uniform pores. Alumina usually has a high pore number density (> 10 ¹⁰ / cm ² ), but only a specific pore size (typically larger than 20 nanometers) is obtained, controlling pore shape and placement Is difficult.

[0006] In summary one embodiment of the invention, the present invention is an implantable analyte sensor, comprising a substrate, electrodes located on the substrate, and the membrane located on the electrode. The membrane is composed of elemental silicon.

In another aspect, the invention relates to a method of making an implantable analyte sensor, comprising covering an electrode with a membrane. The electrodes are on a substrate and the membrane is composed of elemental silicon.

The present invention also relates to implantable analyte sensors that exhibit less than 20% signal drift per day in vivo.

[0009] BRIEF DESCRIPTION following drawings drawings form a part hereof, it is included to further Illustrated certain aspects of the present invention. The invention may be better understood by reference to one or more of these drawings in combination with the detailed description of specific embodiments provided herein.

[0010]   1-9 illustrate a method of making a membrane for use in an embodiment of the present invention.

[0011]   FIG. 10 shows a cross-sectional view of the embedded analyte sensor.

[0012]   FIG. 11 shows an exploded view of the embedded analyte sensor.

[0013]   FIG. 12 shows a cross-sectional view of the embedded analyte sensor.

DETAILED DESCRIPTION FIG. 10 shows a cross-sectional view of one embodiment of the present invention. In the figure, the embedded analyte sensor 2 comprises a substrate 6 on which electrodes 8 and 8 are located. The electrode is covered with the membrane 4. Leads 12 and 12 connect this embedded analyte sensor to an external circuit (
(Not shown). The implantable analyte sensor also includes an outer coat layer 16 and an inner coat layer 14.

FIG. 11 shows an exploded view of one embodiment of the present invention. The inner and outer coat layers have been omitted from the figure for clarity. As shown, the implantable analyte sensor 2 includes electrodes 8 and 8 on the surface of a substrate 6, which electrodes are electrically connected to a microelectronic circuit 10. The microelectronic circuit is electrically connected to leads 12 and 12, which allows the implantable analyte sensor to be electrically connected to external circuitry (not shown). The electrode is covered with the membrane 4.

FIG. 12 shows a cross-sectional view of an embodiment of the invention similar to that shown in FIG. 10, but with a third electrode 8 and a third lead wire 12. There is. Although shown as such, the number of electrodes and the number of leads may be different.

The membrane is composed of a micromachined hard material. Preferably, the membrane is composed of elemental silicon, but other biocompatible hard materials that can be micromachined are also contemplated, such as metals (eg titanium), ceramics (eg
Silica and silicon nitride), and polymers (eg, polytetrafluoroethylene, polymethylmethacrylate, polystyrene, silicone resins, etc.). Micromachining is a method that includes photolithography, such as is used in the semiconductor industry, to remove material from or add material to a substrate. These methods are well known and can be found in the Encyclopedia of Chemical Technology,
Kirk-Othmer, Volume 14, pp. 677-709 (1995); Semiconductor Device Fundam
entals, Robert F. Pierret, Addison-Wesley, 1996; and Microchip Fabrica
tion 3rd edition, Peter Van Zant, McGraw-Hill, 1997. Details of how to fabricate a membrane composed of elemental silicon can be found in the PhD dissertation of Derek James Hansford of the University of California, Berkeley, graduated in the spring semester 1999. Doctor of Philosophy in Engineering-Materials Science and
Mineral Engineering) has been submitted as part of the acquisition requirements).

