Source: https://insight.rpxcorp.com/pat/US20060094944A1
Timestamp: 2019-11-17 22:25:28
Document Index: 272063553

Matched Legal Cases: ['Application No. 60', 'Application No. 60', 'Application No. 60', 'Application No. 60', 'Application No. 60', 'Application No. 60']

Patent US 20060094944A1
Filed: 09/13/2005
1. A transdermal analyte monitoring system comprising:
a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane;
an electrode assembly comprising a plurality of electrodes; and
a processor programmed to implement an error correction method that corrects for drift;
wherein the medium is adapted to react continuously with the analyte, an electrical signal is detected by the electrode assembly, and the electrical signal correlates to an analyte value.
The invention relates to a transdermal analyte monitoring system comprising a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane and an electrode assembly comprising a plurality of electrodes, wherein the medium is adapted to react continuously with the analyte, an electrical signal is detected by the electrode assembly, and the electrical signal correlates to an analyte value. The analyte value may be the flux of the analyte through the biological membrane or the concentration of the analyte in a body fluid of a subject. The medium may comprise a vinyl acetate based hydrogel, an agarose based hydrogel, or a polyethylene glycol diacrylate (PEG-DA) based hydrogel, for example. The surface region of the electrode may comprise pure platinum. The system may include an interference filter located between the biological membrane and the electrode assembly for reducing interference in the system. The system may comprise a processor programmed to implement an error correction method that corrects for sensor drift.
Iontophoresis apparatus and method
US 7,996,077 B2
HYDROGELS FOR VOCAL CORD AND SOFT TISSUE AUGMENTATION AND REPAIR
US 20100055184A1
US 20080058699A1
US 20080311670A1
US 8,214,030 B2
In vivo component measurement method, data processing method for in vivo component measurement, in vivo component measurement apparatus and collection member
US 8,747,316 B2
Methods and systems of matching voice deficits with a tunable mucosal implant to restore and enhance individualized human sound and voice production
US 9,198,568 B2
Sandeep Sidram Karajanagi
US 9,216,188 B2
General Hospital Corporation, Massachusetts Institute of Technology
US 9,549,697 B2
US 9,682,169 B2
Method of determining an analyte concentration
US 10,196,512 B2
Wearable device and health monitoring method
US 10,206,577 B2
US 10,227,485 B2
Apparatus and methods for transdermal sensing of analytes in interstitial fluid and associated data transmission systems
US 10,321,858 B2
Proteadx Inc.
Nicholas F. Warner, Joseph Kost, Sean Moran, Mitragotri Samir S., Scott C. Kellogg, Roode Lauren, Farnham Hannah, Barman Shikha
US 20040039418A1
US 20040087671A1
Noninvasive transdermal systems for detecting an analyte obtained from or underneath skin and methods
US 6,503,198 B1
TECHNICAL CHEMICALS PRODUCTS INC.
Intelligent polymerized crystalline colloidal array carbohydrate sensors
US 20030027240A1
Sonophoretic enhanced transdermal transport
US 6,190,315 B1
SONTRA MEDICAL L.P.
Sliding door with wheel repair kit
US 6,223,471 B1
Ultrasound enhancement of transdermal transport
US 6,234,990 B1
Polyurethane hydrogel drug reservoirs for use in transdermal drug delivery systems, and associated methods of manufacture and use
US 5,902,603 A
Ortho-McNeil Pharmaceutical Incorporated
Supersaturated transdermal drug delivery systems, and methods for manufacturing the same
US 5,906,830 A
Ultrasonic method and apparatus for cosmetic and dermatological applications
US 5,618,275 A
Sonex International Corp.
Gel containment structure
US 5,626,554 A
Adhesive base material
US 5,646,221 A
Programmable apparatus for reducing substance dependency in transdermal drug delivery
US 5,538,503 A
Methods and compositions to enhance epithelial drug transport
US 5,534,496 A
Blood processing for treating blood disease
US 5,401,237 A
Adhesive hydrogels having extended use lives and process for the preparation of same
US 5,405,366 A
Ultrasound therapy system with automatic dose control
US 5,413,550 A
PTI LLC
US 5,315,998 A
Tachibana Katsuro, Tachibana Shunro
Ultrasound-enhanced delivery of materials into and through the skin
US 5,323,769 A
Cygnus Therapeutic Systems, Regents of the University of California
Polyphase fluid extraction process, resulting products and methods of use
US 5,330,756 A
Huffstutler M. Conrad Jr., Steuart Gary M.
Injection instrument with ultrasonic oscillating element
US 5,197,946 A
TACHIBANA SHUNRO, Sumitomo Electric Industries Limited
Method for delivering an active substance topically or percutaneously
US 5,215,520 A
Centre International de Recherches Dermatologiques Galderma
US 5,215,887 A
System for applying ultrasonic waves and a treatment instrument to a body part
US 5,078,144 A
US 5,115,805 A
Method and apparatus for non-invasive monitoring of blood glucose
US 5,119,819 A
Crosslinked hydrogel and method for making same
US 5,120,544 A
HENLEY INTERNATIONAL INC. A CORP. OF TX
Method of providing a substrate with a layer comprising a polyvinyl based hydrogel and a biochemically active material
US 5,134,057 A
Method and therapeutic system for smoking cessation
US 5,135,753 A
PHARMETRIX CORPORATION A CORP OF DE
Dose critical in-vivo detection of anti-cancer drug levels in blood
US 5,001,051 A
Resilient transdermal drug delivery device
US 5,006,342 A
US 5,007,438 A
Tachibana Shunro, Meiji Seika Kaisha Limited
US 5,016,615 A
Control of transport of molecules across tissue using electroporation
US 5,019,034 A
Modified composite electrodes with renewable surface for electrochemical applications and method of making same
US 4,933,062 A
Ultrasound enhancement of transbuccal drug delivery
US 4,948,587 A
US 4,821,740 A
TACHIBANA SHUNRO 1-6-18 KUSAGAE CHUO-KU FUKUOKA-SHI FUKUOKA-KEN JAPAN, Meiji Seika Kaisha Limited
Transdermal drug composition with dual permeation enhancers
US 4,820,720 A
Transdermal detection system
US 4,821,733 A
DERMAL SYSTEMS INTERNATIONAL INC.
DERMAL SYSTEMS INTERNATIONAL
US 4,834,978 A
US 4,860,058 A
6-Halo-1,2,3,4-tetrahydr oquinazoline-4-spiro-4-imidazolidine-2,2'5'-triones useful for the treatment and prophylaxis of diabetic complications
US 4,855,298 A
Transdermal dosimeter
US 4,732,153 A
Method for treating herpes lesions and other infectious skin conditions
US 4,646,725 A
Moasser Manoutchehr
US 4,657,543 A
Transdermal treatment for pain and inflammation with 2-amino-3-aroylbenzeneacetic acids, salts and esters
US 4,683,242 A
Synthetic resin wound dressing and method of treatment using same
US 4,563,184 A
ENQUAY INC.
Korol Bernard
Penetrating topical pharmaceutical compositions containing N-(2-hydroxyethyl) pyrrolidone
US 4,537,776 A
Non-invasive diagnosis method
US 4,457,748 A
Treatment of acne and skin disorders and compositions therefor
US 4,372,296 A
Fahim Mostafa S.
Topical application of medication by ultrasound with coupling agent
US 4,309,989 A
Torsional ultrasonic vibrators
US 4,144,646 A
Lion Hamigaki Kabushiki Kaisha, Kano Denki Kabushiki Kaisha
Method of transmitting ultrasonic impulses to surface using transducer coupling agent
US 4,002,221 A
Selective chemical sensitive FET transducers
US 4,020,830 A
ENHANCING TISSUE PENETRATION OF PHYSIOLOGICALLY ACTIVE STEROIDAL AGENTS WITH DMSO
US 3,711,606 A
Filed 09/02/1970
METHOD AND APPARATUS FOR ULTRASONIC TREATMENT OF LOWER TISSUES SIMULTANEOUS WITH HEATING OF SUBCUTANEOUS, OUTER MUSCLE AND LOWER TISSUES
US 3,828,769 A
Hal C. Mettler
COMPOSITIONS FOR TOPICAL APPLICATION FOR ENHANCING TISSUE PENETRATION OF PHYSIOLOGICALLY ACTIVE AGENTS WITH DMSO
US 3,711,602 A
2. The transdermal analyte monitoring system of claim 1, wherein the analyte comprises glucose.
3. The transdermal analyte monitoring system of claim 2, wherein the medium comprises a hydrogel and glucose oxidase.
4. The transdermal analyte monitoring system of claim 2, wherein the processor is programmed to apply a drift factor D(t) to a blood glucose value X(t) to calculate a drift-corrected blood glucose value Xp(t).
