CA2237102A1 - Device for monitoring changes in analyte concentration - Google Patents
Device for monitoring changes in analyte concentration Download PDFInfo
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- CA2237102A1 CA2237102A1 CA002237102A CA2237102A CA2237102A1 CA 2237102 A1 CA2237102 A1 CA 2237102A1 CA 002237102 A CA002237102 A CA 002237102A CA 2237102 A CA2237102 A CA 2237102A CA 2237102 A1 CA2237102 A1 CA 2237102A1
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- Prior art keywords
- sensor
- analyte
- glucose
- sensor body
- anode
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Classifications
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- 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
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- 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
- C12Q1/54—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions involving glucose or galactose
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B5/00—Measuring for diagnostic purposes; Identification of persons
- A61B5/0002—Remote monitoring of patients using telemetry, e.g. transmission of vital signals via a communication network
- A61B5/0031—Implanted circuitry
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- Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
- Y10—TECHNICAL SUBJECTS COVERED BY FORMER USPC
- Y10S—TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
- Y10S435/00—Chemistry: molecular biology and microbiology
- Y10S435/817—Enzyme or microbe electrode
Abstract
The invention provides an electrochemical sensor system for measuring analyte concentrations in a fluid sample. The invention is particularly useful for measuring analytes such as glucose in a patient. An implantable glucose sensor (18) includes a disc shaped body (20) containing multiple anodes (22) on opposing sides of the body (20). Electrodes (22, 24) are connected to a transmitter (130) which transmits radio signals to an external receiver (134) and computer (138) where data is processed to yield glucose concentration figures.
Description
DEVICE FOR MONITORING CHANGES
IN ANALYTE CONCENTRATION
Field of the Invention The invention relates to electroçhemic~l systems for me~c~rin~
analyte concentration. In particular, the invention involves a sensor including electrodes under a semi-permeable membrane for mol~.lolmg analyte conc~llLI~lions in fluids surrounding the sensor.
Back~round There are many instances when it is necessary to monitor ~he concelll~alion of molecules ("analytes") in a fluid. For example, glucose levelsmust be frequently monitored in persons with diabetes so that a~lopliate doses of insulin can be ~rlminictçred in a timely m~nner Many other analytes are me~cllred commonly in human blood and in other fluids.
A variety of me~ods and devices for me~cllrin~ analytes in fluids have been devised. One such device, referred to as an electrochemical sensor, typically includes oppositely charged electrodes under a semi-pçrme~ble membrane. Depending on what analyte is being monitored, membranes, enzymes and/or other a~plopliate materials are provided around the electrodes so ~hat analyte reaction and transport from the fluid ~ullounding the sensor i controlled. Oxidative and reductive reactions take place at or near the electrodes, thus c~llcin~ electron potentials mç~cllred as ch~n~es in cu.l~,nL
which may be correlated to the concentration of analyte in the fluid.
Electrochemical sensors have been used to me~cllre glucose in human blood for a long time. Most of these sensors are ~lesigne~l to me~cllre glucose in a blood sample which has been drawn or ext~acted from the patient.
For patients such as people with diabetes who must test blood glucose levels as 5 often as several times per day, the regular blood drawing process (typically by finger tip puncture) becomes quite cumbersome, messy and even painfill. The diabetic must car~y special eqllipmçnt for extracting blood. Some p~tient~ fail to test as frequently as they should because of problems associated with the blood extracting process.
Therefore, it has been recognized for a long time that an impl~nte~l glucose sensor would offer the important advantage of avoiding the need for repeated blood extraction. However, there are other problems which must be addressed with an implantable sensor. First, there must be a mech~nicm for accessin~ raw electrical data generated by the sensor under the 15 patient's slcin. Protruding wires are lm(lçcirable because they are cumbersome, prone to c~llcing infection and sometimes painful. Accordingly, it is preferableto incl~l(le a wireless data tr~ncmission (telemetry) device coupled to the sensor in a single impl~nt~ble unit so that no trans-dermal wires are required.
Second, an implanted sensing unit may cause int~rn~l trauma, i.e., 20 bruising or bleeding from the patient's routine movement or contact with his or her environment especially if the sensing unit is large or thick or if it is geometrically shaped with any sharp points or edges.
WO 97/19344 . PCT/US96/18724 Another problem associated with implantable sensors is that over time (days and weeks) a cellular coat tends to develop around the sensor which may evenhl~lly block the analyte of interest from cont~ctinp; the electrodes, thus c~lcin~ the sensor to fail.
For these reasons, and perhaps other reasons, researchers in the field have been lln~lccessfill in their alle,~ to produce an implantable sensor unit which is capable of functioning s~ticf~ctorily for a suficient period of time to justify the expense and inconvenience of producing and surgically imrl~nting the sensing ha~dw~.
A viable implantable glucose sensor should provide reliable performance for at least 1-2 months, preferably three months or more. During its useful life, the device should generate a predictable dose response over a concentration range of approximately 40 to 400 milligrams per deciliter (mg/dl). The device should exhibit a lag time between a concentration change and the resl.lting signal output of less than 20 min~ltes, preferably less than 10 ...i...,les. The sensor should be relatively insensitive to potential i.~ fe~ g substances such as ascorbic acid and acetaminophen. The device should be relatively accurate for at least several days after calibration (stability). Glucose me~llrement with the sensor should be precise to at least within a~p,ox~llately 20 10 mg/dl. The sensor should be incorporated in an implantable unit which is capable of wireless data tr~n~mi~sion, and which is ~limencioned so as to II.i..i...i7P surgical complication and risk of pain, bruising or other in~çrn~ltrauma.
Summary of the Invention The objectives stated above are achievable with the device and system of the present invention which includes a device for electrochemically s~n~in~ ch~n~es in the conce~ Lion of an analyte of i~llelt;~l.
S In one embo-lirnent of the invention, the device inclll~les a sensor body having two opposing sides. Each side of the body includes at least one, erel~bly several, anode(s~ and at least one cathode spaced apart ~om each other and covered by a membrane which is semi-pçrme~le to the analyte of interest. In a p-e~lled sensor design for me~cllnn~ glucose, plural anodes are 10 disposed on two opposing sides of a disc-shaped sensor body. The anodes are covered by an enzyme layer including glucose oxidase and an outer semi-porous membrane layer made of a material such as ParyleneTM ("PPX") or ChronoflexTM AR ("CAR").
In another embodiment of the invention, the sensor body contains 15 a plurality of electrode pairs, each pair including an anode and a cathode. The electrode may take the form of points or lines. In one design linear electrodes are arranged in a "spoke-like" configuration. The electrode pairs preferably aredisposed on both sides of ~e body.
An implantable glucose sensor, according to ffle present 20 invention, may be electrically coupled to a tr~n~mittçr which includes a power source, for example a battery. The tr~n~mitter is capable of converting data signals from the sensor into corresponding radio ~i n~lc A receiver is provided remotely from the sensor for receiving the radio signals. A processor is CA 02237102 1998-0~-08 s connected to the receiver and used to hlLel~l~t the radio ~ ;7 to yield analyte concentration figures.
The present invention also provides a method of m~kin~ an analyte sensor. A substantially disc-shaped body is provided with two opposing sides. At least one cathode and plural anodes are created on each side of the body. A semi-permeable membrane is deposited on the electrodes.
VVhen the method is employed to make a glucose sensor, the enzyme layer including glucose oxidase is created l~etween the anodes and the semi-perme~kle membrane. An interferent retarding layer may be created between the anodes and the enzyme layer.
Description of the Fi~ures Figure 1 is a partially cut-away perspective view of an analyte sensor in accordance with a preferred embodiment of the present invention.
Figure 2 is a cross-sectional view of the sensor shown in Figure 1.
Figure 3 is a top view of an analyte sensor in accordance with a second embodiment of the present invention.
Figure 4A is a top view of an analyte sensor employing linear electrodes in accordance with a third embodiment of the present invention.
Figure 4B is a partial cross-sectional view of the sensor shown in Figure 4A.
Figure 5 is a top view of another analyte sensor in accordance with a fourth embodiment of the present invention.
Figure 6 iS a schematic side view of a glucose sensor inclll~ling an rel elll reklldillg layer.
Figure 7 is a schematic flow chart of an analyte monil(~ g system including sensor, electronics, telemetry and com~u~ g components.
Figure 8 is a flowchart of an analyte monitorin~ system including multiple sensors linked in parallel to the same data acquisition and proces~ing components.
Figure 9 is a top view of an implantable unit including a glucose sensor and radio telemetry device.
Figures 10 and lOA are circuit diagrams illustrating cil~cuill~y employed in glucose sensors of the present invention.
Fi~re 11 is a graph demonstrating the results of an experiment conducted to co~ a~ longevity of single and multiple anode sensors.
Figure 12 is a graph illustrating the results of an experiment conducted to co~ al~ sensor performance pre-implant versus post-explant.
Figure 13 is a graph showing the average glucose dose response and repeatability of eight sensors each of which was coated with PPX.
Figure 14 is a graph showing the average glucose dose response and repeatability (n=3) for a sensor coated with CAR.
Figure 15 is a graph presenting the results of an expeIiment conducted to ~lel~;~...;..e the relative response times (T9Os) for eight sensorseach of which was coated with PPX.
-Figure 16A is a perspective view of a disk-shaped implantable sensor with a cir-iun~~ lial polymer matrix for carrying and slowly rçles-~in~
a fibrotic capsule interference inhibitor.
Figure 16B is a cross-sectional view of the sensor shown in 5 Figure 16A.
Definitions An cle~lrode means an electric conductor, which may be an anode or a ca~ode.
An anode is a positively charged conductor.
A cathode is a negatively charged conductor.
A sensor is a device which detects ch~n~es in analyte concentration in a fluid surrounding the sensor. A sensor includes an anode and a cathode, chemically modified and physically arranged to produce electric signal changes which can be illlel~leted by sensing electronics into analyte IS concent~ation changes over a specified concentration range.
An analyte is a molecule of hllele~L in a fh~id ~ oullding a sensor.
An electrometer is a device which senses small ch~nges in cu.lelll and tr~n~1~tes amps to volts.
A transmitter or radio telemetry device is a device which t~n~mit~ radio ~ip,n~l~
A receiver is a device capable of receiving radio signals from a A body is a housing for supporting and co.~ sensor components.
A semi-permeable membrane or analyte selective coqtin~ is a material which pe~nit~ controlled transfer of an analyte through the mslt~ris-l Interfering substances are molecules in the fluid ~ ding the sensor, which are potentially detectable by the sensor possibly c~ in~ an inaccurate or erroneous analyte concentration determination. An interferent 10 lGl~dillg layer is a material employed in a sensor to either physically or chemically neutralize a potential il~Lelreli~lg substance, thereby ~lcv~nlillg the substance frorn .~lLelr~ g with the desired analyte concenl.aL~on del~....il-~tion.
ChronoflexT~I AR ("CAR") is a trade name for a carbonate based 15 polyurethane available from Polymedica.
ParyleneTM ("PPX") is a trade name for polyparaxylxylene available from Union Carbide.
Description of the Invention We have invented an analyte sensing system including an 20 implantable sensor which exhibits significantly improved performance characteristics over a longer functional life in comparison to prior sensing systems. Our invention has also resulted in improvements which are useful in non-implantable sensors and other sensing applications. The rnodel for CA 02237102 1998-0~-08 illustrating important principles of the present invention, as discussed in detail below, relates to implantable glucose sensors.
Prior implantable glucose sensors do not function s~ f~ctorily over a long enough period to justify the cost and complications of impl~nt~tion.We have observed that increasing the number of anodes, or electrode pairs, or total number of sensors connected in parallel, and by distributing the anodes on~li~elelll sensing faces of one or more sensors, greatly enhances the functionallife spall of an implantable glucose sensing system. Our experiments confirm that redllnd~ncy enhances sensor unit function. Other problems with prior electrochemical glucose sensors relate to electrical drift and instability. The rerllln~l~ncy of ~e present invention, i.e., multiple anodes or multiple sensorsdistributed on multiple faces of one device, appears to significantly reduce such drift. A possible reason for this is that each individual sensing unit may have its own fund~ment~l instability, and that by incorporating multiple sensing units into a single system, an averaging effect tends to cancel out random drift associated with individual sensors.
Figures 1 and 2 illustrate a disc-shaped glucose sensor which has t~,vo opposing faces, each of which has an identical electrode configuration.
