U.S. patent application number 11/223957 was filed with the patent office on 2006-05-04 for system and method for analyte sampling and analysis with error correction.
This patent application is currently assigned to Sontra Medical Corporation. Invention is credited to Han Chuang.
Application Number | 20060094944 11/223957 |
Document ID | / |
Family ID | 36262984 |
Filed Date | 2006-05-04 |
United States Patent
Application |
20060094944 |
Kind Code |
A1 |
Chuang; Han |
May 4, 2006 |
System and method for analyte sampling and analysis with error
correction
Abstract
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.
Inventors: |
Chuang; Han; (Canton,
MA) |
Correspondence
Address: |
HUNTON & WILLIAMS LLP;INTELLECTUAL PROPERTY DEPARTMENT
1900 K STREET, N.W.
SUITE 1200
WASHINGTON
DC
20006-1109
US
|
Assignee: |
Sontra Medical Corporation
|
Family ID: |
36262984 |
Appl. No.: |
11/223957 |
Filed: |
September 13, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11201334 |
Aug 11, 2005 |
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11223957 |
Sep 13, 2005 |
|
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10974963 |
Oct 28, 2004 |
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11201334 |
Aug 11, 2005 |
|
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Current U.S.
Class: |
600/347 ;
600/365 |
Current CPC
Class: |
Y10T 436/144444
20150115; C08L 33/24 20130101; A61L 31/145 20130101; A61B 5/681
20130101; A61B 5/14514 20130101; G01N 27/4166 20130101; A61B
5/14532 20130101; A61B 5/1486 20130101; C08L 2203/02 20130101; G01N
27/3273 20130101; C08L 33/26 20130101; A61L 31/048 20130101 |
Class at
Publication: |
600/347 ;
600/365 |
International
Class: |
A61B 5/05 20060101
A61B005/05; A61B 5/00 20060101 A61B005/00 |
Claims
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.
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*t.sup.3+d*t.sup.2+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*t.sup.3+d*t.sup.2+e*t+f, wherein c, d, e and
f are numerical coefficients calculated to provide a best fit for
D(t) to empirical data.
Description
[0001] 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.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] The present invention relates to non-invasive sampling of
body fluids, and, more particularly, to a system, method, and
device for non-invasive body fluid sampling and analysis.
[0004] 2. Description of the Related Art
[0005] 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,
[0006] the use of iontophoresis, and the use of ultrasound to
sample body fluid, such as blood and interstitial fluid.
[0007] 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.
[0008] Transport of analytes and body fluids can be enhanced
further by the action of a motive force. These motive forces
include, inter alia, sonophoretic, iontophoretic, electromotive,
pressure force, vacuum, electromagnetic motive, thermal force,
magnetic force, chemomotive, capillary action, and osmotic. The use
of active forces provide a means for obtaining fluid for subsequent
analysis.
[0009] 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.
[0010] Another example of a limitation is the amount of energy that
can be applied to the skin in order to create convective flow.
Extraction of usable volume of body fluid has the potential to
cause pain and skin damage with prolonged exposure to ultrasound.
In a similar manner, electro-osmotic extraction of body fluid
through stratum corneum has the potential to cause skin damage due
the need to use high current density. It is evident that there are
limitations to the use of the mentioned extraction methods when
applied to human skin.
SUMMARY OF THE INVENTION
[0011] 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.
[0012] 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.
[0013] 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.
[0014] 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.
[0015] 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.
[0016] 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.
[0017] 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.
[0018] The area of biological membrane may be made permeable using
ultrasound with controlled dosimetry. Extraction of body fluid may
be performed on the area exposed to ultrasound using osmotic
transport. The body fluid may be collected using a receiver. The
receiver may be attached to the biological membrane in a form of a
patch, a wearable reservoir, a membrane, an absorbent strip, a
hydrogel, or an equivalent. The receiver may be analyzed for the
presence of various analytes indicative of blood analytes. The
analysis may comprise the use of electrochemical, biochemical,
optical, fluorescence, absorbance, reflectance, Raman, magnetic,
mass spectrometry, infra-red (IR) spectroscopy measurement methods
and combinations thereof. The receiver may also be attached to a
secondary receiver where the concentration of analyte in the
secondary receiver is continuously maintained substantially lower
than that in the body fluid so the chemical concentration driving
force between body fluid and secondary receiver is maximized. This
may be achieved by chemical reaction or volume for dilution or
similar means. In one embodiment, the receiver and the secondary
receiver may operate on different principles (e.g., osmosis,
dilution, etc.). In another embodiment, the receivers may operate
on the same principle.
[0019] A system for non-invasive body fluid sampling and analysis
is disclosed. According to one embodiment of the present invention,
the system includes a controller that controls the generation of
ultrasound; an ultrasonic applicator that applies the ultrasound to
an area of biological membrane; a receiver that contacts the area
of biological membrane and receives body fluid through and out of
the area of biological membrane; and a meter that interacts with
the receiver and detects the presence of at least one analyte in
the body fluid in the receiver. The receiver may include a membrane
and a medium, such as a hydrogel, a fluid, or a liquid, that is
contained within the membrane.
