U.S. patent number 3,837,339 [Application Number 05/223,077] was granted by the patent office on 1974-09-24 for blood glucose level monitoring-alarm system and method therefor.
This patent grant is currently assigned to Whittaker Corporation. Invention is credited to Sol Aisenberg, Kuo Wei Chang.
United States Patent |
3,837,339 |
Aisenberg , et al. |
September 24, 1974 |
BLOOD GLUCOSE LEVEL MONITORING-ALARM SYSTEM AND METHOD THEREFOR
Abstract
A method and apparatus for monitoring blood glucose levels. In
the preferred embodiment a glucose diffusion-limited fuel cell is
implanted in a living body. The output current of the fuel cell is
proportional to the glucose concentration of the body fluid
electrolyte and is therefore directly indicative of the blood
glucose level. This information is telemetered to an external
receiver which generates an alarm signal whenever the fuel cell
output current exceeds or falls below a predetermined current
magnitude which represents a normal blood glucose level. Valve
means are actuated in response to the telemetered information to
supply glucose or insulin to the monitored living body.
Inventors: |
Aisenberg; Sol (Natick, MA),
Chang; Kuo Wei (Lexington, MA) |
Assignee: |
Whittaker Corporation (Los
Angeles, CA)
|
Family
ID: |
22834925 |
Appl.
No.: |
05/223,077 |
Filed: |
February 3, 1972 |
Current U.S.
Class: |
604/504; 604/66;
600/302; 128/903 |
Current CPC
Class: |
A61B
5/0002 (20130101); A61B 5/0031 (20130101); A61B
5/14532 (20130101); A61M 5/1723 (20130101); H01M
8/08 (20130101); H01M 8/04186 (20130101); Y02E
60/50 (20130101); Y10S 128/903 (20130101) |
Current International
Class: |
A61B
5/00 (20060101); A61M 5/172 (20060101); A61M
5/168 (20060101); H01M 8/04 (20060101); H01M
8/08 (20060101); A61b 005/07 () |
Field of
Search: |
;128/2E,2P,2G,2R,2.5A,2.5M,2.5R,2.1A,2.1E,2.1R,213,419B,419R
;136/86DD,86F,2.6A,2.6E,2.6R ;204/195B ;23/258.5 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
Other References
Updike et al., "Nature," Vol. 214, June 3, 1967, pp. 986-988. .
Drake et al., "Transactions-American Society for Artificial
Internal Organs," Vol. XVI, 1970, pp. 199-205..
|
Primary Examiner: Kamm; William E.
Attorney, Agent or Firm: Birch; Richard J.
Claims
What we claim and desire to secure by Letters Patent of the United
States is:
1. A method for monitoring blood glucose levels comprising the
steps of:
1. exposing a glucose diffusion-limited fuel cell to the body fluid
of a living body;
2. converting the output current generated by said fuel cell into
an electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being proportional to the blood glucose level in
said living body;
3. detecting when said electrical signal characteristic departs
from a predetermined condition; and,
4. actuating an alarm signal whenever said electrical signal
characteristic departs from said predetermined condition.
2. A method for monitoring blood glucose levels comprising the
steps of:
1. exposing a glucose diffusion-limited fuel cell to the body fluid
of a living body;
2. converting the output current generated by said fuel cell into
an electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being proportional to the blood glucose level in
said living body;
3. detecting when said electrical signal characteristic departs
from a predetermined condition;
4. introducing glucose into said living body whenever said
electrical signal characteristic departs from said predetermined
condition in one direction;
5. introducing insulin into said living body whenever said
electrical signal characteristic departs from said predetermined
condition in an opposite direction; and,
6. terminating the introduction of said glucose or insulin when
said electrical signal characteristic returns to said predetermined
condition.
3. A method for monitoring blood glucose levels comprising the
steps of:
1. exposing a glucose diffusion-limited fuel cell to the body fluid
of a living body;
2. converting the output current generated by said fuel cell into
an electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being proportional to the blood glucose level in
said living body;
3. detecting when said electrical signal characteristic departs
from a predetermined condition;
4. introducing insulin into said living body whenever said
electrical signal characteristic departs from said predetermined
condition in one direction; and,
5. terminating the introduction of said insulin when said
electrical signal characteristic returns to said predetermined
condition.