One of the faces can be seen in the partially cut-away perspective view in Figure 1. Sensor 18 includes a disc-shaped body 20. On planar face 21 of sensor 18, four pl~tinllm anodes 22 are symmetrically arranged around centrally disposed silver chloride cathode 24. Each anode 22 is covered by an enzyme layer 25 including the active enzyme glucose oxidase and stabilizing compounds such glutaraldehyde and bovine serum albumin (BSA). A
semipermeable membrane layer 26 covers all of the electrodes and individual enzyme layers. The thickness and porosity of membrane layer 26 is carefully controlled so as to limit diffusion and/or transport of the analyte of i~lle~e~l5 (glucose) from the surrounding fluid into the anode sPn~in~ regions. The meçh~ni~m of selective transport of the analyte of i~ re~L through the membrane may involve one or more of the following principles: molecular size exclusion, simple mass transfer, surface tension phenomena and/or other chemically mediated processes.
A cross-section of sensor 18 is shown in Figure 2. Sensor 18 has a plane of symmetry SS which is normal to the plane of the figure. Under face 31 of sensor 18 anodes 32 are spaced equidistantly apart from ca~ode 34.
Enzyme layers 3~ cover anodes 32. A semipermeable membrane 36, l l~rably PPX or CAR, covers the enzyme layers and electrodes. Each of anodes 22 and 15 32 are connected to a common anode wire 36 which leads out of the sensor for electrical connection to an electrometer. Similarly, each of cathodes 24 and 34 are connected to a common cathode lead 38 which leads out of sensor 18 for electrical connection to the electrometer.
Figure 3 shows an ~ltern~tive embodiment of ~e invention in 20 which a plurality of electrode pairs are presented on bo~ sides of a disc-shaped sensor. Only one side of the sensor is shown in Figure 3. The enzyme and semipermeable membrane layers are removed to permit viewing of the elec~ode confi~lration. Sensor 50 (an "8-in-1 sensor") includes eight electrode pairs 52, only four of which are shown distributed around surface 53 of sensor 50. Each electrode pair 52 includes an anode 54 spaced apart from a cathode 56. Similar to the first embodiment described, all of anodes 54 are linked to a common anode wire (not shown) which ext?n(ls outside the body of sensor 50.
5 All of cathodes 56 are connected to a common cathode wire which e~t~n~l~
outside sensor 50. The anode and cathode wires leading out of sensor 50 are c~ ally connected to an electrometer.
Figures 4A, 4B and S illustrate a different type of aIlode and cathode configuration in which each electrode is exposed along a linear path on the sensor snrf~ce. In Figure 4A, sensor 60 is formed with troughs 62 and 64 that intersect at right angles in the center of the sensor surface. Wi~in trough62 linear anode 66a runs parallel to linear cathode 68a. Similarly, in trough 64linear anode 66b runs parallel to linear cathode 68b. The electrodes are in~ll~tell from each other in the junction area 69 where the troughs intersect.
15 Figure 4B shows a cross section through trough 62 in the junction area of thesensor. Trough 62 has a corresponding trough 70 on the opposite side of sensor 60. In trough 70, linear anode 72a runs parallel to linear cathode 72b. Anodes 66a and 72a are both connected to common anode wire 74. Linear cathodes 68a and 72b are connected to cornmon cathode wire 76. Anode wire 74 and 20 cathode wire 76 lead out of sensor 60 for connection to an electrometer. The troughs are ~l~felably filled with an electrolyte gel.
As shown in Figure 5, the concept of employing linear electrodes across opposing faces of the sensor can be extended to provide more electrode sensing area or "spokes". In Figure 5 sensor 80 is essentially the same as sensor 60 (Figures 4A and 4B) except that it has two additional troughs, each C~ g another pair of parallel linear electrodes. Sensor 80 inçlll(les trough82a, 82b, 82c and 82d, all of which intersect in the center of sensor 80. Each 5 of the troughs 82a-d contains a pair of linear electrodes (anode and cathode) encased in electrolyte gel. All of the linear anodes in sensor 80 are connected to a common anode wire, and all of the linear cathodes are connected to a common cathode wire. Other anode p~q1tern~ which function effectively include a circle, concentric circles or a spiral.
Figure 6 shows schematically a cross section ~hrough multiple layers on one side of a sensor. Sensor 100 includes an electrode configuration similar to the embodiment illustrated in Figure 3. Electrode pairs 102a-102d each include a cathode 104a-104d and an anode 106a-106d, respectively.
An electrolyte gel 107 surrounds the anode-cadlode pairs, ~hus lS providing a faster and more sensitive response to changes in glucose concentration. The gel may be produced from methacrylate compounds or from collagen. For exarnple, a methacrylate compound may be dissolved in an organic solvent and then deposited around the anode-cathode pairs. The solvent is then evaporated. Phosphate buf~ered saline with KCl is then added 20 to the gel to swell the methacrylate compound. In the 8-in-1 embod~ment (Figure 3), ~e electrolyte gel is placed over the surface of the electrode pairsand/or the gel is injected into the drilled cylinder in which the electrode pair is sihl~te~l Similarly, in the radial spoke-type embodiments shown in Figures 4 and 5, the troughs may be filled with electrolyte gel.
Layer 110 is deposited immediately on top of the electrodes for the purpose of ...i~i...i7ing or avoiding ~ e.relcnce due to the presence of 5 illl~. r~. ;..~ substances which may be present in the sample fluid. Enzy~ne layer 112 is deposited on top of interferent rela dillg layer 110. Enzyme layer 112 includes, in the case of a glucose sensor, glucose oxidase, and is applied in a solution of glutaraldehyde and bovine serum albumin (BSA), either by pl~c~ment of a drop over each electrode pair, or by dip-coating the entire 0 sP-n~in,~ unit, or by spin-coating. Semi-permeable membrane 114 is deposited on top of enzyme layer 112 for the purpose of controlling diffusion of glucose from the sample fluid into the electrode region of the sensor. PPX at a thickness of about 3,000-6,000 angstroms works well for this purpose. The plerel~c;d thickness of PPX layer 114 is 4,000-5,000A. Other suitable materials 15 for semi-pennç~hle membrane 114 include CAR and polyul~ll,anes such as TecoflexTM, TechothaneTM, CarbothaneTM and CookTM composite.
A number of interferents which exist in hurnan plasma, can be oxidized at the anode when connected to sensor electronic CilCuiLl~, thus regi~terin~ a cullclll which interferes with the signal of interest, i.e., signal 20 generated due to the presence of glucose. Potential interferents include, for example, uric acid, ascorbic acid and the common ~n~l~çsic drug ace~ ophen. Interferents tend to pass freely through semi-permeable membrane 114 and en_yme layer 112. The compounds can be blocked from re~chin~ ~e electrodes by il~ rel~"g leL~dillg layer 110 which has a pore size big enough to allow diffusion of hydrogen peroxide (H202), a product of glucose oxidation in layer 112, yet small enough to exclll.le compounds such as uric acid, ascorbic acid and acetaminophen from re~chin~ the electrodes.
S A ~.erelled material for the interferent ~elaldi-lg layer 110 is PPX. PPX is a hydrophobic compound which is applied to the substrate in a vacuum deposition chamber. The deposition process can be carefully re~ te~1 to form an i,.le,rel~lt r~l~dillg layer of precise thickness (5,000-8,000A) prior to depositing the enzyme layer. CAR also appears to be a suitable material for use as an illlelrele~ll lc~ling layer.
Paired sensors can also be used to provide an ~ltçrn~te method of avoiding illlelreling cull~nl~ from oxidizable, non-glucose compounds. For example, a first sensor is a standard sensor with glucose oxidase. The first sensor measures glucose and hllelrelillg compounds. A second sensor is the lS same as the first sensor except it does not have glucose oxidase and ~us detects onlythe i~ lre~ g compounds. The m~gni~lcle ofthe ~ nl from the second sensor is subtracted from the magnitude of the current from the first sensor to yield a signal which represents the glucose concentration independent from f l re~ ;.... ........g substance concentrations.
The sensor design~ described above can also be modified so that the electrodes detect fluctuations in oxygen concentration which is relatable toglucose concentration. In this approach, the sensor monitors oxygen disappearance in~te~rl of hydrogen peroxide appearance. First, the polarity is changed so that the pl~*mlm electrodes (previously referred to as "anodes") become negatively charged with respect to the silver chloride (previously ,erel..,d to as the "cathode"), i.e., the pl~timlm becomes the cathode and the silver chloride becomes the anode. Second, a membrane is deposited 5 imme~ tely on top of the cathode and anode which is permeable to oxygen but not to larger molecules. The outer membrane and the enzyme layer remain the same. In ~his configuration, glucose concentration results in a decrease in oxygen concentration at the negatively charged electrode.
Another embodiment of the invention has a modified outer 10 memhrane. It is possible that functional longevity of implantable sensors is limited because the outer membrane tends to become "fouled," i.e., plugged or covered by molecules and/or other cellular materials. Accordingly, one adaptation of the invention employs a changing membrane so that the outer membrane can be renewed over time without disrupting operation of the sensor.
15 In the modified sensor, the outer membrane is a solid sheet which can be moved across the face of the sensor where the electrodes are exposed. For example, the membrane can be transferred from one roller to another roller analogous to the way film is transferred inside a camera. A drive meçh~ni~m such as a small motor may be included in the implantable unit for driving the 20 rollers.
Figure 7 shows schem~tically how an implantable glucose sensor is connected in a glucose monitoring system 120. Electrodes in sensor 122 are polarized by polarizing circuit 124. Polarization of the sensor electrodes may CA 02237l02 l998-05-08 be con~ L or pulsed. Our experiments have shown improvement in sensor p~rfonn~nce stability, i.e., m~ p sensitivity and mi~ drift, when polarization is pulsed. For example, polarization of the sensing electrodes can be pulsed ~lt~rn~tely on and off at intervals of 15 milliseconds. It may also be5 advantageous to ~ltern~te polarization, i.e., switch the charge of each electrode at regular intervals.
Sensor 122 is connected to electrometer 126 which senses small ch~n~s in ~;u~ l and tr~n~l~tes amps to volts. Voltage signals from electrometer 126 are telemet~y conditioned and conveyed to tr~n~mitt~r 130 for 10 radio L~ --ic~ion. All of the components within box 132 are implanted as a single unit in the patient.
Fxtern~lly, radio signals from tr~n~mitter 130, in(lic~tive of glucose concentrations in the patient's blood, are transmitted to receiver 134.
Receiver 134 may be connected to monitor 136 for data monitoring. The same 15 receiver co~ ult;~ or another colll~ul~l 138 may be used to analyze ~e raw data and generate glucose concentration information. A printer 140 connected to col"~ule, 138 generates hard copies of analyzed data.
The concept of including multiple electrode pairs within a single sensor can be çxtçnclerl to an embodiment where separate sensors are implanted 20 and commonly linked to a single electrometer as shown in Figure 8. For çY~mI~le, eight implantable sensors 150 can be implanted in a patient and linkedto a single electrometer 152 and tr~n~mitttor (not shown). Tr~n~mit~e~l signals are received by data acquisition adaptor 154 and acquisition co~ ulel 156. By increasing the nllmber of sensors the overall precision, accuracy and longevity of the system can be greatly enhanced. If one or more anodes (or sensors) fails,the others still provide sufficient data sensing capacity so that the entire unit continlles to perform s~ticf~ctorily. Various algorithms or averaging protocols can be used to process the multiple data streams.
Figure 9 shows scllem~tically the components of an implantable unit in a glucose sensing system. Implantable unit 160 includes disc-shaped glucose sensor 162 which is connected to electrometer and teleme1Ty conditioning package 164 via anode wire 166a and cathode wire 166b. Radio signals derived from the raw current signals are tr~ncmit~e~l from ~n~
element 168.
Circuitnl Figure 10 shows custom circuilly structure employed in a glucose sensing system of the present invention. Shown generally at 210 is a glucose servotr~ncmitter suitable for implementation with the present invention.
Servotr~ncmi1tçr 210 is configured for tr~n~micsion of data which is indicative of a sensed enzy~natic reaction to a remote receiving source for subsequent processin~ the sensing and conveyance of such data being described in detail below.
As shown, servotr~ncmittçr 210 includes a sensor 212 (also referred to as a two-electrode sensor) operatively connected between a voltage rerelence source 214 and an arnplifier circuit 216. The output of circuit 216 isl,~eled at 218 and subsequently provided to a voltage-to-frequency circuit 220, which in the preferred embodiment includes a CMOS 7555 circuit indicated at 220a c~ nfi~lred with a resistive and capacitative network which includes two resistors (R9 and R8) and a capacitor (C2). Utilization of CMOS
for ~le~igning circuit 220 has been found ideal due to its low power con~nmrtion aspects which results in longer battery life. The output tf~rminAl of circuit 220 is connected via line 222 to an AC-coupled trAncmi1t~r 224 (also referred to herein as a ~ )i",;l~er) for trAn~mi~ion of data to an çxl~."~1 receiving source.