[0020] A method for noninvasive body fluid sampling and analysis is
disclosed. According to one embodiment of the present invention,
the method includes the steps of (1) enhancing a permeability level
of an area of biological membrane; (2) attaching a receiver to the
area of biological membrane; (3) extracting an analyte through and
out of the area of biological membrane; (4) collecting the body
fluid in the receiver; and (5) determining a concentration of at
least one analyte in the body fluid.
[0021] A device for noninvasive body fluid sampling and analysis is
disclosed. According to one embodiment of the present invention,
the device includes a receiver that is attached to an area of
biological membrane with an enhanced permeability and receives body
fluid through and out of the area of biological membrane, and a
wearable meter that detects the presence of at least one analyte in
the received body fluid and indicates a concentration of that
analyte. The receiver may include a membrane and a medium, such as
a hydrogel, a fluid, or a liquid, that is contained in the
membrane. The meter may include a processor and a device that
detects the presence of the analyte. The detecting device may
include an electrochemical detector; a biochemical detector; a
fluorescence detector; an absorbance detector; a reflectance
detector; a Raman detector; a magnetic detector; a mass
spectrometry detector; an IR spectroscopy detector; and
combinations thereof.
[0022] According to one embodiment of the present invention,
osmotic forces may be used to sample body fluid from and through a
biological membrane in an on-demand manner. The osmotic agent in
solution, gel, hydrogel, or other form may be applied to the
ultrasound-treated biological membrane using a receiver, such as a
thin liquid reservoir, whenever the concentration of an analyte
needs to be determined for diagnosis and monitoring. The receiver
may be attached to the biological membrane using an adhesive. The
receiver may be attached to the biological membrane for a brief
duration. The solution in the receiver may be subsequently removed
and analyzed for the presence of analytes. In one embodiment, the
receiver may be constructed in the form of a patch. The receiver
may contain a hydrogel and osmotic agent. The receiver may combine
the osmotic agent and the chemical reagents to detect the presence
of the analyte. The reagents may allow the use of electrochemical,
biochemical, optical, fluorescence, absorbance, reflectance, Raman,
magnetic, mass spectrometry, infrared (IR) spectroscopy measurement
methods and combinations thereof to be performed on the
receiver.
[0023] In another embodiment, osmotic forces may be used to sample
body fluid from or through a biological membrane in a periodic or a
continuous manner. The osmotic agent in solution form may be
applied to the ultrasound-treated biological membrane using a thin
receiver, such as a thin liquid reservoir, whenever the
concentration of analyte needs to be determined for diagnosis and
monitoring. The receiver may be attached to biological membrane
using an adhesive. In one embodiment, the receiver may be
constructed in the form of a patch. The receiver may contain a
hydrogel that contains the osmotic agent. The receiver may contain
means for manipulating the intensity and duration of the osmotic
force. The intensity of the osmotic force may be manipulated using
electric field forces, magnetic field forces, electromagnetic field
forces, biochemical reactions, chemicals, molarity adjustment,
adjusting solvents, adjusting pH, ultrasonic field forces,
electro-omostic field forces, iontophoretic field forces,
electroporatic field forces and combinations thereof. The duration
of the osmotic force may be manipulated using electric field
forces, magnetic field forces, electromagnetic field forces,
biochemical reactions, chemicals, molarity adjustment, adjusting
solvents, adjusting pH, ultrasonic field forces, electroomostic
field forces, iontophoretic field forces, electroporatic field
forces and combinations thereof. The receiver may combine the
osmotic agent and the biochemical reagents to detect the presence
of the analyte. The reagents may allow the use of electrochemical,
biochemical, optical, fluorescence, absorbance, reflectance, Raman,
magnetic, mass spectrometry, IR spectroscopy measurement methods
and combinations thereof to be performed on the receiver. The
receiver may also be removed periodically for detection.
[0024] In one embodiment, the intensity, duration, and frequency of
exposure of biological membrane to osmotic forces may be
manipulated by using an electric current to cause a change in the
concentration of the osmotic agent that is in contact with the
ultrasound-exposed biological membrane. The osmotic agent may be a
multi-charged agent that can dissociate into several charged
species. These charged species may be transported using electric
field forces. A membrane may be used to isolate the charged
species. The charged species freely diffuse and combine upon
removal of the electric field force.
[0025] In one embodiment, the intensity, duration, and frequency of
exposure of biological membrane to osmotic forces may be
manipulated by using active forces to cause a change in the
concentration of the osmotic agent that is in contact with the
ultrasound-exposed biological membrane. The osmotic agent may be a
neutral charge agent. The agent may be transported using a variety
of field forces. The field force depends on the constitutive and
colligative properties of the chosen agent. The field force
generates a force necessary to move the osmotic agent toward and
away from the biological membrane surface. The movement of the
osmotic agent modulates the periodic and continuous extraction of
body fluid through the stratum corneum.
[0026] In one embodiment, the intensity, duration, and frequency of
exposure of biological membrane to osmotic forces may be
manipulated by changing the concentration of the osmotic agent that
is in contact with the ultrasound-exposed biological membrane.
Manipulating the volume of the solvent and the volume of the
hydrogel containing the osmotic agent may cause a change in the
concentration of the osmotic agent. The volume of the hydrogel can
be changed by constructing a hydrogel wherein its volume is
sensitive to the concentrations of molecules that can diffuse into
the gel. One example is a hydrogel constructed to be sensitive to
the molecule glucose. The hydrogel volume can also be changed by
manipulating its temperature and by changing the pH of the gel.