4. An in vivo blood glucose level monitoring system comprising:
1. a glucose diffusion-limited fuel cell, said fuel cell being
adapted for implantation in a living body;
2. means for converting the output current generated by said fuel
cell when implanted in a living body, into an electrical signal
having a characteristic which varies in accordance with the
magnitude of the output current, said output current magnitude
being a function of the blood glucose level in said living body;
and,
3. means for monitoring said electrical signal characteristic.
5. The system of claim 4 wherein said fuel cell comprises:
1. at least one permeable membrane which defines a chamber, said
membrane being permeable to body water, oxygen, and glucose;
2. first and second spaced catalyst coated electrodes positioned
within said chamber, said first and second electrodes comprising,
respectively, a cathode electrode and an anode electrode for said
fuel cell;
3. means for glucose diffusion-limiting said fuel cell.
6. An in vivo blood glucose level monitoring system comprising:
1. a glucose diffusion-limited fuel cell, said fuel cell being
adapted for implantation in a living body;
2. means for converting the output current generated by said fuel
cell when implanted in a living body, into an electrical signal
having a characteristic which varies in accordance with the
magnitude of the output current, said output current magnitude
being a function of the blood glucose level in said living
body;
3. means for detecting when said electrical signal characteristic
departs from a predetermined condition; and,
4. alarm signal generating means responsive to said electrical
signal characteristic for generating an alarm signal whenever said
characteristic departs from said predetermined condition.
7. An in vivo blood glucose level monitoring system comprising:
1. a glucose diffusion-limited fuel cell, said fuel cell being
adapted for implantation in a living body; and,
2. means for converting the output current generated by said fuel
cell when implanted in a living body, into an electrical signal
having a characteristic which varies in accordance with the
magnitude of the output current, said output current magnitude
being a function of the blood glucose level in said living
body;
3. means for detecting when said electrical signal characteristic
departs from a predetermined characteristic;
4. means for generating a glucose valve actuation signal whenever
said electrical signal characteristic departs in one direction from
said predetermined condition and an insulin valve actuation signal
whenever said characteristic departs in the opposite direction from
said predetermined condition;
5. first fluid valve means responsive to said glucose valve
actuation signal for supplying glucose to the body;
6. second fluid valve means responsive to said insulin valve
actuation signal for supplying insulin to the body;
7. a source of glucose fluidly coupled to said first fluid valve
means; and,
8. a source of insulin fluidly coupled to said second fluid valve
means.
8. An in vivo blood glucose detecting system comprising:
1. a glucose diffusion-limited fuel cell, said fuel cell being
adapted for implantation in a living body;
2. means for converting the output current generated by said fuel
cell when implanted in a living body, into an electrical signal
having a characteristic which varies in accordance with the
magnitude of the output current, said output current magnitude
being a function of the blood glucose level in said living
body;
2. means for detecting when said electrical signal characteristic
departs from a predetermined condition;
3. means for generating a insulin valve actuation signal whenever
said electrical signal characteristic departs in one direction from
said predetermined condition;
4. fluid valve means responsive to said insulin valve actuation
signal for supplying insulin to the body; and,
5. a source of insulin fluidly coupled to said fluid valve
means.
9. A method for producing an electrical signal representation of
blood glucose levels comprising the steps of:
1. exposing a glucose diffusion-limited fuel cell to the body fluid
of a living body; and,
2. converting the output current generated by said fuel cell into
an electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being proportional to the blood glucose level in
said living body; and,
3. utilizing said electrical signal.
10. A method for monitoring blood glucose levels comprising the
steps of:
1. exposing a glucose diffusion-limited fuel cell to the body fluid
of a living body;
2. converting the output current generated by said fuel cell into
an electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being proportional to the blood glucose level in
said living body; and,
3. monitoring said electrical signal characteristic.
11. A blood glucose level monitoring system comprising:
1. a glucose diffusion-limited fuel cell, said fuel cell being
adapted for exposure to the body fluid of a living body;
2. means for converting the output current generated by said fuel
cell when exposed to the body fluid of a living body, into an
electrical signal having a characteristic which varies in
accordance with the magnitude of the output current, said output
current magnitude being a function of the blood glucose level in
said living body; and,
3. means for monitoring said electrical signal characteristic.
Description
BACKGROUND OF THE INVENTION
This invention relates to glucose monitoring systems in general,
and more particularly to an implantable glucose monitor and alarm
system for use in the measurement and control of blood glucose
levels in diabetics.