5. The transdermal analyte monitoring system of claim 4, wherein the drift factor D(t) is represented by a third order polynomial.
6. The transdermal analyte monitoring system of claim 5, wherein the drift factor D(t) is represented as D(t)=c*t3+d*t2+e*t+f, wherein c, d, e and f are numerical coefficients calculated to provide a best fit for D(t) to empirical data.
7. A method for monitoring an analyte comprising:
positioning a medium with respect to a biological membrane such that the medium can receive an analyte from the biological membrane, wherein an electrode assembly is coupled to the medium;
continuously reacting the analyte with the medium;
detecting an electrical signal with the electrode assembly;
calculating an analyte value based on the electrical signal; and
applying an error correction to the analyte value to correct for drift.
8. The method of claim 7, further comprising pretreating the biological membrane to increase a permeability of the biological membrane.
9. The method of claim 8, wherein the pretreating step comprises applying low frequency ultrasound to the biological membrane.
10. The method of claim 7, wherein the analyte comprises glucose.
11. The method of claim 10, wherein the medium comprises a hydrogel and glucose oxidase.
12. The method of claim 7, wherein the step of applying an error correction comprises applying a drift factor D(t) to a blood glucose value X(t) to calculate a drift-corrected blood glucose value Xp(t).
13. The method of claim 12, wherein the drift factor D(t) is represented by a third order polynomial.
14. The method of claim 13, wherein the drift factor D(t) is represented as D(t)=c*t3+d*t2+e*t+f, wherein c, d, e and f are numerical coefficients calculated to provide a best fit for D(t) to empirical data.
The present application is a divisional of U.S. application Ser. No. 11/201,334, filed Aug. 11, 2005, which is a continuation of U.S. application Ser. No. 10/974,963, filed Oct. 28, 2004, both of which are hereby incorporated by reference in their entireties. The present application is related to the following patent and applications, each of which is incorporated herein by reference it its entirety: U.S. application Ser. No. 09/979,096, filed Mar. 16, 2001; U.S. application Ser. No. 09/868,442, filed Dec. 17, 1999; U.S. Provisional Application No. 60/112,953, filed Dec. 18, 1998; U.S. Provisional Application No. 60/142,941, filed Jul. 12, 1999; U.S. Provisional Application No. 60/142,950, filed Jul. 12, 1999; U.S. Provisional Application No. 60/142,951, filed Jul. 12, 1999; U.S. Provisional Application No. 60/142,975, filed Jul. 12, 1999; U.S. Pat. No. 6,190,315; and U.S. Provisional Application No. 60/070,813, filed Jan. 8, 1998.
Diabetics frequently prick their fingers and forearms to obtain blood in order to monitor their blood glucose concentration. This practice of using blood to perform frequent monitoring can be painful and inconvenient. New, less painful methods of sampling body fluids have been contemplated and disclosed. For example, these painless methods include the use of tiny needles,
the use of iontophoresis, and the use of ultrasound to sample body fluid, such as blood and interstitial fluid.
It has been shown that the application of ultrasound can enhance skin permeability. Examples of such are disclosed in U.S. Pat. Nos. 4,767,402, 5,947,921, and 6,002,961, the disclosures of which are incorporated, by reference, in their entireties. Ultrasound may be applied to the stratum corneum via a coupling medium in order to disrupt the lipid bilayers through the action of cavitation and its bioacoustic effects. The disruption of stratum corneum, a barrier to transport, allows the enhanced diffusion of analyte, such as glucose or drugs, through, into, and out of the skin.
The application of a motive force before, during, and after making the skin permeable has been disclosed in U.S. Pat. Nos. 5,279,543, 5,722,397, 5,947,921, 6,002,961, and 6,009,343, the disclosures of which are incorporated by reference in their entireties. The purpose of using a motive force is to actively extract body fluid and its content out of the skin for the purpose of analysis. As mentioned, active forces, such as vacuum, sonophoresis, and electrosmotic forces, can create convective flow through the stratum corneum. Although these forces can be used for extraction of body fluids, there are certain limitations that may apply when the forces are applied to human skin. For example, a major limitation is the flow and volume of body fluid that can be transported across the stratum corneum. In general, high-pressure force is necessary in order to transport fluid across an enhanced permeable area of stratum corneum. The application of vacuum on skin for an extended period may cause physical separation of the epidermis from the dermis, resulting in bruises and blisters.
Therefore, a need has arisen for a system, method, and device for noninvasive body fluid sampling and analysis that overcomes these and other drawbacks of the related art.
Therefore, a need has arisen for a method of enhancing the permeability of a biological membrane, such as skin, buccal, and nails, for an extended period of time, and a method for extracting body fluid to perform blood, interstitial fluid, lymph, or other body fluid analyte monitoring in a discrete or continuous manner that is noninvasive and practical.
According to one embodiment, the invention relates to a transdermal analyte monitoring system comprising a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane, wherein the medium comprises a hydrogel selected from the group consisting of vinyl acetate based hydrogels, agarose based hydrogels, polyethylene glycol diacrylate (PEG-DA) based hydrogels and mixtures thereof, and an electrode assembly, wherein the medium is adapted to react continuously with the analyte, and wherein an electrical signal is detected by the electrode assembly, and the electrical signal correlates to an analyte value.
According to another embodiment, the invention relates to a transdermal analyte monitoring system comprising a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane, and an electrode assembly comprising a plurality of electrodes, wherein a surface region of at least one of the electrode consists essentially of pure platinum, wherein the medium is adapted to react continuously with the analyte, and wherein an electrical signal is detected by the electrode assembly, and the electrical signal correlates to an analyte value.
According to another embodiment, the invention relates to a transdermal analyte monitoring system comprising a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane, an electrode assembly, and an interference filter located between the biological membrane and the electrode assembly for reducing interference from non-target biological moieties in the transdermal analyte monitoring system.
According to another embodiment, the invention relates to a transdermal analyte monitoring system comprising a medium adapted to interface with a biological membrane and to receive an analyte from the biological membrane, a sensor comprising an electrode assembly, the electrode assembly comprising a plurality of electrodes, and a processor programmed to implement an error correction method that corrects for sensor drift, wherein the medium is adapted to react continuously with the analyte, and wherein an electrical signal is detected by the electrode assembly, and the electrical signal correlates to an analyte value.
A method for non-invasive body fluid sampling and analysis is disclosed. According to one embodiment of the present invention, the method includes the steps of (1) identifying an area of biological membrane having a permeability level; (2) increasing the permeability level of the area of biological membrane; (3) contacting the area of biological membrane with a receiver; (4) extracting body fluid through and out of the area of biological membrane; (5) providing an external force to enhance the body fluid extraction; (6) collecting the body fluid in the receiver; (7) analyzing the collected body fluid for the presence of at least one analyte; and (8) providing the results of the step of analyzing the body fluid.
FIG. 6 depicts the components of a wearable extraction chamber according to one embodiment of the present invention;
FIG. 7 depicts a graph of glucose flux versus blood glucose concentration according to one embodiment of the present invention;
FIG. 8 depicts a flow chart of a method for controlled enhancement of transdermal delivery according to one embodiment of the present invention;
FIG. 9 depicts an apparatus for performing continuous transdermal analyte monitoring according to one embodiment of the present invention;
FIG. 10 is a drawing of the sensor body shown in FIG. 9 from a first view;
FIG. 11 is a drawing of the apparatus shown in FIG. 9 from a second view;
FIG. 12 shows the signal response versus glucose concentration for various hydrogels;
FIG. 13 shows the signal response versus glucose concentration for pure platinum versus platinized carbon as the working electrode;
FIG. 14 shows the current-time profiles of a glucose sensor responding to the addition of hydrogen peroxide using platinum and platinized carbon as the working electrode;
FIG. 15 shows the sensor response to hydrogen peroxide (HP) over acetominophen (AM) and hydrogen peroxide over uric acid (UA) for sensors with and without a Nafion interference filter;
FIG. 16 shows a Clark Error Grid in the absence of an error correction method according to one embodiment of the invention;
FIG. 17 shows a Clark Error Grid after the application of an error correction method according to an embodiment of the invention;
FIG. 18 shows the absorbance spectrum of a standard glucose oxidase solution before and after incorporation into a PEGDA 3.4K gel;
FIG. 19 shows the signal response to glucose of glucose oxidase (GOx) loaded PEG gels of varying molecular weight;
FIG. 20 shows signal response to glucose of 3.4K PEG hydrogel loaded with varying concentrations of GOx;
FIG. 21 shows raw data of the potentiometric signals elicited from PEGDA hydrogels with GOx incorporated in the gel formulation prior to photocrosslinking;
FIG. 22 shows the change in signal between GOx-presoaked versus pre-incorporated hydrogels at different thickness and compositions (PEGDA-nVP, PEGDA);
FIG. 23(a) shows blood glucose versus time utilizing an embodiment of the continuous transdermal analyte monitoring system;
FIG. 23(b) shows a correlation plot of electrode signal in nanoamps versus blood glucose for an embodiment of the invention;
FIG. 24 shows patient data for participants in a clinical study;
FIG. 25 shows a noisy data set from the clinical study;
FIG. 26 shows another data set from the clinical study;
FIG. 27 shows a Clark Error Grid for sensor data from the clinical study according to one embodiment of the invention; and
FIG. 28 shows a Clark Error Grid for sensor data from the clinical study according to another embodiment of the invention.