The characteristic of this membrane is that the pore size is defined, and it has a narrow pore size distribution compared to the standard membrane pore size distribution. Due to the small tolerances of this fabrication method, the pore size can be controlled in the exact diameter range, e.g.
, Or 5 to 20 nm, or even 5 to 15 nm (for example, 12 nm, 18 nm, or
Even at 25 nm) and the deviations are +/- 0.01-20%, +/- 0.1-10 %%,
Or even +/- 1 to 5%. Therefore, molecules above this size can almost certainly be excluded. This is because the pore size distribution is more in the shape of a bowler hat, rather than a bell curve, and thus, for example, 12 nm, 18 nm, 25 nm, or 50 nm.
There are no larger size pores. These membranes can exclude interfering molecules such as proteins. This interfering molecule causes significant drift problems for the sensor if it is implanted inside the body. Signal drift is a change in the magnitude of the signal from the sensor that is independent of changes in analyte concentration. The amount of signal drift is
The size of the signal before drifting is used as the standard. Preferably, the implantable analyte sensor of the present invention exhibits a signal drift of less than 20% per day in vivo, more preferably less than 10% per day in vivo,
Most preferably, it exhibits less than 5% signal drift per day in vivo.

The membranes used in the present invention can be characterized by glucose and albumin diffusion experiments. These experiments will be explained later. Preferably, the membrane has a glucose diffusion experimental result of at least 1 mg / dl in 330 minutes,
More preferably at least 10 mg / dl in 330 minutes, even more preferably 330
Those that are at least 30 mg / dl in minutes, and most preferably at least 60 mg / dl in 330 minutes. Preferably, the membrane has an albumin diffusion test result of 420
At most 0.1 g / dl, more preferably at 420 minutes at most 0.05 g / dl, even more preferably at 420 minutes at most 0.01 g / dl, and most preferably at 420 minutes. At most 0.001 g / dl.

This method of manufacturing a membrane enables a simple and economical manufacturing of a small-sized and embedded type analyte sensor. For example, first make a membrane,
Thereafter, the sensor electrode and the electrical connector can be formed on the substrate. Preferably, the substrate is silicon, but other materials are possible, such as ceramics and polymers. If desired, electronic components such as amplifiers, filters, transmitters, and / or signal preconditioning components can easily be incorporated into this layer. In particular, if the substrate is composed of elemental silicon, well-known integrated circuit technology can be used to place all circuits in a miniaturized form on a single chip.

When reagents are included in the sensor, there are two possible ways to attach the substrate and the membrane: The substrate and membrane are thermally bonded before the reagent is deposited on the electrodes. In this case, one opening is preferably provided in the membrane (since it is produced by a micromachining method, the opening can easily be made in one step of this processing method). If multiple membranes are formed as one piece, and / or multiple substrates are formed as one piece, then after thermal bonding, an additional etching step is used to separate individual membrane / substrate units. be able to. Reagents are deposited through individual openings and the openings are sealed with, for example, a polymer sealant. The individual sensors are then separated and incorporated into a flexible inner coat layer, such as silicone rubber, and then individually coated with an outer coat layer, such as a biocompatible layer.

2. The reagent is deposited on the electrodes before the membrane and substrate are attached. In this case, the enzyme in the reagent may be destroyed, so that thermal bonding cannot be performed. The membrane and substrate are first separated individually and the individual sensors are assembled by adhering one membrane to one substrate using a suitable adhesive such as cyanoacrylate. As a final step, the individual sensors are incorporated into a flexible inner coat layer, such as silicone rubber, and then individually coated with an outer coat layer, such as a biocompatible layer. The sensor can be inserted into the skin using a needle applicator. The control device can usually be left outside the body and connected to the sensor element via electrical wires (leads).

An electrode is formed on the surface of the substrate. The electrodes can be formed from conductive materials such as pure metals or alloys or other materials that are metal conductors by well known semiconductor processing techniques. Examples are aluminum, carbon (eg graphite), cobalt, copper, gallium, gold, indium, iridium, iron, magnesium, mercury (as amalgam), nickel, niobium, osmium, palladium, platinum, rhenium, rhodium, selenium, There are silicon (eg highly doped polycrystalline silicon), silver, tantalum, tin, titanium, tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, and alloys or metal compounds of these elements. Preferably, the electrodes contain gold, platinum, palladium, iridium, or alloys of these metals, because such precious metals and their alloys are inert in biological systems. . The electrodes can be of any thickness, but are preferably 10 nm to 1 mm, more preferably 20 nm.
˜100 μm, or more preferably 25 nm to 1 μm.