It is important for the diabetic patient to maintain normal or near
normal blood glucose levels throughout the day. These levels can be
obtained through appropriate diets, insulin injections and exercise
patterns. However, in order to avoid over or under compensation, it
is desirable for the diabetic patient to know his blood glucose
level in order to take appropriate compensatory action.
Unfortunately, at the present time, continuous blood glucose
measurements can only be performed outside the body. Basically,
such measurements involve the following operations: using a double
lumen cannula, blood is continuously drained from a vein, mixed
with heparin solution and then sent to a dialyasis cell. The
glucose which has been dialyzed out is allowed to react with the
appropriate amount of reagent such as glucose oxidase, glucose
oxidase-HVA-peroxidase mixture, or potassium ferricyanide. The
glucose concentration is then obtained by either
spectropolarimetry, fluoresence or colorimetry depending upon the
reagent used. Blood glucose measurements using this system are
obviously time consuming and inconvenient with respect to an
ambulatory diabetic patient.
It is accordingly, a general object of the present invention to
provide a glucose monitor and alarm system for the measurement and
control of blood glucose levels of diabetics.
It is a specific object of the present invention to provide a
compact, implantable, self-sustained sensor-telemetering system
which is capable of providing measurement of blood glucose
concentrations.
It is another specific object of the invention to provide a system
which releases no harmful materials and which utilizes no toxic
chemical to interact with the blood glucose.
It is a feature of the invention that the implantable monitoring
device requires only minimal electric power.
It is still another feature of the invention that the implantable
device generates only small amounts of heat and consumes very
little glucose and oxygen.
It is still another feature of the invention that the implantable
device can be conveniently calibrated and recalibrated externally
to guard against drift and aging.
It is still another feature of the invention that the implantable
device is insensitive to fluctuations of body oxygen tension and pH
values.
In the accomplishment of these objects the glucose monitor and
alarm system of the present invention utilizes a sensor and
telemetering system which is contained within a small chamber
covered by a membrane into which body water and oxygen and glucose
can freely diffuse. The body fluid within the chamber is in
equilibration with the extra-cellular-extravascular fluid which, is
in turn, in nearly constant equilibration with blood in so far as
glucose level is concerned. Inside the chamber is a fuel cell
comprising two platinum black electrodes (or other catalyst-coated
electrodes) using glucose as a fuel and dissolved oxygen as the
oxidizer. The fuel cell is operated in an essentially glucose
diffusion-limited mode so that the output current of the fuel cell
is proportional to the glucose concentration of the body fluid and
is, therefore, directly indicative of the blood glucose level. This
information is then telemetered to a compact receiver located
externally to the body.
The glucose monitor can be implanted directly in the living body to
be monitored by the system or, alternatively, can be subcutaneously
positioned in the body by means of a hypodermic needle. The system
can also be employed as an external monitoring and alarm
system.
These objects and features and other objects and features of the
present invention will best be understood from a detailed
description of a preferred embodiment thereof, selected for
purposes of illustration, and shown in the accompanying drawings in
which:
FIG. 1 is a diagrammatic view in partial block form showing the
glucose diffusion-limited fuel cell and telemetering circuitry;
FIG. 2 is a polarization curve for a typical fuel cell;
FIG. 3 is a block diagram of the electrical circuitry of the blood
glucose monitoring system; and,
FIG. 4 is a diagrammatic view in partial block form showing the
alarm and control circuitry and the equipment for maintaining a
predetermined blood glucose level in a living body.
The present invention utilizes an implantable fuel cell, indicated
generally by the reference numeral 10 in FIG. 1 to obtain an
electrical indication of the blood glucose level in a living body.
Before discussing the fuel cell 10 in detail, it will be helpful to
briefly review the characteristics of a fuel cell with special
emphasis on the requirements for an in vivo fuel cell.