The preferred embodiments of the present invention and their advantages are best understood by referring to FIGS. 1 through 28 of the drawings, like numerals being used for like and corresponding parts of the various drawings.
As used herein, the term “body fluid” may include blood, interstitial fluid, lymph, and/or analyte. In addition, as used herein, the term “biological membrane” may include tissue, mucous membranes and comified tissues, including skin, buccal, and nails. Further, as used herein, the term “force” may also include force gradients.
In one embodiment, the ultrasound may have an intensity in the range of about 0 to about 100 watt/cm<sup>2</sup>, and preferably in the range of 0 to about 20 watt/cm<sup>2</sup>. Other appropriate intensities may be used as desired.
Referring to FIG. 2, a device for the controlled application of ultrasound to biological membrane to enhance the permeability of a biological membrane according to one embodiment of the present invention is shown. Device 200 includes controller 202, which interfaces with ultrasound applicator 204 by any suitable means, such as a cable. Controller 202 controls the application of ultrasound to the area of biological membrane. In one embodiment, ultrasound or near ultrasound having an intensity in the range of about 0 to about 20 watt/cm<sup>2 </sup>may be generated by controller 202 and ultrasound applicator 204. In one embodiment, the ultrasound may have a frequency of about 20 kHz to about 150 kHz. In another embodiment, the ultrasound may have a frequency of 50 kHz. Other ultrasound frequencies may also be used.
Ultrasound applicator 204 may be provided with cartridge 206, which contains ultrasound coupling solution 208. Cartridge 206 may be made of any material, such as plastic, that may encapsulate ultrasound coupling solution 208. Suitable ultrasound coupling solutions 208 include, but are not limited to, water, saline, alcohols including ethanol and isopropanol (in a concentration range of 10 to 100% in aqueous solution), surfactants such as Triton X-100, SLS, or SDS (preferably in a concentration range of between 0.001 and 10% in aqueous solution), DMSO (preferably in a concentration range of between 10 and 100% in aqueous solution), fatty acids such as linoleic acid (preferably in a concentration range of between 0.1 and 2% in ethanol-water (50:50) mixture), azone (preferably in a concentration range of between 0.1 and 10% in ethanol-water (50:50) mixture), polyethylene glycol in a concentration range of preferably between 0.1 and 50% in aqueous solution, histamine in a concentration range of preferably between 0.1 and 100 mg/ml in aqueous solution, EDTA in a concentration range of preferably between one and 100 mM, sodium hydroxide in a concentration range of preferably between one and 100 mM, sodium octyl sulfate, N-tauroylsarcosine, octyltrimethyl ammoniumbromide, dodecyltrimethyl ammoniumbromide, tetradecyltrimethyl ammoniumbromide, hexadecyltrimethyl ammoniumbromide, dodecylpyridinium chloride hydrate, SPAN 20, BRIJ 30, glycolic acid ethoxylate 4-ter-butyl phenyl ether, IGEPAL CO-210, and combinations thereof.
In one embodiment, target ring 210 may be used to monitor the permeability level of the biological membrane, as disclosed in PCT International Patent Appl'n Ser. No. PCT/US99/30067, entitled “Method and Apparatus for Enhancement of Transdermal Transport,” the disclosure of which is incorporated by reference in its entirety. In such an embodiment, target ring 210 may interface with ultrasound applicator 204.
Referring to FIG. 4, a device for the continuous extraction and analysis of body fluid to infer analyte concentrations according to another embodiment of the present invention is provided. As shown in the figure, a biological membrane site on the forearm, the abdomen, or thigh may be exposed to ultrasound; other biological membrane sites, such as those on the back, may also be used. Receiver 402, which may be similar to receiver 214, may contact the ultrasound exposed biological membrane site to perform continuous extraction of body fluid. In one embodiment, receiver 402 may contain a medium, such as a hydrogel layer, that may incorporate an osmotic agent, such as sodium chloride. The hydrogel is formulated to contain phosphate buffered saline (PBS), with the saline being sodium chloride in the concentration range of 0.01 M to 10 M. The hydrogel may be buffered at pH 7.
The detection methods and results may be performed and presented to the user by meter 404, which may be similar in function to meter 212, discussed above. In one embodiment, meter 404 may be wearable. For example, as depicted in the figure, meter 404 may be worn in a manner similar to the way a wristwatch is worn. Meter 404 may also be worn on a belt, in a pocket, etc.
The following example does not limit the present invention in any way, and is intended to illustrate an embodiment of the present invention.
FIG. 6 describes the components of wearable extraction chamber 600. Four extraction chambers were placed on each sonicated site of the human volunteer. Thin circular foam chamber 602 was constructed using foam MED 5636 Avery Dennison ( 7/16″ ID× 11/8″ OD). Foam chambers 602 were attached concentrically to the sonicated biological membrane sites using double-sided adhesive (Adhesive Arcade 8570, 7/16″ ID×⅞″ OD) attached to one side of element 602. The other side of foam chamber 602 was attached concentrically to double-sided adhesive 604 (Adhesive Arcade 8570, 7/16″ ID×⅞″ OD). Thin transparent lid 606 was made of 3M Polyester 1012 ( 11/8″× 11/8″). Double-sided adhesive 604 permitted thin transparent lid 606 to be attached to foam chamber 602 after placement of liquid into the inner diameter of foam chamber 602 when attached to biological membrane. Thin transparent lid 606 acted as a lid to prevent liquid from leaking out of the extraction chamber, and to allow the extraction chambers to be wearable for an extended period of time.
Solutions were collected and analyzed for glucose concentration using high-pressure liquid chromatography. The results of the HPLC concentration were normalized for the injection amount and the total solution volume, and were reported as glucose flux (Q<sub>g</sub>), the mass of glucose that crossed the sonicated site per unit time per unit area. Body fluid glucose concentrations (C<sub>bg</sub>) were obtained by testing capillary blood obtained from a lanced finger in a Bayer Glucometer Elite meter. It was hypothesized that Q<sub>g </sub>would be linearly proportional to C<sub>bg</sub>. FIG. 7 shows a graph of Q<sub>g </sub>versus C<sub>bg</sub>. Unexpectedly, Q<sub>g </sub>from the sonicated sites exposed to 1 M NaCl correlated to C<sub>bg </sub>much more strongly than Q<sub>g </sub>from the sonicated sites exposed to 0.15 M NaCl.
According to another aspect of the present invention, an apparatus and method for regulating the degree of skin permeabilization through a feedback system is provided. This apparatus and method may be similar to what has been described above, with the addition of further regulation of the degree of skin permeabilization. Feedback control as a method of monitoring the degree of skin permeability is described in more detail in U.S. application Ser. No. 09/868,442, entitled “Methods and Apparatus for Enhancement of Transdermal Transport,” which is hereby incorporated by reference in its entirety. In this embodiment, the application of the skin permeabilizing device is terminated when desired values of parameters describing skin conductance are achieved. As the discussion proceeds with regard to FIG. 8, it should be noted that the descriptions above may be relevant to this description.
Referring to FIG. 8, a flowchart of the method is provided. In step 802, a first, or source, electrode is coupled in electrical contact with a first area of skin where permeabilization is required. The source electrode does not have to make direct contact with the skin. Rather, it may be electrically coupled to the skin through the medium that is being used to transmit ultrasound. In one embodiment, where an ultrasound-producing device is used as the skin permeabilizing device, the ultrasonic transducer and horn that will be used to apply the ultrasound doubles as the source electrode through which electrical parameters of the first area of skin may be measured and is coupled to the skin through a saline solution used as an ultrasound medium. In another embodiment, a separate electrode is affixed to the first area of skin and is used as the source electrode. In still another embodiment, the housing of the device used to apply ultrasound to the first area of skin is used as the source electrode. The source electrode can be made of any suitable conducting material including, for example, metals and conducting polymers.
Next, in step 804, a second, or counter, electrode is coupled in electrical contact with a second area of skin at another chosen location. This second area of skin can be adjacent to the first area of skin, or it can be distant from the first area of skin. The counter electrode can be made of any suitable conducting material including, for example, metals and conducting polymers.