There must be at least two electrodes. The number of electrodes is 2 to 1000, or 3 to 2
The number may be 00, or 3 to 99. Individual electrode sets (2 or 3 electrodes) can be placed separately in individual chambers, each covered by the membrane described above. In addition, different reagents may be attached to individual electrode sets (two or three electrodes), so that at least two types (for example, 3 ~
An implantable analyte sensor that can measure 100 or 4 to 20 different analytes is possible.

The remaining individual parts of the implantable analyte sensor are well known to those skilled in the art,
For example, US Pat. Nos. 5,387,327; 5,411,647; and 5,476,776; and PCT International Publication WO 91/15993; WO 94/20602; WO 96/06947; and WO 97 /.
19344.

Although illustrated with both leads and microelectronic circuitry, these components are optional and are included if desired. The microelectronic circuit may include some or all of the electrical components that normally lie outside the implantable analyte sensor, such as a microprocessor, amplifier, or power supply. Also, if the microelectronic circuit also includes a transmitter, or other device that wirelessly transmits information, such as a laser, which emits light through the skin, it need not include leads. On the contrary,
There may be no microelectronic circuitry, in which case the leads directly electrically connect the electrodes to external electrical components.

There may be one or more inner coat layers if desired. The inner coat layer is believed to serve the function of controlling diffusion. Examples of the inner coat layer include cellulose acetate, polyurethane, polyallylamine (PAL), polyaziridine (PAZ),
And silicon-containing polymers. PC with some concrete examples
T-published WO 98/17995, WO 98/13685 and WO 96/06947, and US Pat.
650,547 and 5,165,407.

There may be one or more outer coat layers if desired. The implantable analyte sensor of the present invention is intended to be used in vivo, preferably subcutaneously in a mammal such as human, dog or mouse. The outer coat layer functions to improve the biocompatibility of the implantable analyte sensor. Examples of the outer coating layer include Nafion, polyurethane, polytetrafluoroethylene (PTFE), poly (ethylene oxide) (PEO), and a copolymer of 2-methacryloyloxyethylphosphorylcholine and n-butyl methacrylate (MPC )
There are various membranes. Some specific examples are described in PCT Publication WO 96/06947, and Medical Progress through Technology, Nishida et al., 21: 91-103 (1995).

The electrodes may be coated with a reagent. Reagents are optional and are used to provide electrochemical probes for specific analytes. The reagent can be as simple as a single enzyme, eg glucose oxidase or glucose hydrogenase for glucose detection. The enzyme may be immobilized or "wired" as described in PCT Publication WO 96/06947. The reagent may contain a mediator in order to increase the sensitivity of the sensor. The starting reagents are the reactants or components of the reagents and are usually mixed together in liquid form prior to application to the electrodes. The liquid then evaporates, leaving the reagent in solid form.
The choice of a specific reagent depends on the specific analyte or the analyte to be measured and is well known to those skilled in the art. For example, as an example of a glucose measuring reagent, 62.2 mg of polyethylene oxide (average molecular weight 100 to 900 kilodalton), 3.3 mg of NATROSOL 250M, 41.5 mg of AVICEL RC-591 F per gram of the reagent are used.
, 89.4 mg monobasic potassium phosphate, 157.9 mg dibasic potassium phosphate, 437.3 mg potassium ferricyanide, 46.0 mg sodium succinate, 148.0 mg trehalose, 2.6 mg TRITON X-100 surfactant, and 2,000 Mention may be made of those containing ~ 9,000 units of enzyme activity. This enzyme is prepared as an enzyme solution from 12.5 mg of coenzyme PQQ and 1.12 million units of apoenzyme of quinoprotein glucose dehydrogenase to make a quinoprotein glucose dehydrogenase solution. This reagent is WO 9
9/30152, pages 7-10, which is hereby incorporated by reference.