Discussing the above servotr~n~mi~t~r in more detail, circuit 210 is configured for detecting electrons which are generated duling an enzymatic reaction, and conveying data which is representative of such detected electrons to an e~t~rnAl source for subsequent processing. More specifically, sensor 212 inclll~les two electrodes, a cathode 212a and an anode 212b. Ca~ode 212a is connected to voltage reference source or circuit 214, and anode 212b is connected to amplifier circuit 216. Voltage reference circuit 214 is made up of three resistors R4, R5, and R6 and a 1.2-volt Zener diode Zl. Resistor R6 is connected at one end to a negative voltage potential, and at the other end to diode Zl and resistor R4. The other end of resistor R4 is connected to resistor R5, which in turn is connected to diode Zl as shown. The common node between resistors R4 and R5 is connected to cathode 212a.
Anode 212b is connected via resistors R3, R2, to the inverting te. ,~ Al of amplifier 216a, and a capacitor Cl is connected between resistors R3, R2 and ground. The non-inverting tçrminAl of arnplifier 216a is tied to ground. A resistor Rl is connected between the output of amplifier 216a and its inverting termin~l to provide negative fee~lb~ck The output of amplifier 216a is connected to the non-inverting lf...~ of amplifier 218a, the output of which is connected to the illv~llhlg tçrmin~l in a voltage follower configuration for buffering the output of amplifier 216a. A resistor R7 is connected between the output of amplifier 218a and the trigger t~rmin~l 2 of the CMOS 7555 circuit.
The CMOS 7555 is configured, with its ~ttçn(l~nt resistive and c~p~cit~tive network, as a voltage-to-frequency converter whose output frequency is proportional to its input control voltage. Referring more specifically to the 7555, it may be seen that reset terminal 4 is connected to tertnin~l 8, both of which are connected to a voltage potential which may be referred to as VCC. A resistor R8 is connected between reset terrnin~l 4 and discharge tçrmin~l 7. A resistor R9 is connected between discharge tçrmins~l 7 and the threshold tennin~l 6. A capacitor C2 is connected between trigger t~rrnin~l 2 and ground. Output tçrmin~l 3 is connected to the AC-coupled tr~n~ e. 224 for tr~n~mi~ion of data to a remote location for processin~
Discussing the operation of the above-described glucose servotr~n~mitter, it will be understood that voltage ~ lce circuit 214 develops a potential of -0.6 volts which is used by sensor 212 to cause electrons produced in the vicinity of the sensor to flow, in the form of a generated current, with amplifier circuit 216, which includes operational amplifier 216a configured for feedback as described above. The output of amplifier 216a is a voltage which is buffered at 218 by operational amplifier 218a, the voltage output of which controls ~equency for the tngger te~nin:~l of the CMOS 7555 through resister R7 and frequency selection c~ y C2, R8, and R9. The output tt~rmin~l 3 of the CMOS 7555 is connected, via line 222, to 5 tr~n~mitter 224 for tr~n~mi~ion to an extern~l source.
It will be appreciated that the above-described 7555 configuration c~llv~lL~ ~he ou~ut of buffer 218 into a frequency which is detr....;I.ed by lhevoltage at threshold tt rmin~l 6. The 7555 serves two functions in ~e above configuration which are necessary for the tr~n~mi~sion of sensed data to a remote location for processing. First, the 7555 provides a 15-msec pulse to key tr~n~ 224, ~ereby hlrnin~ it on and off in accordance with prac~ces which will be understood by those of skill in the aIt. Second, the 7555 is operable for voltage-to-frequency conversion, which is a me~ rement of sensor response. This dual function enables the afor~men1ioned data tr~n~mi~sion in a manner which will be understood by those of skill in the art.
Preferred component values (resistive and capacitative values) for the above-described servotr~n~mit~er 210 are as follows~ for voltage ence circuit 214: R4 = 1 meg ohm; R5 = 4.7 meg ohm; and, R6--470 kohm; (2) for amplifier circuit 216: Rl = 500 meg ohm; R2 and R3 = 499 kohm; and, Cl = 10 pf; (3) for converter circuit 220: R9 = 180 kohm; R8 = 1 meg ohm; and, C2 = 1 microfarad; and, (4) R7 = 4.7 meg ohm.
The above system is referred to as a "two-electrode" system bec~llse of the fact that two electrodes are utilized (the anode and the cathode) CA 02237102 1998-0~-08 in the sensing of electrons produced during a particular enzymatic reaction.
Another system which is suitable for sensing produced electrons and conveying data relative to such sensed electrons is a so-called "three-electrode" system which is shown in Figure lOA and described briefly below.
In Figure lOA, like or similar elements of the three-electrode glucose serVotr~n~m~ r 210 are labeled to correspond with the two-electrode elementc appearmg in Figure 10. The Figure shows a sensor 212, a voltage ~cfelellce source 214, a voltage-to-frequency converter circuit 220, and a tr~n~mittçr 224. Voltage refclcllce source 214, voltage-to-frequency converter circuit 220, and tr~n~mitter 224 will not be described because the operation of those elements is the same as, or similar to the operation of such elements as they appear in Figure 10.
Sensor 212 in Figure lOA varies somewhat from its Figure 10 coulllc~ l. Such variations take into account some observations regarding current and voltage control which have been made with respect to the two-electrode system described above, and improve somewhat, the control of such parameters. The three electrode sensor, set forth at 212, includes a counter electrode 212a (which may be formed from silver), a common return electrode 212b (also referred to as a worlcing electrode and which may be formed from pl~1imlm), and a voltage probe 212c which may also be termed the lcre~cnce electrode (and which also may be formed from pl~tinllm). Two operational amplifiers 212d, 212e are provided and operatively coupled to the electrodes as show4 in a configuration which provides greater current and voltage control which, in turn, assists in m~ .g the integrity of the electrodes' sensitivity and the ability of the same to detect a produced cu~ which is indicative of an er~natic reaction. The control is effectuated in a clamped, controlled m~nner.
The three electrode sensor 212 is shown in Figure 10A.
Amplifier 212d m~int~in~ a voltage which is the sarne as the reference voltage of -0.6 volts between the reference and working electrodes 212c, 212b respectively. This is accomplished by varying the current at the counter electrode 212a which is in the feedback loop of amplifier 212d. Amplifier 212e 10 m~int~in~ the working electrode 212b at virtual ground converting the C-LIlc.to an output voltage, which is buffered at 212f and provided to CMOS 7555 co~ l circuit 220 for conversion from a voltage to a frequency (in a rn~nner described above), the converter circuit thereafter triggering tr~n~mi~t~?r 224 is a pre~ e....il~ble fashion to transmit sensed data indicative of an enzyma~c 15 reaction to a remote location for processing.
Decreasin~ Fibrotic Capsule Interference One of the primary reasons why a subcutaneously-implanted sensor eventually loses its ability to measure the concentration of arl analyte of interest is that a collagenous capsule forms around the sensor. The capsule 20 eventually loses vascularity and becomes thick and fibrous, ~ereby subst~n*~lly blocking the sensor from accessing the analyte present in blood.
There are at least two promising approaches for ~
fibrotic capsule illle,rerence with analyte detection, thereby ~xtt~.n~lin~ longevity CA 02237102 1998-0~-08 of an implanted sensor. First, it is possible to prevent or retard capsule form~tion by slow controlled release of certain collagen deposition inhibitors.
Drugs which inhibit collagen formation can be incorporated in a polymer matrix which allows slow release of the drug locally to achieve the desired 5 effect without c~llcing adverse distant systemic effects in the animal or hnm~n For example, collagen inhibitors which can be used for this purpose include corticosteroids such as ~lex~methasone, relaxin and ~;~mm~ t~lr~r~n. A
efelled polymer material for car~ying and controlling slow release of the drug is polydimethylsiloxane. Corticosteroids can be impregnated in a 10 polydimethylsiloxane matrix so as to provide relatively long-term, slow rele~e of the corticosteroids in the surrounding tissue. It is important, however, that~le~r~methasone be released in small doses in order to avoid iatrogenic Cushing's syndrome, which is a serious illness caused from systemic excess of corticosteroids. If corticosteroids are released from a sensor for a prolonged 15 period, for example, more than two weeks, we recommend that a patient's serum be tested in order to confirrn that adverse systemic effects are avoided.
Another approach for minimi7:in~ fibrotic capsule i~le-r~;~ellce with sensor performance, i.e., increasing sensor longevity, is to promote vascularity in the capsule so that the sensor can continue to have access to 20 blood analytes. Accordingly, vascular growth factors can be incorporated in amatrix around the sensor so that the grow~ factors are slowly released into the ou~lding tissue. The released growth factors enhance capillary growth in the coll~gennus capsule which forms around the implanted sensor. Retention of capillary perfusion by the capsule enhances sensor function by continuously providing the sensor access to the patient's blood analyte. Examples of c~pill~ growdl factors include vascular endothelial growth factor (VEGF) and endothelial cell growth factor (ECGF). Polymer materials which are capable of S slowly releasing polypeptide factors such as ECGF and VEGF include poly-l-lactic acid and poly glycolic lactic acid. As with ~e steroid approach, the grow~ factor dosage, i.e., quantity and rate of release, must be carefully controlled so that the growth factor's effect is local, not systemic.
A method of employing steroids or growth factors to ~ e or 10 avoid fibrotic capsule interference with sensor performance, is to provide for the active agent's slow release from the perimeter of ~e disk sensor. For example, as shown in Figure 16A, glucose sensor 300 has a carrier layer or matrix 302 such as a tape made of or cont~inin~ polydimethylsiloxane g~-~te-l with dexamethasone. Tape 302 is attached to outer perimeter edge 304 of disk-shaped housing or body 306 of sensor 300. The width of tape 302 is subst~nti~lly the same dimension as the width of edge 304, i.e., thiclmess ofhousing 306, so that the steroid is released on or near both faces of ~e sensor.Time-release steroid compositions have been lltili7e~1 in the past for other puIposes. For example, U.S. Patent No. 5,265, 608 to Lee et al., the 20 entire content of which is hereby incorporated by reference, discloses a steroid elll~lin~ electrode in which ~lex~methasone is incorporated in a polymer matrix which permits slow controlled release of the steroid to control infl~mm~tion~
init~tiQn and swelling in connection with a device such as a pacçm~ker.
However, no one has previously employed a time release corticosteroid matrix for inhibiting collagen formation on an implantable analyte sensor.
Sensin~ Other Analytes With minor modifications, the sensor design~ described above 5 may be used to detect analytes other than glucose. By ch~n ing the specific type of enzyme which covers the anode, the sensor can be used to me~nre many compounds. Several examples appear in Table 1 below.
ANALYTE ENZYME
glucose glucose oxidase glucose hexose oxidase lactate lactate oxidase l-methionine l-amino acid oxidase l-phenylalanine l-amino acid oxidase d-aap~le d-amino acid oxidase d-~h~ te d-amino acid oxidase urate urate oxidase ethyl alcohol alcohol oxidase methyl alcohol alcohol oxidase cholesterol cholesterol oxidase ascorbic acid ascorbate oxidase In addition to measuring analytes in body fluids, sensors of the present invention can be used to measure the concentration of substrates in other fluids, for example, fruit and vegetable juices, wine, yogurt, etc.
Construction of a Glucose Sensor A preferred sensor is constructed of epoxy resin in a disc shape, 1.3-cen~imet~ors in diameter and 0.2-centimeters in height. Four 36-gauge pk77inllm wires le....i~ e peripherally on one face of the disc (in holes drilled in 5 the resin) and service hydrogen peroxide-sçn~in~ anodes. A solid silver cylinder, 0.7-centimeters outside diameter (the cathode), is secured by epoxy resin in the center of the disc. A layer of silver chloride can be deposited onto the surface of the silver by one of several processes. The sensor is ~lc;rel~blydouble-sided, which may be, for example, two of the four anode sensors 10 c~mfi~77red "back-to-back", m~7kin~ a sensor composed of four anodes and one cathode on each face of the sensor.
Anode and cathode recording wires terminate in an amplifier and polarizing voltage source. An electrometer converts the current signal to a voltage signal and applies a con~t~nt pol~ri~ing voltage of 0.60 V to the 15 electrodes. Output from the amplifier is routed both to a digital volt meter (Micronta 22-185A, Tandy Corp., Fortworth, TX 76102) and to a strip chart continuous chart recorder (Gould Inshuments Model No. 11-2963-300, Valley View, OH 44126). The signal can also be routed directly into a cc,~ uler by use of a data acquisition board. All of these electrical components can be 20 mini~hlri7ecl without ~ltering the function.
A working 8-anode sensor (which has been demonstrated to respond to peroxide) then is selected for testing. The sensor is s~ntle-l first with 600 and 1500 grit wet-or-dry, then followed by a polishing with 2000 grit wet-or-dry. The sensor is rinsed thoroughly in a stream of deionized water (DW) followed by blow drying in a cold nitrogen stream. The sensor is then immersed in an acetone bath and vigorously twirled for 20 seconds to remove any solvents or oils from the surface. The sensor is withdrawn from the 5 acetone bath and is immediately rinsed in a DW stream. The sensor is again blown dry in a cold nitrogen stream, and contimles to dry in room air for another 30 minlltes-If it is desired to include an intel~r~nt reLarding layer, then alayer of PPX (or other suitable membrane material) approximately 5,000-10 8,000A thick is deposited directly on top of the anodes before depositing anyenzyme.