When the two electrodes are properly positioned, in step 806, an initial conductivity between the two electrodes is measured. This may be accomplished by applying an electrical signal to the patch of skin through the electrodes. In one embodiment, the electrical signal supplied may have sufficient intensity so that the electrical parameter of the skin can be measured, but have a suitably low intensity so that the electrical signal does not cause permanent damage to the skin, or any significant electrophoresis effect for the substance being delivered. In one embodiment, a 10 Hz AC source is used to create a voltage differential between the source electrode and the counter electrode. The voltage supplied should not exceed 500 mV, and preferably not exceed 100 mV, or there will be a risk of damaging the skin. In another embodiment, an AC current source is used. The current source may also be suitably limited. The initial conductivity measurement is made after the source has been applied using appropriate circuitry. In one embodiment, a resistive sensor is used to measure the impedance of the patch of skin at 10 Hz. In another embodiment, a 1 kHz source is used. Sources of other frequencies are also possible.
In step 808, a skin permeabilizing device is applied to the skin at the first site. Any suitable device that increases the permeability of the skin may be used: In one embodiment, ultrasound is applied to the skin at the first site. According to one embodiment, ultrasound having a frequency of 20 kHz and an intensity of about 10 W/cm<sup>2 </sup>is used to enhance the permeability of the patch of skin to be used for transdermal transport.
In step 810, the conductivity between the two sites is measured. The conductivity may be measured periodically, or it may be measured continuously. The monitoring measurements are made using the same electrode set up that was used to make the initial conductivity measurement.
In step 812, mathematical analysis and/or signal processing may be performed on the time-variance of skin conductance data. Experiments were performed on human volunteers according to the procedure above, with ultrasound used as the method of permeabilization. Ultrasound was applied until the subjects reported pain. Skin conductivity was measured once every second during ultrasound exposure. After plotting the conductance data, the graph resembled a sigmoidal curve. The conductance data was in a general sigmoidal curve equation:<FORM>C=C<sub>i</sub>+(C<sub>f</sub>−C<sub>i</sub>)/(1+e<sup>S(t−t*)</sup>) </FORM>where:
C is current;
C<sub>i </sub>is current at t=0;
C<sub>f </sub>is the final current;
S is a sensitivity constant;
t* is the exposure time required to achieve an inflection point; and
t is the time of exposure.
Referring again to FIG. 8, in step 814, the parameters describing the kinetics of skin conductance changes are calculated. These parameters include, inter alia, skin impedance, the variation of skin impedance with time, final skin impedance, skin impedance at inflection time, final current, exposure time to achieve the inflection time, etc.
In step 816, the skin permeabilizing device applied in step 808 is terminated when desired values of the parameters describing skin conductance are achieved. For instance, when the skin conductance increases to a certain value, the permeabilizing device may be deactivated. Alternatively, when the rate of change in the value of skin conductance is a maximum, the permeabilizing device may be deactivated. Additional details of the method for regulating the degree of skin permeabilization are disclosed in the aforementioned U.S. application Ser. No. 09/868,442.
A preferred embodiment of a continuous transdermal glucose monitoring system and method is described in connection with FIGS. 9-11. As discussed above, the term “body fluid” may include blood, interstitial fluid, lymph, and/or analyte. Body fluids include, for example, both complete fluids as well as molecular and/or ionic components thereof. Preferred embodiments of the invention may involve extraction and measurement of just the analyte.
FIG. 9 is a drawing of a continuous glucose monitoring system according to an exemplary embodiment of the invention. In this embodiment, the system includes a sensor assembly generally including a sensor body 901 and a backing plate 902 as well as other components as described herein. The sensor body may include electrodes, as shown in FIG. 10, on its surface for electrochemical detection of analytes or reaction products that are indicative of analytes. A thermal transducer 903, which may be housed in a housing with a shape that corresponds to that of the sensor body 901, is located between the sensor body 901 and the backing plate 902. Electrochemical sensors, such as hydrogen peroxide sensors, can be sensitive to temperature fluctuation. The thermal transducer 903 may be used to normalize and report only those changes attributed to a change in analyte or analyte indicator. An adhesive disc 904 may be attached to the side of the sensor body 901 that faces the thermal transducer 903. An adhesive ring 905 may be attached to the side of the sensor body 901 that is opposite the adhesive disc 904. The cut-out center portion of the adhesive ring 905 preferably exposes some or all of the sensor components on the sensor body 901. The adhesive ring 905 and adhesive disc 904 may have a shape that corresponds to that of the sensor body as shown in FIG. 9. A hydrogel disc 906 may be positioned within the cut-out center portion of the adhesive ring 905 adjacent a surface of the sensor body 901. During operation, the sensor assembly may be positioned adjacent a permeable region 907 of a user's skin as shown by the dashed line in FIG. 9. The sensor assembly may be attached to a potentiostat recorder 908, which may include a printed circuit board 911, by way of a flexible connecting cable 909. The connecting cable 909 preferably attaches to the potentiostat recorder 908 using a connector 910 that facilitates removal and attachment of the sensor assembly.
The system shown in FIG. 9 can be used to carry out continuous monitoring of an analyte such as glucose as follows. First, a region of skin on the user is made permeable using, for example, sonication as described above. The sensor assembly, such as that shown in FIG. 9, is then attached to the permeable region 907 of skin so that the hydrogel disc 906 is in fluid communication with the permeable skin. An analyte may be extracted through the permeable region 907 of the user's skin so that it is in contact with the hydrogel disc 906 of the sensor assembly. For example, an analyte such as glucose may be transported by diffusion into the hydrogel disc 906 where it can contact glucose oxidase. The glucose can then react with glucose oxidase present in the hydrogel disc 906 to form gluconic acid and hydrogen peroxide. Next, the hydrogen peroxide is transported to the surface of the electrode in the sensor body 901 where it is electrochemically oxidized. The current produced in this oxidation is indicative of the rate of hydrogen peroxide being produced in the hydrogel, which is related to the amount of glucose flux through the skin (the rate of glucose flow through a fixed area of the skin ). The glucose flux through the skin is proportional to the concentration of glucose in the blood of the user. The signal from the sensor assembly can thus be utilized to continuously monitor the blood glucose concentration of a user by displaying blood glucose concentration on the potentiostat 908 in a continuous, real-time manner.
Detailed views of a preferred embodiment of the sensor body 901 are shown in FIG. 10. The sensor body 901 includes a body layer 1007 upon which leads 1004, 1005, and 1006 are patterned. The leads may be formed, for example, by coating metal over the body layer 1007 in the desired locations. A working electrode 1001, is typically located at the center of the sensor body 901. The working electrode 1001 may comprise pure platinum, platinized carbon, glassy carbon, carbon nanotube, mezoporous platinum, platinum black, paladium, gold, or platinum-iridium, for example. The working electrode 1001 may be patterned over lead 1006 so that it is in electrical contact with the lead 1006. A counter electrode 1002, preferably comprising carbon, may be positioned about the periphery of a portion of the working electrode 1001, as shown in FIG. 10. The counter electrode 1002 may be patterned over lead 1005 so that it is in electrical contact with the lead 1005. A reference electrode 1003, preferably comprising Ag/AgCl, may be positioned about the periphery of another portion of the working electrode 1001 as shown in FIG. 10. The electrodes 1001, 1002, and 1003 can be formed to roughly track the layout of the electrical leads 1006, 1005, 1004, respectively, that are patterned in the sensing area of the device. The electrodes 1001, 1002, and 1003 may be screen printed over the electrical leads 1006, 1005, 1004, respectively. The leads can be pattered, using screen printing or other methods known in the art, onto the sensor body 901 in a manner that permits electrical connection to external devices or components. For example, the leads may form a 3X connector pin lead including leads 1004, 1005, and 1006 at the terminus of an extended region of the sensor body as shown in FIG. 10. A standard connector may then be used to connect the sensor electrodes to external devices or components.
The electrochemical sensor utilizes the working electrode 1001, the counter electrode 1002, and the reference electrode 1003 to measure the rate hydrogen peroxide or glucose is being generated in the hydrogel. The electrochemical sensor is preferably operated in potentiostat mode during continuous glucose monitoring. In potentiostat mode, the electrical potential between the working and reference electrodes of a three-electrode cell are maintained at a preset value. The current between the working electrode and the counter electrode is measured. The sensor is maintained in this mode as long as the needed cell voltage and current do not exceed the current and voltage limits of the potentiostat. In the potentiostat mode of operation, the potential between the working and reference electrode may be selected to achieve selective electrochemical measurement of a particular analyte or analyte indicator. Other operational modes can be used to investigate the kinetics and mechanism of the electrode reaction occurring on the working electrode surface, or in electroanalytical applications. For instance, according to an electrochemical cell mode of operation, a current may flow between the working and counter electrodes while the potential of the working electrode is measured against the reference electrode. It will be appreciated by those skilled in the art that the mode of operation of the electrochemical sensor may be selected depending on the application.