A fuel cell is an electrochemical energy conversion device composed
of a nonconsumable anode and cathode, an electrolyte, and suitable
arrangements and controls for maintaining selective environments
for a fuel anode and an oxidant cathode. Fundamentally, any
oxidation-reduction reaction is a fuel cell candidate; the
practicality, however, depends primarily on the reaction rate. The
most efficient and highly refined fuel cell system known to date is
the human body which uses enzymes to catalyze the oxidation of food
(fuel) in an electrolyte (body or cell fluid), producing
energy--some of which is electrical. By providing different kinds
of active catalysts such as platinum, palladium and nickel, certain
carbohydrates (glucose, for instance), plentiful in the human body,
which contain aldehyde (or similar groups) can be activated at low
temperatures in a fuel cell to generate electricity. A metallic
catalyst impregnated on the electrode surface will promote the
reaction of glucose with water, absorbing electrons and releasing
hydrogen ions. This fuel rich electrode constitutes, therefore, the
anode of the fuel cell. If, in addition, an identical
catalyst-coated electrode supplied with oxygen is introduced into
the same electrolyte solution, OH ions will be released and a
potential difference can be detected across these electrodes. The
latter oxygen rich electrode is naturally the cathode of the fuel
cell, and the generated voltage is essentially a constant,
characteristic of the fuel used, while the current flowing in leads
connecting the electrodes is closely related to the fuel
concentration near the anode. Based on this principle, an implanted
fuel cell can be considered for the measurement of glucose level of
body fluid or blood.
However, simply implanting two catalyst-coated electrodes into the
body will not yield any electrical output due to the lack of
asymmetry. Provision must be made for alteration of conditions near
the electrodes and this can be accomplished by placing the
electrodes at different locations within the body to achieve a
selective electrode environment. Although electrical energy can
thus be obtained, such a system cannot be used to measure glucose
concentration. The transport of ions within the porous electrode
and electrolyte, will be rate limiting, and the internal cell
resistance will be high. The present invention eliminates this
problem by utilizing a fuel cell in which the electrical output of
the fuel cell is glucose diffusion-limited.
The basic construction of the glucose fuel cell sensor 10 and its
associated circuitry are shown in FIG. 1. The fuel cell 10 is
principally a controlled-diffusion device which employs artificial
membranes and coating materials of varying thickness and
characteristics to vary the diffusion rate of glucose relative to
that of oxygen on the basis of molecular size, mobility, or
solubility in the membrane and coating materials.
The fuel cell 10 and its associated microcircuitry 12 are encased
in an external membrane 14 constructed of newly-developed inert
high dialysis rate membrane materials (such as those developed by
Union Carbide, G.E., and DuPont) which permit free transmission of
oxygen, glucose and similar size compounds but impede the diffusion
of large, more complex macromolecules, such as proteins,
polysaccharides, cholesterols, etc. The membrane 14 defines a
chamber 16 within which are positioned two spaced transition metal
catalyst coated electrodes 18 and 20 that comprise an anode
electrode and a cathode electrode, respectively. The anodic and
cathodic reactions are:
C.sub.6 h.sub.12 o.sub.6 (glucose)+H.sub.2 O .sup.Pt C.sub.6
H.sub.12 O.sub.7 (Gluconic acid)+2H.sup.++2e
(Anodic) (1)
and
1/2 O.sub.2 +H.sub.2 O+2e .sup.Pt
20H.sup.- (Cathodic) (2) and the overall reaction is
1/2 O.sub.2 +C.sub.6 H.sub.12 O.sub.6 .fwdarw.C.sub.6 H.sub.12
O.sub.7
(Overall) (3)
The cathode which has a larger surface area is covered with a thin
layer of an artificial membrane 22 which allows free passage of
water, oxygen, etc., but strongly resists the diffusion of glucose
so that it serves as an oxygen electrode. The smaller anode 18 is
covered with a relatively thick layer of porous plastic material 24
which impedes the diffusion of glucose and protects the
catalyst(platinum) from poisoning and from the physical, chemical,
and biological harassment of the body. The body fluid within the
membrane encased chamber 16 constitutes an electrolyte 26 for the
fuel cell. Alternatively, an anion, a cation exchange membrane, or
a combination of the two ion exchange membranes can be interposed
directly between the fuel and the oxygen electrodes to serve as a
solid electrolyte as well as a partition for the fuel and the
oxygen half cells. When both ion exchange membranes are
simultaneously displayed in parallel, cell performance is generally
improved, noise and signal drift reduced, and problems associated
with water accumulation or starvation in the oxygen half cell can
be avoided.