The sanded, cleaned and dried sensor (with or without i~ ,relelll ,e~dillg layer) is enzyme activated with a Glucose Oxidase (GO) ---Bovine Serum Albumin (BSA)--- Glutaraldehyde (GA) matrix prepared from mixin~
two parts GO+BSA (20mg GO + 5mg BSA with 0.5 gram DW) plus one part GA (2.5% GA diluted with DW). Approximately 2.5 ~11 of this solution is applied via pipette directly to each anode. The solution is allowed to dry in room air for one hour. The sensor is then immersed in DW for 15 minl!tes to remove excess GA, rinsed briefly in a DW stream, and blown dry in a nitrogen 20 st~eam. The sensor continlles to dry in room air for one hour, after which spin-coating with PU (Tecoflex, Tecothane, Cook composite, or CAR) or vapor deposition w~th PPX (thickness = 3000-SOOOA) is carried out.
Further mini~tllrization of the glucose sensor, as described above, will not adversely effect performance of the unit.
Testin~. Connectin~ and Implantin~ Sensors Sensors m~mlf~ctured as described above, are tested the day after S they are made by applying a polarizing voltage of 600 mV. The voltage output should stabilize after a one to two hour immersion in a tempc.alule controlled PBS solution (37~C) in the laboratory water bath. The sensor is tested in st~n-l~rd glucose solutions prepared by adding glucose to PBS so that ~he resnlting test solutions (G=glucose), are concentrated in mg/dl as follows:
G(0), G(100), G(200), G(300), G~400) and G(500); and in millimolar concentrations as follows: G(0), G(5.6), G(l l. l), G(16.7), G(22.2) and G(27.8). The first data point is collected while the sensor is still immersed inPBS and represents the baseline output. After noting the output value, t~e sensor is moved to G(100) for ten minlltes. The process of measuring ~e speed 15 with which the sensor responds to the increase in glucose allows calculation of T90 (defined T90 below). The sensor is moved to the G(200) standard and dle ten minute output value is collected from this standard. All of the following uul~uls are collected in ascending order in the same m~nner.
An implantable sensor has to satisfy three criteria: (1) it must 20 have a T90 of less than three min~ltes; (2) it must be dose responsive in theglucose concenl~lion range of 40-400 mg/dl; and (3) it must have adequate sensitivity. The T90 value is measured by using the continuous sensor readout provided by the data acquisition system. The point at which 90% of ~he m5 xi~ -- output is reached (after ch~n~ing from the zero glucose level to the 100 mg/dl level) is recorded as the T90.
A sensor that is acceptable for implant must also be dose responsive, preferably substantially linear over the glucose concentration of 40-5 400 mg/dl. Minor to moderate non-linearities can be mathematically corrected to allow estim~*on of glucose level from sensor output data.
If a sensor meets all the previous criteria, it is attached to a tr~n~mit~r. For example, a suitable transmitter may be obtained from Mini-Mitter which has a custom-built interface circuit between the tr~n~mitter and 10 the sensor. The tr~n~mi1ter should have a battery pack which is fully charged.
The sensor can be implanted in the body of ~nim~l~ or hnm~n~.
The sensor can be implanted subcutaneously, in an artery or vein, uscularly~ intraperitoneally, in the brain or cerebrospinally. The prerelled location is subcutaneous. The sensor can also be used in vitro, for 15 example, in a laboratory to measure glucose concentration or other substrates or analytes in a liquid media.
The transmitter and sensor package are tested in vitro the day of the planned implant procedure. If the results are s~ti~f~ctory (T90 less than 3 les, s~ti~f~ctorily dose-responsive, adequate sensitivity), then the unit is 20 sterili~e~l, rinsed in sterile saline, and implanted subcutaneously in the recipient (after the a~rol~liate preparation and anesthesia procedures).
Experiments Experiment 1 We compared the performance of sensors with one anode to the performance of sensors with four anodes. Twelve one-sensor anode sensors 5 were constructed subst~nti~lly as described above. These sensors were sirnilarto the ones shown in Figure 1 except they only included one anode in~te~d of four, and they only had electrodes on one side of the disc-shaped sensor. All sensors in ~is experiment were dip-coated with polyurethane (Cook Composites) instead of parylene. Twenty-four-anode sensors were constructed 10 the same as the one-anode sensors except that they included four anodes on one face of the sensor subst~nti~lly as shown in Figure 1.
The sensors were implanted in rats. Glucose dose response data was collected for each of the sensors at frequent time points after impl~nt~tionun~l the given sensor failed to perform s~ti~f~ctorily. For each sensor, the last 15 check point at which the sensor performed adequately, d~Lt;~ ed the functional life of that sensor.
Figure 11 shows the results of this experiment. The average longevity for the one-anode sensors was about 4 days. In contrast, the average longevity for the four-anode sensors was about 28 days. This is a highly 20 significant improvement in the functional life of an implanted glucose sensor, which we attribute to the increased number of anodes.
CA 02237l02 l998-0~-08 Experiment 2 The purpose of this experiment was to determine in vitro the perfonn~nce capability in sensors which had failed in vivo. In this experiment eight of the four anode sensors used in Experiment 1 were tested before 5 implantation (pre-implant), and then tested again after eventually failing to perform and being removed (post-explant) from the rat.
The results of the ~x~tlilllent are shown in Fig~re 12. In Figure 12 (and Figures 13 and 14), the "Normal Range" includes glucose concentrations which are typically observed in the normal population. The 10 "Dynamic Range" includes the Normal Range plus abnormally high and low glucose concentrations which should be measurable with a glucose sensor. The results show that in vitro the sensors performed as well post-explant as they did pre-implant. This result demonstrates that failure of the sensors in vivo is notdue to inactivation or loss of the glucose oxidase enzyme. We noted that over 15 time in vivo a cellular coat tends to envelop the sensor. Before performing dle post-explant testing on the sensors, the coats were removed. This suggests that the cellular coat which develops around the sensor may be involved with c;vt;nlual sensor failure. Since the cellular coat is relatively non-uniform, it is possible to theorize that one of the reasons why longevity is increased with 20 multiple anodes is that the probability of m~in~ining one or more anodes under a portion of the coat which is minim~l enough so that the sensor still performs,is increased by increasing the number or surface area of sensing anodes.
Experiment 3 II1 Experimçnt~ 1 and 2, the sensors were dip-coated in polyurethane (Cook Composites). We subsequently discovered that ~ irO~ y and overall performance of the sensors can be improved by using PPX as the S outer coat or semipermeable membrane. The purpose of this experiment was to llemo~ e glucose dose response and repeatability for eight sensors, each of which was coated with PPX at a ~hickness of a~ro~ lately 3800A. As shown in Figure 13, we observed a dose response appro~rhing linearity in the useful me~nrement range. Test repeatability was also improved with the PPX coated 10 sensors, as shown by the smaller standard deviation margins in colll~ison to those shown in Figure 12.
Experiment 4 This experiment was similar to Experiment 3 except instead of using PPX as the semipermeable outer membrane, CAR was used. Eight-15 percent CAR was spin-coated over the sulface of the sensor for 2.5 ~ es at 4,000 RPM. The sensor was tested in vitro at various glucose concentrations in 3 l~ccessive runs. The data is shown in Figure 14. The dose response over the useful me~cllrement range approached linearity with a higher slope in c~ pA.icon to slopes obtained with PPX and dip-coated sensors. We also 20 noted a relatively small standard deviation on repeat tests with the CAR coated sensors.
EXperiTnent 5 This experiment was performed in vitro with PPX coated eight anode (four on each side) sensors to dele~ e how rapidly the sensors respond to çh~n~,es in glucose concentration (T90). Six sensors were constructed with PPX outer coats of 3000-5000A. Results of this experiment are shown in Figure 15. Each of the sensors responded with a T90, i.e., tirne to reach 90-percent of ~ ;m~te current output for a given change in glucose concentration, in less than one minute. This is a faster response time than we had observed previously with polyurethane dip-coated sensors.
IN ANALYTE CONCENTRATION
Field of the Invention The invention relates to electroçhemic~l systems for me~c~rin~
analyte concentration. In particular, the invention involves a sensor including electrodes under a semi-permeable membrane for mol~.lolmg analyte conc~llLI~lions in fluids surrounding the sensor.
Back~round There are many instances when it is necessary to monitor ~he concelll~alion of molecules ("analytes") in a fluid. For example, glucose levelsmust be frequently monitored in persons with diabetes so that a~lopliate doses of insulin can be ~rlminictçred in a timely m~nner Many other analytes are me~cllred commonly in human blood and in other fluids.
A variety of me~ods and devices for me~cllrin~ analytes in fluids have been devised. One such device, referred to as an electrochemical sensor, typically includes oppositely charged electrodes under a semi-pçrme~ble membrane. Depending on what analyte is being monitored, membranes, enzymes and/or other a~plopliate materials are provided around the electrodes so ~hat analyte reaction and transport from the fluid ~ullounding the sensor i controlled. Oxidative and reductive reactions take place at or near the electrodes, thus c~llcin~ electron potentials mç~cllred as ch~n~es in cu.l~,nL
which may be correlated to the concentration of analyte in the fluid.
Electrochemical sensors have been used to me~cllre glucose in human blood for a long time. Most of these sensors are ~lesigne~l to me~cllre glucose in a blood sample which has been drawn or ext~acted from the patient.
For patients such as people with diabetes who must test blood glucose levels as 5 often as several times per day, the regular blood drawing process (typically by finger tip puncture) becomes quite cumbersome, messy and even painfill. The diabetic must car~y special eqllipmçnt for extracting blood. Some p~tient~ fail to test as frequently as they should because of problems associated with the blood extracting process.
Therefore, it has been recognized for a long time that an impl~nte~l glucose sensor would offer the important advantage of avoiding the need for repeated blood extraction. However, there are other problems which must be addressed with an implantable sensor. First, there must be a mech~nicm for accessin~ raw electrical data generated by the sensor under the 15 patient's slcin. Protruding wires are lm(lçcirable because they are cumbersome, prone to c~llcing infection and sometimes painful. Accordingly, it is preferableto incl~l(le a wireless data tr~ncmission (telemetry) device coupled to the sensor in a single impl~nt~ble unit so that no trans-dermal wires are required.
Second, an implanted sensing unit may cause int~rn~l trauma, i.e., 20 bruising or bleeding from the patient's routine movement or contact with his or her environment especially if the sensing unit is large or thick or if it is geometrically shaped with any sharp points or edges.
WO 97/19344 . PCT/US96/18724 Another problem associated with implantable sensors is that over time (days and weeks) a cellular coat tends to develop around the sensor which may evenhl~lly block the analyte of interest from cont~ctinp; the electrodes, thus c~lcin~ the sensor to fail.
For these reasons, and perhaps other reasons, researchers in the field have been lln~lccessfill in their alle,~ to produce an implantable sensor unit which is capable of functioning s~ticf~ctorily for a suficient period of time to justify the expense and inconvenience of producing and surgically imrl~nting the sensing ha~dw~.
A viable implantable glucose sensor should provide reliable performance for at least 1-2 months, preferably three months or more. During its useful life, the device should generate a predictable dose response over a concentration range of approximately 40 to 400 milligrams per deciliter (mg/dl). The device should exhibit a lag time between a concentration change and the resl.lting signal output of less than 20 min~ltes, preferably less than 10 ...i...,les. The sensor should be relatively insensitive to potential i.~ fe~ g substances such as ascorbic acid and acetaminophen. The device should be relatively accurate for at least several days after calibration (stability). Glucose me~llrement with the sensor should be precise to at least within a~p,ox~llately 20 10 mg/dl. The sensor should be incorporated in an implantable unit which is capable of wireless data tr~n~mi~sion, and which is ~limencioned so as to II.i..i...i7P surgical complication and risk of pain, bruising or other in~çrn~ltrauma.
Summary of the Invention The objectives stated above are achievable with the device and system of the present invention which includes a device for electrochemically s~n~in~ ch~n~es in the conce~ Lion of an analyte of i~llelt;~l.
S In one embo-lirnent of the invention, the device inclll~les a sensor body having two opposing sides. Each side of the body includes at least one, erel~bly several, anode(s~ and at least one cathode spaced apart ~om each other and covered by a membrane which is semi-pçrme~le to the analyte of interest. In a p-e~lled sensor design for me~cllnn~ glucose, plural anodes are 10 disposed on two opposing sides of a disc-shaped sensor body. The anodes are covered by an enzyme layer including glucose oxidase and an outer semi-porous membrane layer made of a material such as ParyleneTM ("PPX") or ChronoflexTM AR ("CAR").
