CA2406878A1 - Implantable analyte sensor - Google Patents
Implantable analyte sensor Download PDFInfo
- Publication number
- CA2406878A1 CA2406878A1 CA002406878A CA2406878A CA2406878A1 CA 2406878 A1 CA2406878 A1 CA 2406878A1 CA 002406878 A CA002406878 A CA 002406878A CA 2406878 A CA2406878 A CA 2406878A CA 2406878 A1 CA2406878 A1 CA 2406878A1
- Authority
- CA
- Canada
- Prior art keywords
- membrane
- analyte sensor
- electrodes
- implantable analyte
- glucose
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Abandoned
Links
Classifications
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B5/00—Measuring for diagnostic purposes; Identification of persons
- A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue
- A61B5/1486—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue using enzyme electrodes, e.g. with immobilised oxidase
- A61B5/14865—Measuring characteristics of blood in vivo, e.g. gas concentration, pH value; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid, cerebral tissue using enzyme electrodes, e.g. with immobilised oxidase invasive, e.g. introduced into the body by a catheter or needle or using implanted sensors
-
- C—CHEMISTRY; METALLURGY
- C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
- C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
- C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
- C12Q1/001—Enzyme electrodes
- C12Q1/005—Enzyme electrodes involving specific analytes or enzymes
- C12Q1/006—Enzyme electrodes involving specific analytes or enzymes for glucose
Abstract
An implantable analyte sensor includes a substrate, electrodes on the substrate, and a membrane on the electrodes. The membrane can comprise elemental silicon and has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
Description
Implantable Analyte Sensor BACKGROUND
The present invention relates to implantable analyte sensors.
Several implantable glucose sensors have been developed. Examples include those described in U.S. Patent numbers 5,387,327; 5,411,647; and 5,476,776; as well as those described in PCT International Publication numbers WO 91/15993; WO 94/20602;
WO
96/06947; and WO 97/ 19344. The implantable glucose sensors usually include a polymer substrate, with metal electrodes printed on the surface of the substrate. A
biocompatible membrane covers the electrodes, allowing glucose to reach the electrodes, while excluding other molecules, such as proteins. Electrochemistry, often with the aid of enzymes at the electrodes, is used to determine the quantity of glucose present. The glu-cose sensor is implanted into a patient, and the electrodes may be attached via wires that pass out of the patient's body to external circuitry that controls the electrodes, measures and reports the glucose concentration. Alternatively, all or part of this external circuitry may be miniaturized and included in the implantable glucose sensor. A
transmitter, such as that described in WO 97/ 19344, may even be included in the implantable glucose sensor, completely eliminating the need for leads that pass out of the patient.
A problem associated with an amperometric glucose sensor is unstable signals.
This may result from degradation of the enzyme from interaction with protein, leakage of the enzyme, and/or fouling of the electrode. The usual way to overcome this is to use the above described biocompatible membrane, or a coating. However, several problems are also associated with these membranes. For example, Nafion-based biosensor membranes exhibit cracking, flaking, protein adhesion, and calcium deposits.
Mineralization of polymer-based membranes occurs in the biological environment, resulting in cracking and changes in permeabilit<~. The tortuous porosity associated with polymer membranes has also been shown to be important in membrane stability and mineralization in vivo. Biological components, which enter pores or voids in the mate rial, cause metabolic shadows, which are loci for ion and calcium accumulation. This situation, coupled with the fact that mineral deposits have been known to propagate surface fractures in polymeric membranes, presents a potentially serious problem for implantable glucose sensors.
In polymer membranes the pore size distribution usually follows some kind of prob-ability distribution (e.g. gaussian), which leaves a finite probability for large proteins to eventually transfer through the membrane. Drift may be caused by this leakage or inadequate diffusion properties, and events at the body-sensor interface such as bio-fouling and protein adsorption, encapsulation with fibrotic tissue, and degradation of the device material over time.
Currently, membranes with nominal pore sizes as small as 20 nm are available.
Even so, the filtration at these dimensions is far from absolute. The most common filters are polymeric membranes for med from a solvent-casting process, which result in a pore size distribution with variations as large as 30%. The use of ion-track etching to form mem-branes (e.g. MILLPORE ISO PORE) produces a much tighter pore size distribution (~10%). However, these membranes have low porosities (<109 pores/cm2), limited pore sizes, and the pores are randomly distributed across the surface. Porous alumina (e.g.
WHATMAN) has also been used to achieve uniform pores. Although the aluminas typically have higher pore densities (>10~°/cm'), only certain pore sizes (typically greater than 20 nanometers ) can be achieved and the pore configurations and arrangements are difficult to control.
BRIEF SUMMARY
In one aspect, the present invention is an implantable analyte sensor, comprising a sub-strate, electrodes on the substrate, and a membrane on the electrodes. The membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffu-sion test result of at most 0.1 g/dl in 420 min. and can comprise elemental silicon.
In another aspect, the present invention relates to a method of making an implantable analyte sensor, comprising covering electrodes with a membrane. The electrodes are on a substrate, and the membrane has a glucose diffusion test result of at least 1 mg/dl in WO 01/68901 CA 02406878 2002-09-13 pCT~P01/03027 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
The mem-brane can comprise elemental silicon.
BRIEF DESCRIPTION OF THE DRAWINGS
The following drawings form part of the present specification and are included to further demonstrate 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 presented herein:
Figured-9 illustrate the process of making a membrane for use in an embodiment of the present mvenrion;
Figure 10 shows a cut-away view of an implantable analyte sensor;
Figure 11 shows an exploded view of an implantable analyte sensor; and Figure 12 shows a cut-away view of an implantable analyte sensor.
DETAILED DESCRIPTION
Figure 10 shows a cut away view of an embodiment of the present invention. In the figure, an implantable analWe sensor 2 includes a substrate 6 on which are electrodes 8 and 8. The electrode, are cowered with a membrane 4. Leads 12 and 12 allow for elec-trically connecting the implantable analyte sensor to external circuitry (not shown). The implantable analyte sensor also includes an external coating 16 and an internal coating 14.
Figure 11 shows an exploded view of an embodiment of the present invention.
The internal and external coatings are not included in the figure for clarity. As shown in the figure, the implantable analyte sensor 2 includes the electrodes 8 and 8 on the substrate 6 surface, which are electrically connected with microelectronic circuitry 10.
The micro-electronic circuitry is electrically connected to leads 12 and 12, which allow for electri-tally connecting the implantable analyte sensor to external circuitry (not shown). The electrodes are covered with the membrane 4.
Figure 12 shows a cut away view of an embodiment of the present invention similar to that shown in Figure 10, except for the presence of a third electrode 8 and a third lead 12. Although so illustrated, the number of electrodes may be different from the number of leads.
The membrane is composed of a hard material that has been micromachined.
Prefera-bly, the membrane comprises elemental silicon, but other hard, biocompatable materials that can be micromachined are possible, such as metals (for example titanium), ceram-ics (for example, silica or silicon nitride), and polymers (such as polytetrafluoroethylene, polymethylmethacrylate, polystyrenes and silicones). Micromachining is a process that includes photolithography, such as that used in the semiconductor industry, to remove material from, or add material too, a substrate. These techniques are well known, and are described, in Encyclopedia of Chemical Technology, Kirk-Othmer, Volume 14, pp.
677-709 ( 1995); Semiconductor Device Fundamentals, Robert F. Pierret, Addison-~~~esley, 1996; and Mlicrochip Fabrication 3rd. edition, Peter Van Zant, McGraw-Hill, 1997. A detailed fabrication method for a membrane comprising elemental silicon is described in the dissertation of Derek James Hansford, submitted in partial satisfaction of the requirements for the degree of Doctor of Philosophy in Engineering-Materials Science and hfineral Engineering in the Graduate Division of the University of Califor-nia, Berkeley, submitted in the spring of 1999.
A special property of the membrane is a defined pore size, which has a small size distri-bution compared to the size distribution of standard membranes. Due to tight toler-ances in the manufacturing process, the pore size can be controlled at precise diameters, for example 1 to 50 nm, or ~ to 20 nm, or even 5 to 15 nm (such as 12 nm, 18 nm or even 25 nm), with a variation of +/- 0.01-20%, +/-0.1-10% or even +/-1-5%.
Therefore molecules above this size can be excluded with high certainty, since the size distribution has the shape of a top hat, rather than a bell curve, and hence pore sizes above, for example 12 nm, 18 nm, 25 nm or 50 nm are not present. These membranes can exclude interfering molecules, such as proteins, which could otherwise cause major drift prob-lems of the sensor, when the sensor is implanted in vivo. Signal drift is a change in the magnitude of the signal from a sensor which is unrelated to changes in analyte concen-tration. The amount of signal drift is based on the magnitude of the signal prior to the drifting. Preferably, the implantable analyte sensors of the present invention exhibit a signal drift of less than 20% per day in vivo, more preferably less than 10%
per day in vivo, most preferably less than 5% per day in vivo.
Membranes for use in the present invention may be characterized by a glucose diffusion test and an albumin diffusion test. These tests are described below.
Preferably, the membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., more pref erably at least 10 mg/dl in 330 min., even more preferably at least 30 mg/dl in 330 min., 5 and most preferably at least 60 mg/dl in 330 min. Preferably, the membrane has an albumin diffusion test result of at most 0.1 g/dl in 420 min., more preferably at most 0.05 g/dl in 420 min., even more preferably at most 0.01 g/dl in 420 min., and most pref erably at most 0.001 g/dl in 420 min.
The manufacturing process of the membranes may allow a simple and economical pro-duction of small, implantable analyte sensors. For example, the membranes can be first manufactured, and then on a substrate, the electrodes for the sensor and the electrical connectors can be formed. Preferably, the substrate is silicon, but other materials are possible, such as ceramics, or polymers. If desired, electronic components, for example, amplifiers, filters, transmitters and/or signal preconditioning components, can easily be incorporated in this layer. In particular, if the substrate comprises elemental silicon, well known integrated circuit technology may be used to place all the circuitry in miniaturized form on a single chip.
There are two possible approaches to attach the substrate and the membrane, when a reagent is included in the sensor:
1. The substrate and the membrane are thermally bonded before the reagent is depos-ited on the electrodes. In this case, an opening, preferably in the membrane is pro-vided (since this may be manufactured with a micromachining process, an opening is easily generated during one of the processing steps). In the case where multiple membranes are formed as a single piece, and or multiple substrates are formed as a single piece, after thermal bonding, a further etching step may be used to separate the individual membrane/substrate units. The reagent is deposited through the in-dividual openings and the openings are sealed using, for example a polymer sealant.
The individual sensors are then separated, incorporated into a flexible, inner coat-ing, for example silicone rubber, and individually coated with an outer coating, such as a biocompatible layer.
2. The reagent is deposited on the electrodes before the membrane and substrate are attached. In this case, thermal bonding is not possible, since the enzyme in the rea-gent would be destroyed. The individual membranes and substrates are first sepa-rated and the individual sensors are assembled by bonding one membrane with one substrate using a suitable bonding agent, for example, cyanoacrylate. As a final step, the individual sensors are incorporated into a flexible, inner coating, for example silicone rubber, and individually coated with an outer coating, such as a biocom-patible layer. The sensor can be inserted into the skin using a needle applicator. The control unit typically remains outside the body and can be connected to the sensor element through electrical wires (leads).
The electrodes are formed on the surface of the substrate. They may be formed by well known semiconductor processing techniques, from conductive materials, such as pure metals or alloys, or other materials which are metallic conductors. Examples include aluminum, carbon (such as graphite), cobalt, copper, gallium, gold, indium, iridium, iron, lead, magnesium, mercury (as an amalgam), nickel, niobium, osmium, palladium, platinum, rhenium, rhodium, selenium, silicon (such as highly doped polycrystalline silicon), silver, tantalum, tin, titanium, tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, and alloys or metallic compounds of these elements.