Preferably, a platinized anode is used which catalyzes the
dehydrogenation of the aldehyde group of the glucose molecules that
have diffused through the anode coating 24 and impinged upon the
platinum surface. This electrode is therefore the glucose or the
fuel electrode. Because the cathode or oxygen electrode 20 is
larger, and because oxygen is lighter and smaller and therefore has
a larger diffusion coefficient, and further because the diffusion
of glucose to the anode is impeded, the rate of oxygen molecules
arriving at the cathode can be arranged in such a way that it is
always larger than the rate of glucose impingement on the anode
surface. As a result, the current that can be drawn from the fuel
cell 10 is proportional to the diffusion or arrival rate of glucose
molecules and hence to the concentration of glucose in the body
fluid, and, in turn, to the concentration of glucose in blood.
To ensure a glucose diffusion-limited fuel cell, an ample supply of
oxygen must be maintained and the following condition must be
satisfied at all times and at all possible glucose levels:
(D.sub.oc N.sub.o A.sub.c /.delta..sub.c) > .OMEGA./2 (D.sub.ga
N.sub.g A.sub.a /.delta..sub.a) (4)
where .OMEGA. is the effective steric factor of the anodic
reaction, A.alpha. is the area of electrode .alpha. N, the number
density of species .beta. in the body fluid,
D.sub..delta..sub..alpha., the diffusion coefficient for transport
of species .delta. through the surface layer of electrode .alpha.
and .delta..alpha., the thickness of surface layer of electrode
.alpha.. Subscripts a, c, g, and o pertain to the anode 18, the
cathode 20 and the glucose and oxygen, respectively.
The standard open circuit voltage of the glucose fuel cell is
approximately 0.85 volt and it is a constant, essentially
characteristic of the overall reaction expressed by Eq. (3). This
voltage can be evaluated based on known electrochemical constants
and is approximately equal to the sum of the theoretical E.M.F. of
the participating anodic and cathodic reactions. The electrodes
must be arranged in close proximity to one another so that the
diffusion of electrode ionic products H+ and OH.sup.- is not the
rate limiting process in the generation of electrical power.
The terminal voltage of a fuel cell depends on its current, and a
typical voltage versus current commonly referred to as the
polarization curve is given in FIG. 2. Since the glucose
concentration is only proportional to the output current that can
be sustained by the fuel cell, load resistance 28 must be very
small so that the current will correspond to the value at the tail
of the polarization curve (i.e., in the concentration polarization
regime). Typically the resistance of load resistor 28 is in the
range of 0 to 10 ohms.
Although a platinum black anode is preferred for use in the glucose
fuel cell 10, other Group VIII transition metals can also function
satisfactorily as fuel (glucose) electrode catalysts. These metals
(palladium, nickel, platinum) are active catalysts for
heterogeneous hydrogenation-dehydrogenation reactions. Their
catalytic properties can be explained by their electron receiving
capacity and by the fact that they are capable of forming covalent
bonds with fuels through the metal d-band during the electrode
reaction. This also explains why nontransition group metals, whose
d-orbitals are completely filled, are not catalytic. The limited
catalytic activity of the metals other than Group VII, notably the
Group I.sub.B metals (gold, silver, copper, etc.) is attributed to
a d-s promotion that gives then d-orbital vacancies.
While the selection of (anodic) fuel electrode catalysts is
relatively limited, the choice for (cathodic) oxygen electrode
catalysts is considerably broader. In contrast to their performance
as fuel catalysts, the Group I.sub.B metals to their oxides are at
least as active oxygen catalysts as the Group VIII metals, except
perhaps that the path of oxygen reduction is different. The
reduction has been postulated as yielding (1) hydroxyl ions (Eq. 2)
or (2) perhydroxyl ion and a hydroxyl ion,
O.sub.2 +h.sub.2 o+2e.fwdarw.O.sub.2 H.sup.-+OH.sup.- (5)
It has been established through chronopotentiometric studies that
the reduction on platinum proceeds according to Eq. (2) in both
acid and alkaline electrolytes. For this reason, the use of
platinum as oxygen electrode catalysts is favored for more
efficient utilization of oxygen.
Finally, it is important to note that it may be more advantageous
to employ metals (or metal oxides) such as gold, silver, etc. as
the cathode (oxygen electrode) catlayst to achieve asymmetry which
is indispensible for generation of electric output. These materials
are good oxygen electrode catalysts but poor glucose catalysts
(relative to platinum, palladium and nickel). The necessary
asymmetry or electrode selectivity can be achieved through one or
all of the following schemes: (1) differential electrode area; (2)
control of diffusion rates by different surface coatings; (3)
dissimilar electrode materials.