[0027] A receiver that is attached to an area of biological
membrane with an enhanced permeability and receives body fluid
through and out of the area of biological membrane is disclosed.
According to one embodiment of the present invention, the receiver
includes a first grid; a medium layer comprising at least one
agent; a membrane that induces a concentration gradient barrier for
the at least one agent; a counter grid; an oxidase layer; a
detection layer; and a voltage source that provides a potential
difference between the first grid and the counter grid. The body
fluid, which may include blood, interstitial fluid, analyte, and
lymph, may flow out of, or through, the biological membrane, to the
detector layer via the first grid, the counter grid, and the
oxidase layer.
[0028] It is a technical advantage of the present invention that a
system, method, and device for non-invasive sampling and analysis
of body fluids is disclosed. It is another technical advantage of
the present invention that a concentration of an analyte may be
measured continuously or periodically.
BRIEF DESCRIPTION OF THE DRAWINGS
[0029] For a more complete understanding of the present invention,
the objects and advantages thereof, reference is now made to the
following descriptions taken in connection with the accompanying
drawings in which:
[0030] FIG. 1 is a flowchart depicting a method for non-invasive
body fluid sampling according to one embodiment of the present
invention;
[0031] FIG. 2 depicts a device for controlled application of
ultrasound to a biological membrane to enhance the permeability of
the biological membrane according to one embodiment of the present
invention;
[0032] FIG. 3 depicts the components to perform discrete extraction
and measurement of body fluid to infer analyte concentrations
according to one embodiment of the present invention;
[0033] FIG. 4 depicts the components to perform continuous
extraction and measurement of body fluid to infer analyte
concentrations according to one embodiment of the present
invention;
[0034] FIG. 5 depicts an approach to periodic monitoring of an
analyte by performing periodic osmotic extractions of body fluid
according to one embodiment of the present invention;
[0035] FIG. 6 depicts the components of a wearable extraction
chamber according to one embodiment of the present invention;
[0036] FIG. 7 depicts a graph of glucose flux versus blood glucose
concentration according to one embodiment of the present
invention;
[0037] FIG. 8 depicts a flow chart of a method for controlled
enhancement of transdermal delivery according to one embodiment of
the present invention;
[0038] FIG. 9 depicts an apparatus for performing continuous
transdermal analyte monitoring according to one embodiment of the
present invention;
[0039] FIG. 10 is a drawing of the sensor body shown in FIG. 9 from
a first view;
[0040] FIG. 11 is a drawing of the apparatus shown in FIG. 9 from a
second view;
[0041] FIG. 12 shows the signal response versus glucose
concentration for various hydrogels;
[0042] FIG. 13 shows the signal response versus glucose
concentration for pure platinum versus platinized carbon as the
working electrode;
[0043] 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;
[0044] 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;
[0045] FIG. 16 shows a Clark Error Grid in the absence of an error
correction method according to one embodiment of the invention;
[0046] FIG. 17 shows a Clark Error Grid after the application of an
error correction method according to an embodiment of the
invention;
[0047] FIG. 18 shows the absorbance spectrum of a standard glucose
oxidase solution before and after incorporation into a PEGDA 3.4K
gel;
[0048] FIG. 19 shows the signal response to glucose of glucose
oxidase (GOx) loaded PEG gels of varying molecular weight;
[0049] FIG. 20 shows signal response to glucose of 3.4K PEG
hydrogel loaded with varying concentrations of GOx;
[0050] FIG. 21 shows raw data of the potentiometric signals
elicited from PEGDA hydrogels with GOx incorporated in the gel
formulation prior to photocrosslinking;
[0051] FIG. 22 shows the change in signal between GOx-presoaked
versus pre-incorporated hydrogels at different thickness and
compositions (PEGDA-nVP, PEGDA);
[0052] FIG. 23(a) shows blood glucose versus time utilizing an
embodiment of the continuous transdermal analyte monitoring
system;
[0053] FIG. 23(b) shows a correlation plot of electrode signal in
nanoamps versus blood glucose for an embodiment of the
invention;
[0054] FIG. 24 shows patient data for participants in a clinical
study;
[0055] FIG. 25 shows a noisy data set from the clinical study;
[0056] FIG. 26 shows another data set from the clinical study;
[0057] FIG. 27 shows a Clark Error Grid for sensor data from the
clinical study according to one embodiment of the invention;
and
[0058] FIG. 28 shows a Clark Error Grid for sensor data from the
clinical study according to another embodiment of the
invention.
DETAILED DESCRIPTION OF THE INVENTION
[0059] 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.
[0060] 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.
[0061] Although the present invention may be described in
conjunction with human applications, veterinary applications are
within the contemplation and the scope of the present
invention.
[0062] Referring to FIG. 1, a flowchart depicting a method for
non-invasive body fluid sampling and analysis according to one
embodiment of the present invention is provided. In step 102, the
permeability of an area of biological membrane is enhanced. In one
embodiment, the area of biological membrane may be located on the
volar forearm of a mammalian subject. In another embodiment, the
area of biological membrane may be located on a thigh of a
mammalian subject. In yet another embodiment, the area of
biological membrane may be located on the abdomen. In still another
embodiment, the area of biological membrane may be located on the
back. Other body locations may also be used.