Preferably, the electrodes include gold, platinum, palladium, iridium, or alloys of these metals, since such noble metals and their alloys are unreactive in biological systems. The electrodes may be any thickness, but preferably are 10 nm to 1 mm, more preferably, 20 nm to 100 ~tm, or even 25 nm to 1 ym.
At least two electrodes must be present. The number of electrodes may be 2-1000, or 3-200, or even 3-99. Individual electrode sets (2 or 3 electrodes) may be separated into individual chambers, each covered with the membrane. Furthermore, individual elec-trode sets (2 or 3 electrodes] may each have a different reagent, allowing for an im-plantable analyte sensor that can measure at least two, such as 3-100, or 4-20, different analytes.
The remaining individual part of the implantable analyte sensors are well known to those of ordinary skill in the art, and are described, for example, in U.S.
Patent numbers 5,387,327; 5,411,647; and 5,476,776; as well as in PCT International Publication num-bers WO 91/15993; VVO 94/20602; VVO 96/06947; and WO 97/19344.
Although illustrated with both leads and microelectronic circuitry, these components are optional. The microelectronic circuitry may include some or all of the electrical components normally external to the implantable analyte sensor, such as a microproc-essor, an amplifier, or a power supply. If the microelectronic circuitry also includes a transmitter, or another device for sending information wirelessly, such as a laser which emits light through the skin, then there is no need to include the leads.
Alternatively, the microelectronic circuitry may not be present, in which case the lead will directly electrically connect the electrodes with external electrical components.
Optionally, one or more internal coatings may be present. The internal coating may function to regulate diffusion. Examples of internal coatings include cellulose acetate, polyurethanes, polyallylamines (PAL), polyaziridine (PAZ), and silicon-containing polymers. Some specific examples are described in PCT Publications WO
98/17995, WO
98/13685 and WO 96/06947, and in U.S. Patent Nos. 4,650,547 and 5,165,407.
Optionally, one or more external coatings may be present. The implantable analyte sensors of the present invention are intended to be used in vivo, preferably subcutane-ously in mammals, such as humans, dogs or mice. The external coatings function to improve the biocompatibility of the implantable analyte sensor. Examples of external coatings include nafion, polyurethanes, polytetrafluoroethylenes (PTFE), poly (ethylene oxide) (PEO), and 2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate (MPC) membranes. Some specific examples are described in PCT Publication WO 96/06947, and in "Medical Progress through Technology", Nishida et al. 21:
(1995).
The electrodes may be coated with a reagent. The reagent is optional, and may be used to provide electrochemical probes for specific analytes. The reagent may be as simple as a single enzyme, such as glucose oxidase or glucose hydrogenase for the detection of glu-cose. The enzyme may be immobilized or "wired" as described in PCT Publication WO
96/06947. The reagents may optionally also include a mediator, to enhance sensitivity of the sensor. The starting reagents are the reactants or components of the reagent, and are often compounded together in liquid form before application to the electrodes.
The liquid may then evaporate, leaving the reagent in solid form. The choice of specific rea-gent depends on the specific analyte or analytes to be measured, and are well known to those of ordinary skill in the art. For example, a reagent for measurement of glucose can contain 62.2 mg polyethylene oxide (mean molecular weight of 100-900 kilodaltons), 3.3 mg NATROSOL 250 M, 41.5 mg AVICEL RC-591 F, 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 to 9,000 units of enzyme activity per gram of reagent. The enzyme is prepared as an enzyme solution from 12.5 mg coenzyme PQQ and 1.21 million units of the apoen-zyme of quinoprotein glucose dehydrogenase, forming a solution of quinoprotein glu-cose dehydrogenase. This reagent is described in WO 99/30152, pages 7-10, hereby in-corporated by referece.
Other non-limiting examples of enzymes and optional mediators that may be used in measuring particular analytes in the present invention are listed below in Table 1.
Analyte Enzymes Mediator Additional Mediator (Oxidized Form) Glucose Glucose Dehydro-Ferricyanide genase and Diapho-rase Glucose Glucose-Dehydroge-Ferricyanide vase Cholesterol (Quinoprotein) Ferricyanide2,6-Dimethyl-1,4-Cholesterol Benzoquinone Esterase and Cholesterol 2,5-Dichloro-1,4-Oxidase Benzoquinone or Phenazine Ethosulfate HDL CholesterolCholesterol Ferricyanide2,6-Dimethyl-1,4-Esterase and Cholesterol Benzoquinone Oxidase 2,5-Diehloro-1,4-Benzoquinone or Phenazine Ethosulfate TriglyceridesLipoprotein FerricyanidePhenazine Methosul-Lipase, or C~lycerol Kinase,Phenazine fate and Glycerol-3-Phos-Ethosulfate phate Oxidase Lactate Lactate OxidaseFerricyanide2,6-Dichloro-1,4-Benzoquinone Lactate Lactate Dehydroge-Ferricyanide nase and DiaphorasePhenazine Ethosulfate, or Phenazine Methosulfate Lactate Diaphorase FerricyanidePhenazine Ethosul-Dehydrogenase fate, or Phenazine Methosulfate uvate Alco amore Bilirubin Bilirubin Oxidase 1-Methoxv-Phenazine Methosulfate Uric Acid Uricase Ferricyanide In some of the examples shown in Table 1, at least one additional enzyme is used as a reaction catalyst. Also, some of the examples shown in Table 1 may utilize an additional mediator, which facilitates electron transfer to the oxidized form of the mediator. The additional mediator may be provided to the reagent in lesser amount than the oxidized form of the mediator. While the above assays are described, it is appreciated that a vari-ety of electrochemical assays may be conducted in accordance with this disclosure.
Formation of membrane The following describes how to make a membrane for use in the present invention, based on the description from the dissertation of Derek James Hansford, submitted in partial satisfaction of the requirements for the degree of Doctor of Philosophy in Engi-neering-Materials Science and Mineral Engineering in the Graduate Division of the University of California, Berkeley, submitted in the spring of 1999.
Other membranes, made from other material, may also be used. This specific method relies upon a buried nitride etch stop layer.
The buried nitride etch stop layer acts as an etchant stop during the formation of nano-meter scale pores. The buried nitride etch stop layer facilitates three-dimensional control of the pore structure, and facilitates the formation of pores less than 50 nanometers in diameter. Moreover, these pores can be uniformly formed across the entire wafer.
Preferably, the first step in the fabrication protocol is to etch a support ridge structure into a substrate. The ridges provide mechanical rigidity to the subsequently formed membrane structure.
A low stress silicon nitride (LSN or nitride), which operates as an etch stop layer, is then deposited on the substrate using low pressure chemical vapor depositions (LPCVD). In one embodiment, 0.4 ~m of nitride was used. The resultant structure is shown in Figure 1. Figure 1 illustrates a substrate 20 with a nitride etch stop layer 22 formed thereon.
The base structural layer (base layer) of the membrane is deposited on top of the stop layer 22. Since the etch stop layer 22 is thin, the structural layer is deposited down into the support ridges formed in the substrate 20. In one embodiment, 5 ~m of polysilicon is used as the base layer. Figure 2 illustrates the base layer 24 positioned on the etch stop layer 22. Low stress silicon nitride may also be used as the base layer, in which case it operates 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 pores. Masks, such as those used in traditional semiconductor processing, may be used 5 to define the pores. For example, the holes may be etched through the polysilicon by chlorine plasma, with a thermally grown oxide layer used as a mask. In this step, it is important to make sure the etching goes completely through the base layer 24, so a 10-15% overetch is preferably used. It is useful to note that the buried nitride etch stop 22 acts as an etch stop for the plasma etching of a silicon base layer 24.
Otherwise, if the 10 plasma punched through the nitride, tighter control of the etch step would have to be exercised to prevent the complete removal of the nitride under the plug layer (to prevent removal in the final 1~OH etch). Figure 3 illustrates the result of this processing. In particular, the figure illustrates holes 26 formed in the base layer 24, but terminating in the nitride etch stop layer 22.
Pore sacrificial oxide is subsequently grown on the base layer 24. Figure 4 illustrates a sacrificial oxide 28 positioned on the base layer 24.
The sacrificial oxide thickness determines the pore size in the final membrane, so con-trol of this step is critical to reproducible membranes. This is accomplished by the thermal oxidation of the base layer 24 (e.g., a growth temperature of between 850-950°C
for approximately one hour with a ten minute anneal). Naturally, many techniques may be used to form a controlled thickness sacrificial layer. For example, a thermally evapo-rated tungsten film may be used as a sacrificial layer for polymer membranes and selec-tively removed with hydrogen peroxide. The basic requirement of the sacrificial layer is the ability to control the thickness with high precision across the entire wafer. Thermal oxidation of both polysilicon and nitride allows the control of the sacrificial layer thick-ness of less than 5% across the entire wafer. Limitations on this control arise from local inhomogeneities in the base layer, such as the initial thickness of the native oxide (espe-cially for polysilicon) the grain size or the density, and the impurity concentrations.
To mechanically connect the base layer 24 with the plug layer (necessary to maintain the pore spacing between layers), anchor points were defined in the sacrificial oxide layer 26.
In the present design, this is accomplished by using the same mask shifted from the pore holes by 1 ~tm diagonally. This produced anchors in one or two corners of each pore hole, which provides the desired mechanical connection between the structural layers while opening the pore area as much as possible. Figure 5 illustrates anchors 30 formed via this process.
A plug structural layer is subsequently deposited to fill in the holes 26.
This step has been implemented by depositing 1.5 ~m of polysilicon. The resultant plug layer 32 is shown in Figure 6.
To open the pores at the surface, the plug layer 32 is planarized down to the base layer, leaving the final structure with the plug layer only in the pore hole openings, as shown in Figure 7.
The method of planariration depends on the material used as the plug material.
For the hard micro-fabrication materials (polysilicon and nitride), chemical mechanical polish-ing was used for planarization. The other materials studied were roughly planarized using a plasma etch, with a quick wet chemical smoothing. This technique has the advantage that, assuming it is not etched by the plasma used, the base layer is not affected, but has the disadvantage of the need for controlled etch timing to avoid com-pletely etching the plugs themselves.
At this point, the membrane is ready for release, so a protective layer 34 is deposited on the wafer (completely covering both sides of the wafer). The requirements of the pro-tective layer 34 are that it be impervious to the silicon etch (KOH for these studies) and that it be removed without removing the plug 32 or base 24 structural layers.
For poly-silicon and nitride structural layers, a thin nitride layer is used as the protective layer (nitride is not etched at all by KOH and dissolves slowly in HF). For polymeric struc-tural materials, silicon is used as a protective layer, due to the processing temperature necessary for nitride deposition (835° C).
The backside etch windows were etched in the protective layer, exposing the silicon in desired areas, and then the entire structure was placed in an 80°C KOH
bath until the silicon wafer substrate 20 is etched up to the membrane base layer 24 (as evidenced by the smooth buried etch stop layer). Figure 8 illustrates the resultant aperture 36 formed in the substrate 20.
At this point, the buried nitride layer 22, the sacrificial oxide layer 34, and plug layer 32 are removed by etching in HF or SF6/oxygen plasma. The resultant membrane 4 with nanometer scale pores is shown in Figure 9.
Characterization of membranes The purpose of the membranes is to allow the analyte of interest (such as glucose) to diffuse through the membrane, while excluding large molecules (such as proteins).
Therefore, two important characteristics of the membranes are glucose diffusion and albumin diffusion. All tests are carried out at room temperature (25°C).
The following is a glucose diffusion test:
Diffusion of glucose is measured using a mini diffusion chamber constructed around the membranes. The diffusion chamber, fabricated out of acrylic, consists of two compart-ments A and B with fixed volumes of 2 ml, separated by the desired membrane, sealed with o-rings, and screwed together.