Although the open circuit equilibrium cell potential is insensitive
to the glucose concentration, the rate of charging generally varies
with the glucose level. Hence, by periodically discharging the
cell, the glucose concentration, can alternavitely be determined by
measuring the rate of potential rise. Other modes of operation of
the fuel cell sensor include measurements of temperature rise due
to glucose oxidation, the change in pH value as a result of the
formation of gluconic acid, and the reduction of oxygen tension as
a result of O.sub.2 consumption, which may be caused by catalytic
action of electrodes.
The implantable glucose monitor requires non-autogenous materials
for long-time subdermal contact with the human body. These
materials must, therefore, be non-antagonistic to the environment
into which they are placed. Recent advances in biomaterials
research, motivated by the development of artificial kidney, lung,
heart and other organs, have resulted in a number of new materials
whose biological compatibility has clearly been demonstrated. These
include "Silastic", silicone rubbers, "Teflon," polyethylene,
cellulose, semipermeable hollow fibers, collagens and etc. Since
these materials can generally be synthesized into different forms
with different porosity and selectivity, they are ideally suited
for the present invention.
The output current from the fuel cell 10 is amplified by a current
amplifier 30 which has a low input impedance and therefore measures
the short circuit current which is proportional to the glucose
diffusion rate, which is in turn, proportional to the blood glucose
concentration. The amplified output current is converted to a
frequency by a current-to-frequency converter 32. The blood glucose
level information now in frequency form, is transmitted by
transmitter 34 to an external receiver 36.
The detailed circuitry employed in the implantable glucose
sensor-monitor alarm system is shown in FIG. 3. Looking at FIG. 3,
the output current from fuel cell glucose detector 10 is applied to
the current amplifier 30. The amplified current output from
amplifier 30 is used to charge a small integrating capacitor 38 to
much higher voltages than normally are obtainable from the glucose
cell sensor 10. The integrating capacitor 38 together with a
low-power electronic device, such as a unijunction transistor 40 is
used to provide short pulses in the range of 1 kilohertz. Since the
unijunction transistor draws power only when it is switching, the
power requirements for this circuit are quite low.
The frequency of oscillation of the UJT is directly proportional to
the current from the glucose fuel cell sensor 10. By the use of a
current-to-frequency converter technique, the blood glucose level
information can be processed to a form which is much more readily
transmitted to the external receiver 36. The output pulses of the
UJT oscillator 40 are used to trigger a silicon control rectifier
switch 42 to drive a resonant LC network 44 which is tuned to about
1 megahertz. The shock excited resonant circuit 44 generates bursts
of 1 megahertz rf energy at a repetition rate of about 1 kilohertz.
The output from the shock excited circuit 44 is directly coupled to
an rf radiating plate 46.
The use of the 1 megahertz carrier extends the transmitting range
outside of the body and it also simplifies the receiver tuning to
reduce spurious signals. The repetition rate of the rf carrier
contains the information with respect to the glucose concentration.
The use of brief bursts of rf energy with a low duty cycle of about
10 percent or less reduces the average power requirements of the rf
transmitter with a concommitant reduction in the required battery
size or an extension of the battery life.
In its preferred form, the telemetering system of the present
invention utilizes pulse-code modulation. However, it should be
understood that the PCM mode is merely illustrative and that the
other modulation modes can be employed to telemeter the blood
glucose level information to an external receiver.
The repetition rate at normal glucose levels is chosen to be
approximately 1 kilohertz with provision for a dynamic operating
range of a factor .+-.10. In this way, the UJT oscillator 40 is
able to operate from 100 hertz up to 10 kilohertz with a nominal
value of 1 kilohertz. These parameters permit telemetering of
information about the glucose condition from a value of 1/10 the
normal to 10 times the normal level.
The telemetering accuracy of this system is quite high. Under a
worst case situation corresponding to 100 hertz, the reception of
the signal for 1 second will be sufficient to give a reading with
an accuracy of .+-.1 count corresponding to an accuracy of .+-.1%.
The accuracy of the system at higher repetition rates is obviously
much greater and is far greater than one really needs for the
overall system. This provides a satisfactory margin of reserve
while at the same time keeping the electronic system reliable and
compact.