Non-limiting examples of other enzymes and additional mediators that can be used in the present invention to measure specific analytes are listed below in Table 1.

[0031]

[Table 1]

In some of the examples shown in Table 1, at least one additional enzyme is used as the reaction catalyst. Also, in some of the examples shown in Table 1, additional mediators may be used, which facilitate electron transfer to the oxidized form of the mediator. The additional mediator is added to the reagent in lesser amount than the oxidized mediator. Although the above assays are described, it will be appreciated that a variety of electrochemical assays can be performed in accordance with the present disclosure.

Formation of Membrane The following describes how to make the membrane used in the present invention, which was submitted in the spring semester of 1999 by Derek James Hansford of the University of California, Berkeley Graduate School.
The doctoral dissertation (which was submitted as part of the requirements for obtaining a doctoral degree in engineering in Materials Science and Mineral Engineering) is based on the description given in the above.

Other membranes made from other materials can also be used. This unique method relies on a buried nitride etch stop layer.

The buried nitride film etch stop layer acts as an etchant stop when forming nanometer scale holes. The buried nitride etch stop layer facilitates three-dimensional control of the pore structure and facilitates the formation of pores with diameters less than 50 nanometers. Moreover, these holes can be formed uniformly over the entire surface of the wafer.

Preferably, the first step in this fabrication protocol is to etch the supporting ridge structure into the substrate. This ridge imparts mechanical rigidity to the subsequently formed membrane structure.

Then, a low stress chemical vapor deposition (LPCVD) method is used to deposit a low stress silicon nitride film (LSN or nitride film) acting as an etch stop layer on the substrate. In one embodiment, a 0.4 μm nitride film was used. The resulting structure is shown in FIG. FIG. 1 shows a substrate 20 and a nitride film etching stop layer 22 formed thereon.

On the stop layer 22, a base structure layer (base layer) of the membrane is deposited.
The etch stop layer 22 is thin so that the structural layer is deposited up to the supporting ridges formed in the substrate 20. In one embodiment, 5 μm polysilicon was used as the base layer. FIG. 2 illustrates the base layer 24 disposed on the etch stop layer 22. The low stress silicon nitride film can also be used as the base layer, in which case it acts as its own etch stop layer.

The next processing step is to etch holes in the base layer 24 to define the shape of the holes. Masks such as those used in conventional semiconductor processing can be used to define the holes. For example, a thermally grown oxide layer may be used as a mask and a chlorine plasma may be used to etch holes in the polysilicon. In this step, it is important to make sure that the etching completely penetrates the base layer 24, so it is preferable to perform 10-15% overetching.
It is worth noting that the buried nitride etch stop layer 22 acts as an etch stop for the plasma etching of the silicon base layer 24. Otherwise, if the plasma punctures the nitride, to prevent complete removal of the nitride underneath the plug layer (prevent it from being removed by the final KOH etch). Therefore, it is thought that more precise control of the etching process needs to be performed. FIG. 3 illustrates the result of this processing method. In particular, this figure shows a hole formed in the base layer 24 but ending in the nitride etch stop layer 22.
26 is an illustration of 26.

Next, a hole sacrificial oxide is grown on the base layer 24. FIG. 4 illustrates the sacrificial oxide 28 disposed on the base layer 24.

The thickness of this sacrificial oxide determines the pore size of the final membrane. Therefore, control of this process is extremely important for reproducible membranes. This is the base layer
24 thermal oxidation (eg, growth temperature of 850-950 ° C. for about 1 hour and 10 minutes of annealing). Of course, various methods can be used to form a sacrificial layer of controlled thickness. For example, a thermally evaporated tungsten film may be used as a sacrificial layer for the polymeric membrane, which can be selectively removed with hydrogen peroxide. A fundamental requirement of the sacrificial layer is the ability to control its thickness with high precision over the entire wafer surface. Thermal oxidation of both polysilicon and nitride allows sacrificial layer thickness control of less than 5% over the entire wafer. The limits of this control come from local inhomogeneities in the base layer, such as the initial thickness of the native oxide (especially for polysilicon), grain size or density, and impurity concentration.