Glucose is measured on either side of the membrane using the diffusion chamber by means of a quantitative enzymatic assay (TRINDERT~~, SIGMA) and colorometric reading via a spectrophotometer. Starting glucose concentrations for all tests were 6,666 mg/dl and 0.0 mg/dl in chambers A and B, respectively. Samples of 0.1 ml are taken from the diffusion chamber and 10 ftl of that are added to 3 ml of glucose reagent in a cuvette, and mixed gently by inversion. Each tube is incubated for 18 minutes at room temperature and then readings are taken at a wavelength of 505 nm. The reagent is linear up to 750 mg/dl. The diffusion chamber itself is attached to a motor for stirring in order to minimize boundary layer effects (diffusion resistance at the liquid/membrane interface). In order to ensure wetting of the pores, the receptor cell is first filled with phosphate buffer saline (PBS) for fifteen minutes before the filling of the donor cell. The donor cell is filled with solutions of glucose in PBS in varying concentrations.
The following is an albumin diffusion test:
Albumin is also measured on either side of the membrane using the same diffusion chamber as in the glucose diffusion test. Albumin diffusion and/or exclusion is first measured and quantified using Albumin BCP (bromocresol purple, SIGMA).
Starting albumin concentrations for all tests are 4 g/dl and 0.0 mg/dI in chambers A
and B, respectively. A sample of 0.1 ml is taken at time zero and at the end of the diffusion period (time = 330 minutes). An aliquot of 300 ~l is then added to 3 ml of the reagent and absorbence is read at 600 nm. Reagent plus deionized water is used as the blank. The BCP assay is linear up to 6g/dl but is not accurate below 1 g/dl. For the small con-centration of albumin that might be present in chamber A, the presence of any protein in chamber B is measured using the Bradford Method (MICRO PROTEIN KIT, SIGMA). This method quantitates the binding of Coomassie brilliant blue to an unknown protein and compares this binding to that of different amounts of a standard protein. Albumin is used as a standard protein. This method quantifies 1 to 100 micro-grams protein using a standard curve, with sensitivity down to 10 mg/dl or 0.1 g/dl pro-tein. The absorbance is measured at 595 nm.
Analysis of membranes Diffusion of glucose was measured for three types of membranes: silicon micromachi-ned membranes (average pore size = 0.0245 microns), WHATMAN ANODISC mem-branes (average pore size = .02 microns), and MF-MILLIPORE mixed cellulose acetate and nitrate membrane (average pore size = 0.025 microns).
The results from the albumin test are shown in the table below.
WHATMAN MILLIPORE silicon (micromachined) time albumin conc. albumin conc. albumin conc.
(g/dL) (g/dL) (g/dL) 420 MinØ25'c'0.05 0.2b'0.01 0b'0.001 The presence of albumin does not seem to impede passage of glucose through the mem-branes, nor slow down glucose transport. No detectable amounts of albumin diffuse through the micromachined membrane. The same membrane, however, shows glucose diffusion. The micromachined membranes are able to achieve complete exclusion of albumin (to within the limits of detection), while allowing glucose diffusion.
Comparing diffusion rates with that of commercially available membranes, the micromachined membranes have glucose diffusion properties comparable to MIILLIPORE and alumina WHATMAN membranes with similar pore sizes.
The passage of albumin through the micromachined membrane is measured by looking at the change of albumin concentration in chamber A and chamber B over time.
Using the BCP assay, there are no detectable traces of albumin in chamber B.
However, the amount of albumin in chamber B may have been below the limits of detectability of this assay system. Therefore, the Bradford Method was also employed. Using this microas-say, again no detectable amounts of albumin were found in chamber B for the micro-machined membrane, but small amounts of protein were found in chamber B using both the MILLIPORE and WHATMAN membranes. The amounts of albumin detected after 420 minutes in chamber B were approximately 0.25 g/dI and 0.20 g/dI
albumin for the MILLIPORE and WHATh-lAN membranes, respectively.
Glucose does diffuse through micromachined membranes at a rate comparable to com-mercially available membranes. At the same time, albumin is excluded from passage. In mixed solutions of glucose and albumin, only glucose diffuses through the micromachi-ned membranes.
The present invention relates to implantable analyte sensors.
Several implantable glucose sensors have been developed. Examples include those described in U.S. Patent numbers 5,387,327; 5,411,647; and 5,476,776; as well as those described in PCT International Publication numbers WO 91/15993; WO 94/20602;
WO
96/06947; and WO 97/ 19344. The implantable glucose sensors usually include a polymer substrate, with metal electrodes printed on the surface of the substrate. A
biocompatible membrane covers the electrodes, allowing glucose to reach the electrodes, while excluding other molecules, such as proteins. Electrochemistry, often with the aid of enzymes at the electrodes, is used to determine the quantity of glucose present. The glu-cose sensor is implanted into a patient, and the electrodes may be attached via wires that pass out of the patient's body to external circuitry that controls the electrodes, measures and reports the glucose concentration. Alternatively, all or part of this external circuitry may be miniaturized and included in the implantable glucose sensor. A
transmitter, such as that described in WO 97/ 19344, may even be included in the implantable glucose sensor, completely eliminating the need for leads that pass out of the patient.
A problem associated with an amperometric glucose sensor is unstable signals.
This may result from degradation of the enzyme from interaction with protein, leakage of the enzyme, and/or fouling of the electrode. The usual way to overcome this is to use the above described biocompatible membrane, or a coating. However, several problems are also associated with these membranes. For example, Nafion-based biosensor membranes exhibit cracking, flaking, protein adhesion, and calcium deposits.
Mineralization of polymer-based membranes occurs in the biological environment, resulting in cracking and changes in permeabilit<~. The tortuous porosity associated with polymer membranes has also been shown to be important in membrane stability and mineralization in vivo. Biological components, which enter pores or voids in the mate rial, cause metabolic shadows, which are loci for ion and calcium accumulation. This situation, coupled with the fact that mineral deposits have been known to propagate surface fractures in polymeric membranes, presents a potentially serious problem for implantable glucose sensors.
In polymer membranes the pore size distribution usually follows some kind of prob-ability distribution (e.g. gaussian), which leaves a finite probability for large proteins to eventually transfer through the membrane. Drift may be caused by this leakage or inadequate diffusion properties, and events at the body-sensor interface such as bio-fouling and protein adsorption, encapsulation with fibrotic tissue, and degradation of the device material over time.
Currently, membranes with nominal pore sizes as small as 20 nm are available.
Even so, the filtration at these dimensions is far from absolute. The most common filters are polymeric membranes for med from a solvent-casting process, which result in a pore size distribution with variations as large as 30%. The use of ion-track etching to form mem-branes (e.g. MILLPORE ISO PORE) produces a much tighter pore size distribution (~10%). However, these membranes have low porosities (<109 pores/cm2), limited pore sizes, and the pores are randomly distributed across the surface. Porous alumina (e.g.
WHATMAN) has also been used to achieve uniform pores. Although the aluminas typically have higher pore densities (>10~°/cm'), only certain pore sizes (typically greater than 20 nanometers ) can be achieved and the pore configurations and arrangements are difficult to control.
BRIEF SUMMARY
In one aspect, the present invention is an implantable analyte sensor, comprising a sub-strate, electrodes on the substrate, and a membrane on the electrodes. The membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffu-sion test result of at most 0.1 g/dl in 420 min. and can comprise elemental silicon.
In another aspect, the present invention relates to a method of making an implantable analyte sensor, comprising covering electrodes with a membrane. The electrodes are on a substrate, and the membrane has a glucose diffusion test result of at least 1 mg/dl in WO 01/68901 CA 02406878 2002-09-13 pCT~P01/03027 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
The mem-brane can comprise elemental silicon.
BRIEF DESCRIPTION OF THE DRAWINGS
The following drawings form part of the present specification and are included to further demonstrate 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 presented herein:
Figured-9 illustrate the process of making a membrane for use in an embodiment of the present mvenrion;
Figure 10 shows a cut-away view of an implantable analyte sensor;
Figure 11 shows an exploded view of an implantable analyte sensor; and Figure 12 shows a cut-away view of an implantable analyte sensor.
DETAILED DESCRIPTION
Figure 10 shows a cut away view of an embodiment of the present invention. In the figure, an implantable analWe sensor 2 includes a substrate 6 on which are electrodes 8 and 8. The electrode, are cowered with a membrane 4. Leads 12 and 12 allow for elec-trically connecting the implantable analyte sensor to external circuitry (not shown). The implantable analyte sensor also includes an external coating 16 and an internal coating 14.
Figure 11 shows an exploded view of an embodiment of the present invention.
The internal and external coatings are not included in the figure for clarity. As shown in the figure, the implantable analyte sensor 2 includes the electrodes 8 and 8 on the substrate 6 surface, which are electrically connected with microelectronic circuitry 10.
The micro-electronic circuitry is electrically connected to leads 12 and 12, which allow for electri-tally connecting the implantable analyte sensor to external circuitry (not shown). The electrodes are covered with the membrane 4.
Figure 12 shows a cut away view of an embodiment of the present invention similar to that shown in Figure 10, except for the presence of a third electrode 8 and a third lead 12. Although so illustrated, the number of electrodes may be different from the number of leads.
The membrane is composed of a hard material that has been micromachined.
Prefera-bly, the membrane comprises elemental silicon, but other hard, biocompatable materials that can be micromachined are possible, such as metals (for example titanium), ceram-ics (for example, silica or silicon nitride), and polymers (such as polytetrafluoroethylene, polymethylmethacrylate, polystyrenes and silicones). Micromachining is a process that includes photolithography, such as that used in the semiconductor industry, to remove material from, or add material too, a substrate. These techniques are well known, and are described, in Encyclopedia of Chemical Technology, Kirk-Othmer, Volume 14, pp.
677-709 ( 1995); Semiconductor Device Fundamentals, Robert F. Pierret, Addison-~~~esley, 1996; and Mlicrochip Fabrication 3rd. edition, Peter Van Zant, McGraw-Hill, 1997. A detailed fabrication method for a membrane comprising elemental silicon is described in the dissertation of Derek James Hansford, submitted in partial satisfaction of the requirements for the degree of Doctor of Philosophy in Engineering-Materials Science and hfineral Engineering in the Graduate Division of the University of Califor-nia, Berkeley, submitted in the spring of 1999.
A special property of the membrane is a defined pore size, which has a small size distri-bution compared to the size distribution of standard membranes. Due to tight toler-ances in the manufacturing process, the pore size can be controlled at precise diameters, for example 1 to 50 nm, or ~ to 20 nm, or even 5 to 15 nm (such as 12 nm, 18 nm or even 25 nm), with a variation of +/- 0.01-20%, +/-0.1-10% or even +/-1-5%.
Therefore molecules above this size can be excluded with high certainty, since the size distribution has the shape of a top hat, rather than a bell curve, and hence pore sizes above, for example 12 nm, 18 nm, 25 nm or 50 nm are not present. These membranes can exclude interfering molecules, such as proteins, which could otherwise cause major drift prob-lems of the sensor, when the sensor is implanted in vivo. Signal drift is a change in the magnitude of the signal from a sensor which is unrelated to changes in analyte concen-tration. The amount of signal drift is based on the magnitude of the signal prior to the drifting. Preferably, the implantable analyte sensors of the present invention exhibit a signal drift of less than 20% per day in vivo, more preferably less than 10%
per day in vivo, most preferably less than 5% per day in vivo.
Membranes for use in the present invention may be characterized by a glucose diffusion test and an albumin diffusion test. These tests are described below.