Since the glucose fuel cell 10 and associated telemetering
components shown in FIG. 3 are implanted in the body, it is
desirable to provide external adjustment of the electronics within
the telemetering system without requiring surgical techniques. This
can be accomplished by the inclusion of a tiny adjustable
potentiometer 48 to which is attached a small magnetic bar 50 which
can be readily rotated by means of a permanent magnet (not shown)
located outside of the body. By properly positioning the external
permanent magnet and rotating it the required number of turns, it
will be possible to adjust the multi-turn potentiometer 48 to
change the calibration set points of the telemetering system. This
can be done most conveniently in the gain-control portion of the
current amplifier 30 as shown in FIG. 3.
A variety of options are available with respect to supplying power
to the electronic portion of the glucose monitoring system.
Referring to FIG. 3, power can be obtained from an internal battery
52. If desired, provision can be made for external recharging of
the implanted battery by means of magnetic coupling through the
skin. If this mode of operation is employed, a magnetically powered
battery recharger, indicated generally by the reference numeral 54,
is included within the implanted glucose monitor. Power for the
battery recharger 54 is provided by magnetic coupling from an
external electromagnet 56.
In the basic mode of operation where the current from the fuel cell
sensor 10 is amplified by current amplifier 30, converted to a
frequency by a UJT oscillator 40 and used to shock-excite a
resonant LC circuit 44, the average power of the electronics is
quite low in comparison to the peak rf power transmitted by the
internal shock-excited oscillator. This is because the energy
stored within a capacitor is periodically dumped into the
shock-excited LC circuit 44 and flows for only a short time (on the
order of 10 or 20 microseconds). The duty cycle of this oscillator
is low so that the average power is quite small compared to the
peak power transmitted.
Assuming that the rf radiating plate 46 and the receiver 36 (FIG.
4) normally will be separated by approximately 10 feet and that the
normal range of a transmitter at milliwatt power levels is
approximately 10 feet, it can be seen that the shock-excited
oscillator 44 will have the desired 10 foot radiation range while
requiring peak powers on the order of 10 milliwatts or average
powers on the order of 1 milliwatt--even allowing for relatively
low efficiency in a shock-excited oscillator. Since the power
requirements of a uni-junction oscillator are negligble except
during short periods of time when it acts as a current switch
transferring the charge in the capacitor to the resonant LC
network, the main power requirement for the system will be from the
current amplifier 30. Using off-the-shelf integrated circuits, the
current amplifier can be designed to have an average power
requirement of about 5 milliwatts. It will, therefore, be
appreciated that the information gathering, processing and
transmitting electronics of the glucose monitoring system will
require an average power of about 10 milliwatts, taking into
account the appropriate duty cycles and a continuous transmission
of the one kilohertz modulated carrier wave.
Further reduction in the average power requirements can be obtained
by providing intermittent telemetering of the glucose level
information. This is accomplished by including a relatively simple
uni-junction, low-powered clock oscillator, indicated generally by
the reference numeral 58. The clock oscillator circuit comprises a
15 minute UJT clock 60, a 5 second clock 62, a normally OFF
flip-flop 64 and a FET switch 66. The clock oscillator turns on the
measuring, amplifying and telemetering circuits once every 15
minutes. If the telemetering system is turned on for a period of
approximately 5 seconds during each 15 minute interval, this
corresponds to a duty cycle of one part in 180 and the average
power requirement is reduced by a factor 180. Although there will
be some minor increases in the power requirements due to the
additional electronics, even if this doubled the average power
requirements for the electronics, the overall savings will be at
least a factor of 90, which is almost 100-fold reduction of power
requirements or a 100-fold extension of the life of the battery 52,
assuming that the shelf life of the battery is not the basic
constraint.
Under certain circumstances, it may be desirable to provide a
manual override control for the clock oscillator 58. This can be
accomplished by providing a magnetically actuated reed relay 68
which bypasses the 15 minute UJT clock 60 and actuates the normally
OFF control flip-flop. When FF 64 is in the ON condition, the
output thereof biases FET switch 66 into conduction thereby
applying power from battery 52 to the power bus 70.