In another embodiment of the invention, the sensor body contains 15 a plurality of electrode pairs, each pair including an anode and a cathode. The electrode may take the form of points or lines. In one design linear electrodes are arranged in a "spoke-like" configuration. The electrode pairs preferably aredisposed on both sides of ~e body.
An implantable glucose sensor, according to ffle present 20 invention, may be electrically coupled to a tr~n~mittçr which includes a power source, for example a battery. The tr~n~mitter is capable of converting data signals from the sensor into corresponding radio ~i n~lc A receiver is provided remotely from the sensor for receiving the radio signals. A processor is CA 02237102 1998-0~-08 s connected to the receiver and used to hlLel~l~t the radio ~ ;7 to yield analyte concentration figures.
The present invention also provides a method of m~kin~ an analyte sensor. A substantially disc-shaped body is provided with two opposing sides. At least one cathode and plural anodes are created on each side of the body. A semi-permeable membrane is deposited on the electrodes.
VVhen the method is employed to make a glucose sensor, the enzyme layer including glucose oxidase is created l~etween the anodes and the semi-perme~kle membrane. An interferent retarding layer may be created between the anodes and the enzyme layer.
Description of the Fi~ures Figure 1 is a partially cut-away perspective view of an analyte sensor in accordance with a preferred embodiment of the present invention.
Figure 2 is a cross-sectional view of the sensor shown in Figure 1.
Figure 3 is a top view of an analyte sensor in accordance with a second embodiment of the present invention.
Figure 4A is a top view of an analyte sensor employing linear electrodes in accordance with a third embodiment of the present invention.
Figure 4B is a partial cross-sectional view of the sensor shown in Figure 4A.
Figure 5 is a top view of another analyte sensor in accordance with a fourth embodiment of the present invention.
Figure 6 iS a schematic side view of a glucose sensor inclll~ling an rel elll reklldillg layer.
Figure 7 is a schematic flow chart of an analyte monil(~ g system including sensor, electronics, telemetry and com~u~ g components.
Figure 8 is a flowchart of an analyte monitorin~ system including multiple sensors linked in parallel to the same data acquisition and proces~ing components.
Figure 9 is a top view of an implantable unit including a glucose sensor and radio telemetry device.
Figures 10 and lOA are circuit diagrams illustrating cil~cuill~y employed in glucose sensors of the present invention.
Fi~re 11 is a graph demonstrating the results of an experiment conducted to co~ a~ longevity of single and multiple anode sensors.
Figure 12 is a graph illustrating the results of an experiment conducted to co~ al~ sensor performance pre-implant versus post-explant.
Figure 13 is a graph showing the average glucose dose response and repeatability of eight sensors each of which was coated with PPX.
Figure 14 is a graph showing the average glucose dose response and repeatability (n=3) for a sensor coated with CAR.
Figure 15 is a graph presenting the results of an expeIiment conducted to ~lel~;~...;..e the relative response times (T9Os) for eight sensorseach of which was coated with PPX.
-Figure 16A is a perspective view of a disk-shaped implantable sensor with a cir-iun~~ lial polymer matrix for carrying and slowly rçles-~in~
a fibrotic capsule interference inhibitor.
Figure 16B is a cross-sectional view of the sensor shown in 5 Figure 16A.
Definitions An cle~lrode means an electric conductor, which may be an anode or a ca~ode.
An anode is a positively charged conductor.
A cathode is a negatively charged conductor.
A sensor is a device which detects ch~n~es in analyte concentration in a fluid surrounding the sensor. A sensor includes an anode and a cathode, chemically modified and physically arranged to produce electric signal changes which can be illlel~leted by sensing electronics into analyte IS concent~ation changes over a specified concentration range.
An analyte is a molecule of hllele~L in a fh~id ~ oullding a sensor.
An electrometer is a device which senses small ch~nges in cu.lelll and tr~n~1~tes amps to volts.
A transmitter or radio telemetry device is a device which t~n~mit~ radio ~ip,n~l~
A receiver is a device capable of receiving radio signals from a A body is a housing for supporting and co.~ sensor components.
A semi-permeable membrane or analyte selective coqtin~ is a material which pe~nit~ controlled transfer of an analyte through the mslt~ris-l Interfering substances are molecules in the fluid ~ ding the sensor, which are potentially detectable by the sensor possibly c~ in~ an inaccurate or erroneous analyte concentration determination. An interferent 10 lGl~dillg layer is a material employed in a sensor to either physically or chemically neutralize a potential il~Lelreli~lg substance, thereby ~lcv~nlillg the substance frorn .~lLelr~ g with the desired analyte concenl.aL~on del~....il-~tion.
ChronoflexT~I AR ("CAR") is a trade name for a carbonate based 15 polyurethane available from Polymedica.
ParyleneTM ("PPX") is a trade name for polyparaxylxylene available from Union Carbide.
Description of the Invention We have invented an analyte sensing system including an 20 implantable sensor which exhibits significantly improved performance characteristics over a longer functional life in comparison to prior sensing systems. Our invention has also resulted in improvements which are useful in non-implantable sensors and other sensing applications. The rnodel for CA 02237102 1998-0~-08 illustrating important principles of the present invention, as discussed in detail below, relates to implantable glucose sensors.
Prior implantable glucose sensors do not function s~ f~ctorily over a long enough period to justify the cost and complications of impl~nt~tion.We have observed that increasing the number of anodes, or electrode pairs, or total number of sensors connected in parallel, and by distributing the anodes on~li~elelll sensing faces of one or more sensors, greatly enhances the functionallife spall of an implantable glucose sensing system. Our experiments confirm that redllnd~ncy enhances sensor unit function. Other problems with prior electrochemical glucose sensors relate to electrical drift and instability. The rerllln~l~ncy of ~e present invention, i.e., multiple anodes or multiple sensorsdistributed on multiple faces of one device, appears to significantly reduce such drift. A possible reason for this is that each individual sensing unit may have its own fund~ment~l instability, and that by incorporating multiple sensing units into a single system, an averaging effect tends to cancel out random drift associated with individual sensors.
Figures 1 and 2 illustrate a disc-shaped glucose sensor which has t~,vo opposing faces, each of which has an identical electrode configuration.
One of the faces can be seen in the partially cut-away perspective view in Figure 1. Sensor 18 includes a disc-shaped body 20. On planar face 21 of sensor 18, four pl~tinllm anodes 22 are symmetrically arranged around centrally disposed silver chloride cathode 24. Each anode 22 is covered by an enzyme layer 25 including the active enzyme glucose oxidase and stabilizing compounds such glutaraldehyde and bovine serum albumin (BSA). A
semipermeable membrane layer 26 covers all of the electrodes and individual enzyme layers. The thickness and porosity of membrane layer 26 is carefully controlled so as to limit diffusion and/or transport of the analyte of i~lle~e~l5 (glucose) from the surrounding fluid into the anode sPn~in~ regions. The meçh~ni~m of selective transport of the analyte of i~ re~L through the membrane may involve one or more of the following principles: molecular size exclusion, simple mass transfer, surface tension phenomena and/or other chemically mediated processes.
A cross-section of sensor 18 is shown in Figure 2. Sensor 18 has a plane of symmetry SS which is normal to the plane of the figure. Under face 31 of sensor 18 anodes 32 are spaced equidistantly apart from ca~ode 34.
Enzyme layers 3~ cover anodes 32. A semipermeable membrane 36, l l~rably PPX or CAR, covers the enzyme layers and electrodes. Each of anodes 22 and 15 32 are connected to a common anode wire 36 which leads out of the sensor for electrical connection to an electrometer. Similarly, each of cathodes 24 and 34 are connected to a common cathode lead 38 which leads out of sensor 18 for electrical connection to the electrometer.
Figure 3 shows an ~ltern~tive embodiment of ~e invention in 20 which a plurality of electrode pairs are presented on bo~ sides of a disc-shaped sensor. Only one side of the sensor is shown in Figure 3. The enzyme and semipermeable membrane layers are removed to permit viewing of the elec~ode confi~lration. Sensor 50 (an "8-in-1 sensor") includes eight electrode pairs 52, only four of which are shown distributed around surface 53 of sensor 50. Each electrode pair 52 includes an anode 54 spaced apart from a cathode 56. Similar to the first embodiment described, all of anodes 54 are linked to a common anode wire (not shown) which ext?n(ls outside the body of sensor 50.
5 All of cathodes 56 are connected to a common cathode wire which e~t~n~l~
outside sensor 50. The anode and cathode wires leading out of sensor 50 are c~ ally connected to an electrometer.
Figures 4A, 4B and S illustrate a different type of aIlode and cathode configuration in which each electrode is exposed along a linear path on the sensor snrf~ce. In Figure 4A, sensor 60 is formed with troughs 62 and 64 that intersect at right angles in the center of the sensor surface. Wi~in trough62 linear anode 66a runs parallel to linear cathode 68a. Similarly, in trough 64linear anode 66b runs parallel to linear cathode 68b. The electrodes are in~ll~tell from each other in the junction area 69 where the troughs intersect.
15 Figure 4B shows a cross section through trough 62 in the junction area of thesensor. Trough 62 has a corresponding trough 70 on the opposite side of sensor 60. In trough 70, linear anode 72a runs parallel to linear cathode 72b. Anodes 66a and 72a are both connected to common anode wire 74. Linear cathodes 68a and 72b are connected to cornmon cathode wire 76. Anode wire 74 and 20 cathode wire 76 lead out of sensor 60 for connection to an electrometer. The troughs are ~l~felably filled with an electrolyte gel.
As shown in Figure 5, the concept of employing linear electrodes across opposing faces of the sensor can be extended to provide more electrode sensing area or "spokes". In Figure 5 sensor 80 is essentially the same as sensor 60 (Figures 4A and 4B) except that it has two additional troughs, each C~ g another pair of parallel linear electrodes. Sensor 80 inçlll(les trough82a, 82b, 82c and 82d, all of which intersect in the center of sensor 80. Each 5 of the troughs 82a-d contains a pair of linear electrodes (anode and cathode) encased in electrolyte gel. All of the linear anodes in sensor 80 are connected to a common anode wire, and all of the linear cathodes are connected to a common cathode wire. Other anode p~q1tern~ which function effectively include a circle, concentric circles or a spiral.
Figure 6 shows schematically a cross section ~hrough multiple layers on one side of a sensor. Sensor 100 includes an electrode configuration similar to the embodiment illustrated in Figure 3. Electrode pairs 102a-102d each include a cathode 104a-104d and an anode 106a-106d, respectively.
An electrolyte gel 107 surrounds the anode-cadlode pairs, ~hus lS providing a faster and more sensitive response to changes in glucose concentration. The gel may be produced from methacrylate compounds or from collagen. For exarnple, a methacrylate compound may be dissolved in an organic solvent and then deposited around the anode-cathode pairs. The solvent is then evaporated. Phosphate buf~ered saline with KCl is then added 20 to the gel to swell the methacrylate compound. In the 8-in-1 embod~ment (Figure 3), ~e electrolyte gel is placed over the surface of the electrode pairsand/or the gel is injected into the drilled cylinder in which the electrode pair is sihl~te~l Similarly, in the radial spoke-type embodiments shown in Figures 4 and 5, the troughs may be filled with electrolyte gel.
Layer 110 is deposited immediately on top of the electrodes for the purpose of ...i~i...i7ing or avoiding ~ e.relcnce due to the presence of 5 illl~. r~. ;..~ substances which may be present in the sample fluid. Enzy~ne layer 112 is deposited on top of interferent rela dillg layer 110. Enzyme layer 112 includes, in the case of a glucose sensor, glucose oxidase, and is applied in a solution of glutaraldehyde and bovine serum albumin (BSA), either by pl~c~ment of a drop over each electrode pair, or by dip-coating the entire 0 sP-n~in,~ unit, or by spin-coating. Semi-permeable membrane 114 is deposited on top of enzyme layer 112 for the purpose of controlling diffusion of glucose from the sample fluid into the electrode region of the sensor. PPX at a thickness of about 3,000-6,000 angstroms works well for this purpose. The plerel~c;d thickness of PPX layer 114 is 4,000-5,000A. Other suitable materials 15 for semi-pennç~hle membrane 114 include CAR and polyul~ll,anes such as TecoflexTM, TechothaneTM, CarbothaneTM and CookTM composite.