Preferably, the membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., more pref erably at least 10 mg/dl in 330 min., even more preferably at least 30 mg/dl in 330 min., 5 and most preferably at least 60 mg/dl in 330 min. Preferably, the membrane has an albumin diffusion test result of at most 0.1 g/dl in 420 min., more preferably at most 0.05 g/dl in 420 min., even more preferably at most 0.01 g/dl in 420 min., and most pref erably at most 0.001 g/dl in 420 min.
The manufacturing process of the membranes may allow a simple and economical pro-duction of small, implantable analyte sensors. For example, the membranes can be first manufactured, and then on a substrate, the electrodes for the sensor and the electrical connectors can be formed. Preferably, the substrate is silicon, but other materials are possible, such as ceramics, or polymers. If desired, electronic components, for example, amplifiers, filters, transmitters and/or signal preconditioning components, can easily be incorporated in this layer. In particular, if the substrate comprises elemental silicon, well known integrated circuit technology may be used to place all the circuitry in miniaturized form on a single chip.
There are two possible approaches to attach the substrate and the membrane, when a reagent is included in the sensor:
1. The substrate and the membrane are thermally bonded before the reagent is depos-ited on the electrodes. In this case, an opening, preferably in the membrane is pro-vided (since this may be manufactured with a micromachining process, an opening is easily generated during one of the processing steps). In the case where multiple membranes are formed as a single piece, and or multiple substrates are formed as a single piece, after thermal bonding, a further etching step may be used to separate the individual membrane/substrate units. The reagent is deposited through the in-dividual openings and the openings are sealed using, for example a polymer sealant.
The individual sensors are then separated, incorporated into a flexible, inner coat-ing, for example silicone rubber, and individually coated with an outer coating, such as a biocompatible layer.
2. The reagent is deposited on the electrodes before the membrane and substrate are attached. In this case, thermal bonding is not possible, since the enzyme in the rea-gent would be destroyed. The individual membranes and substrates are first sepa-rated and the individual sensors are assembled by bonding one membrane with one substrate using a suitable bonding agent, for example, cyanoacrylate. As a final step, the individual sensors are incorporated into a flexible, inner coating, for example silicone rubber, and individually coated with an outer coating, such as a biocom-patible layer. The sensor can be inserted into the skin using a needle applicator. The control unit typically remains outside the body and can be connected to the sensor element through electrical wires (leads).
The electrodes are formed on the surface of the substrate. They may be formed by well known semiconductor processing techniques, from conductive materials, such as pure metals or alloys, or other materials which are metallic conductors. Examples include aluminum, carbon (such as graphite), cobalt, copper, gallium, gold, indium, iridium, iron, lead, magnesium, mercury (as an amalgam), nickel, niobium, osmium, palladium, platinum, rhenium, rhodium, selenium, silicon (such as highly doped polycrystalline silicon), silver, tantalum, tin, titanium, tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, and alloys or metallic compounds of these elements.
Preferably, the electrodes include gold, platinum, palladium, iridium, or alloys of these metals, since such noble metals and their alloys are unreactive in biological systems. The electrodes may be any thickness, but preferably are 10 nm to 1 mm, more preferably, 20 nm to 100 ~tm, or even 25 nm to 1 ym.
At least two electrodes must be present. The number of electrodes may be 2-1000, or 3-200, or even 3-99. Individual electrode sets (2 or 3 electrodes) may be separated into individual chambers, each covered with the membrane. Furthermore, individual elec-trode sets (2 or 3 electrodes] may each have a different reagent, allowing for an im-plantable analyte sensor that can measure at least two, such as 3-100, or 4-20, different analytes.
The remaining individual part of the implantable analyte sensors are well known to those of ordinary skill in the art, and are described, for example, in U.S.
Patent numbers 5,387,327; 5,411,647; and 5,476,776; as well as in PCT International Publication num-bers WO 91/15993; VVO 94/20602; VVO 96/06947; and WO 97/19344.
Although illustrated with both leads and microelectronic circuitry, these components are optional. The microelectronic circuitry may include some or all of the electrical components normally external to the implantable analyte sensor, such as a microproc-essor, an amplifier, or a power supply. If the microelectronic circuitry also includes a transmitter, or another device for sending information wirelessly, such as a laser which emits light through the skin, then there is no need to include the leads.
Alternatively, the microelectronic circuitry may not be present, in which case the lead will directly electrically connect the electrodes with external electrical components.
Optionally, one or more internal coatings may be present. The internal coating may function to regulate diffusion. Examples of internal coatings include cellulose acetate, polyurethanes, polyallylamines (PAL), polyaziridine (PAZ), and silicon-containing polymers. Some specific examples are described in PCT Publications WO
98/17995, WO
98/13685 and WO 96/06947, and in U.S. Patent Nos. 4,650,547 and 5,165,407.
Optionally, one or more external coatings may be present. The implantable analyte sensors of the present invention are intended to be used in vivo, preferably subcutane-ously in mammals, such as humans, dogs or mice. The external coatings function to improve the biocompatibility of the implantable analyte sensor. Examples of external coatings include nafion, polyurethanes, polytetrafluoroethylenes (PTFE), poly (ethylene oxide) (PEO), and 2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate (MPC) membranes. Some specific examples are described in PCT Publication WO 96/06947, and in "Medical Progress through Technology", Nishida et al. 21:
(1995).
The electrodes may be coated with a reagent. The reagent is optional, and may be used to provide electrochemical probes for specific analytes. The reagent may be as simple as a single enzyme, such as glucose oxidase or glucose hydrogenase for the detection of glu-cose. The enzyme may be immobilized or "wired" as described in PCT Publication WO
96/06947. The reagents may optionally also include a mediator, to enhance sensitivity of the sensor. The starting reagents are the reactants or components of the reagent, and are often compounded together in liquid form before application to the electrodes.
The liquid may then evaporate, leaving the reagent in solid form. The choice of specific rea-gent depends on the specific analyte or analytes to be measured, and are well known to those of ordinary skill in the art. For example, a reagent for measurement of glucose can contain 62.2 mg polyethylene oxide (mean molecular weight of 100-900 kilodaltons), 3.3 mg NATROSOL 250 M, 41.5 mg AVICEL RC-591 F, 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 to 9,000 units of enzyme activity per gram of reagent. The enzyme is prepared as an enzyme solution from 12.5 mg coenzyme PQQ and 1.21 million units of the apoen-zyme of quinoprotein glucose dehydrogenase, forming a solution of quinoprotein glu-cose dehydrogenase. This reagent is described in WO 99/30152, pages 7-10, hereby in-corporated by referece.
Other non-limiting examples of enzymes and optional mediators that may be used in measuring particular analytes in the present invention are listed below in Table 1.
Analyte Enzymes Mediator Additional Mediator (Oxidized Form) Glucose Glucose Dehydro-Ferricyanide genase and Diapho-rase Glucose Glucose-Dehydroge-Ferricyanide vase Cholesterol (Quinoprotein) Ferricyanide2,6-Dimethyl-1,4-Cholesterol Benzoquinone Esterase and Cholesterol 2,5-Dichloro-1,4-Oxidase Benzoquinone or Phenazine Ethosulfate HDL CholesterolCholesterol Ferricyanide2,6-Dimethyl-1,4-Esterase and Cholesterol Benzoquinone Oxidase 2,5-Diehloro-1,4-Benzoquinone or Phenazine Ethosulfate TriglyceridesLipoprotein FerricyanidePhenazine Methosul-Lipase, or C~lycerol Kinase,Phenazine fate and Glycerol-3-Phos-Ethosulfate phate Oxidase Lactate Lactate OxidaseFerricyanide2,6-Dichloro-1,4-Benzoquinone Lactate Lactate Dehydroge-Ferricyanide nase and DiaphorasePhenazine Ethosulfate, or Phenazine Methosulfate Lactate Diaphorase FerricyanidePhenazine Ethosul-Dehydrogenase fate, or Phenazine Methosulfate uvate Alco amore Bilirubin Bilirubin Oxidase 1-Methoxv-Phenazine Methosulfate Uric Acid Uricase Ferricyanide In some of the examples shown in Table 1, at least one additional enzyme is used as a reaction catalyst. Also, some of the examples shown in Table 1 may utilize an additional mediator, which facilitates electron transfer to the oxidized form of the mediator. The additional mediator may be provided to the reagent in lesser amount than the oxidized form of the mediator. While the above assays are described, it is appreciated that a vari-ety of electrochemical assays may be conducted in accordance with this disclosure.
Formation of membrane The following describes how to make a membrane for use in the present invention, based on the description from the dissertation of Derek James Hansford, submitted in partial satisfaction of the requirements for the degree of Doctor of Philosophy in Engi-neering-Materials Science and Mineral Engineering in the Graduate Division of the University of California, Berkeley, submitted in the spring of 1999.
Other membranes, made from other material, may also be used. This specific method relies upon a buried nitride etch stop layer.
The buried nitride etch stop layer acts as an etchant stop during the formation of nano-meter scale pores. The buried nitride etch stop layer facilitates three-dimensional control of the pore structure, and facilitates the formation of pores less than 50 nanometers in diameter. Moreover, these pores can be uniformly formed across the entire wafer.
Preferably, the first step in the fabrication protocol is to etch a support ridge structure into a substrate. The ridges provide mechanical rigidity to the subsequently formed membrane structure.
A low stress silicon nitride (LSN or nitride), which operates as an etch stop layer, is then deposited on the substrate using low pressure chemical vapor depositions (LPCVD). In one embodiment, 0.4 ~m of nitride was used. The resultant structure is shown in Figure 1. Figure 1 illustrates a substrate 20 with a nitride etch stop layer 22 formed thereon.
The base structural layer (base layer) of the membrane is deposited on top of the stop layer 22. Since the etch stop layer 22 is thin, the structural layer is deposited down into the support ridges formed in the substrate 20. In one embodiment, 5 ~m of polysilicon is used as the base layer. Figure 2 illustrates the base layer 24 positioned on the etch stop layer 22. Low stress silicon nitride may also be used as the base layer, in which case it operates 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 pores. Masks, such as those used in traditional semiconductor processing, may be used 5 to define the pores. For example, the holes may be etched through the polysilicon by chlorine plasma, with a thermally grown oxide layer used as a mask. In this step, it is important to make sure the etching goes completely through the base layer 24, so a 10-15% overetch is preferably used. It is useful to note that the buried nitride etch stop 22 acts as an etch stop for the plasma etching of a silicon base layer 24.
Otherwise, if the 10 plasma punched through the nitride, tighter control of the etch step would have to be exercised to prevent the complete removal of the nitride under the plug layer (to prevent removal in the final 1~OH etch). Figure 3 illustrates the result of this processing. In particular, the figure illustrates holes 26 formed in the base layer 24, but terminating in the nitride etch stop layer 22.
Pore sacrificial oxide is subsequently grown on the base layer 24. Figure 4 illustrates a sacrificial oxide 28 positioned on the base layer 24.
The sacrificial oxide thickness determines the pore size in the final membrane, so con-trol of this step is critical to reproducible membranes. This is accomplished by the thermal oxidation of the base layer 24 (e.g., a growth temperature of between 850-950°C
for approximately one hour with a ten minute anneal). Naturally, many techniques may be used to form a controlled thickness sacrificial layer. For example, a thermally evapo-rated tungsten film may be used as a sacrificial layer for polymer membranes and selec-tively removed with hydrogen peroxide. The basic requirement of the sacrificial layer is the ability to control the thickness with high precision across the entire wafer. Thermal oxidation of both polysilicon and nitride allows the control of the sacrificial layer thick-ness of less than 5% across the entire wafer. Limitations on this control arise from local inhomogeneities in the base layer, such as the initial thickness of the native oxide (espe-cially for polysilicon) the grain size or the density, and the impurity concentrations.
To mechanically connect the base layer 24 with the plug layer (necessary to maintain the pore spacing between layers), anchor points were defined in the sacrificial oxide layer 26.