The sensor assembly described generally in relation to FIG. 9 is show in expanded detail from another angle in FIG. 11. The sensor body 901, which is covered on each side by adhesive disc 904 and adhesive ring 905, is shown in relation to the backing plate 902. The hydrogel disc 906 may be positioned in such a manner that it will face toward the user after folding over onto the backing plate 902 as shown in FIG. 9. The sensor body may be connected to the backing plate 902 using standard connectors such as a SLIM/RCPT connector 1301 with a latch that mates with a corresponding connector interface that is mounted onto the backing plate 902.
The sensor assembly shown in FIGS. 9-11 may be incorporated into any one of a number of detection devices. For instance, this sensor assembly may be incorporated into the receiver of FIG. 4 to provide for discrete and/or continuous glucose monitoring. Additionally, the sensor assembly may be connected to a display or computing device through a wireless connection or any other means for electrical connection in addition to the cable 909.
Continuous glucose monitoring as described herein can be achieved without accumulation of a certain volume of body fluid in a reservoir before measuring the concentration of the withdrawn fluid. Continuous glucose monitoring is capable of measuring the blood concentration of glucose without relying on accumulation of body fluids in the sensor device. In continuous glucose monitoring, for instance, one may prefer to minimize accumulation of both glucose and hydrogen peroxide in the hydrogel so that the current measured by the electrochemical sensor is reflective of the glucose flux through the permeable region of skin in real-time. This advantageously permits continuous real-time transdermal glucose monitoring.
According to another aspect of the invention, a step of skin hydration may be employed prior to or concurrently with increasing the porosity of the skin (e.g. by applying ultrasound) to improve the continuous transdermal analyte monitoring. Skin hydration prior to or concurrently with increasing the porosity, and prior to attaching the sensor may improve sensor performance by establishing or stabilizing liquid pathways between the skin and the sensor, improving the moisture balance over the sensor-skin interface, and/or continuing to maintain ample water to the hydrogel to maintain enzyme activity. The skin hydration procedure can be performed, for example, by applying a hydrating agent to the target skin site. The hydrating agent may be applied in combination with a delipidation or cleansing agent. Where both hydrating and cleansing agents are utilized, they may be applied in a single application using a single solution. Alternatively, the cleansing agent and the hydrating agent can be applied using successive application of two different solutions. In one aspect, one or both solutions are applied using a pad applicator. In another aspect, the solution can be held in contact with the skin by positioning it in the bellows of a sonication device or another device that might function to hold a liquid in contact with skin.
In one embodiment, a glycerin/water prep pad solution may be prepared for skin hydration. The following batch formulation can be used to prepare the glycerin/water prep pad solution. 300.00 grams of glycerin 99% USP is added to the first container. 2.70 grams of Nipagin M (methylparaben), 0.45 grams of Nipasol M (propylparaben), and 30.00 grams of benzyl alcohol NF are dissloved in a second container and then added to the first container. The glycerin and benzyl alcohol solutions are then mixed in the first container until the solution clears. 1133.85 grams of deionized water is then added to the solution in the first container and mixed until homogeneous. 1.50 grams of Potassium Sorbate NF is added to the solution in the first container and mixed until homogenious. 1.50 grams of Glydant 2000 is then added to the solution in the first container and mixed until homogenious. Lastly, 30.00 grams of deionized water is added to the solution in the first container and mixed until homogeneous.
In one embodiment, a 1 3/16″ prep pad is utilized. Preferably the prep pads are composed of 70% polypropylene/30% cellulose. In one embodiment, the prep pad has a width that ranges from 1 1/16″ to 1 5/16″. In one embodiment, the thickness of the prep pad is 21-29 mils. In another embodiment, the thickness of the prep pad is 26-34 mils. In one embodiment the prep pad has a basis weight of 1.43-1.87 g/yd using ATM#102. In another embodiment, the prep pad has a basis weight of 1.72-2.24 g/yd using ATM#102. Preferably, the prep pad is utilized with a prep pad solution, such as the prep pad solution above, to hydrate a biological membrane before increasing its porosity.
According to another aspect of the invention, the working electrode 1001 of FIG. 10 may include a surface layer of pure platinum. The pure platinum working electrode 1001 may be screen printed or otherwise coated onto the surface of a lead 1006. Using pure platinum as the working electrode can enhance sensitivity and increase the rate of conversion of hydrogen peroxide. This can provide advantages for continuous transdermal glucose monitoring as the conversion of hydrogen peroxide is preferably fast to prevent its accumulation, which may cause positive sensor drift and/or enzyme deactivation. In transdermal glucose sensing applications, pure platinum can offer advantages over traditional platinized carbon materials.
One advantage that pure platinum can offer relative to platinized carbon is an enhanced sensitivity to glucose concentration. FIG. 13 shows the glucose sensitivity of both pure platinum and platinized carbon. As shown by this comparison, the glucose sensitivity of pure platinum is about 2.9 times that of platinized carbon. The glucose sample size used to generate the data of FIG. 13 was 2 microliters.
Another advantage that pure platinum can offer relative to platinized carbon is enhanced sensitivity to hydrogen peroxide. FIG. 14 shows the hydrogen peroxide sensitivity of both pure platinum and platinized carbon. Specifically, FIG. 14 shows the current-time profiles of a glucose sensor responding to the addition of hydrogen peroxide (sometimes referred to as a hydrogen peroxide “challenge”) using platinum and platinized carbon as the working electrode. As shown by this comparison, the hydrogen peroxide sensitivity of pure platinum is about 5 times that of platinized carbon.
Another advantage that pure platinum can offer relative to platinized carbon is a higher success rate for glucose monitoring. The percentage success rate for glucose monitoring using pure platinum was 83% versus 60% for platinized carbon (correlation coefficient R<sup>2</sup>>=0.5 as the passing criteria). R refers to the correlation between conventional whole blood glucose measurements and measurements of blood glucose using the system of FIG. 9. R is calculated by comparing the continuous data from the system of FIG. 9 with discrete whole blood measurements (taken every 20 minutes). A linear regression analysis is run on the two data sets to generate an R value. The correlation between sensor signal and blood glucose levels using pure platinum was R<sup>2</sup>=0.87 versus R<sup>2</sup>=0.71 for platinized carbon.
According to another aspect of the invention, a protective interference filter can be provided to reduce or even eliminate interference effects from unwanted electrochemical oxidation and/or biofouling. One form of interference, for example, involves the production of unwanted anodic signal by electrochemical oxidation of ascorbic acid, uric acid, and/or acetaminophen, which can all be oxidized electrochemically at voltage levels applied in glucose monitoring. Another form of interference can involve biofouling, which can occur when biological species deposit on a sensor surface thereby limiting the sensor's free access to analyte or deactivating its functionality by reacting with the electrode. It is generally advantageous to reduce or eliminate the effects of interfering species through the use of an interference filter since many of these species may be present in body fluids during glucose monitoring.
According to an exemplary embodiment of the invention, the interference filter comprises a Nafion film coated onto one or more surfaces of the sensor assembly. Other interference filter materials such as (3-mercaptopropyl)trimethylsilane, cellulose acetate, electropolymerized films such as 1,8-diaminonapthaline and phenylenediamine, PTFE or other hydrophobic, Nylon or other hydrophylic membranes may be used. Nafion is a biocompatible anionic fluoropolymer that can be coated on sensor surfaces as a protective layer against physiological interferents and biofouling based on hydrophobicity, charge selection, and size exclusion, for example. Nafion is available from Aldrich Chemical of Milwaukee, Wis. A Nafion film may be coated directly on the surface of at least the working electrode 1001 of the sensor body 901. Alternatively, a Nafion film may be coated on an outer surface of the sensor assembly such as the hydrogel layer 906. In general, one or more interference filter layers may be provided between the working electrode surface and any other layer or on the outermost surface of the sensor assembly that contacts the user's skin during operation.
A Nafion layer can be conveniently coated on a sensor surface using a micropipette, for example, or by dip-coating the sensor in aqueous or organic Nafion solution followed by air drying for several hours before use. FIG. 15 shows the effect of a Nafion coating on the sensor response to glucose relative to the interferents acetaminophen and uric acid. The plot shows the hydrodynamic sensor response to 0.294 mM of hydrogen peroxide (HP) over acetominophen (AM) and uric acid (UA) in phosphate buffered saline with 0.5 V of applied voltage. The amperometric current produced by acetaminophen and uric acid is greatly reduced for a sensor coated with Nafion relative to an uncoated sensor. Thus, Nafion can significantly improve the analyte/interferent signal ratio.