A number of interferents which exist in hurnan plasma, can be oxidized at the anode when connected to sensor electronic CilCuiLl~, thus regi~terin~ a cullclll which interferes with the signal of interest, i.e., signal 20 generated due to the presence of glucose. Potential interferents include, for example, uric acid, ascorbic acid and the common ~n~l~çsic drug ace~ ophen. Interferents tend to pass freely through semi-permeable membrane 114 and en_yme layer 112. The compounds can be blocked from re~chin~ ~e electrodes by il~ rel~"g leL~dillg layer 110 which has a pore size big enough to allow diffusion of hydrogen peroxide (H202), a product of glucose oxidation in layer 112, yet small enough to exclll.le compounds such as uric acid, ascorbic acid and acetaminophen from re~chin~ the electrodes.
S A ~.erelled material for the interferent ~elaldi-lg layer 110 is PPX. PPX is a hydrophobic compound which is applied to the substrate in a vacuum deposition chamber. The deposition process can be carefully re~ te~1 to form an i,.le,rel~lt r~l~dillg layer of precise thickness (5,000-8,000A) prior to depositing the enzyme layer. CAR also appears to be a suitable material for use as an illlelrele~ll lc~ling layer.
Paired sensors can also be used to provide an ~ltçrn~te method of avoiding illlelreling cull~nl~ from oxidizable, non-glucose compounds. For example, a first sensor is a standard sensor with glucose oxidase. The first sensor measures glucose and hllelrelillg compounds. A second sensor is the lS same as the first sensor except it does not have glucose oxidase and ~us detects onlythe i~ lre~ g compounds. The m~gni~lcle ofthe ~ nl from the second sensor is subtracted from the magnitude of the current from the first sensor to yield a signal which represents the glucose concentration independent from f l re~ ;.... ........g substance concentrations.
The sensor design~ described above can also be modified so that the electrodes detect fluctuations in oxygen concentration which is relatable toglucose concentration. In this approach, the sensor monitors oxygen disappearance in~te~rl of hydrogen peroxide appearance. First, the polarity is changed so that the pl~*mlm electrodes (previously referred to as "anodes") become negatively charged with respect to the silver chloride (previously ,erel..,d to as the "cathode"), i.e., the pl~timlm becomes the cathode and the silver chloride becomes the anode. Second, a membrane is deposited 5 imme~ tely on top of the cathode and anode which is permeable to oxygen but not to larger molecules. The outer membrane and the enzyme layer remain the same. In ~his configuration, glucose concentration results in a decrease in oxygen concentration at the negatively charged electrode.
Another embodiment of the invention has a modified outer 10 memhrane. It is possible that functional longevity of implantable sensors is limited because the outer membrane tends to become "fouled," i.e., plugged or covered by molecules and/or other cellular materials. Accordingly, one adaptation of the invention employs a changing membrane so that the outer membrane can be renewed over time without disrupting operation of the sensor.
15 In the modified sensor, the outer membrane is a solid sheet which can be moved across the face of the sensor where the electrodes are exposed. For example, the membrane can be transferred from one roller to another roller analogous to the way film is transferred inside a camera. A drive meçh~ni~m such as a small motor may be included in the implantable unit for driving the 20 rollers.
Figure 7 shows schem~tically how an implantable glucose sensor is connected in a glucose monitoring system 120. Electrodes in sensor 122 are polarized by polarizing circuit 124. Polarization of the sensor electrodes may CA 02237l02 l998-05-08 be con~ L or pulsed. Our experiments have shown improvement in sensor p~rfonn~nce stability, i.e., m~ p sensitivity and mi~ drift, when polarization is pulsed. For example, polarization of the sensing electrodes can be pulsed ~lt~rn~tely on and off at intervals of 15 milliseconds. It may also be5 advantageous to ~ltern~te polarization, i.e., switch the charge of each electrode at regular intervals.
Sensor 122 is connected to electrometer 126 which senses small ch~n~s in ~;u~ l and tr~n~l~tes amps to volts. Voltage signals from electrometer 126 are telemet~y conditioned and conveyed to tr~n~mitt~r 130 for 10 radio L~ --ic~ion. All of the components within box 132 are implanted as a single unit in the patient.
Fxtern~lly, radio signals from tr~n~mitter 130, in(lic~tive of glucose concentrations in the patient's blood, are transmitted to receiver 134.
Receiver 134 may be connected to monitor 136 for data monitoring. The same 15 receiver co~ ult;~ or another colll~ul~l 138 may be used to analyze ~e raw data and generate glucose concentration information. A printer 140 connected to col"~ule, 138 generates hard copies of analyzed data.
The concept of including multiple electrode pairs within a single sensor can be çxtçnclerl to an embodiment where separate sensors are implanted 20 and commonly linked to a single electrometer as shown in Figure 8. For çY~mI~le, eight implantable sensors 150 can be implanted in a patient and linkedto a single electrometer 152 and tr~n~mitttor (not shown). Tr~n~mit~e~l signals are received by data acquisition adaptor 154 and acquisition co~ ulel 156. By increasing the nllmber of sensors the overall precision, accuracy and longevity of the system can be greatly enhanced. If one or more anodes (or sensors) fails,the others still provide sufficient data sensing capacity so that the entire unit continlles to perform s~ticf~ctorily. Various algorithms or averaging protocols can be used to process the multiple data streams.
Figure 9 shows scllem~tically the components of an implantable unit in a glucose sensing system. Implantable unit 160 includes disc-shaped glucose sensor 162 which is connected to electrometer and teleme1Ty conditioning package 164 via anode wire 166a and cathode wire 166b. Radio signals derived from the raw current signals are tr~ncmit~e~l from ~n~
element 168.
Circuitnl Figure 10 shows custom circuilly structure employed in a glucose sensing system of the present invention. Shown generally at 210 is a glucose servotr~ncmitter suitable for implementation with the present invention.
Servotr~ncmi1tçr 210 is configured for tr~n~micsion of data which is indicative of a sensed enzy~natic reaction to a remote receiving source for subsequent processin~ the sensing and conveyance of such data being described in detail below.
As shown, servotr~ncmittçr 210 includes a sensor 212 (also referred to as a two-electrode sensor) operatively connected between a voltage rerelence source 214 and an arnplifier circuit 216. The output of circuit 216 isl,~eled at 218 and subsequently provided to a voltage-to-frequency circuit 220, which in the preferred embodiment includes a CMOS 7555 circuit indicated at 220a c~ nfi~lred with a resistive and capacitative network which includes two resistors (R9 and R8) and a capacitor (C2). Utilization of CMOS
for ~le~igning circuit 220 has been found ideal due to its low power con~nmrtion aspects which results in longer battery life. The output tf~rminAl of circuit 220 is connected via line 222 to an AC-coupled trAncmi1t~r 224 (also referred to herein as a ~ )i",;l~er) for trAn~mi~ion of data to an çxl~."~1 receiving source.
Discussing the above servotr~n~mi~t~r in more detail, circuit 210 is configured for detecting electrons which are generated duling an enzymatic reaction, and conveying data which is representative of such detected electrons to an e~t~rnAl source for subsequent processing. More specifically, sensor 212 inclll~les two electrodes, a cathode 212a and an anode 212b. Ca~ode 212a is connected to voltage reference source or circuit 214, and anode 212b is connected to amplifier circuit 216. Voltage reference circuit 214 is made up of three resistors R4, R5, and R6 and a 1.2-volt Zener diode Zl. Resistor R6 is connected at one end to a negative voltage potential, and at the other end to diode Zl and resistor R4. The other end of resistor R4 is connected to resistor R5, which in turn is connected to diode Zl as shown. The common node between resistors R4 and R5 is connected to cathode 212a.
Anode 212b is connected via resistors R3, R2, to the inverting te. ,~ Al of amplifier 216a, and a capacitor Cl is connected between resistors R3, R2 and ground. The non-inverting tçrminAl of arnplifier 216a is tied to ground. A resistor Rl is connected between the output of amplifier 216a and its inverting termin~l to provide negative fee~lb~ck The output of amplifier 216a is connected to the non-inverting lf...~ of amplifier 218a, the output of which is connected to the illv~llhlg tçrmin~l in a voltage follower configuration for buffering the output of amplifier 216a. A resistor R7 is connected between the output of amplifier 218a and the trigger t~rmin~l 2 of the CMOS 7555 circuit.
The CMOS 7555 is configured, with its ~ttçn(l~nt resistive and c~p~cit~tive network, as a voltage-to-frequency converter whose output frequency is proportional to its input control voltage. Referring more specifically to the 7555, it may be seen that reset terminal 4 is connected to tertnin~l 8, both of which are connected to a voltage potential which may be referred to as VCC. A resistor R8 is connected between reset terrnin~l 4 and discharge tçrmin~l 7. A resistor R9 is connected between discharge tçrmins~l 7 and the threshold tennin~l 6. A capacitor C2 is connected between trigger t~rrnin~l 2 and ground. Output tçrmin~l 3 is connected to the AC-coupled tr~n~ e. 224 for tr~n~mi~ion of data to a remote location for processin~
Discussing the operation of the above-described glucose servotr~n~mitter, it will be understood that voltage ~ lce circuit 214 develops a potential of -0.6 volts which is used by sensor 212 to cause electrons produced in the vicinity of the sensor to flow, in the form of a generated current, with amplifier circuit 216, which includes operational amplifier 216a configured for feedback as described above. The output of amplifier 216a is a voltage which is buffered at 218 by operational amplifier 218a, the voltage output of which controls ~equency for the tngger te~nin:~l of the CMOS 7555 through resister R7 and frequency selection c~ y C2, R8, and R9. The output tt~rmin~l 3 of the CMOS 7555 is connected, via line 222, to 5 tr~n~mitter 224 for tr~n~mi~ion to an extern~l source.
It will be appreciated that the above-described 7555 configuration c~llv~lL~ ~he ou~ut of buffer 218 into a frequency which is detr....;I.ed by lhevoltage at threshold tt rmin~l 6. The 7555 serves two functions in ~e above configuration which are necessary for the tr~n~mi~sion of sensed data to a remote location for processing. First, the 7555 provides a 15-msec pulse to key tr~n~ 224, ~ereby hlrnin~ it on and off in accordance with prac~ces which will be understood by those of skill in the aIt. Second, the 7555 is operable for voltage-to-frequency conversion, which is a me~ rement of sensor response. This dual function enables the afor~men1ioned data tr~n~mi~sion in a manner which will be understood by those of skill in the art.
Preferred component values (resistive and capacitative values) for the above-described servotr~n~mit~er 210 are as follows~ for voltage ence circuit 214: R4 = 1 meg ohm; R5 = 4.7 meg ohm; and, R6--470 kohm; (2) for amplifier circuit 216: Rl = 500 meg ohm; R2 and R3 = 499 kohm; and, Cl = 10 pf; (3) for converter circuit 220: R9 = 180 kohm; R8 = 1 meg ohm; and, C2 = 1 microfarad; and, (4) R7 = 4.7 meg ohm.
The above system is referred to as a "two-electrode" system bec~llse of the fact that two electrodes are utilized (the anode and the cathode) CA 02237102 1998-0~-08 in the sensing of electrons produced during a particular enzymatic reaction.
Another system which is suitable for sensing produced electrons and conveying data relative to such sensed electrons is a so-called "three-electrode" system which is shown in Figure lOA and described briefly below.
In Figure lOA, like or similar elements of the three-electrode glucose serVotr~n~m~ r 210 are labeled to correspond with the two-electrode elementc appearmg in Figure 10. The Figure shows a sensor 212, a voltage ~cfelellce source 214, a voltage-to-frequency converter circuit 220, and a tr~n~mittçr 224. Voltage refclcllce source 214, voltage-to-frequency converter circuit 220, and tr~n~mitter 224 will not be described because the operation of those elements is the same as, or similar to the operation of such elements as they appear in Figure 10.
Sensor 212 in Figure lOA varies somewhat from its Figure 10 coulllc~ l. Such variations take into account some observations regarding current and voltage control which have been made with respect to the two-electrode system described above, and improve somewhat, the control of such parameters. The three electrode sensor, set forth at 212, includes a counter electrode 212a (which may be formed from silver), a common return electrode 212b (also referred to as a worlcing electrode and which may be formed from pl~1imlm), and a voltage probe 212c which may also be termed the lcre~cnce electrode (and which also may be formed from pl~tinllm). Two operational amplifiers 212d, 212e are provided and operatively coupled to the electrodes as show4 in a configuration which provides greater current and voltage control which, in turn, assists in m~ .g the integrity of the electrodes' sensitivity and the ability of the same to detect a produced cu~ which is indicative of an er~natic reaction. The control is effectuated in a clamped, controlled m~nner.
The three electrode sensor 212 is shown in Figure 10A.
Amplifier 212d m~int~in~ a voltage which is the sarne as the reference voltage of -0.6 volts between the reference and working electrodes 212c, 212b respectively. This is accomplished by varying the current at the counter electrode 212a which is in the feedback loop of amplifier 212d. Amplifier 212e 10 m~int~in~ the working electrode 212b at virtual ground converting the C-LIlc.to an output voltage, which is buffered at 212f and provided to CMOS 7555 co~ l circuit 220 for conversion from a voltage to a frequency (in a rn~nner described above), the converter circuit thereafter triggering tr~n~mi~t~?r 224 is a pre~ e....il~ble fashion to transmit sensed data indicative of an enzyma~c 15 reaction to a remote location for processing.