In the present design, this is accomplished by using the same mask shifted from the pore holes by 1 ~tm diagonally. This produced anchors in one or two corners of each pore hole, which provides the desired mechanical connection between the structural layers while opening the pore area as much as possible. Figure 5 illustrates anchors 30 formed via this process.
A plug structural layer is subsequently deposited to fill in the holes 26.
This step has been implemented by depositing 1.5 ~m of polysilicon. The resultant plug layer 32 is shown in Figure 6.
To open the pores at the surface, the plug layer 32 is planarized down to the base layer, leaving the final structure with the plug layer only in the pore hole openings, as shown in Figure 7.
The method of planariration depends on the material used as the plug material.
For the hard micro-fabrication materials (polysilicon and nitride), chemical mechanical polish-ing was used for planarization. The other materials studied were roughly planarized using a plasma etch, with a quick wet chemical smoothing. This technique has the advantage that, assuming it is not etched by the plasma used, the base layer is not affected, but has the disadvantage of the need for controlled etch timing to avoid com-pletely etching the plugs themselves.
At this point, the membrane is ready for release, so a protective layer 34 is deposited on the wafer (completely covering both sides of the wafer). The requirements of the pro-tective layer 34 are that it be impervious to the silicon etch (KOH for these studies) and that it be removed without removing the plug 32 or base 24 structural layers.
For poly-silicon and nitride structural layers, a thin nitride layer is used as the protective layer (nitride is not etched at all by KOH and dissolves slowly in HF). For polymeric struc-tural materials, silicon is used as a protective layer, due to the processing temperature necessary for nitride deposition (835° C).
The backside etch windows were etched in the protective layer, exposing the silicon in desired areas, and then the entire structure was placed in an 80°C KOH
bath until the silicon wafer substrate 20 is etched up to the membrane base layer 24 (as evidenced by the smooth buried etch stop layer). Figure 8 illustrates the resultant aperture 36 formed in the substrate 20.
At this point, the buried nitride layer 22, the sacrificial oxide layer 34, and plug layer 32 are removed by etching in HF or SF6/oxygen plasma. The resultant membrane 4 with nanometer scale pores is shown in Figure 9.
Characterization of membranes The purpose of the membranes is to allow the analyte of interest (such as glucose) to diffuse through the membrane, while excluding large molecules (such as proteins).
Therefore, two important characteristics of the membranes are glucose diffusion and albumin diffusion. All tests are carried out at room temperature (25°C).
The following is a glucose diffusion test:
Diffusion of glucose is measured using a mini diffusion chamber constructed around the membranes. The diffusion chamber, fabricated out of acrylic, consists of two compart-ments A and B with fixed volumes of 2 ml, separated by the desired membrane, sealed with o-rings, and screwed together.
Glucose is measured on either side of the membrane using the diffusion chamber by means of a quantitative enzymatic assay (TRINDERT~~, SIGMA) and colorometric reading via a spectrophotometer. Starting glucose concentrations for all tests were 6,666 mg/dl and 0.0 mg/dl in chambers A and B, respectively. Samples of 0.1 ml are taken from the diffusion chamber and 10 ftl of that are added to 3 ml of glucose reagent in a cuvette, and mixed gently by inversion. Each tube is incubated for 18 minutes at room temperature and then readings are taken at a wavelength of 505 nm. The reagent is linear up to 750 mg/dl. The diffusion chamber itself is attached to a motor for stirring in order to minimize boundary layer effects (diffusion resistance at the liquid/membrane interface). In order to ensure wetting of the pores, the receptor cell is first filled with phosphate buffer saline (PBS) for fifteen minutes before the filling of the donor cell. The donor cell is filled with solutions of glucose in PBS in varying concentrations.
The following is an albumin diffusion test:
Albumin is also measured on either side of the membrane using the same diffusion chamber as in the glucose diffusion test. Albumin diffusion and/or exclusion is first measured and quantified using Albumin BCP (bromocresol purple, SIGMA).
Starting albumin concentrations for all tests are 4 g/dl and 0.0 mg/dI in chambers A
and B, respectively. A sample of 0.1 ml is taken at time zero and at the end of the diffusion period (time = 330 minutes). An aliquot of 300 ~l is then added to 3 ml of the reagent and absorbence is read at 600 nm. Reagent plus deionized water is used as the blank. The BCP assay is linear up to 6g/dl but is not accurate below 1 g/dl. For the small con-centration of albumin that might be present in chamber A, the presence of any protein in chamber B is measured using the Bradford Method (MICRO PROTEIN KIT, SIGMA). This method quantitates the binding of Coomassie brilliant blue to an unknown protein and compares this binding to that of different amounts of a standard protein. Albumin is used as a standard protein. This method quantifies 1 to 100 micro-grams protein using a standard curve, with sensitivity down to 10 mg/dl or 0.1 g/dl pro-tein. The absorbance is measured at 595 nm.
Analysis of membranes Diffusion of glucose was measured for three types of membranes: silicon micromachi-ned membranes (average pore size = 0.0245 microns), WHATMAN ANODISC mem-branes (average pore size = .02 microns), and MF-MILLIPORE mixed cellulose acetate and nitrate membrane (average pore size = 0.025 microns).
The results from the albumin test are shown in the table below.
WHATMAN MILLIPORE silicon (micromachined) time albumin conc. albumin conc. albumin conc.
(g/dL) (g/dL) (g/dL) 420 MinØ25'c'0.05 0.2b'0.01 0b'0.001 The presence of albumin does not seem to impede passage of glucose through the mem-branes, nor slow down glucose transport. No detectable amounts of albumin diffuse through the micromachined membrane. The same membrane, however, shows glucose diffusion. The micromachined membranes are able to achieve complete exclusion of albumin (to within the limits of detection), while allowing glucose diffusion.
Comparing diffusion rates with that of commercially available membranes, the micromachined membranes have glucose diffusion properties comparable to MIILLIPORE and alumina WHATMAN membranes with similar pore sizes.
The passage of albumin through the micromachined membrane is measured by looking at the change of albumin concentration in chamber A and chamber B over time.
Using the BCP assay, there are no detectable traces of albumin in chamber B.
However, the amount of albumin in chamber B may have been below the limits of detectability of this assay system. Therefore, the Bradford Method was also employed. Using this microas-say, again no detectable amounts of albumin were found in chamber B for the micro-machined membrane, but small amounts of protein were found in chamber B using both the MILLIPORE and WHATMAN membranes. The amounts of albumin detected after 420 minutes in chamber B were approximately 0.25 g/dI and 0.20 g/dI
albumin for the MILLIPORE and WHATh-lAN membranes, respectively.
Glucose does diffuse through micromachined membranes at a rate comparable to com-mercially available membranes. At the same time, albumin is excluded from passage. In mixed solutions of glucose and albumin, only glucose diffuses through the micromachi-ned membranes.
Claims (19)
1. An implantable analyte sensor, comprising:
(a) a substrate, (b) electrodes on said substrate, and (c) a membrane on said electrodes, wherein said membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
(a) a substrate, (b) electrodes on said substrate, and (c) a membrane on said electrodes, wherein said membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
2. The irnplantable analyte sensor of claim 1, wherein said membrane comprises elemental silicon.
3. The implantable analyte sensor of Claim 1 or 2, wherein said glucose diffusion test result is at least 60 nig/dl in 330 min., and said albumin diffusion test result is at mostØ001 g/dl in 420 min.
4. The implantable analyte sensor of Claim 1 or 2, further comprising:
(d) microelectronic circuitry electrically connected to said electrodes.
(d) microelectronic circuitry electrically connected to said electrodes.
5. The implantable analyte sensor of Claim 1 or 2, further comprising:
(e) leads electrically connected to said electrodes.
(e) leads electrically connected to said electrodes.
6. The implantable analyte sensor of Claim 4, further comprising:
(e) leads electrically connected to said electrodes, wherein said leads are electrically connected to said electrodes via said microelectronic circuitry.
(e) leads electrically connected to said electrodes, wherein said leads are electrically connected to said electrodes via said microelectronic circuitry.
7. The implantable analyte sensor of Claim 4, wherein said microelectronic circuitry comprises a transmitter and a power supply.
8. The implantable analyte sensor of Claim 1 or 2, further comprising:
(f) a coating surrounding said substrate and said membrane.
(f) a coating surrounding said substrate and said membrane.
9. The implantable analyte sensor of Claim 8, wherein said coating comprises an internal coating and an external coating.
10. The implantable analyte sensor of Claim 1 or 2, wherein said substrate comprises elemental silicon.
11. The implantable analyte sensor of Claim 1 or 2, wherein said membrane is prepared by micromachining.
12. The implantable analyte sensor of Claim 1 or 2, wherein the implantable analyte sensor is a glucose sensor.
13. A method of making an implantable analyte sensor, comprising:
covering electrodes with a membrane;
wherein said electrodes are on a substrate, and said membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
covering electrodes with a membrane;
wherein said electrodes are on a substrate, and said membrane has a glucose diffusion test result of at least 1 mg/dl in 330 min., and an albumin diffusion test result of at most 0.1 g/dl in 420 min.
14. The method of Claim 13, wherein said membrane comprises elemental silicon.
15. The method of Claim 13 or 14, further comprising:
forming said membrane by micromachining elemental silicon.
forming said membrane by micromachining elemental silicon.
16. The method of Claim 13 or I4, further comprising:
surrounding said membrane and said substrate with a coating.
surrounding said membrane and said substrate with a coating.
17. The method of Claim 13 or 14, wherein said glucose diffusion test result is at least 60 mg/dl in 330 min., and said albumin diffusion test result is at most 0.0001 g/dl in 420 mm.