[0063] In general, several techniques may be used to enhance the
permeability of the biological membrane, such as creating physical
micropores, physically disrupting the lipid bilayers, chemically
modifying the lipid bilayers, physically disrupting the stratum
corneum, and chemically modifying the stratum corneum. The creation
of micropores, or the disruption thereof, may be achieved by
physical penetration using a needle, a microneedle, a silicon
microneedle, a laser, a laser in combination with an absorbing dye,
a heat source, an ultrasonic needle, an ultrasonic transducer,
cryogenic ablation, RF ablation, photo-acoustic ablation, and
combinations thereof.
[0064] In a preferred embodiment, ultrasound may be applied to the
area of biological membrane to enhance its permeability. Ultrasound
is generally defined as sound at a frequency of greater than about
20 kHz. Therapeutic ultrasound is typically between 20 kHz and 5
MHz. Near ultrasound is typically about 10 kHz to about 20 kHz. It
should be understood that in addition to ultrasound, near
ultrasound may be used in embodiments of the present invention.
[0065] In general, ultrasound, or near ultrasound, is preferably
applied to the area of biological membrane at a frequency
sufficient to cause cavitation and increase the permeability of the
biological membrane. In one embodiment, ultrasound may be applied
at a frequency of from about 10 kHz to about 500 kHz. In another
embodiment, ultrasound may be applied at a frequency of from about
20 kHz to about 150 kHz. In yet another embodiment, the ultrasound
may be applied at 50 kHz. Other frequencies of ultrasound may be
applied to enhance the permeability level of the biological
membrane.
[0066] In one embodiment, the ultrasound may have an intensity in
the range of about 0 to about 100 watt/cm.sup.2, and preferably in
the range of 0 to about 20 watt/cm.sup.2. Other appropriate
intensities may be used as desired.
[0067] Techniques for increasing the permeability of a biological
membrane are disclosed in U.S. Pat. No. 6,190,315 to Kost et al.,
the disclosure of which is hereby incorporated by reference in its
entirety.
[0068] In step 104, body fluid is extracted through or out of the
area of biological membrane. In one embodiment, an external force,
such as an osmotic force, may assist in the extraction. In one
embodiment, the osmotic force may be controlled before, during, and
after the permeability of the biological membrane is enhanced.
[0069] In one embodiment, the osmotic force may be generated by the
application of an osmotic agent to the area of biological membrane.
The osmotic agent may be in the form of an element, a molecule, a
macromolecule, a chemical compound, or combinations thereof. The
osmotic agent may also be combined with a liquid solution, a
hydrogel, a gel, or an agent having a similar function.
[0070] In step 106, the magnitude, intensity, and duration of the
external force may be regulated by at least one additional first
energy and/or force. In one embodiment, the first additional energy
and/or force may be applied to control and regulate the movement
and function of the osmotic agent for extraction of body fluid
through and out of the biological membrane. The first additional
energy and/or force may be provided in the form of heat, a
temperature force, a pressure force, an electromotive force, a
mechanical agitation, ultrasound, iontophoresis, an electromagnetic
force, a magnetic force, a photothermal force, a photoacoustic
force, and combinations thereof. The effect of an electric field
and ultrasound on transdermal drug delivery is disclosed in U.S.
Pat. No. 6,041,253, the disclosure of which is incorporated, by
reference, in its entirety.
[0071] In one embodiment, if the first additional energy and/or
force is provided by ultrasound, the frequency of the ultrasound
may be provided at a different frequency than the frequency used to
enhance the permeability of the biological membrane. In one
embodiment, the frequency of the first additional energy/force
ultrasound may be higher than the frequency of the permeability
enhancing ultrasound.
[0072] In step 108, the body fluid may be collected in a receiver.
In one embodiment, the receiver may be contacted with the
biological membrane in a form of a patch, a wearable reservoir, a
membrane, an absorbent strip, a hydrogel, or a structure that
performs an equivalent function. Other types and configurations of
receivers may be used.
[0073] In one embodiment, the receiver may be provided with a
secondary receiver having an analyte concentration that is
continuously maintained to be substantially lower than the analyte
concentration in the body fluid, so the chemical concentration
driving force between body fluid and secondary receiver is
maximized. This may be achieved by chemical reaction or volume for
dilution or similar means.
[0074] In one embodiment, a second external energy/force may be
applied between the first receiver and the secondary receiver. In
one embodiment, the second external energy/force may be different
(e.g., a different type of external force) from the first external
energy/force. In another embodiment, the second external
energy/force may be the same (e.g., the same type of external
force) as the first external energy/force. The first and second
external energy/force may vary in type, duration, and intensity,
and may be controlled through different additional energy and/or
forces.
[0075] In step 110, the collected body fluid may be analyzed. In
one embodiment, the analysis may include the use of appropriate
methods, such as electrochemical, biochemical, optical,
fluorescence, absorbance, reflectance, Raman, magnetic, mass
spectrometry, infra-red (IR) spectroscopy measurement, and
combinations thereof.
[0076] In one embodiment, multiple analytes may be analyzed
simultaneously, in parallel, or in series. The results from these
multiple analyses may be used in combination with algorithms, for
example, to increase the accuracy, or precision, or both, of the
analysis and measurements.