In various embodiments of the invention described herein, hydrogels can be used as part of the analyte monitoring system. Hydrogels constitute an important class of biomaterials utilized for medical and biotechnological applications such as in contact lenses, biosensors, linings for artificial implants and drug delivery devices. FIGS. 9 and 11 show a preferred hydrogel disc 906 in relation to the sensor assembly. The hydrogel disc 906 may be located over the sensor body 901 within the cutout center portion of the adhesive ring 905 of the sensor assembly. The continuous transdermal analyte monitoring system may utilize one or more of the preferred hydrogel materials described below. Classes of hydrogel materials that may be used in exemplary embodiments of the invention include: agarose based hydrogels, polyethylene glycol diacrylate (PEG-DA) based hydrogels, and vinyl acetate based hydrogels, for example. Following a general description of these gels are examples detailing the procedures used to produce and/or characterize the various hydrogels.
Agarose based hydrogels can offer advantages for continuous transdermal analyte monitoring. For instance, agarose can offer one or more of the following features: good response to glucose and hydrogen peroxide due to its high water content, high enzyme loading, good bio-compatibility, and excellent permeation and diffusion properties. In addition, agarose hydrogels may offer cleanliness, low cost, and/or ease of preparation.
An agarose gel may be formed, for example, from 1-20% agarose in buffer solution containing 0-1 M sodium or potassium phosphate, 0-1 M sodium chloride, 0-1 M potassium chloride, 0-2 M lactic acid, surfactant such as 0-1 M Triton X-100, Tween 80 or sodium lauryl sulfate, and any other biocompatible components. Loading of glucose oxidase in agarose hydrogel can be up to 0-20% (by weight), for example, by soaking the solid hydrogel in concentrated glucose oxidase solution, or alternatively by mixing concentrated glucose oxidase powder or solution with agarose solution during its melting stage (15-65° C.), followed by cooling and gelling at lower temperature (40° C. or lower).
PEG based hydrogels can offer several advantages for continuous transdermal analyte monitoring. Structurally, PEG is highly hydrophilic and presents a high degree of solvation in aqueous solvents. The preferential solvation of PEG molecules can effectively exclude proteins from the PEG chain volume, thereby protecting the surface from bio-fouling by proteins. An advantage that can be provided by chemically crosslinked PEG-based hydrogels is that their physical and chemical properties can be modulated by varying the molecular weight of the PEG chains and varying the initiator concentration. For example, increasing the molecular weight of the polyethylene oxide backbone increases the network mesh size. The release of a bioactive molecule such as an enzyme can be controlled by control of the network density. Therefore, a hydrogel comprised of PEGs of molecular weight 8000 daltons would have a higher rate of release of an entrapped drug than a hydrogel comprised of PEGs of molecular weight 3.3K. Furthermore, ionic moieties can be incorporated into the hydrogels to impart added functionalities such as bioadhesiveness, etc. For example, hyaluronic acid or polyacrylic acid can be added to the PEG macromer prior to crosslinking to create bioadhesive hydrogels. In another example, an ionic character can be imparted to the crosslinked hydrogels to provide molecular interaction with entrapped drugs to slow down rates of release of drug from the matrix.
PEG-hydrogels used in biosensors can provide one or more of the following features: (a) a biocompatible, non-biofouling surface appropriate for long-term exposure to biological fluids without compromise of sensor function, (b) a reservoir for glucose oxidase, (c) a matrix that can be incorporated with ionic moieties to enhance entrapment of glucose oxidase, (d) a matrix that can be modulated in terms of its physical and chemical properties (network density, swelling) by varying the molecular weight of the backbone and (e) a matrix that can be rendered bioadhesive by addition of ionic excipients such as chitosan gluconate, polyacrylic acid, poly(amidoamine), poly(ethyleneimine) and hyaluronic acid.
Vinyl acetate based hydrogels, such as n-vinylpyrolidone/vinyl acetate copolymer, can exhibit features such as transparency, tackiness, non-toxicity, flexibility, and/or hydrophobicity. Vinyl acetate based hydrogels typically have a good ability to retain moisture and entrap enzymes such as glucose oxidase, biocompatibility, and tackiness to skin to improve skin-sensor coupling. A glucose flux sensor using n-vinylpyrolidone/vinyl acetate copolymer as the hydrogel material shows good performance in tracking the plasma glucose levels of a patient with diabetes during a glucose clamping study.
The following examples set forth exemplary hydrogels that can be used with transdermal analyte monitoring according to embodiments of the present invention.
Vinyl acetate based hydrogels for use with glucose monitoring can be prepared as follows. A 1:1 mixture of n-vinylpyrolidone and vinyl acetate can be polymerized by ultraviolet radiation using 0-0.5% Irgacure as the photoinitiator. A non-woven plastic scrim (such as Delstar product# RB0707-50P) is used to provide mechanic support. The hydrogel's equilibrium water content is 20-95% with its aqueous composition containing 0-1 M sodium or potassium phosphate, 0-1 M sodium chloride, 0-1 M potassium chloride, 0-2 M lactic acid, surfactant such as 0-1 M Triton X-100, Tween 80 or sodium lauryl sulfate, and any other biocompatible components. Glucose oxidase can be loaded by soaking the solid hydrogel layer in concentrated glucose oxidase solution for a period of time.
A particular example of a vinyl acetate based hydrogel was made with the following constituents: 15% n-vinylpyrolidone, 15% vinyl acetate, 0.05% Irgacure, 0.05 M potassium phosphate, 0.30 M sodium chloride, 0.025 M potassium chloride, 0.5 M lactic acid, 0.1% Triton X-100, 0.5% GOx, and the remaining composition is water, approximately 65%
The continuous transdermal analyte monitoring system according to an exemplary embodiment of the present invention was used to reliably predict hypoglycemia (blood glucose<70 mg/dl) with 96% specificity and 77% sensitivity using a vinyl acetate hydrogel. In a study, thirty six glucose flux biosensors (3 per patient) were placed on the skin of twelve adults with either Type 1 or Type 2 diabetes. Patient data for participants in the study are shown in FIG. 24. Blood glucose measurements were collected over an eight hour period. These measurements included collecting current versus time data from the patients using a continuous transdermal analyte monitoring system as described herein. The blood glucose of each patient was rapidly increased or decreased through the administration of insulin or glucose intravenously in a controlled manner at a rate of change two times greater than that usually experienced by patients with diabetes. Specifically, the ranges tested were 35-372 mg/dl blood glucose, with a rate of glucose concentration decrease of 21 mg/(dl*min) and rate of glucose concentration increase of 11 mg/(dl*min). As a control, blood glucose measurements were collected from an intravenous catheter. A total of 2039 sensor-blood glucose data pairs from 29 data sets were generated. Five of the data sets had significant noise as shown in FIG. 25. The typical data set, however, kept noise below an excessive level as shown, for example, in FIG. 26. The data sets were analyzed with both an individually optimized algorithm and an independent algorithm, and the results are shown in FIGS. 27 and 28, respectively. The individually optimized algorithm used each data set's optimal lag time and baseline for data analysis. The independent algorithm was developed from a separate glucose clamping study, from which a single lag time value and a single baseline value were found, then were used in the algorithm for data analysis. As will be described below in connection with FIG. 17, an additional algorithm can also be utilized to compensate for temperature change and sensor drift. Completed data sets from the glucose biosensors showed a 90 percent (R=0.9) correlation to blood glucose measurements obtained via intravenous catheter over a period of 8 hours. Ninety six percent of the sensor-blood glucose pairs fell within the A+B regions in the Clark Error Grid. Seventy seven percent (164 out of 212) hypoglycemic events (BG<70 mg/dL) were successfully predicted. Sonication treatment (using Sonoprep) averaged 15 seconds and the glucose sensor required only 89±6 minutes on average to break in. No pain or irritation was reported during the sonication procedure. Accordingly, the glucose biosensor was able to reliably predict hypoglycemia (blood glucose<70 mg/dl) with 96% specificity and 77% sensitivity.
Agarose based hydrogels for use with glucose monitoring were prepared as follows. 0.0116 g of sodium chloride, 0.015 g of potassium chloride, 0.0348 g of dibasic potassium phosphate and 0.002 g of Triton X-100 were dissolved in 10 mL of water. The pH of the solution was adjusted to 7.0 using 0.5 M hydrochloric acid with the aid of a pH meter. The solution was diluted with water to 20 mL. This was Solution A. 0.2 g of agarose powder was mixed and dispersed in Solution A. Agarose was heated and dissolved until boiling in a water bath. This was Solution B. Solution B was allowed to cool down to 35° C. 0.01 g of glucose oxidase powder was completely mixed and dissolved in Solution B. This was Solution C. Solution C was cast and filled onto a warm, flat mold surface. The mold was transferred to room temperature or lower to form gels.