18. The method of Claim 13 or 14, wherein said membrane is prepared by micromachining.
19. The method of Claim 13 or 14, wherein the implantable analyte sensor is a glucose sensor.
Applications Claiming Priority (3)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US09/528,277 | 2000-03-17 | ||
| US09/528,277 US6405066B1 (en) | 2000-03-17 | 2000-03-17 | Implantable analyte sensor |
| PCT/EP2001/003027 WO2001068901A2 (en) | 2000-03-17 | 2001-03-16 | Implantable analyte sensor |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA2406878A1 true CA2406878A1 (en) | 2001-09-20 |
Family
ID=24104988
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002406878A Abandoned CA2406878A1 (en) | 2000-03-17 | 2001-03-16 | Implantable analyte sensor |
Country Status (6)
| Country | Link |
|---|---|
| US (1) | US6405066B1 (en) |
| EP (1) | EP1267705A2 (en) |
| JP (1) | JP2003526491A (en) |
| AU (1) | AU2001239304A1 (en) |
| CA (1) | CA2406878A1 (en) |
| WO (1) | WO2001068901A2 (en) |
Families Citing this family (201)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US8527026B2 (en) | 1997-03-04 | 2013-09-03 | Dexcom, Inc. | Device and method for determining analyte levels |
| US6862465B2 (en) | 1997-03-04 | 2005-03-01 | Dexcom, Inc. | Device and method for determining analyte levels |
| US7192450B2 (en) | 2003-05-21 | 2007-03-20 | Dexcom, Inc. | Porous membranes for use with implantable devices |
| US6001067A (en) | 1997-03-04 | 1999-12-14 | Shults; Mark C. | Device and method for determining analyte levels |
| US6949816B2 (en) | 2003-04-21 | 2005-09-27 | Motorola, Inc. | Semiconductor component having first surface area for electrically coupling to a semiconductor chip and second surface area for electrically coupling to a substrate, and method of manufacturing same |
| US8346337B2 (en) | 1998-04-30 | 2013-01-01 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8480580B2 (en) | 1998-04-30 | 2013-07-09 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US9066695B2 (en) | 1998-04-30 | 2015-06-30 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8465425B2 (en) | 1998-04-30 | 2013-06-18 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8688188B2 (en) | 1998-04-30 | 2014-04-01 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8974386B2 (en) | 1998-04-30 | 2015-03-10 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US6175752B1 (en) | 1998-04-30 | 2001-01-16 | Therasense, Inc. | Analyte monitoring device and methods of use |
| WO2001036321A1 (en) * | 1999-11-17 | 2001-05-25 | The Regents Of The University Of California | Apparatus and method for forming a membrane with nanometer scale pores |
| RU2278612C2 (en) * | 2000-07-14 | 2006-06-27 | Лайфскен, Инк. | Immune sensor |
| US6560471B1 (en) | 2001-01-02 | 2003-05-06 | Therasense, Inc. | Analyte monitoring device and methods of use |
| US7310543B2 (en) | 2001-03-26 | 2007-12-18 | Kumetrix, Inc. | Silicon microprobe with integrated biosensor |
| EP1397068A2 (en) | 2001-04-02 | 2004-03-17 | Therasense, Inc. | Blood glucose tracking apparatus and methods |
| US6702857B2 (en) | 2001-07-27 | 2004-03-09 | Dexcom, Inc. | Membrane for use with implantable devices |
| US20030032874A1 (en) | 2001-07-27 | 2003-02-13 | Dexcom, Inc. | Sensor head for use with implantable devices |
| US6663615B1 (en) * | 2001-09-04 | 2003-12-16 | The Ohio State University | Dual stage microvalve and method of use |
| WO2003038851A1 (en) * | 2001-11-01 | 2003-05-08 | Massachusetts Institute Of Technology | Organic field emission device |
| US8364229B2 (en) | 2003-07-25 | 2013-01-29 | Dexcom, Inc. | Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise |
| US7613491B2 (en) | 2002-05-22 | 2009-11-03 | Dexcom, Inc. | Silicone based membranes for use in implantable glucose sensors |
| US8260393B2 (en) | 2003-07-25 | 2012-09-04 | Dexcom, Inc. | Systems and methods for replacing signal data artifacts in a glucose sensor data stream |
| US10022078B2 (en) | 2004-07-13 | 2018-07-17 | Dexcom, Inc. | Analyte sensor |
| US9247901B2 (en) | 2003-08-22 | 2016-02-02 | Dexcom, Inc. | Systems and methods for replacing signal artifacts in a glucose sensor data stream |
| US8010174B2 (en) | 2003-08-22 | 2011-08-30 | Dexcom, Inc. | Systems and methods for replacing signal artifacts in a glucose sensor data stream |
| US9282925B2 (en) | 2002-02-12 | 2016-03-15 | Dexcom, Inc. | Systems and methods for replacing signal artifacts in a glucose sensor data stream |
| CA2501965C (en) * | 2002-09-11 | 2012-04-10 | The Regents Of The University Of Michigan | Ultrafiltration membrane, device, bioartificial organ, and methods |
| US20040064050A1 (en) * | 2002-09-20 | 2004-04-01 | Jun Liu | System and method for screening tissue |
| DE60336834D1 (en) | 2002-10-09 | 2011-06-01 | Abbott Diabetes Care Inc | FUEL FEEDING DEVICE, SYSTEM AND METHOD |
| US7727181B2 (en) | 2002-10-09 | 2010-06-01 | Abbott Diabetes Care Inc. | Fluid delivery device with autocalibration |
| US7993108B2 (en) | 2002-10-09 | 2011-08-09 | Abbott Diabetes Care Inc. | Variable volume, shape memory actuated insulin dispensing pump |
| AU2003303597A1 (en) | 2002-12-31 | 2004-07-29 | Therasense, Inc. | Continuous glucose monitoring system and methods of use |
| US7679407B2 (en) | 2003-04-28 | 2010-03-16 | Abbott Diabetes Care Inc. | Method and apparatus for providing peak detection circuitry for data communication systems |
| US7875293B2 (en) | 2003-05-21 | 2011-01-25 | Dexcom, Inc. | Biointerface membranes incorporating bioactive agents |
| US8066639B2 (en) | 2003-06-10 | 2011-11-29 | Abbott Diabetes Care Inc. | Glucose measuring device for use in personal area network |
| US7097788B2 (en) * | 2003-06-30 | 2006-08-29 | The Board Of Trustees Of The University Of Illinois | Conducting inks |
| US7722536B2 (en) * | 2003-07-15 | 2010-05-25 | Abbott Diabetes Care Inc. | Glucose measuring device integrated into a holster for a personal area network device |
| EP1648298A4 (en) | 2003-07-25 | 2010-01-13 | Dexcom Inc | Oxygen enhancing membrane systems for implantable devices |
| US9763609B2 (en) | 2003-07-25 | 2017-09-19 | Dexcom, Inc. | Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise |
| JP2007500336A (en) | 2003-07-25 | 2007-01-11 | デックスコム・インコーポレーテッド | Electrode system for electrochemical sensors |
| WO2005019795A2 (en) | 2003-07-25 | 2005-03-03 | Dexcom, Inc. | Electrochemical sensors including electrode systems with increased oxygen generation |
| US8282549B2 (en) | 2003-12-09 | 2012-10-09 | Dexcom, Inc. | Signal processing for continuous analyte sensor |
| US8160669B2 (en) | 2003-08-01 | 2012-04-17 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US8761856B2 (en) | 2003-08-01 | 2014-06-24 | Dexcom, Inc. | System and methods for processing analyte sensor data |
| US7276029B2 (en) | 2003-08-01 | 2007-10-02 | Dexcom, Inc. | System and methods for processing analyte sensor data |
| US9135402B2 (en) | 2007-12-17 | 2015-09-15 | Dexcom, Inc. | Systems and methods for processing sensor data |
| US8886273B2 (en) | 2003-08-01 | 2014-11-11 | Dexcom, Inc. | Analyte sensor |
| US7494465B2 (en) | 2004-07-13 | 2009-02-24 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US8275437B2 (en) | 2003-08-01 | 2012-09-25 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US7933639B2 (en) | 2003-08-01 | 2011-04-26 | Dexcom, Inc. | System and methods for processing analyte sensor data |
| US7519408B2 (en) | 2003-11-19 | 2009-04-14 | Dexcom, Inc. | Integrated receiver for continuous analyte sensor |
| US7591801B2 (en) | 2004-02-26 | 2009-09-22 | Dexcom, Inc. | Integrated delivery device for continuous glucose sensor |
| US8369919B2 (en) | 2003-08-01 | 2013-02-05 | Dexcom, Inc. | Systems and methods for processing sensor data |
| US7774145B2 (en) | 2003-08-01 | 2010-08-10 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US20190357827A1 (en) | 2003-08-01 | 2019-11-28 | Dexcom, Inc. | Analyte sensor |
| US20100168657A1 (en) | 2003-08-01 | 2010-07-01 | Dexcom, Inc. | System and methods for processing analyte sensor data |
| US20140121989A1 (en) | 2003-08-22 | 2014-05-01 | Dexcom, Inc. | Systems and methods for processing analyte sensor data |
| US7920906B2 (en) | 2005-03-10 | 2011-04-05 | Dexcom, Inc. | System and methods for processing analyte sensor data for sensor calibration |
| EP1681992B2 (en) † | 2003-09-30 | 2015-03-04 | Roche Diagnostics GmbH | Sensor with increased biocompatibility |
| US7388488B2 (en) * | 2003-10-30 | 2008-06-17 | Peter Lupoli | Method and system for storing, retrieving, and managing data for tags |
| US7956742B2 (en) | 2003-10-30 | 2011-06-07 | Motedata Inc. | Method and system for storing, retrieving, and managing data for tags |
| US9247900B2 (en) | 2004-07-13 | 2016-02-02 | Dexcom, Inc. | Analyte sensor |
| US8423114B2 (en) | 2006-10-04 | 2013-04-16 | Dexcom, Inc. | Dual electrode system for a continuous analyte sensor |
| US11633133B2 (en) | 2003-12-05 | 2023-04-25 | Dexcom, Inc. | Dual electrode system for a continuous analyte sensor |
| US8364231B2 (en) | 2006-10-04 | 2013-01-29 | Dexcom, Inc. | Analyte sensor |
| US8774886B2 (en) | 2006-10-04 | 2014-07-08 | Dexcom, Inc. | Analyte sensor |
| US7553625B2 (en) | 2003-12-22 | 2009-06-30 | John Wayne Cancer Institute | Method and apparatus for in vivo collection of circulating biological components |
| US20050153309A1 (en) * | 2003-12-22 | 2005-07-14 | David Hoon | Method and apparatus for in vivo surveillance of circulating biological components |
| US8948836B2 (en) * | 2003-12-26 | 2015-02-03 | Medtronic Minimed, Inc. | Implantable apparatus for sensing multiple parameters |
| EP1718198A4 (en) | 2004-02-17 | 2008-06-04 | Therasense Inc | Method and system for providing data communication in continuous glucose monitoring and management system |
| US8808228B2 (en) | 2004-02-26 | 2014-08-19 | Dexcom, Inc. | Integrated medicament delivery device for use with continuous analyte sensor |
| US8277713B2 (en) * | 2004-05-03 | 2012-10-02 | Dexcom, Inc. | Implantable analyte sensor |
| CA2572455C (en) * | 2004-06-04 | 2014-10-28 | Therasense, Inc. | Diabetes care host-client architecture and data management system |
| US20060020192A1 (en) | 2004-07-13 | 2006-01-26 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US8565848B2 (en) | 2004-07-13 | 2013-10-22 | Dexcom, Inc. | Transcutaneous analyte sensor |
| AU2005295106B2 (en) | 2004-10-12 | 2012-03-15 | Bayer Healthcare Llc | Concentration determination in a diffusion barrier layer |
| US7545272B2 (en) | 2005-02-08 | 2009-06-09 | Therasense, Inc. | RF tag on test strips, test strip vials and boxes |
| CA2601441A1 (en) | 2005-03-21 | 2006-09-28 | Abbott Diabetes Care Inc. | Method and system for providing integrated medication infusion and analyte monitoring system |
| US20070066138A1 (en) * | 2005-04-05 | 2007-03-22 | The Ohio State University Research Foundation | Diffusion Delivery Systems and Methods of Fabrication |