[0077] In one embodiment, the receiver may be removed from contact
with the biological membrane in order to analyze the collected body
fluid. In another embodiment, the receiver may remain in contact
with the biological membrane as the collected body fluid is
analyzed.
[0078] 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 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.
[0079] In addition, controller 202 may include a display, such as a
LCD or a LED display, in order to convey information to the user as
required. Controller 202 may also include a user interface as is
known in the art.
[0080] 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.
[0081] In one embodiment, the coupling medium may also include a
chemical enhancer. Transport enhancement may be obtained by adding
capillary permeability enhancers, for example, histamine, to the
coupling medium. The concentration of histamine in the coupling
medium may be in the range of between 0.1 and 100 mg/ml. These
agents may be delivered across the biological membrane during
application of ultrasound and may cause local edema that increases
local fluid pressure and may enhance transport of analytes across
the biological membrane. In addition, the occurrence of free fluid
due to edema may induce cavitation locally so as to enhance
transport of analytes across the biological membrane.
[0082] In one embodiment, cartridge 206 may be pierced when
inserted into ultrasound applicator 204, and ultrasound coupling
solution 208 may be transferred to a chamber (not shown).
[0083] A target identifying device, such as target ring 210, may be
attached to the area of biological membrane that will have its
permeability increased. Target ring 210 may be attached to the area
of biological membrane by a transdermal adhesive (not shown). In
one embodiment, target ring 210 may have the transdermal adhesive
pre-applied, and may be disposed after each use. In another
embodiment, target ring 210 may be reusable.
[0084] Target ring 210 may be made of any suitable material,
including plastic, ceramic, rubber, foam, etc. In general, target
ring 210 identifies the area of biological membrane for
permeability enhancement and body fluid extraction. In one
embodiment, target ring 210 may be used to hold receiver 214 in
contact with the biological membrane after the permeability of the
biological membrane has been increased.
[0085] 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.
[0086] Ultrasound applicator 204 may be applied to target ring 210
and activated to expose ultrasound coupling solution 208 to the
biological membrane. Controller 202 controls ultrasound applicator
204 to transmit ultrasound through ultrasound coupling solution
208. During ultrasound exposure, controller 202 may monitor changes
in biological membrane permeability, and may display this
information to the user.
[0087] Controller 202 may cease, or discontinue, the application of
ultrasound once a predetermined level of biological membrane
permeability is reached. This level of permeability may be
preprogrammed, or it may be determined in real-time as the
ultrasound is applied. The predetermined level of permeability may
be programmed for each individual due to biological membrane
differences among individuals.
[0088] After the predetermined level of permeability is reached,
ultrasound coupling solution 208 may be vacuated from chamber (not
shown) into cartridge 206, which may then be discarded. In another
embodiment, ultrasound coupling solution 208 may be vacuated into a
holding area (not shown) in ultrasound applicator 204, and later
discharged. Ultrasound applicator 204 may then be removed from
target ring 210.
[0089] Referring to FIG. 3, an device for the analysis of body
fluid according to one embodiment of the present invention is
provided. Receiver 214 may be placed into target ring 210 to
perform a discrete, or on-demand, extraction of body fluid through
and/or out of the biological membrane. Receiver 214 may contain a
medium, such as a hydrogel layer, that incorporates an osmotic
agent. In one embodiment, the hydrogel may be formulated to contain
phosphate buffered saline (PBS), with the saline being sodium
chloride having a concentration range of about 0.01 M to about 10
M. The hydrogel may be buffered at pH 7. Other osmotic agents may
also be used in place of, or in addition to, sodium chloride.
Preferably, these osmotic agents are non-irritating, non-staining,
and non-immunogenic. Examples of such osmotic agents include, inter
alia, lactate and magnesium sulfate.
[0090] In another embodiment, receiver 214 may include a fluid or
liquid medium, such as water or a buffer, that is contained within
a semi-permeable membrane. Receiver 214 may also include a spongy
material, such as foam.
[0091] Receiver 214 may be applied to the biological membrane to
contact the ultrasound exposed biological membrane. In one
embodiment, receiver 214 may be applied to the biological membrane
for a time period sufficient to collect an amount of body fluid
sufficient for detection. In another embodiment, receiver 214 may
be applied to the biological membrane for a sufficient time period
to collect a predetermined amount of body fluid. In yet another
embodiment, receiver 214 may be applied to the biological membrane
for a predetermined time. In one embodiment, the contact between
receiver 214 and the biological membrane may last for 15 minutes or
less. In another embodiment, the contact between receiver 214 and
the biological membrane may last for 5 minutes or less. In still
another embodiment, the contact between receiver 214 and the
biological membrane may last for 2 minutes or less. The actual
duration of contact may depend on the sensitivity of the detection
method used for analysis.
[0092] In one embodiment, the medium of receiver 214 may contain at
least one reagent (not shown) in order to detect the presence of
certain analytes in the body fluid that has been extracted from or
through the biological membrane. In one embodiment, the hydrogel
layer of receiver 214 may contain the reagents, and the reagents
may be attached to the hydrogel by ionic and/or covalent means, or
may be immobilized by gel entrapment. The reagents may also be
arranged as an adjacent layer to the hydrogel wherein the analyte
from the body fluid that has been extracted into the hydrogel can
diffuse into and react to generate by-products. The by-products may
then be detected using electrochemical, biochemical, optical,
fluorescence, absorbance, reflectance, Raman, magnetic, mass
spectrometry, IR spectroscopy measurement methods and combinations
thereof.