Decreasin~ Fibrotic Capsule Interference One of the primary reasons why a subcutaneously-implanted sensor eventually loses its ability to measure the concentration of arl analyte of interest is that a collagenous capsule forms around the sensor. The capsule 20 eventually loses vascularity and becomes thick and fibrous, ~ereby subst~n*~lly blocking the sensor from accessing the analyte present in blood.
There are at least two promising approaches for ~
fibrotic capsule illle,rerence with analyte detection, thereby ~xtt~.n~lin~ longevity CA 02237102 1998-0~-08 of an implanted sensor. First, it is possible to prevent or retard capsule form~tion by slow controlled release of certain collagen deposition inhibitors.
Drugs which inhibit collagen formation can be incorporated in a polymer matrix which allows slow release of the drug locally to achieve the desired 5 effect without c~llcing adverse distant systemic effects in the animal or hnm~n For example, collagen inhibitors which can be used for this purpose include corticosteroids such as ~lex~methasone, relaxin and ~;~mm~ t~lr~r~n. A
efelled polymer material for car~ying and controlling slow release of the drug is polydimethylsiloxane. Corticosteroids can be impregnated in a 10 polydimethylsiloxane matrix so as to provide relatively long-term, slow rele~e of the corticosteroids in the surrounding tissue. It is important, however, that~le~r~methasone be released in small doses in order to avoid iatrogenic Cushing's syndrome, which is a serious illness caused from systemic excess of corticosteroids. If corticosteroids are released from a sensor for a prolonged 15 period, for example, more than two weeks, we recommend that a patient's serum be tested in order to confirrn that adverse systemic effects are avoided.
Another approach for minimi7:in~ fibrotic capsule i~le-r~;~ellce with sensor performance, i.e., increasing sensor longevity, is to promote vascularity in the capsule so that the sensor can continue to have access to 20 blood analytes. Accordingly, vascular growth factors can be incorporated in amatrix around the sensor so that the grow~ factors are slowly released into the ou~lding tissue. The released growth factors enhance capillary growth in the coll~gennus capsule which forms around the implanted sensor. Retention of capillary perfusion by the capsule enhances sensor function by continuously providing the sensor access to the patient's blood analyte. Examples of c~pill~ growdl factors include vascular endothelial growth factor (VEGF) and endothelial cell growth factor (ECGF). Polymer materials which are capable of S slowly releasing polypeptide factors such as ECGF and VEGF include poly-l-lactic acid and poly glycolic lactic acid. As with ~e steroid approach, the grow~ factor dosage, i.e., quantity and rate of release, must be carefully controlled so that the growth factor's effect is local, not systemic.
A method of employing steroids or growth factors to ~ e or 10 avoid fibrotic capsule interference with sensor performance, is to provide for the active agent's slow release from the perimeter of ~e disk sensor. For example, as shown in Figure 16A, glucose sensor 300 has a carrier layer or matrix 302 such as a tape made of or cont~inin~ polydimethylsiloxane g~-~te-l with dexamethasone. Tape 302 is attached to outer perimeter edge 304 of disk-shaped housing or body 306 of sensor 300. The width of tape 302 is subst~nti~lly the same dimension as the width of edge 304, i.e., thiclmess ofhousing 306, so that the steroid is released on or near both faces of ~e sensor.Time-release steroid compositions have been lltili7e~1 in the past for other puIposes. For example, U.S. Patent No. 5,265, 608 to Lee et al., the 20 entire content of which is hereby incorporated by reference, discloses a steroid elll~lin~ electrode in which ~lex~methasone is incorporated in a polymer matrix which permits slow controlled release of the steroid to control infl~mm~tion~
init~tiQn and swelling in connection with a device such as a pacçm~ker.
However, no one has previously employed a time release corticosteroid matrix for inhibiting collagen formation on an implantable analyte sensor.
Sensin~ Other Analytes With minor modifications, the sensor design~ described above 5 may be used to detect analytes other than glucose. By ch~n ing the specific type of enzyme which covers the anode, the sensor can be used to me~nre many compounds. Several examples appear in Table 1 below.
ANALYTE ENZYME
glucose glucose oxidase glucose hexose oxidase lactate lactate oxidase l-methionine l-amino acid oxidase l-phenylalanine l-amino acid oxidase d-aap~le d-amino acid oxidase d-~h~ te d-amino acid oxidase urate urate oxidase ethyl alcohol alcohol oxidase methyl alcohol alcohol oxidase cholesterol cholesterol oxidase ascorbic acid ascorbate oxidase In addition to measuring analytes in body fluids, sensors of the present invention can be used to measure the concentration of substrates in other fluids, for example, fruit and vegetable juices, wine, yogurt, etc.
Construction of a Glucose Sensor A preferred sensor is constructed of epoxy resin in a disc shape, 1.3-cen~imet~ors in diameter and 0.2-centimeters in height. Four 36-gauge pk77inllm wires le....i~ e peripherally on one face of the disc (in holes drilled in 5 the resin) and service hydrogen peroxide-sçn~in~ anodes. A solid silver cylinder, 0.7-centimeters outside diameter (the cathode), is secured by epoxy resin in the center of the disc. A layer of silver chloride can be deposited onto the surface of the silver by one of several processes. The sensor is ~lc;rel~blydouble-sided, which may be, for example, two of the four anode sensors 10 c~mfi~77red "back-to-back", m~7kin~ a sensor composed of four anodes and one cathode on each face of the sensor.
Anode and cathode recording wires terminate in an amplifier and polarizing voltage source. An electrometer converts the current signal to a voltage signal and applies a con~t~nt pol~ri~ing voltage of 0.60 V to the 15 electrodes. Output from the amplifier is routed both to a digital volt meter (Micronta 22-185A, Tandy Corp., Fortworth, TX 76102) and to a strip chart continuous chart recorder (Gould Inshuments Model No. 11-2963-300, Valley View, OH 44126). The signal can also be routed directly into a cc,~ uler by use of a data acquisition board. All of these electrical components can be 20 mini~hlri7ecl without ~ltering the function.
A working 8-anode sensor (which has been demonstrated to respond to peroxide) then is selected for testing. The sensor is s~ntle-l first with 600 and 1500 grit wet-or-dry, then followed by a polishing with 2000 grit wet-or-dry. The sensor is rinsed thoroughly in a stream of deionized water (DW) followed by blow drying in a cold nitrogen stream. The sensor is then immersed in an acetone bath and vigorously twirled for 20 seconds to remove any solvents or oils from the surface. The sensor is withdrawn from the 5 acetone bath and is immediately rinsed in a DW stream. The sensor is again blown dry in a cold nitrogen stream, and contimles to dry in room air for another 30 minlltes-If it is desired to include an intel~r~nt reLarding layer, then alayer of PPX (or other suitable membrane material) approximately 5,000-10 8,000A thick is deposited directly on top of the anodes before depositing anyenzyme.
The sanded, cleaned and dried sensor (with or without i~ ,relelll ,e~dillg layer) is enzyme activated with a Glucose Oxidase (GO) ---Bovine Serum Albumin (BSA)--- Glutaraldehyde (GA) matrix prepared from mixin~
two parts GO+BSA (20mg GO + 5mg BSA with 0.5 gram DW) plus one part GA (2.5% GA diluted with DW). Approximately 2.5 ~11 of this solution is applied via pipette directly to each anode. The solution is allowed to dry in room air for one hour. The sensor is then immersed in DW for 15 minl!tes to remove excess GA, rinsed briefly in a DW stream, and blown dry in a nitrogen 20 st~eam. The sensor continlles to dry in room air for one hour, after which spin-coating with PU (Tecoflex, Tecothane, Cook composite, or CAR) or vapor deposition w~th PPX (thickness = 3000-SOOOA) is carried out.
Further mini~tllrization of the glucose sensor, as described above, will not adversely effect performance of the unit.
Testin~. Connectin~ and Implantin~ Sensors Sensors m~mlf~ctured as described above, are tested the day after S they are made by applying a polarizing voltage of 600 mV. The voltage output should stabilize after a one to two hour immersion in a tempc.alule controlled PBS solution (37~C) in the laboratory water bath. The sensor is tested in st~n-l~rd glucose solutions prepared by adding glucose to PBS so that ~he resnlting test solutions (G=glucose), are concentrated in mg/dl as follows:
G(0), G(100), G(200), G(300), G~400) and G(500); and in millimolar concentrations as follows: G(0), G(5.6), G(l l. l), G(16.7), G(22.2) and G(27.8). The first data point is collected while the sensor is still immersed inPBS and represents the baseline output. After noting the output value, t~e sensor is moved to G(100) for ten minlltes. The process of measuring ~e speed 15 with which the sensor responds to the increase in glucose allows calculation of T90 (defined T90 below). The sensor is moved to the G(200) standard and dle ten minute output value is collected from this standard. All of the following uul~uls are collected in ascending order in the same m~nner.
An implantable sensor has to satisfy three criteria: (1) it must 20 have a T90 of less than three min~ltes; (2) it must be dose responsive in theglucose concenl~lion range of 40-400 mg/dl; and (3) it must have adequate sensitivity. The T90 value is measured by using the continuous sensor readout provided by the data acquisition system. The point at which 90% of ~he m5 xi~ -- output is reached (after ch~n~ing from the zero glucose level to the 100 mg/dl level) is recorded as the T90.
A sensor that is acceptable for implant must also be dose responsive, preferably substantially linear over the glucose concentration of 40-5 400 mg/dl. Minor to moderate non-linearities can be mathematically corrected to allow estim~*on of glucose level from sensor output data.
If a sensor meets all the previous criteria, it is attached to a tr~n~mit~r. For example, a suitable transmitter may be obtained from Mini-Mitter which has a custom-built interface circuit between the tr~n~mitter and 10 the sensor. The tr~n~mi1ter should have a battery pack which is fully charged.
The sensor can be implanted in the body of ~nim~l~ or hnm~n~.
The sensor can be implanted subcutaneously, in an artery or vein, uscularly~ intraperitoneally, in the brain or cerebrospinally. The prerelled location is subcutaneous. The sensor can also be used in vitro, for 15 example, in a laboratory to measure glucose concentration or other substrates or analytes in a liquid media.
The transmitter and sensor package are tested in vitro the day of the planned implant procedure. If the results are s~ti~f~ctory (T90 less than 3 les, s~ti~f~ctorily dose-responsive, adequate sensitivity), then the unit is 20 sterili~e~l, rinsed in sterile saline, and implanted subcutaneously in the recipient (after the a~rol~liate preparation and anesthesia procedures).
Experiments Experiment 1 We compared the performance of sensors with one anode to the performance of sensors with four anodes. Twelve one-sensor anode sensors 5 were constructed subst~nti~lly as described above. These sensors were sirnilarto the ones shown in Figure 1 except they only included one anode in~te~d of four, and they only had electrodes on one side of the disc-shaped sensor. All sensors in ~is experiment were dip-coated with polyurethane (Cook Composites) instead of parylene. Twenty-four-anode sensors were constructed 10 the same as the one-anode sensors except that they included four anodes on one face of the sensor subst~nti~lly as shown in Figure 1.
The sensors were implanted in rats. Glucose dose response data was collected for each of the sensors at frequent time points after impl~nt~tionun~l the given sensor failed to perform s~ti~f~ctorily. For each sensor, the last 15 check point at which the sensor performed adequately, d~Lt;~ ed the functional life of that sensor.
Figure 11 shows the results of this experiment. The average longevity for the one-anode sensors was about 4 days. In contrast, the average longevity for the four-anode sensors was about 28 days. This is a highly 20 significant improvement in the functional life of an implanted glucose sensor, which we attribute to the increased number of anodes.
CA 02237l02 l998-0~-08 Experiment 2 The purpose of this experiment was to determine in vitro the perfonn~nce capability in sensors which had failed in vivo. In this experiment eight of the four anode sensors used in Experiment 1 were tested before 5 implantation (pre-implant), and then tested again after eventually failing to perform and being removed (post-explant) from the rat.
The results of the ~x~tlilllent are shown in Fig~re 12. In Figure 12 (and Figures 13 and 14), the "Normal Range" includes glucose concentrations which are typically observed in the normal population. The 10 "Dynamic Range" includes the Normal Range plus abnormally high and low glucose concentrations which should be measurable with a glucose sensor. The results show that in vitro the sensors performed as well post-explant as they did pre-implant. This result demonstrates that failure of the sensors in vivo is notdue to inactivation or loss of the glucose oxidase enzyme. We noted that over 15 time in vivo a cellular coat tends to envelop the sensor. Before performing dle post-explant testing on the sensors, the coats were removed. This suggests that the cellular coat which develops around the sensor may be involved with c;vt;nlual sensor failure. Since the cellular coat is relatively non-uniform, it is possible to theorize that one of the reasons why longevity is increased with 20 multiple anodes is that the probability of m~in~ining one or more anodes under a portion of the coat which is minim~l enough so that the sensor still performs,is increased by increasing the number or surface area of sensing anodes.