| WO2006110193A2 (en) | 2005-04-08 | 2006-10-19 | Dexcom, Inc. | Cellulosic-based interference domain for an analyte sensor |
| US8744546B2 (en) | 2005-05-05 | 2014-06-03 | Dexcom, Inc. | Cellulosic-based resistance domain for an analyte sensor |
| US20060249381A1 (en) * | 2005-05-05 | 2006-11-09 | Petisce James R | Cellulosic-based resistance domain for an analyte sensor |
| US8112240B2 (en) | 2005-04-29 | 2012-02-07 | Abbott Diabetes Care Inc. | Method and apparatus for providing leak detection in data monitoring and management systems |
| US7768408B2 (en) | 2005-05-17 | 2010-08-03 | Abbott Diabetes Care Inc. | Method and system for providing data management in data monitoring system |
| US7620437B2 (en) | 2005-06-03 | 2009-11-17 | Abbott Diabetes Care Inc. | Method and apparatus for providing rechargeable power in data monitoring and management systems |
| EP1896839A2 (en) * | 2005-06-30 | 2008-03-12 | MC3, Inc. | Analyte sensors and compositions for use therein |
| CN102440785A (en) | 2005-08-31 | 2012-05-09 | 弗吉尼亚大学专利基金委员会 | Sensor signal processing method and sensor signal processing device |
| US7756561B2 (en) | 2005-09-30 | 2010-07-13 | Abbott Diabetes Care Inc. | Method and apparatus for providing rechargeable power in data monitoring and management systems |
| US7583190B2 (en) | 2005-10-31 | 2009-09-01 | Abbott Diabetes Care Inc. | Method and apparatus for providing data communication in data monitoring and management systems |
| US7766829B2 (en) | 2005-11-04 | 2010-08-03 | Abbott Diabetes Care Inc. | Method and system for providing basal profile modification in analyte monitoring and management systems |
| EP2004796B1 (en) * | 2006-01-18 | 2015-04-08 | DexCom, Inc. | Membranes for an analyte sensor |
| US8344966B2 (en) | 2006-01-31 | 2013-01-01 | Abbott Diabetes Care Inc. | Method and system for providing a fault tolerant display unit in an electronic device |
| US7885698B2 (en) | 2006-02-28 | 2011-02-08 | Abbott Diabetes Care Inc. | Method and system for providing continuous calibration of implantable analyte sensors |
| EP1991110B1 (en) | 2006-03-09 | 2018-11-07 | DexCom, Inc. | Systems and methods for processing analyte sensor data |
| US8219173B2 (en) | 2008-09-30 | 2012-07-10 | Abbott Diabetes Care Inc. | Optimizing analyte sensor calibration |
| US8226891B2 (en) | 2006-03-31 | 2012-07-24 | Abbott Diabetes Care Inc. | Analyte monitoring devices and methods therefor |
| US7630748B2 (en) | 2006-10-25 | 2009-12-08 | Abbott Diabetes Care Inc. | Method and system for providing analyte monitoring |
| US7620438B2 (en) | 2006-03-31 | 2009-11-17 | Abbott Diabetes Care Inc. | Method and system for powering an electronic device |
| US8583205B2 (en) * | 2008-03-28 | 2013-11-12 | Abbott Diabetes Care Inc. | Analyte sensor calibration management |
| US9675290B2 (en) | 2012-10-30 | 2017-06-13 | Abbott Diabetes Care Inc. | Sensitivity calibration of in vivo sensors used to measure analyte concentration |
| US9392969B2 (en) | 2008-08-31 | 2016-07-19 | Abbott Diabetes Care Inc. | Closed loop control and signal attenuation detection |
| US8224415B2 (en) | 2009-01-29 | 2012-07-17 | Abbott Diabetes Care Inc. | Method and device for providing offset model based calibration for analyte sensor |
| US7920907B2 (en) | 2006-06-07 | 2011-04-05 | Abbott Diabetes Care Inc. | Analyte monitoring system and method |
| US8135548B2 (en) | 2006-10-26 | 2012-03-13 | Abbott Diabetes Care Inc. | Method, system and computer program product for real-time detection of sensitivity decline in analyte sensors |
| US8579853B2 (en) | 2006-10-31 | 2013-11-12 | Abbott Diabetes Care Inc. | Infusion devices and methods |
| US20080121045A1 (en) * | 2006-11-29 | 2008-05-29 | Cole Matthew C | Multiplexed sensor array |
| WO2008086477A1 (en) | 2007-01-10 | 2008-07-17 | The Regents Of The University Of Michigan | Ultrafiltration membrane, device, bioartificial organ, and related methods |
| US8930203B2 (en) | 2007-02-18 | 2015-01-06 | Abbott Diabetes Care Inc. | Multi-function analyte test device and methods therefor |
| US8732188B2 (en) | 2007-02-18 | 2014-05-20 | Abbott Diabetes Care Inc. | Method and system for providing contextual based medication dosage determination |
| US8123686B2 (en) | 2007-03-01 | 2012-02-28 | Abbott Diabetes Care Inc. | Method and apparatus for providing rolling data in communication systems |
| EP2146627B1 (en) | 2007-04-14 | 2020-07-29 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in medical communication system |
| EP2137637A4 (en) * | 2007-04-14 | 2012-06-20 | Abbott Diabetes Care Inc | Method and apparatus for providing data processing and control in medical communication system |
| ES2784736T3 (en) | 2007-04-14 | 2020-09-30 | Abbott Diabetes Care Inc | Procedure and apparatus for providing data processing and control in a medical communication system |
| EP2146625B1 (en) | 2007-04-14 | 2019-08-14 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in medical communication system |
| US9204827B2 (en) | 2007-04-14 | 2015-12-08 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in medical communication system |
| US7928850B2 (en) | 2007-05-08 | 2011-04-19 | Abbott Diabetes Care Inc. | Analyte monitoring system and methods |
| US8461985B2 (en) | 2007-05-08 | 2013-06-11 | Abbott Diabetes Care Inc. | Analyte monitoring system and methods |
| US8665091B2 (en) | 2007-05-08 | 2014-03-04 | Abbott Diabetes Care Inc. | Method and device for determining elapsed sensor life |
| US8456301B2 (en) | 2007-05-08 | 2013-06-04 | Abbott Diabetes Care Inc. | Analyte monitoring system and methods |
| US9125548B2 (en) | 2007-05-14 | 2015-09-08 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8239166B2 (en) | 2007-05-14 | 2012-08-07 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8103471B2 (en) | 2007-05-14 | 2012-01-24 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8560038B2 (en) | 2007-05-14 | 2013-10-15 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8260558B2 (en) | 2007-05-14 | 2012-09-04 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8600681B2 (en) | 2007-05-14 | 2013-12-03 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US8444560B2 (en) | 2007-05-14 | 2013-05-21 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US10002233B2 (en) | 2007-05-14 | 2018-06-19 | Abbott Diabetes Care Inc. | Method and apparatus for providing data processing and control in a medical communication system |
| US20200037875A1 (en) | 2007-05-18 | 2020-02-06 | Dexcom, Inc. | Analyte sensors having a signal-to-noise ratio substantially unaffected by non-constant noise |
| EP2152350A4 (en) | 2007-06-08 | 2013-03-27 | Dexcom Inc | Integrated medicament delivery device for use with continuous analyte sensor |
| CN103251414B (en) | 2007-06-21 | 2017-05-24 | 雅培糖尿病护理公司 | Device for detecting analyte level |
| JP5680960B2 (en) | 2007-06-21 | 2015-03-04 | アボット ダイアベティス ケア インコーポレイテッドＡｂｂｏｔｔ Ｄｉａｂｅｔｅｓ Ｃａｒｅ Ｉｎｃ． | Health care device and method |
| US8160900B2 (en) | 2007-06-29 | 2012-04-17 | Abbott Diabetes Care Inc. | Analyte monitoring and management device and method to analyze the frequency of user interaction with the device |
| EP2017350A1 (en) * | 2007-07-19 | 2009-01-21 | F. Hoffmann-La Roche AG | Electrochemical sensor with covalent-bound enzyme |
| US8834366B2 (en) | 2007-07-31 | 2014-09-16 | Abbott Diabetes Care Inc. | Method and apparatus for providing analyte sensor calibration |
| EP2027812A1 (en) * | 2007-08-24 | 2009-02-25 | F. Hoffman-la Roche AG | Method for manufacturing a micro-dialysis catheter and micro-dialysis catheter manufactured accordingly |
| EP4098177A1 (en) | 2007-10-09 | 2022-12-07 | DexCom, Inc. | Integrated insulin delivery system with continuous glucose sensor |
| US8417312B2 (en) | 2007-10-25 | 2013-04-09 | Dexcom, Inc. | Systems and methods for processing sensor data |
| US9839395B2 (en) | 2007-12-17 | 2017-12-12 | Dexcom, Inc. | Systems and methods for processing sensor data |
| US20090164239A1 (en) | 2007-12-19 | 2009-06-25 | Abbott Diabetes Care, Inc. | Dynamic Display Of Glucose Information |
| US20090247856A1 (en) * | 2008-03-28 | 2009-10-01 | Dexcom, Inc. | Polymer membranes for continuous analyte sensors |
| US8583204B2 (en) | 2008-03-28 | 2013-11-12 | Dexcom, Inc. | Polymer membranes for continuous analyte sensors |
| US8682408B2 (en) | 2008-03-28 | 2014-03-25 | Dexcom, Inc. | Polymer membranes for continuous analyte sensors |
| US11730407B2 (en) | 2008-03-28 | 2023-08-22 | Dexcom, Inc. | Polymer membranes for continuous analyte sensors |
| US7826382B2 (en) | 2008-05-30 | 2010-11-02 | Abbott Diabetes Care Inc. | Close proximity communication device and methods |
| US8591410B2 (en) | 2008-05-30 | 2013-11-26 | Abbott Diabetes Care Inc. | Method and apparatus for providing glycemic control |
| US8924159B2 (en) | 2008-05-30 | 2014-12-30 | Abbott Diabetes Care Inc. | Method and apparatus for providing glycemic control |
| WO2010027771A1 (en) | 2008-08-27 | 2010-03-11 | Edwards Lifesciences Corporation | Analyte sensor |
| US9326707B2 (en) | 2008-11-10 | 2016-05-03 | Abbott Diabetes Care Inc. | Alarm characterization for analyte monitoring devices and systems |
| US8103456B2 (en) | 2009-01-29 | 2012-01-24 | Abbott Diabetes Care Inc. | Method and device for early signal attenuation detection using blood glucose measurements |
| US8560082B2 (en) | 2009-01-30 | 2013-10-15 | Abbott Diabetes Care Inc. | Computerized determination of insulin pump therapy parameters using real time and retrospective data processing |
| US20100198034A1 (en) | 2009-02-03 | 2010-08-05 | Abbott Diabetes Care Inc. | Compact On-Body Physiological Monitoring Devices and Methods Thereof |
| EP2410910A4 (en) | 2009-03-27 | 2014-10-15 | Dexcom Inc | Methods and systems for promoting glucose management |
| US9226701B2 (en) | 2009-04-28 | 2016-01-05 | Abbott Diabetes Care Inc. | Error detection in critical repeating data in a wireless sensor system |
| WO2010129375A1 (en) | 2009-04-28 | 2010-11-11 | Abbott Diabetes Care Inc. | Closed loop blood glucose control algorithm analysis |
| WO2010127187A1 (en) | 2009-04-29 | 2010-11-04 | Abbott Diabetes Care Inc. | Method and system for providing data communication in continuous glucose monitoring and management system |
| WO2010138856A1 (en) | 2009-05-29 | 2010-12-02 | Abbott Diabetes Care Inc. | Medical device antenna systems having external antenna configurations |
| EP2456351B1 (en) | 2009-07-23 | 2016-10-12 | Abbott Diabetes Care, Inc. | Real time management of data relating to physiological control of glucose levels |
| WO2011014851A1 (en) | 2009-07-31 | 2011-02-03 | Abbott Diabetes Care Inc. | Method and apparatus for providing analyte monitoring system calibration accuracy |
| CA2765712A1 (en) | 2009-08-31 | 2011-03-03 | Abbott Diabetes Care Inc. | Medical devices and methods |
| WO2011026148A1 (en) | 2009-08-31 | 2011-03-03 | Abbott Diabetes Care Inc. | Analyte monitoring system and methods for managing power and noise |