[0093] The detection methods may be performed by meter 212. Meter
212 may include a processor (not shown) and a display, such as an
LCD display. Other suitable displays may be provided.
[0094] In one embodiment, meter 212 may provide an interface that
allows information be downloaded to an external device, such as a
computer. Such an interface may allow the connection of interface
cables, or it may be a wireless interface.
[0095] Meter 212 may be configured to determine body fluid glucose
concentration by incorporating glucose oxidase in the medium of
receiver 214. In one embodiment, glucose from extracted body fluid
may react with glucose oxidase to generate hydrogen peroxide.
Hydrogen peroxide may be detected by the oxidation of hydrogen
peroxide at the surface of electrodes incorporated into receiver
214. The oxidation of hydrogen peroxide transfers electrons onto
the electrode surface which generates a current flow that can be
quantified using a potentiostat, which may be incorporated into
meter 212. A glucose concentration proportional to the
concentration of hydrogen peroxide may be calculated, and the
result may be reported to the user via a display. Various
configurations of electrodes and reagents, known to those of
ordinary skill in the art, may be incorporated to perform detection
and analysis of glucose and other analytes.
[0096] Meter 212 may also be configured to simultaneously measure
the concentration of an analyte, such as glucose, where the body
fluid concentration is expected to fluctuate, and an analyte, like
creatinine or calcium, where the body fluid concentration is
expected to remain relatively stable over minutes, hours, or days.
An analyte concentration, which may be determined by an algorithm
that takes into account the relative concentrations of the
fluctuating and the more stable analyte, may be reported to the
user via a display.
[0097] In another embodiment, meter 212 may analyze multiple
analytes simultaneously, in parallel, or in series. The results
from these multiple analyses may be used in combination with
algorithms, for example, to increase the accuracy, or precision, or
both, of the analysis and measurements.
[0098] Receiver 214 may be discarded after the extraction and
measurement steps. In another embodiment, receiver 214 may be
reused. In one embodiment, receiver 214 may be cleaned, sanitized,
etc. before it may be reused. Various configurations of electrodes
and reagents, known to those of ordinary skill in the art, may be
incorporated to perform detection and analysis of glucose and other
analytes.
[0099] 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.
[0100] Other osmotic agents may also be used in place of, or in
addition to, sodium chloride. These osmotic agents are preferably
non-irritating, non-staining, and non-immunogenic. Examples of
these other osmotic agents may include, inter alia, lactate and
magnesium sulfate. Receiver 402 may be applied to contact the
ultrasound exposed biological membrane. In one embodiment, the
duration of this contact may be 12-24 hours, or more. In another
embodiment, other durations of contact, including substantially
shorter durations, and substantially longer durations, may be used
as desired.
[0101] In another embodiment, receiver 402 may include a fluid or
liquid medium, such as water or a buffer, that is contained within
a semi-permeable membrane. Receiver 402 may also include a spongy
material, such as foam.
[0102] In one embodiment, the medium of receiver 402 may contain at
least one reagent (not shown) that detects the presence of analytes
in the body fluid that has been extracted thorough and out of the
biological membrane. In one embodiment, the hydrogel layer of
receiver 402 may contain reagents that may be attached by ionic and
covalent means to the hydrogel, or may be immobilized by gel
entrapment. The reagents may also be arranged as an adjacent layer
to the hydrogel wherein the analyte from the body fluid that has
been extracted into the hydrogel may diffuse into and react to
generate by-products. The by-products may be detected using
electrochemical, biochemical, optical, fluorescence, absorbance,
reflectance, Raman, magnetic, mass spectrometry, IR spectroscopy
measurement methods and combinations thereof.
[0103] 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.
[0104] Meter 404 may incorporate power and electronics to control
the periodic extraction of body fluid, to detect analyte, and to
present the analyte concentration in a continuous manner. Meter 404
may contain electronics and software for the acquisition of sensor
signals, and may perform signal processing, and may store analysis
and trending information.
[0105] In one embodiment, meter 404 may provide an interface that
allows information be downloaded to an external device, such as a
computer. Such an interface may allow the connection of interface
cables, or it may be a wireless interface.
[0106] Meter 404 may be configured to determine body fluid glucose
concentration by incorporating glucose oxidase in the medium. In
one embodiment, glucose from extracted body fluid may react with
glucose oxidase to generate hydrogen peroxide. Hydrogen peroxide
may be detected by the oxidation of hydrogen peroxide at the
surface of electrodes incorporated into receiver 402. The oxidation
of hydrogen peroxide transfers electrons onto the electrode surface
which generates a current flow that can be quantified using a
potentiostat, which may be incorporated into meter 404. A glucose
concentration proportional to the concentration of hydrogen
peroxide may be calculated and the result may be reported to the
user via a display. Various configurations of electrodes and
reagents, known to those of ordinary skill in the art, may be
incorporated to perform detection and analysis of glucose and other
analytes.