Experiment 3 II1 Experimçnt~ 1 and 2, the sensors were dip-coated in polyurethane (Cook Composites). We subsequently discovered that ~ irO~ y and overall performance of the sensors can be improved by using PPX as the S outer coat or semipermeable membrane. The purpose of this experiment was to llemo~ e glucose dose response and repeatability for eight sensors, each of which was coated with PPX at a ~hickness of a~ro~ lately 3800A. As shown in Figure 13, we observed a dose response appro~rhing linearity in the useful me~nrement range. Test repeatability was also improved with the PPX coated 10 sensors, as shown by the smaller standard deviation margins in colll~ison to those shown in Figure 12.
Experiment 4 This experiment was similar to Experiment 3 except instead of using PPX as the semipermeable outer membrane, CAR was used. Eight-15 percent CAR was spin-coated over the sulface of the sensor for 2.5 ~ es at 4,000 RPM. The sensor was tested in vitro at various glucose concentrations in 3 l~ccessive runs. The data is shown in Figure 14. The dose response over the useful me~cllrement range approached linearity with a higher slope in c~ pA.icon to slopes obtained with PPX and dip-coated sensors. We also 20 noted a relatively small standard deviation on repeat tests with the CAR coated sensors.
EXperiTnent 5 This experiment was performed in vitro with PPX coated eight anode (four on each side) sensors to dele~ e how rapidly the sensors respond to çh~n~,es in glucose concentration (T90). Six sensors were constructed with PPX outer coats of 3000-5000A. Results of this experiment are shown in Figure 15. Each of the sensors responded with a T90, i.e., tirne to reach 90-percent of ~ ;m~te current output for a given change in glucose concentration, in less than one minute. This is a faster response time than we had observed previously with polyurethane dip-coated sensors.
Claims (60)
1. A device for electrochemically sensing changes in the concentration of an analyte of interest comprising an implantable sensor body having a pair of spaced facial expanses, each expanse of the body having at least one analyte-sensing electrode pair including an anode and a cathode spaced apart from each other and covered by a membrane which is semi-permeable to the analyte of interest, and an enzyme composition covering each anode of each analyte-sensing electrode pair, wherein the enzyme composition reacts specifically with the analyte of interest.
2. The device of claim 1 wherein the analyte of interest is glucose.
3. The device of claim 2 wherein the enzyme composition includes glucose oxidase and is located between the anode and the membrane.
4. The device of claim 1 wherein the membrane comprises polyparaxylxylene.
5. The device of claim 1 wherein each expanse of the sensor body includes a plurality of anodes.
6. The device of claim 5 wherein each expanse of the sensor body has a plurality of electrode pairs, each electrode pair including one anode and one cathode.
7. The device of claim 6 wherein each expanse of the sensor body has four electrode pairs.
8. The device of claim 7 wherein the sensor body is disk-shaped.
9. The device of claim 3 wherein each expanse of the sensor body has an interferent retarding layer between the anode and the enzyme layer.
10. A device for electrochemically sensing changes in the concentration of an analyte of interest comprising an implantable sensor body including a plurality of analyte-sensing electrode pairs, each electrode pair including an anode and a cathode, and a membrane comprising polyparaxylxylene or a carbonate-based polyurethane covering the anodes.
11. The device of claim 10 further comprising a layer including glucose oxidase disposed between the anode and the membrane, wherein the device senses changes in glucose concentration of a fluid surrounding the device.
12. The device of claim 10 wherein the sensor body has two opposing sides, each side of the sensor body having at least one anode and at least one cathode.
13. The device of claim 12 wherein each side of the sensor body has a plurality of anodes.
14. The device of claim 13 wherein each side of the sensor body has a plurality of cathodes, each cathode being paired with one of the anodes.
15. A device for electrochemically sensing the concentration of an analyte of interest comprising an implantable sensor body supporting and defining multiple analyte detection sites, each site having an analyte-sensing electrode pair, each electrode pair including an anode and a cathode for generating signals indicative of analyte concentration at one of the sites on the sensor body.
16. The device of claim 15 wherein the sensor body has two opposing sides, the electrode pairs being distributed on both sides of the sensor body.
17. The device of claim 16 wherein the sensor body is disk-shaped.
18. The device of claim 16 wherein each side of the sensor body has at least four electrode pairs.
19. The device of claim 15 wherein the anodes lead to a common anode conductor and the cathodes lead to a common cathode conductor.
20. The device of claim 15 wherein the analyte of interest is glucose and each anode is covered with an enzyme layer comprising glucose oxidase.
21. An implantable device for electrochemically sensing changes in the concentration of an analyte of interest, and transmitting signals indicative of the concentration changes, comprising a transmitter including a power source, a sensor electrically coupled to the transmitter, the sensor including a disk-shaped body having two opposing sides, each side of the body having and a cathode and a plurality of anodes, whereby the combined transmitter and sensor can be implanted in a mammal for wireless transmission of data indicative of analyte concentration to an external receiver.
22. The device of claim 21 wherein the analyte of interest is glucose.
23. The device of claim 22 further comprising an enzyme layer comprising glucose oxidase covering the anodes, and a membrane semi-permeable to glucose covering the enzyme layer.
24. The device of claim 21 further comprising an amplifier and an electrometer, the cathodes and anodes from the sensor being connected to the amplifier and the electrometer converting current signals into voltage signals before transmitting corresponding signals to an external processing device.
25. The device of claim 21 further comprising analog-to-digital converter connected to the sensor for converting analog signals indicative of current changes into digital signals prior to transmitting corresponding data to an external receiver.
26. An analyte concentration monitoring system comprising a sensor including a body having two opposing sides, each side of the body having at least one cathode, plural anodes and a semi-permeable membrane covering the anodes, the sensor being capable of generating analog data signals indicative of analyte concentration in a fluid surrounding the sensor, a transmitter including a power source, the transmitter being electrically coupled to the sensor and capable of converting the data signals into corresponding radio transmission signals a receiver for receiving the radio transmission signals at a remote location.
27. The system of claim 26 further comprising a processor connected to the receiver for interpreting and converting the radio transmission signals into analyte concentration information.
28. A method of making an implantable analyte sensor comprising providing a body having two opposing sides, creating at least one cathode and plural anodes on both sides of the body, and depositing a semi-permeable membrane on the cathodes and anodes.
29. The method of claim 28 wherein the analyte is glucose, further comprising depositing an enzyme layer including glucose oxidase on the anodes before the step of depositing the semi-permeable membrane.
30. The method of claim 28 comprising depositing a layer of polyparaxylxylene or Chronoflex~ AR on the enzyme layer.
31. The method of claim 28 comprising forming the body in the shape of a disk.
32. The method of claim 29 further comprising electrically coupling the sensor to a radio transmitter.
33. The method of claim 32 further comprising implanting the sensor and transmitter into a mammal, sensing glucose concentration changes, transmitting corresponding radio signals to a remote receiver, and processing and interpreting the radio signals into glucose concentration data.
34. A system for monitoring analyte concentrations in the blood of a mammal comprising an organization of implantable sensors, each sensor including an anode and a cathode covered by a membrane which is semi-permeable to the analyte.
35 The system of claim 34 wherein the analyte is glucose, further comprising an enzyme layer between the anodes and the membrane.
36. The system of claim 34 further comprising a body substantially containing all of the sensors.
37. The system of claim 36 wherein the body is substantially disk shaped and has two opposing sides, the sensors being disposed on both sides of the body.
38. The system of claim 37 wherein each side of the body has four sensors.
39. The system of claim 34 further comprising of plurality of bodies, each sensor being provided in a separate body.
40. The system of claim 34 wherein all anodes lead to a common anode conductor and all cathodes lead to a common cathode conductor.
41. A device for electrochemically sensing changes in the concentration of an analyte of interest comprising a wholly subcutaneously implantable sensor body, an analyte detection mechanism contained within the sensor body, and a fibrotic coat interference inhibitor in close proximity to and exterior to the sensor body for decreasing interference with analyte detection due to formation of a fibrotic coat around the sensor body.
42. The device of claim 41 wherein the analyte is glucose.
43. The device of claim 41 wherein the inhibitor has the property of decreasing the rate of collagen deposition on the sensor body.
44. The device of claim 41 wherein the inhibitor has the property of increasing vascularity in a fibrotic coat.
45. The device of claim 41 wherein the inhibitor is a steroid.
46. The device of claim 41 wherein the inhibitor is a vascular growth factor.
47. The device of claim 45 wherein the inhibitor is selected from the group consisting of dexamethasone, relaxin and gamma interferon, or mixtures thereof.
48. The device of claim 46 wherein the inhibitor is selected from the group consisting of vascular endothelial growth factor, endothelial cell growth factor, and mixtures thereof.
49. The device of claim 41 wherein the inhibitor is incorporated in a polymer matrix which provides controlled relatively constant release of the inhibitor over time into tissue surrounding the sensor body.
50. The device of claim 49 wherein the polymer matrix includes a material selected from the group consisting of polydimethylsiloxane, poly-1-lactic acid, polyglycolic lactic acid and mixtures thereof.
51. The device of claim 50 wherein the polymer matrix is in the form of a tape impregnated with the inhibitor.
52. The device of claim 41 wherein the sensor body is substantially disk shaped.
53. The device of claim 52 wherein the sensor body has a peripheral edge, the inhibitor being impregnated in a time-release polymer matrix associated with the peripheral edge of the sensor body.
54. The device of claim 41 wherein the inhibitor is dexamethasone.
55. The device of claim 49 wherein the inhibitor comprises dexamethasone and the polymer matrix comprises polydimethylsiloxane.
56. The device of claim 41 wherein the sensor body has a pair of spaced facial expanses, each expanse of the sensor body including at least one anode and at least one cathode spaced apart from each other and covered by a membrane which is semipermeable to the analyte of interest.
57. The device of claim 56 wherein the analyte of interest is glucose.
58. The device of claims 1,2,3,4,5,6,7,8 or 9 further comprising a fibrotic coat interference inhibitor in close proximity to the sensor body for decreasing interference with analyte detection due to formation of a fibrotic coat around the sensor body.
59. The device of claim 58 wherein the inhibitor is a corticosteroid or a growth factor.
60. The device of claim 41 wherein the effective dose of inhibitor is large enough to significantly decrease interference with analyte detection caused by fibrotic capsule formation, but small enough to avoid adverse systemic effects.
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US08/561,972 US5711861A (en) | 1995-11-22 | 1995-11-22 | Device for monitoring changes in analyte concentration |
| US08/561,972 | 1995-11-22 |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA2237102A1 true CA2237102A1 (en) | 1997-05-29 |
Family
ID=24244276
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002237102A Abandoned CA2237102A1 (en) | 1995-11-22 | 1996-11-21 | Device for monitoring changes in analyte concentration |
Country Status (8)
| Country | Link |
|---|---|
| US (3) | US5711861A (en) |
| EP (1) | EP0877931B1 (en) |
| JP (1) | JP2000500380A (en) |
| AT (1) | ATE403860T1 (en) |
| AU (1) | AU1022697A (en) |
| CA (1) | CA2237102A1 (en) |
| DE (1) | DE69637629D1 (en) |
| WO (1) | WO1997019344A1 (en) |
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- 1996-11-21 EP EP96940582A patent/EP0877931B1/en not_active Expired - Lifetime
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- 1996-11-21 WO PCT/US1996/018724 patent/WO1997019344A1/en active Application Filing
- 1996-11-21 JP JP9519908A patent/JP2000500380A/en active Pending
- 1996-11-21 AT AT96940582T patent/ATE403860T1/en not_active IP Right Cessation
- 1996-11-21 DE DE69637629T patent/DE69637629D1/en not_active Expired - Fee Related
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2000
- 2000-11-28 US US09/724,918 patent/US6466810B1/en not_active Expired - Fee Related
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| US6212416B1 (en) | 2001-04-03 |
| AU1022697A (en) | 1997-06-11 |
| DE69637629D1 (en) | 2008-09-18 |
| EP0877931A1 (en) | 1998-11-18 |
| WO1997019344A1 (en) | 1997-05-29 |
| US6466810B1 (en) | 2002-10-15 |
| EP0877931A4 (en) | 1999-08-11 |
| US5711861A (en) | 1998-01-27 |
| JP2000500380A (en) | 2000-01-18 |
| ATE403860T1 (en) | 2008-08-15 |
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