FIG. 12 shows sensor signal response as a function of glucose concentration for two types of agarose hydrogels relative to a polyethylene oxide polymer, and a n-vinyl pyrolidone/vinyl acetate copolymer. It can be seen from FIG. 12 that agarose offers improved signal response relative to polyethylene oxide polymer and n-vinyl pyrolidone/vinyl acetate copolymer.
Agarose based hydrogels for use with glucose monitoring can also be prepared as follows. Mix and disperse 0.2 g of agarose powder in water. Heat and dissolve agarose until boiling in a water bath. Cast and fill the solution onto a warm, flat mold surface. Transfer the mold to room temperature or lower to form gels. Dissolve 0.01 g of glucose oxidase powder in Solution A to form Solution D. Soak the gel in Solution D overnight or longer to ensure sufficient loading of glucose oxidase in the gel.
PEG-diacrylate (PEGDA) hydrogels utilized in glucose monitoring were prepared according to the following procedures.
10% weight/volume (“w/v”) solutions of (100 mg/ml) PEG2K-diacrylate, PEG3.4K-diacrylate and PEG8K-diacrylate (SunBio, Korea) were prepared in 0.01M phosphate buffered saline (PBS), pH 7.4 (ultrapure, Spectrum Chemicals, Gardena, Calif.). The solutions all contained Irgacure 2959 (Ciba Specialty Chemicals, Tarrytown, N.Y.) as the photoinitiator. Irgacure concentrations were varied to determine the effect of photoinitiator concentration on gel strength. Similarly, the polymer molecular weights were varied (2K, 3.4K, 8K) to determine the effect of molecular weight on the strength of the gelled network. As used herein, the notation “PEG2K” refers to PEG having a molecular weight of 2,000, etc.
100 mg of dry polymer was weighed into a scintillation vial. 900 μl of phosphate buffered saline (PBS) containing 500 ppm of Irgacure 2959 was added to the vial and the final weight of the solution was recorded. The vial was screw-capped and the vial swirled gently to dissolve the PEGDA. The gel solution was stored in the drawer (in the dark) for 5 minutes to ensure homogeneity. 900 μl of the gel solution was placed between two glass plates (250μ spacers) and clamped. The glass assembly containing the polymer solution was placed under a UV Blak-Ray lamp, at an intensity of 15-20 mW/cm<sup>2 </sup>and photo-crosslinked between 5-30 minutes. The gel was removed carefully from the glass and weighed before transferring to 10 ml of PBS in a plastic petri dish. After removal from the glass plates, the hydrogels were placed in approximately 10 ml of PBS. The hydrogels were then qualitatively assessed for bulk gel properties such as brittleness, gel strength and photo-yellowing as a function of molecular weight and initiator concentration.
The following procedure was used to measure the equilibrium hydration of the gels. The gels were weighed after curing was complete. The initial weight of the gel was obtained, post wiping gently with a Kim-wipe. 10 ml of PBS was added to the petri dish containing the gels. The petri dishes were placed on an orbital shaker. The buffer was replaced at pre-determined time intervals. The retrieved buffer solutions were saved to analyze for residual Irgacure. At each time interval, the gel was wiped dry with a Kim Wipe and weighed. The percent swelling (% hydration) was calculated by the change in total weight as compared to the initial weight of the gel.
By qualitative assessment, the gels varied in gel strength in the following order (strongest gel to weakest gel): PEG8K>PEG3.4K>PEG2K. Gel strength was ascertained by degree of pliability, ease of handling, and brittleness. Gel strength was also noted to vary with concentration of the photoinitiator, with higher concentrations yielding hydrogels that were hard and brittle. Photoyellowing from Irgacure photoinitiation was noted in hydrogels in the following order (most photoyellowing to least photoyellowing): 5000 ppm>2500 ppm>1500 ppm>500 ppm. The photoinitiator concentration of 500 ppm and a PEGDA molecular weight of 8K resulted in the highest gel strength.
The following procedures were performed to incorporate glucose oxidase (GOx) into the gels. First, the gels were tested for residual Irgacure 2959. Next a glucose oxidase solution was prepared. The glucose oxidase was then loaded into the PEGDA hydrogels. The glucose oxidase concentration in the gels was measured. Lastly, the bioactivity of the gels was measured. The following describes these steps in detail.
The hydrogels were washed twice with buffer until there was no detectable residual Irgacure extracted from the hydrogels. The wash solutions were scanned on the UV-Vis from 200-400 nm, for the presence of Irgacure 2959. Non-detectable levels of Irgacure were determined to be an absorbance at 280 nm<0.010, equivalent to 0.13 ppm, as compared to a 25 ppm Irgacure solution that had an absorbance of 1.8 at 280 nm.
An LPT buffer solution was prepared by mixing 5% w/v glucose oxidase in PBS solution with 0.25 M lactic acid and 0.05% Triton X-100. This was accomplished by adding 0.5 grams of GOx to a total volume of 10 ml of a stock solution comprised of 0.25 M lactic acid and 0.05% Triton X-100 dissolved in PBS. The solution was kept at 4° C.
PEGDA hydrogels comprised of varying PEG molecular weights (2K, 3.4K, 8K) were soaked in the glucose oxidase solution. The gels were soaked for overnight or longer at 4° C., but no more than seven days.
Glucose oxidase concentrations were measured by the Bradford Assay, a method commonly used to determine concentrations of solubilized protein. The method involves addition of an acidic blue dye (Coomassie Brilliant Blue G-250) to a protein solution. The dye binds primarily to basic and aromatic amino acid residues, especially arginine, with the absorption maximum shifting from 465 nm to 595 nm with complete dye-protein binding. The molar extinction coefficient of the dye-protein complex has been determined to be constant over a 10-fold concentration range; therefore, Beer-Lambert's Law can be utilized to accurately determine concentrations of protein. A standard curve of glucose oxidase solutions at concentrations 0.125%, 0.25%, 0.375%, 0.5% and 2.5% w/v was obtained by UV-Vis Spectroscopy at 595 nm after treatment of the standard solutions and the gel fragments with standard Bradford protein assay dye procedure. See Bradford Assay, BioRad Laboratories Brochure. A linear correlation of 0.999 was obtained for the standard curve. GOx incorporation in the hydrogels was determined in the following method: (a) a piece of gel was soaked in 4 ml LPT solution containing 1 ml of protein assay dye concentrate, (b) A piece of GOx-soaked then dyed (Coomassie dye) hydrogel was sandwiched between two glass cuvettes, (c) a non-GOx soaked and dyed hydrogel was used in the reference cell, (d) The gels were scanned from 400-800 nm and (e) the concentration of GOx incorporated in the hydrogels were calculated from Beer Lambert's Law: A=εbc, where A=absorbance, ε=molar extinction coefficient, b=path length and c=concentration of the analyte. Concentrations of glucose oxidase incorporated in 2K, 3.4K and 8K molecular weight PEG hydrogels were determined. FIG. 18(a) is a UV-Vis spectrum of a standard glucose oxidase solution. FIG. 18(b) is an UV-Vis spectrum of Coomassie-bound glucose oxidase. The concentration in the gels is approximately 0.6%.
Electrochemical sensors were used to test the enzymatic activity of the hydrogel-incorporated GOx. Prior to the placement on sensor, the PEGDA hydrogels are cut to the diameter of the sensor surface and rinsed briefly in LPT to remove surface residual GOx. Solutions of glucose (0.25 and 0.50 mg/dl) in PBS were used as the standard test solutions and solutions of hydrogen peroxide (20 and 55 M) in PBS were used as the positive controls. Hydrogen peroxide, the reaction product of glucose and GOx, produced amperometric current, which was recorded by a potentiostat connecting to the sensor. Therefore, positive sensor signal in response to a glucose challenge (addition of glucose) indicates that the incorporated enzyme was bioactive, while a positive sensor signal in response to a hydrogen peroxide challenge (addition of hydrogen peroxide) indicates that the eletrochemical sensor is functioning. PEGDA hydrogels with incorporated GOx were tested for peak signal strength and baseline stability. These tests demonstrate that all hydrogels (2K, 3.4K, 8K) contain bioactive GOx, and that 2K and 3.4K are advantageous for signal strength and baseline stability (See FIGS. 19-20). FIG. 19 shows the signal response to glucose of glucose oxidase loaded PEG gels of varying molecular weight. FIG. 19 demonstrates that the PEG gels contain bioactive GOx and that 2K and 3.4K molecular weight PEG hydrogels are advantageous for signal strength and baseline stability. FIG. 20 shows signal response to glucose of PEG3.4K-diacrylate hydrogel loaded with varying concentrations of GOx in the gel as well as for GOx immobilized on the sensor surface. The label “n” in FIGS. 19-20 corresponds to the number of data sets that were taken with respect to each condition tested. FIG. 21 shows the raw data of the potentiometric signals elicited from PEGDA hydrogels with GOx incorporated in the gel formulation prior to photocrosslinking. The data from FIG. 21 demonstrates that hydrogels with a thickness of 400 μm had significant non-Gaussian peak shapes and tailing relative to gels at 200 μm, which is indicative of slow diffusion of glucose and hydrogen peroxide through the hydrogel. FIG. 22 shows the change in signal between GOx-presoaked versus pre-incorporated, i.e., preloaded, hydrogels at different gel thickness and gel compositions (PEGDA-nVP, PEGDA). Among the variations of gels tested were PEGDA hydrogels at varied thickness (200 μm, 400 μm) and PEGDA-nVP at 200 μm. The data from FIG. 22 demonstrates that the GOx incorporated in the hydrogels is bioactive. Baseline stability was acceptable for all formulations and signals were not compromised.