| WO2011026147A1 (en) | 2009-08-31 | 2011-03-03 | Abbott Diabetes Care Inc. | Analyte signal processing device and methods |
| US9320461B2 (en) | 2009-09-29 | 2016-04-26 | Abbott Diabetes Care Inc. | Method and apparatus for providing notification function in analyte monitoring systems |
| EP2482724A2 (en) | 2009-09-30 | 2012-08-08 | Dexcom, Inc. | Transcutaneous analyte sensor |
| US8185181B2 (en) | 2009-10-30 | 2012-05-22 | Abbott Diabetes Care Inc. | Method and apparatus for detecting false hypoglycemic conditions |
| WO2011091336A1 (en) * | 2010-01-22 | 2011-07-28 | Abbott Diabetes Care Inc. | Method and apparatus for providing notification in analyte monitoring systems |
| WO2011112753A1 (en) | 2010-03-10 | 2011-09-15 | Abbott Diabetes Care Inc. | Systems, devices and methods for managing glucose levels |
| US8361196B2 (en) * | 2010-04-09 | 2013-01-29 | Inficon Gmbh | Gas-selective membrane and method of its production |
| US8635046B2 (en) | 2010-06-23 | 2014-01-21 | Abbott Diabetes Care Inc. | Method and system for evaluating analyte sensor response characteristics |
| US10092229B2 (en) | 2010-06-29 | 2018-10-09 | Abbott Diabetes Care Inc. | Calibration of analyte measurement system |
| WO2012048168A2 (en) | 2010-10-07 | 2012-04-12 | Abbott Diabetes Care Inc. | Analyte monitoring devices and methods |
| CA3177983A1 (en) | 2011-02-28 | 2012-11-15 | Abbott Diabetes Care Inc. | Devices, systems, and methods associated with analyte monitoring devices and devices incorporating the same |
| US10136845B2 (en) | 2011-02-28 | 2018-11-27 | Abbott Diabetes Care Inc. | Devices, systems, and methods associated with analyte monitoring devices and devices incorporating the same |
| WO2012142502A2 (en) | 2011-04-15 | 2012-10-18 | Dexcom Inc. | Advanced analyte sensor calibration and error detection |
| WO2013066873A1 (en) | 2011-10-31 | 2013-05-10 | Abbott Diabetes Care Inc. | Electronic devices having integrated reset systems and methods thereof |
| US9622691B2 (en) | 2011-10-31 | 2017-04-18 | Abbott Diabetes Care Inc. | Model based variable risk false glucose threshold alarm prevention mechanism |
| JP6443802B2 (en) | 2011-11-07 | 2018-12-26 | アボット ダイアベティス ケア インコーポレイテッドＡｂｂｏｔｔ Ｄｉａｂｅｔｅｓ Ｃａｒｅ Ｉｎｃ． | Analyte monitoring apparatus and method |
| US9317656B2 (en) | 2011-11-23 | 2016-04-19 | Abbott Diabetes Care Inc. | Compatibility mechanisms for devices in a continuous analyte monitoring system and methods thereof |
| US8710993B2 (en) | 2011-11-23 | 2014-04-29 | Abbott Diabetes Care Inc. | Mitigating single point failure of devices in an analyte monitoring system and methods thereof |
| EP2890297B1 (en) | 2012-08-30 | 2018-04-11 | Abbott Diabetes Care, Inc. | Dropout detection in continuous analyte monitoring data during data excursions |
| US9968306B2 (en) | 2012-09-17 | 2018-05-15 | Abbott Diabetes Care Inc. | Methods and apparatuses for providing adverse condition notification with enhanced wireless communication range in analyte monitoring systems |
| WO2014052136A1 (en) | 2012-09-26 | 2014-04-03 | Abbott Diabetes Care Inc. | Method and apparatus for improving lag correction during in vivo measurement of analyte concentration with analyte concentration variability and range data |
| US9788765B2 (en) | 2012-09-28 | 2017-10-17 | Dexcom, Inc. | Zwitterion surface modifications for continuous sensors |
| US20140107450A1 (en) * | 2012-10-12 | 2014-04-17 | Dexcom, Inc. | Sensors for continuous analyte monitoring, and related methods |
| US20140213866A1 (en) | 2012-10-12 | 2014-07-31 | Dexcom, Inc. | Sensors for continuous analyte monitoring, and related methods |
| US11109779B2 (en) | 2012-12-28 | 2021-09-07 | Senseonics, Incorporated | Chemical modification of analyte permeable membrane for enhanced oxidative stability |
| AU2013370204B2 (en) | 2012-12-28 | 2019-02-14 | Senseonics, Incorporated | Analyte permeable membrane systems for oxidative and optical stability |
| US10433773B1 (en) | 2013-03-15 | 2019-10-08 | Abbott Diabetes Care Inc. | Noise rejection methods and apparatus for sparsely sampled analyte sensor data |
| US9474475B1 (en) | 2013-03-15 | 2016-10-25 | Abbott Diabetes Care Inc. | Multi-rate analyte sensor data collection with sample rate configurable signal processing |
| US9694190B2 (en) | 2013-03-15 | 2017-07-04 | Lawrence Livermore National Security, Llc | Method to pattern <10 micrometer conducting and passivating features on 3D substrates for implantable devices |
| US10076285B2 (en) | 2013-03-15 | 2018-09-18 | Abbott Diabetes Care Inc. | Sensor fault detection using analyte sensor data pattern comparison |
| US11229382B2 (en) | 2013-12-31 | 2022-01-25 | Abbott Diabetes Care Inc. | Self-powered analyte sensor and devices using the same |
| US20170185748A1 (en) | 2014-03-30 | 2017-06-29 | Abbott Diabetes Care Inc. | Method and Apparatus for Determining Meal Start and Peak Events in Analyte Monitoring Systems |
| US20150289788A1 (en) | 2014-04-10 | 2015-10-15 | Dexcom, Inc. | Sensors for continuous analyte monitoring, and related methods |
| EP3319518A4 (en) | 2015-07-10 | 2019-03-13 | Abbott Diabetes Care Inc. | System, device and method of dynamic glucose profile response to physiological parameters |
| US20170188905A1 (en) | 2015-12-30 | 2017-07-06 | Dexcom, Inc. | Biointerface layer for analyte sensors |
| WO2017214550A1 (en) | 2016-06-09 | 2017-12-14 | Haimachek, Inc. | Collector for detection and reversible capturing of cells from body fluids in vivo |
| US11596330B2 (en) | 2017-03-21 | 2023-03-07 | Abbott Diabetes Care Inc. | Methods, devices and system for providing diabetic condition diagnosis and therapy |
| US11331022B2 (en) | 2017-10-24 | 2022-05-17 | Dexcom, Inc. | Pre-connected analyte sensors |
| US11382540B2 (en) | 2017-10-24 | 2022-07-12 | Dexcom, Inc. | Pre-connected analyte sensors |
Family Cites Families (25)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US4592824A (en) * | 1985-09-13 | 1986-06-03 | Centre Suisse D'electronique Et De Microtechnique S.A. | Miniature liquid junction reference electrode and an integrated solid state electrochemical sensor including the same |
| GB8529300D0 (en) * | 1985-11-28 | 1986-01-02 | Ici Plc | Membrane |
| US4894339A (en) * | 1985-12-18 | 1990-01-16 | Seitaikinouriyou Kagakuhin Sinseizogijutsu Kenkyu Kumiai | Immobilized enzyme membrane for a semiconductor sensor |
| JPS63182559A (en) * | 1987-01-24 | 1988-07-27 | Kanzaki Paper Mfg Co Ltd | Production of enzyme electrode |
| DE3875149T2 (en) | 1987-03-27 | 1993-02-11 | Isao Karube | MINIATURIZED BIO-SENSOR WITH MINIATURIZED OXYGEN ELECTRODE AND ITS PRODUCTION PROCESS. |
| US5200051A (en) | 1988-11-14 | 1993-04-06 | I-Stat Corporation | Wholly microfabricated biosensors and process for the manufacture and use thereof |
| US5431160A (en) * | 1989-07-19 | 1995-07-11 | University Of New Mexico | Miniature implantable refillable glucose sensor and material therefor |
| FR2652736A1 (en) | 1989-10-06 | 1991-04-12 | Neftel Frederic | IMPLANTABLE DEVICE FOR EVALUATING THE RATE OF GLUCOSE. |
| US5165407A (en) | 1990-04-19 | 1992-11-24 | The University Of Kansas | Implantable glucose sensor |
| US5593852A (en) | 1993-12-02 | 1997-01-14 | Heller; Adam | Subcutaneous glucose electrode |
| US5773270A (en) | 1991-03-12 | 1998-06-30 | Chiron Diagnostics Corporation | Three-layered membrane for use in an electrochemical sensor system |
| US5322063A (en) | 1991-10-04 | 1994-06-21 | Eli Lilly And Company | Hydrophilic polyurethane membranes for electrochemical glucose sensors |
| US5387327A (en) | 1992-10-19 | 1995-02-07 | Duquesne University Of The Holy Ghost | Implantable non-enzymatic electrochemical glucose sensor |
| ZA938555B (en) | 1992-11-23 | 1994-08-02 | Lilly Co Eli | Technique to improve the performance of electrochemical sensors |
| GB9304306D0 (en) | 1993-03-03 | 1993-04-21 | Univ Alberta | Glucose sensor |
| NL9401260A (en) | 1993-11-12 | 1995-06-01 | Cornelis Johannes Maria Van Ri | Membrane for microfiltration, ultrafiltration, gas separation and catalysis, method for manufacturing such a membrane, mold for manufacturing such a membrane, as well as various separation systems comprising such a membrane. |
| US5497772A (en) | 1993-11-19 | 1996-03-12 | Alfred E. Mann Foundation For Scientific Research | Glucose monitoring system |
| US5985328A (en) | 1994-03-07 | 1999-11-16 | Regents Of The University Of California | Micromachined porous membranes with bulk support |
| US5494562A (en) * | 1994-06-27 | 1996-02-27 | Ciba Corning Diagnostics Corp. | Electrochemical sensors |
| DE4427921C2 (en) | 1994-08-06 | 2002-09-26 | Forschungszentrum Juelich Gmbh | Chemical sensors, especially biosensors, based on silicon |
| US5786439A (en) | 1996-10-24 | 1998-07-28 | Minimed Inc. | Hydrophilic, swellable coatings for biosensors |
| US5711861A (en) * | 1995-11-22 | 1998-01-27 | Ward; W. Kenneth | Device for monitoring changes in analyte concentration |
| US6001067A (en) | 1997-03-04 | 1999-12-14 | Shults; Mark C. | Device and method for determining analyte levels |
| US6119028A (en) | 1997-10-20 | 2000-09-12 | Alfred E. Mann Foundation | Implantable enzyme-based monitoring systems having improved longevity due to improved exterior surfaces |
| US6201980B1 (en) * | 1998-10-05 | 2001-03-13 | The Regents Of The University Of California | Implantable medical sensor system |
-
2000
- 2000-03-17 US US09/528,277 patent/US6405066B1/en not_active Expired - Lifetime
-
2001
- 2001-03-16 WO PCT/EP2001/003027 patent/WO2001068901A2/en not_active Application Discontinuation
- 2001-03-16 AU AU2001239304A patent/AU2001239304A1/en not_active Abandoned
- 2001-03-16 CA CA002406878A patent/CA2406878A1/en not_active Abandoned
- 2001-03-16 EP EP01913889A patent/EP1267705A2/en not_active Withdrawn
- 2001-03-16 JP JP2001567382A patent/JP2003526491A/en active Pending
Also Published As
| Publication number | Publication date |
|---|---|
| US6405066B1 (en) | 2002-06-11 |
| WO2001068901A3 (en) | 2001-12-27 |
| AU2001239304A1 (en) | 2001-09-24 |
| EP1267705A2 (en) | 2003-01-02 |
| JP2003526491A (en) | 2003-09-09 |
| WO2001068901A2 (en) | 2001-09-20 |
Similar Documents
| Publication | Publication Date | Title |
|---|---|---|
| US6405066B1 (en) | Implantable analyte sensor | |
| CA2406814A1 (en) | Implantable analyte sensor | |
| JP5568298B2 (en) | Method for performing an electrochemical reaction in an analyte sensor | |
| US9163273B2 (en) | Biosensors and methods for making and using them | |
| US9804114B2 (en) | Sensor head for use with implantable devices | |
| JP3950174B2 (en) | Glucose sensor | |
| JP4763777B2 (en) | Enzyme sensor comprising a cover membrane layer coated with a hydrophilic polymer | |
| CA2658607C (en) | Analyte sensors and methods for making and using them | |
| KR20090004674A (en) | Enzyme electrode and enzyme sensor | |
| Zhu et al. | An overview of Si-based biosensors |
Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| EEER | Examination request | ||
| FZDE | Discontinued |