[0107] In one embodiment, meter 404 may also be configured to
simultaneously measure concentration of an analyte, such as
glucose, where the body fluid concentration is expected to
fluctuate, and an analyte, like creatinine or calcium, where the
body fluid concentration is expected to remain relatively stable
over minutes, hours, or days. An analyte concentration, which may
be determined by an algorithm that takes into account the relative
concentrations of the fluctuating and the more stable analyte, may
be reported to the user via a display.
[0108] In another embodiment, meter 404 may analyze multiple
analytes simultaneously, in parallel, or in series. The results
from these multiple analyses may be used in combination with
algorithms, for example, to increase the accuracy, or precision, or
both, of the analysis and measurements.
[0109] In another embodiment, receiver 402 may be removed from
contact with the biological membrane for analysis by meter 404.
Receiver 402 may be put in contact with the biological membrane
after such analysis.
[0110] Meter 404 may provide analyte readings to the user in a
periodic or a continuous manner. For example, in one embodiment, in
continuous monitoring of the analyte glucose, glucose concentration
may be displayed to the user every 30 minutes, more preferably
every 15 minutes, most preferable every 5 minutes, or even more
frequently. In another embodiment, the glucose concentration may be
displayed continuously. The period may depend on the sensitivity
and method of analyte detection. In continuous glucose monitoring,
in one embodiment, glucose detection may be performed by an
electrochemical method using electrodes and reagents incorporated
into receiver 402 and detection and analysis performed by meter
404. During the measurement period, osmotic extraction of body
fluid may be performed continuously by the hydrogel layer of
receiver 402. Body fluid may accumulate in the hydrogel of receiver
402. Glucose in body fluid diffuses to react with glucose oxidase
and is converted into hydrogen peroxide. The hydrogen peroxide is
consumed by poising the working electrode with respect to a
reference electrode. During the resting period, hydrogen peroxide
accumulates and is consumed or destroyed before the measuring
period. The magnitude of the working potential can be applied to
rapidly consume the build up of hydrogen peroxide.
[0111] Referring to FIG. 5, an approach to periodic monitoring of
an analyte by performing periodic osmotic extractions of body fluid
according to another embodiment of the present invention is shown.
The osmotic extraction intensity and frequency may be manipulated
by using an osmotic agent that dissociates into multiple charged
species, and an electrical potential may be used to move the
concentration of charges toward and away from biological membrane
surface 550. Receiver 500 may include grid, mesh, or screen 504;
medium 506, which may be a hydrogel layer; membrane 508; counter
grid, mesh, or screen 510; oxidase layer 512; and detection layer
514. Grid 504 and counter grid 510 may be connected to voltage
source 516. Membrane 508 may be a semi-permeable membrane that is
used to induce a concentration gradient barrier for the osmotic
agent contained in medium 506. The preferable osmotic agent may
contain negative and positive species or counter ions. Manipulating
the concentration of charged species at the boundary adjacent to
the stratum corneum of the ultrasound-exposed biological membrane
may provide periodic extraction of body fluid.
[0112] In one embodiment, receiver 500 may make contact with the
skin though contact medium 502, which may be a hydrogel, or other
suitable medium.
[0113] The concentration of the charged species may be manipulated
by applying a potential difference between grid 504 and counter
grid 510 using voltage source 516. In one embodiment, the potential
difference may be of a magnitude that is sufficient to manipulate
the osmotic agent. The polarity of the grid may also be changed to
transport charges toward and away from biological membrane surface
550. Grid 504 and counter grid 510 may be configured with optimum
porosity as to allow body fluid and/or analyte to travel out of
stratum corneum, through grid 504, through grid 510, and into
oxidase layer 512, and ultimately to detection layer 514. Oxidase
layer 512 may be used with an appropriate catalyst, or enzyme, to
confer specificity of analyte detection. Detection layer 514 may
include working and reference electrodes (not shown) that allow for
the detection of the by-products of oxidase layer 512 to quantify
the concentration of the desired analyte of detection.
EXAMPLE 1
[0114] The following example does not limit the present invention
in any way, and is intended to illustrate an embodiment of the
present invention.
[0115] The following is a description of experiments which
implemented painless extraction, collection, and analysis of body
fluid to determine body fluid glucose concentration in a human
using a hyperosmotic extraction fluid and comparing this condition
with iso-osmotic extraction fluid, in accordance with one
embodiment of the present invention. Although body fluid glucose
concentration serves as an example to demonstrate feasibility,
other analytes are within the contemplation of the present
invention. In addition, multiple analytes may be measured and/or
analyzed simultaneously, in parallel, or in series, and results
from these multiple measurements may be used in combination with
algorithms, for example, to increase the accuracy or precision or
both of measurements. As may be recognized by one of ordinary skill
in the art, these steps may be automated and implemented with the
device described above.
[0116] Four sites on the volar forearm of a human volunteer were
treated with ultrasound using the device described in FIG. 2. The
ultrasound transducer and its housing were placed on the volar
forearm of the volunteer with enough pressure to produce a good
contact between the skin and the outer transducer housing, and to
prevent leaking. The area surrounding the transducer was then
filled with a coupling medium of sodium dodecyl sulfate and silica
particles in phosphate-buffered saline (PBS). Ultrasound was
briefly applied (5-30 s), the transducer apparatus was removed from
the biological membrane, and the skin was rinsed with tap water and
dried.