The following describes ex vivo glucose testing on a patient with diabetes using GOx loaded PEGDA hydrogel in a complete sensor assembly. The ultrasonic skin permeation procedure, sensing mechanism, sensor configuration and protocols for clinical trials are described in Chuang H, Taylor E, and Davison T., “Clinical Evaluation of a Continuous Minimally Invasive Glucose Flux Sensor Placed Over Ultrasonically Permeated Skin,” Diabetes Technology & Therapeutics, 6:21-30 (2004). In this clinical trial, PEGDA3.4K and pure platinum were used as the hydrogel and sensor materials, respectively.
Glucose sensor function using PEGDA hydrogel is shown in FIGS. 23(a)-(b). FIG. 23(a) shows an example of sensor signal (nA) responding continuously to changes of blood glucose (BG) levels in a glucose-clamping clinical study over a period of seven hours. The corresponding nA-BG correlation plot shown in FIG. 23b has a Perason's correlation coefficient R=0.9476 (R<sup>2 </sup>square=0.8979), revealing excellent sensor's function to monitor BG levels. Use of GOx loaded PEGDA hydrogel enables successful, continuous transdermal glucose monitoring.
PEG-diacrylate-n-vinyl pyrrolidone-GOx hydrogels (PEGDA-NVP) for use with glucose monitoring were prepared according to the following procedures. PEGDA-NVP are slightly cationic, which provides ionic interaction that retains GOx. Incorporating GOx within the hydrogel prior to crosslinking also contributes to physical entrapment of GOx in the matrix. PEGDa-NVP hydrogels were prepared and characterized according to the following procedure.
100 mg of dry polymer was weighed into a tared scintillation vial. 500 μl PBS containing 1000 ppm of Irgacure 2959, 250 μl of 20% GOx in PBS, and 150 μl of 2% n-vinyl pyrrolidone (“n-VP”) was added to the vial and the final weight of the solution was recorded. The vial was screw-capped and the vial swirled gently to dissolve the PEGDA. The gel solution was stored in the drawer (in the dark) for 5 minutes to ensure homogeneity. 900 μl of the gel solution was placed between two glass plates (200μ spacers) and clamped. The glass assembly containing the polymer solution was placed under an UV Blak-Ray lamp, at an intensity of 15-20 mW/cm<sup>2 </sup>and cured for 5 minutes. The gel was removed carefully from the glass and weighed before transferring to 10 ml of LPT in a plastic petri dish.
The 200 micron hydrogels were transparent, easy to handle, pliable with considerable gel strength, as assessed qualitatively. Water content of the hydrogels were approximately 90%. The GOx was incorporated in the hydrogels prior to crosslinking, resulting in semi-interpenetrating networks. The hydrogels retained their yellow color (due to the GOx), post hydration. This indicated higher retention of the enzyme within the hydrogel.
Bioactivity of the incorporated enzyme was determined by potentiometry. This experiment demonstrated that glucose oxidase incorporated with PEG diacrylate-n-vinyl pyrrolidone hydrogels is bioactive and chemically compatible with the hydrogel delivery system. Data in FIGS. 21-22 demonstrate that GOx incorporated within the hydrogels are bioactive and functional.
PEG-diacrylate/Polyethyleneimine (PEGDA-PEI) hydrogels for use with glucose monitoring can be prepared according to the following procedures. PEGDA-PEI are cationic hydrogels. Polyethyleneimine (branched, or dendrimer, Sigma Chemicals) can be incorporated within PEG diacrylate hydrogels to impart cationic character. A cationic hydrogel can ionically interact with slightly anionic glucose oxidase to provide a controlled release reservoir for the enzyme. A solution comprised of 0.3-0.5% PEI, 10% PEGDA, 500 ppm Irgacure 2959 and 5% glucose oxidase can be photocrosslinked with a BlakRay UV light, as described in previous sections. Incorporation of the highly cationic PEI can provide a high-binding substrate for GOx resulting in enhanced retention of the enzyme in the matrix. Furthermore, the highly cationic character of the hydrogels can provide the added functionality of bioadhesivity to the skin. Other cationic, bioadhesive macromolecules that can be incorporated into PEGDA hydrogels are chitosan, polyamidoamine, poly(n-vinyl pyrrolidone), etc.
According to another aspect of the invention, an error correction method can be utilized to correct for sensor drift in a measured blood glucose value as a function of time. FIG. 16 shows a Clark Error Grid without the error correction method to correct for sensor drift. The data in FIG. 16 were taken from ten ex vivo tests on diabetic subjects in a clinical trial. The different data labels indicate data from different patients. FIG. 17 shows the Clark Error Grid after application of the error correction method to correct sensor drift. The data in FIG. 17 were taken from ten ex vivo tests on diabetic subjects in a clinical trial. The error correction method is described below.
The sensor signal, Y, as a function of time, t, is related to the sensor sensitivity, m, blood glucose value, X, and a constant offset value, b, according to the following linear relationship:<FORM>Y=mX(t)+b </FORM>
The above equation can be rearranged, and the blood glucose value can be conveniently predicted with a single point calibration protocol as follows:<FORM>X(t)=(Y−b)/m, and m=(Yc−b)/Xrc(t) </FORM>
The value of sensor sensitivity, m, can be found from each ex vivo study using the sensor's current reading Yc and a standard reference blood glucose value Xrc(t) at the sensor calibration time point. When comparing subsequent blood glucose value, X(t), with corresponding standard reference blood glucose value Xr(t), it is found that a drift factor D(t) can be computed at different points as follows:<FORM>D(t)=Xr(t)/X(t) </FORM>
By plotting D(t) vs. time, t, from a bulk number of successful ex vivo studies, a best fit for the D(t) vs. t plot was a third order polynomial function, which can be represented as follows:<FORM>D(t)=c*t<sup>3</sup>+d*t<sup>2</sup>+e*t+f </FORM>where c, d, e, f are numerical coefficients calculated to provide the best fit for the D(t) vs. t data to the above third order polynomial. The use of a third order polynomial is, however, exemplary and other methods of representing the drift factor such as an algorithm fitting the drift data to an exponential function, or utilizing a direct look-up table method can also be utilized.
To predict a drift-corrected blood glucose value Xp(t) at time t, one can simply multiply X(t) by D(t) as follows:<FORM>Xp(t)=X(t)*D(t)=X(t)*(c*t<sup>3</sup>+d*t<sup>2</sup>+e*t+f) </FORM>
This equation represents an error correction method, and its utility may be appreciated by a comparison of the Clark Error Grid where the algorithm is not applied (FIG. 16) versus where it is applied (FIG. 17). The negative bias and wide scattering of data pairs in FIG. 16 is effectively corrected, and as a result all data points fall in the clinically relevant A and B regions in the Clark Error Grid, as shown in FIG. 17. This error correction method may be applied to data generated using the continuous transdermal analyte monitoring system according an exemplary embodiment of the present invention.
Sontra Medical Corporation (Echo Therapeutics Incorporated)
A61L 31/048 : obtained by reactions only ...
A61L 31/145 : Hydrogels or hydrocolloids
C08L 2203/02 : for biomedical use
C08L 33/24 : Homopolymers or copolymers ...
C08L 33/26 : Homopolymers or copolymers ...
Current Assignee: Echo Therapeutics Incorporated
Sponsoring Entity: SONTRA MEDICAL, L.P.
System, Method, And Device For Non Invasive Body Fluid Sampling And Analysis
Sponsoring Entity: Sontra Medical, Inc. (Echo Therapeutics Incorporated)
Sponsoring Entity: Echo Therapeutics Incorporated