Anchor points were defined in the sacrificial oxide layer 28 to mechanically connect the base layer 24 with the plug layers (required to maintain hole spacing between the layers). In this design,
This is achieved by using the same mask, which is diagonally offset from the holes by 1 μm. This provided anchors at one or two corners of each hole, which provided the desired mechanical connection between the layers of the structure while opening as much hole area as possible. FIG. 5 illustrates an anchor 30 formed by this method.

Next, a plug structure layer is deposited to fill the hole 26. This step was performed by depositing 1.5 μm of polysilicon. The resulting plug layer 32 is shown in FIG.

In order to make holes in the surface, the plug layer 32 is planarized up to the base layer,
, Leaving the final structure with the plug layer only in the hole openings.

The method of planarization depends on the material used as the plug material. For hard microfabrication materials (polysilicon and nitride films), chemical mechanical polishing was used for planarization. Among the other materials examined, plasma etching is used as a quick wet chemical smoothing method.
) Was used to roughen and flatten. This method has the advantage that the base layer is not affected, assuming that the plasma used does not etch the base layer.
There is a drawback in that it is necessary to control the etching timing in order to avoid completely etching the plug itself.

At this point, the membrane is ready for release and a protective layer 34 is deposited on the wafer (completely covering both sides of the wafer). The essential requirements for the protective layer 34 are
The protective layer is not affected by the etching of silicon (KOH in these experiments) and the protective layer can be removed without removing the plug 32 or the structural layer of the base 24. In the case of polysilicon and nitride structure layers, a thin nitride layer is used as a protective layer (nitride film is not etched in KOH at all and dissolves slowly in HF).
In the case of a polymer structural material, silicon is used as a protective layer in view of the processing temperature (835 ° C.) required for vapor deposition of a nitride film.

A backside etching window was provided on the protective layer by etching. Expose the silicon to the desired extent, then place the entire structure in a 80 ° C. KOH bath to etch the silicon wafer substrate 20 (as evidenced by the smooth buried etch stop) to the base layer 24 of the membrane. did. FIG. 8 illustrates the resulting aperture 36 formed in the substrate 20.

At this point, the buried nitride film layer 22, the sacrificial oxide layer 34, and the plug layer 32 are exposed to HF or
It is removed by etching in SF ₆ / oxygen plasma. The resulting membrane 4 with nanometer scale pores is shown in FIG.

Characterization of the Membrane The purpose of the membrane is to disperse analytes of interest (eg glucose) through the membrane while at the same time eliminating large molecules (eg proteins). Therefore, two important properties of the membrane are glucose diffusion and albumin diffusion. All experiments are performed at room temperature (25 ° C).

The following is a glucose diffusion experiment: Glucose diffusion is measured using a mini-diffusion chamber assembled around a membrane. Made from acrylic resin, this diffusion chamber consists of two compartments A and B with a constant volume of 2 mL, separated by the desired membrane, sealed with an O-ring, and screw-like. It is screwed in by turning it to.

Using this diffusion chamber, an enzymatic quantitative assay (TRINDER ™, SIGMA
) And a spectrophotometric colorimetric measurement to measure glucose on both sides of the membrane. The starting glucose concentration for all experiments was 6,666 mg / dl and 0.0 mg / dl in chambers A and B, respectively. A 0.1 ml sample was taken from the diffusion chamber, 10 μl of which was added to 3 ml glucose reagent in a cuvette and gently mixed by inversion. Incubate each tube for 18 minutes at room temperature and
Measure at 5 nm. Reagents are linear up to 750 mg / dl. The diffusion chamber itself is mounted on a motor for agitation in order to minimize the boundary layer effect (diffusion resistance at the liquid / membrane interface). To ensure that the pores are moistened, first fill the recipient cells with phosphate buffered saline (PBS) for 15 minutes, then
Fill donor cell. The donor cell is filled with various concentrations of glucose solution in PBS.