Looking now at FIG. 4, there is shown in diagrammatic and partial
block form the external portion of the glucose monitoring-alarm
system of the present invention. The radio frequency energy
radiated by the implanted radiation plate 46 (FIG. 3) is received
by the external receiver 36. The modulation frequency is extracted
from modulated rf carrier by extractor 72 and converted to a
voltage by converter 74. The output voltage from converter 74
represents the blood glucose level in the monitored living body.
This voltage is then inputted to a voltage comparator 76 which
compares the blood glucose level input voltage with a voltage or a
range of voltages which represent a normal or desired blood glucose
concentration.
In the preferred embodiment, two adjustable voltage reference
levels 78 and 80 also are inputted to the voltage comparator. These
two reference voltages define the acceptable range for the voltage
output from the comparator 76. If the output voltage from the
frequency-to-voltage converter (which represents the blood glucose
concentration in the monitored living body) falls within the range
of voltages defined by the two reference levels, no output is
generated by the voltage comparator. Normally, the two voltage
reference levels are selected to correspond to the points at which
glucose or insulin must be supplied to the monitored living body in
order to maintain a normal blood glucose level. For purposes of
illustration, the relative voltage levels can be considered only in
terms of positive voltages with the glucose voltage level being the
most positive. Therefore, it can be seen that if the output voltage
from the frequency-to-voltage converter 74 exceeds the glucose
voltage reference level, the voltage comparator will produce a
glucose output signal on lead 82. Conversely, if the voltage output
from the converter falls below the insulin voltage reference level,
the voltage comparator will produce an insulin output signal on
lead 84. The output leads from voltage comparator are inputted to
an OR circuit 86 which in turn is connected to a suitable alarm
means 88. Various types of alarm means can be employed including
visual, audible, and/or a physical stimulus to the monitored living
body. The alarm means will actuate whenever the output voltage from
the frequency to-voltage converter falls outside of the normal
range of voltages established by the glucose and insulin reference
input voltages.
The glucose and insulin output signals from the voltage comparitor
can be used to actuate corresponding electrically actuated fluid
valves 90 and 92, respectively. These valves control, respectively,
the flow of glucose and insulin from corresponding reservoirs 94
and 96 to the monitored living body, thus providing a fully closed
loop system. Obviously, the insulin and the glucose reservoir and
dispensing system can also be placed inside the living body with
the amplified signal of the fuel cell glucose sensor feeding
directly to the voltage comparator to actuate the appropriate fluid
valves.
It is well known that in an electrochemical sensing device, the
activity of the platinized surface will degrade in time, resulting
in a decrease in sensitivity and reproducibility of the signal
output. The electrode catalyst of the present glucose sensor can be
rejuvenated to maintain its activity so as to eliminate or to
reduce the frequency of recalibration after implanation. The
rejuvenation is achieved by the use of a short duration cycle of
negative and positive potential pulses to maintain a highly active,
oxide-free fuel anode.
In the operation of the fuel cell glucose sensor, the platinum
surface of the anode may slowly degrade by the external oxidation
of the surface. These oxides inhibit the glucose oxidation reaction
and decrease the available surface sites on the anode active toward
the oxidation of the aldehyde glucose. Also, if oxides are present
on the surface of the anode, a fraction of the glucose presented to
the anode for oxidation will be consumed in the chemical reduction
of the oxide film, so that the total amount of glucose present will
not be sensed by the anode since no electrons are donated to the
anode in the chemical oxide reduction process.
This problem is eliminated by frequently actuating the platinized
electrode by an electrochemical pulsing technique such as described
in U.S. Pat. No. 3,509,034, issued April 28, 1970 for
PULSE-ACTIVATED POLAROGRAPHIC HYDROGEN DETECTOR. The anode is
cycled from anodic to cathodic, going from oxygen evolution to
hydrogen evolution, by means of a third biased electrode (not
shown). The potential pulses are short-duration square waves
generated at 20 second intervals. The anodic-cathodic polarization
cycle is carried out about three times and is always terminated on
the cathodic part of the cycle thereby reducing the platinum oxide
surface to a highly active, disordered surface of platinum.
It will be appreciated from the preceding description that the
glucose monitor and sensor of the present invention provides an
accurate means for determining glucose levels in vivo, either by
direct implantation or by subcutaneous insertions or in vitro.
Having described in detail a preferred embodiment of our invention,
it will be apparent to those skilled in the art that numerous
modifications can be made therein without departing from the scope
of the invention as defined in the following claims.
* * * * *