[0117] 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.times.
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.times.7/8'' 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.times.7/8'' OD). Thin transparent lid 606
was made of 3M Polyester 1012 ( 11/8''.times. 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.
[0118] Each extraction chamber was alternately filled with 100
.mu.l of extraction solution for 15 min and 100 .mu.l hydration
solution for 10-40 min. Extraction solution was PBS; on two sites
the PBS contained additional NaCl to bring the total concentration
of NaCl to 1 M. Hydration solution was PBS for all sites.
[0119] 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), the mass of glucose that crossed the sonicated site
per unit time per unit area. Body fluid glucose concentrations
(C.sub.bg) were obtained by testing capillary blood obtained from a
lanced finger in a Bayer Glucometer Elite meter. It was
hypothesized that Q.sub.g would be linearly proportional to
C.sub.bg. FIG. 7 shows a graph of Q.sub.g versus C.sub.bg.
Unexpectedly, Q.sub.g from the sonicated sites exposed to 1 M NaCl
correlated to C.sub.bg much more strongly than Q.sub.g from the
sonicated sites exposed to 0.15 M NaCl.
[0120] 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.
[0121] 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.
[0122] 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.
[0123] 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.
[0124] 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 is used to enhance the
permeability of the patch of skin to be used for transdermal
transport.
[0125] 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.
[0126] 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:
C=C.sub.i+(C.sub.f-C.sub.i)/(1+e.sup.S(t-t*)) where: [0127] C is
current; [0128] C.sub.i is current at t=0; [0129] C.sub.f is the
final current; [0130] S is a sensitivity constant; [0131] t* is the
exposure time required to achieve an inflection point; and [0132] t
is the time of exposure.
[0133] 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.
[0134] 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.
[0135] 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.
[0136] 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.
[0137] 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.
[0138] 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.
[0139] 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.
[0140] 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.
[0141] 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.
[0142] 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.
[0143] 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.
[0144] 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.
[0145] 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.
[0146] 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.
[0147] 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.
[0148] 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.
[0149] 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>=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=0.87 versus R.sup.2=0.71 for platinized carbon.
[0150] 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.
[0151] 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.
[0152] 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.
[0153] 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.
[0154] 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.
[0155] 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.degree. C.), followed by
cooling and gelling at lower temperature (40.degree. C. or
lower).
[0156] 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.
[0157] 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.
[0158] 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.
[0159] The following examples set forth exemplary hydrogels that
can be used with transdermal analyte monitoring according to
embodiments of the present invention.
EXAMPLE 2
[0160] 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.
[0161] 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%
[0162] 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.
EXAMPLE 3
[0163] 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.degree. 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.
[0164] 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.
EXAMPLE 4
[0165] 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.
EXAMPLE 5
[0166] PEG-diacrylate (PEGDA) hydrogels utilized in glucose
monitoring were prepared according to the following procedures.
[0167] 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.
[0168] 100 mg of dry polymer was weighed into a scintillation vial.
900 .mu.l 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 .mu.l of the gel solution was placed between two glass plates
(250.mu. 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 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.
[0169] 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.
[0170] 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.
[0171] 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.
[0172] 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.
[0173] 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.degree. C.
[0174] 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.degree. C., but no
more than seven days.
[0175] 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=.epsilon.bc, where A=absorbance, .epsilon.=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%.
[0176] 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 .mu.m had significant non-Gaussian peak shapes and tailing
relative to gels at 200 .mu.m, 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 .mu.m, 400 .mu.m) and PEGDA-nVP at 200 .mu.m. 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.
[0177] 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.
[0178] 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 square=0.8979),
revealing excellent sensor's function to monitor BG levels. Use of
GOx loaded PEGDA hydrogel enables successful, continuous
transdermal glucose monitoring.
EXAMPLE 6
[0179] 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.
[0180] 100 mg of dry polymer was weighed into a tared scintillation
vial. 500 .mu.l PBS containing 1000 ppm of Irgacure 2959, 250 .mu.l
of 20% GOx in PBS, and 150 .mu.l 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 .mu.l of the gel
solution was placed between two glass plates (200.mu. 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 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.
[0181] 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.
[0182] 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.
EXAMPLE 7
[0183] 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.
[0184] 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.
[0185] 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: Y=mX(t)+b
[0186] The above equation can be rearranged, and the blood glucose
value can be conveniently predicted with a single point calibration
protocol as follows: X(t)=(Y-b)/m, and m=(Yc-b)/Xrc(t)
[0187] 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: D(t)=Xr(t)/X(t)
[0188] 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: D(t)=c*t.sup.3+d*t.sup.2+e*t+f 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.
[0189] To predict a drift-corrected blood glucose value Xp(t) at
time t, one can simply multiply X(t) by D(t) as follows:
Xp(t)=X(t)*D(t)=X(t)*(c*t.sup.3+d*t.sup.2+e*t+f)
[0190] 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.
[0191] Other embodiments and uses of the invention will be apparent
to those skilled in the art from consideration of the specification
and practice of the invention disclosed herein. All references
cited herein, including all U.S. and foreign patents and patent
applications, are specifically and entirely hereby incorporated
herein by reference. It is intended that the specification and
examples be considered exemplary only, with the true scope and
spirit of the invention indicated by the following claims.
* * * * *