The following are albumin diffusion experiments: Similar to glucose diffusion experiments, the same diffusion chamber is used to measure albumin on both sides of the membranen. Albumin diffusion and / or exclusion is first measured and quantified using Albumin BCP (bromocresol purple, SIGMA). Starting albumin concentrations for all experiments were in chambers A and B respectively.
4 g / dl and 0.0 mg / dl. Zero time and end of diffusion period (time = 330 minutes)
Take a 0.1 ml sample with. Then add a 300 μL aliquot to 3 mL of reagent and measure the absorbance at 600 nm. As a blank, a reagent prepared by adding deionized water is used. The BCP assay was linear up to 6 g / dl but not accurate below 1 g / dl. For low concentrations of albumin that may be present in chamber A, use the Bradford Method (MICRO PROTEIN KIT, SIGMA) for chamber B
The presence of protein in was determined. This method quantifies the binding of Coomassie brilliant blue to unknown proteins and compares the amount of binding to the amount of binding to various amounts of standard protein. Albumin is used as a standard protein. This method quantifies 1-100 micrograms of protein using a calibration curve with sensitivities up to 10 mg / dl or 0.1 g / dl of protein. Absorbance is measured at 595 nm.

Membrane Analysis Glucose diffusion was measured on three types of membranes: silicon micromachined membranes (mean pore size = 0.0245 micron), WHATMAN ANODISC membranes (mean pore size = 0.02 micron), and MF-MILLIPORE acetic acid nitric acid mixture. Cellulose membrane (average pore size = 0.025 micron).

[0054]   The results of the albumin experiments are shown in the table below.

[0055]

The presence of albumin is not expected to impede glucose passage through the membrane or slow the transport of glucose. The amount of albumin that diffuses through the micromachined membrane is undetectable. However, the same membrane shows diffusion of glucose. Micromachined membranes allow the diffusion of glucose while achieving complete exclusion of albumin (up to the limit of detection). Comparing the diffusion rate to that of commercial membranes, the micromachined membranes have MILLIPORE membranes and alumina WHATMAN with similar pore sizes.
It has glucose diffusion properties comparable to membranes.

The passage of albumin through the micromachined membrane is measured by looking at the change in albumin concentration in chamber A and chamber B over time. A BCP assay was used, but there is no detectable trace of albumin in chamber B. However, it is possible that the amount of albumin in chamber B is below the detection limit of this assay system. So I also adopted the Bradford method. Using this microassay, no detectable amount of albumin was again found in chamber B on the micromachined membrane, but MILLIPORE or WHA
When using the TMAN membrane, a small amount of protein was found in chamber B in both cases. The amount of albumin detected in chamber B after 420 minutes was determined by MILLIPORE
And about WHATMAN membrane were about 0.25 g / dL and 0.20 g / dL, respectively.

Glucose, of course, diffuses through micromachined membranes at rates similar to commercially available membranes. At the same time, the passage of albumin is excluded. In a mixed glucose and albumin solution, only glucose diffuses through the micromachined membrane.

[Brief description of drawings]

[Figure 1]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Fig. 2]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 3]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 4]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 5]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 6]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 7]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 8]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 9]   It is a figure which shows the manufacturing method of the membrane used by embodiment of this invention.

[Figure 10]   It is sectional drawing of an embedded type analyte sensor.

FIG. 11   FIG. 6 is an exploded view of the embedded analyte sensor.

[Fig. 12]   It is sectional drawing of an embedded type analyte sensor.

[Procedure for Amendment] Submission for translation of Article 34 Amendment of Patent Cooperation Treaty

[Submission date] February 22, 2001 (2001.2.22)

[Procedure Amendment 1]

[Document name to be amended] Statement

[Name of item to be amended] Claims

[Correction method] Change

[Contents of correction]

[Claims]

15. The method of claim 11, wherein the implantable analyte sensor is a glucose sensor.

[Procedure amendment]

[Submission date] October 17, 2002 (2002.10.17)

[Procedure Amendment 1]

[Document name to be amended] Statement

[Name of item to be amended] Claims

[Correction method] Change

[Contents of correction]

[Claims]

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Claims (27)

[Claims] 1. A substrate comprising: (a) a substrate; (b) an electrode on the substrate; and (c) a membrane on the electrode, the membrane comprising elemental silicon. Embedded analyte sensor that does. 2. The implantable analyte sensor of claim 1, further comprising (d) a microelectronic circuit electrically connected to the electrodes. 3. The embedded analyte sensor according to claim 1, further comprising: (e) a lead wire electrically connected to the electrode. 4. (e) further comprising a lead wire electrically connected to said electrode, said lead wire being electrically connected to said electrode via said microelectronic circuit. The embedded analyte sensor described in 2. 5. The implantable analyte sensor of claim 2, wherein the microelectronic circuit includes a transmitter and a power supply. 6. (f) a coat layer surrounding the substrate and the membrane,
The implantable analyte sensor of claim 1, further comprising: 7. The embedded analyte sensor according to claim 6, wherein the coat layer includes an inner coat layer and an outer coat layer. 8. The embedded analyte sensor according to claim 1, wherein the substrate is composed of elemental silicon. 9. The membrane is manufactured by micromachining.
Embedded analyte sensor described in. 10. The implantable analyte sensor of claim 1, wherein the implantable analyte sensor is a glucose sensor. 11. A method of making an embedded analyte sensor comprising covering an electrode with a membrane, wherein the electrode is on a substrate and the membrane is composed of elemental silicon. And the above method. 12. The method of claim 11, further comprising forming the membrane by micromachining elemental silicon. 13. The method of claim 11, further comprising surrounding the membrane and the substrate with a coat layer. 14. The membrane is manufactured by micromachining.
The method according to 1. 15. The method of claim 11, wherein the implantable analyte sensor is a glucose sensor. 16. An embedded analyte sensor comprising an electrode on a substrate and a membrane covering the electrode, the improvement being the embedded analyte sensor, wherein the membrane is composed of elemental silicon. . 17. A method comprising: (1) means for electrochemically measuring an analyte; and (2) a membrane on the means, wherein the membrane is composed of elemental silicon. Embedded analyte sensor. 18. A substrate comprising: (a) a substrate; (b) an electrode on the substrate; and (c) a membrane on the electrode, the membrane being manufactured by micromachining. Embedded analyte sensor that does. 19. A substrate comprising: (a) a substrate; (b) an electrode on the substrate; and (c) a membrane on the electrode, which exhibits a signal drift of less than 20% per day in vivo. Embedded type analyte sensor featuring. 20. The implantable analyte sensor of claim 19, wherein the sensor exhibits less than 10% signal drift per day in vivo. 21. The implantable analyte sensor of claim 19, wherein the sensor exhibits less than 5% signal drift per day in vivo. 22. The implantable analyte sensor of claim 19, further comprising (d) a microelectronic circuit electrically connected to the electrodes. 23. The embedded analyte sensor according to claim 19, further comprising: (e) a lead wire electrically connected to the electrode. 24. The method further comprising (e) a lead wire electrically connected to the electrode, the lead wire being electrically connected to the electrode via the microelectronic circuit. 22. The embedded analyte sensor described in 22. 25. The implantable analyte sensor of claim 22, wherein the microelectronic circuit includes a transmitter and a power supply. 26. The embedded analyte sensor according to claim 19, further comprising (f) a coat layer surrounding the substrate and the membrane. 27. The coat layer includes an inner coat layer and an outer coat layer.
The embedded analyte sensor according to claim 26.