U.S. patent application number 14/570800 was filed with the patent office on 2015-09-24 for implantable biosensor and methods of use thereof.
The applicant listed for this patent is Optoelectronics Systems Consulting, Inc.. Invention is credited to Diane Burgess, Faquir Jain, Fotios Papadimitrakopoulos.
Application Number | 20150265182 14/570800 |
Document ID | / |
Family ID | 38969901 |
Filed Date | 2015-09-24 |
United States Patent
Application |
20150265182 |
Kind Code |
A1 |
Jain; Faquir ; et
al. |
September 24, 2015 |
Implantable Biosensor and Methods of Use Thereof
Abstract
Disclosed herein is an analyte sensing device capable of
continuously monitoring metabolic levels of a plurality of
analytes. The device comprises an external unit, which, for
example, could be worn around the wrist like a wristwatch or could
be incorporated into a cell phone or PDA device, and an implantable
sensor platform that is suitable, for example, for implantation
under the skin. The external device and the internal device are in
wireless communication. In one embodiment, the external device and
the internal device are operationally linked by a feedback system.
In one embodiment, the internal device is encapsulated in a
biocompatible coating capable of controlling the local tissue
environment in order to prevent/minimize inflammation and fibrosis,
promote neo-angiogenesis and wound healing and this facilitate
device functionality.
Inventors: |
Jain; Faquir; (Storrs,
CT) ; Papadimitrakopoulos; Fotios; (West Hartford,
CT) ; Burgess; Diane; (Storrs, CT) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Optoelectronics Systems Consulting, Inc. |
Storrs |
CT |
US |
|
|
Family ID: |
38969901 |
Appl. No.: |
14/570800 |
Filed: |
December 15, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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11862866 |
Sep 27, 2007 |
8914090 |
|
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14570800 |
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Current U.S.
Class: |
600/302 ;
606/129 |
Current CPC
Class: |
A61B 5/14546 20130101;
A61B 5/1495 20130101; A61B 2017/00367 20130101; A61B 5/0017
20130101; A61B 5/742 20130101; A61B 2034/2057 20160201; A61B
17/3468 20130101; A61B 17/3403 20130101; A61B 5/7264 20130101; A61B
5/14539 20130101; A61B 34/20 20160201; A61B 5/14865 20130101; A61B
2560/0219 20130101; A61B 5/076 20130101; A61B 5/0031 20130101; A61B
2560/0214 20130101; A61B 2560/0475 20130101; A61B 5/1459 20130101;
A61B 1/04 20130101; A61B 5/14532 20130101 |
International
Class: |
A61B 5/07 20060101
A61B005/07; A61B 5/1495 20060101 A61B005/1495; A61B 19/00 20060101
A61B019/00; A61B 1/04 20060101 A61B001/04; A61B 17/34 20060101
A61B017/34; A61B 5/1486 20060101 A61B005/1486; A61B 5/00 20060101
A61B005/00 |
Claims
1. An analyte sensing device comprising: an external control unit
and an implantable sensor platform in wireless optical two-way
operable communication, wherein the implantable sensor platform can
pass though a bore of a needle, wherein the implantable sensor
platform comprises, in operable communication, a photovoltaic
device to receive optical power from the external control unit to
serve as a power source for powering said implantable sensor
platform, an optical receiver for detecting signals produced by the
external control unit, a plurality of sensor elements deposited on
a surface of the implantable sensor platform and operable for
sensing one or more analytes, wherein the plurality of sensor
elements have one or more working electrodes, a reference electrode
and a counter electrode in contact with the surface they are
deposited on such that the electrodes do not delaminate when
exposed to body fluids, an interfacing circuit, for providing
operating parameters to the electrodes of the plurality of sensor
elements and controlled feedback for the operation of the plurality
of sensor elements, wherein the plurality of sensor elements
generates a sensor output signal having a sensor output signal
magnitude proportional to the amount of analyte present, wherein
the interfacing circuit comprises at least one potentiostat, a
signal processing circuit interfaced with the sensor output signal,
wherein the signal processing circuit converts the sensor output
signal of the plurality of sensor elements to digital pulses having
a pulse frequency, wherein the pulse frequency is determined by the
sensor output signal magnitude and wherein changes in the pulse
frequency are proportional to changes in the analyte levels,
wherein the digital pulses are transmitted to the external unit, a
switching mode selector configured to cause the implantable sensor
platform to perform at least one of an initialization function, a
power level check function, a potentiostat circuit reconfiguration
function for analyte level measurement, an implantable sensor
selection function, and an implantable sensor calibration function,
one or more optical components for facilitating wavelength
selection, transmission and/or reflection, and a biocompatible
coating surrounding at least a portion of the implantable sensor
platform, wherein said biocompatible coating comprises a drug and
is designed to control the timed release of the drug, wherein the
external control unit comprises, in operable communication, an
optical source suitable for powering the photovoltaic device of the
implantable sensor platform, a receiver suitable for receiving one
or more digital pulses from the implantable sensor platform and
processing the one or more digital pulses to determine the analyte
levels, an optical transmitter suitable to transmit one or more
optical pulses to the optical receiver of the implantable sensor
platform, wherein the optical pulse relays instructions to the
switching mode selector to cause the implantable sensor platform to
perform at least one of the initialization function, the power
level check function, the potentiostat circuit reconfiguration
function, the sensor selection for analyte level measurement
function, the implantable sensor selection function and the
implantable sensor calibration function, an integrated circuit for
processing and displaying the analyte levels, wherein the
integrated circuit is in operable communication with the receiver,
a microcontroller comprising a program code, programmable memory,
and means to display output and communicate with other devices,
means of interfacing with the receiver, optical source and optical
transmitter, to establish an operable communication with the
implantable sensor platform, a power supply to power the external
unit, one or more optical components providing wavelength
selection, transmission or reflection functions, and a miniaturized
camera to align the implantable sensor platform with the optical
components of the external control unit.
2. The analyte sensing device of claim 1, wherein the implantable
sensor platform comprises a first sub-chip, a second sub-chip and a
third sub-chip in operable communication, wherein the first
sub-chip includes, the photovoltaic device, a transmitter
configured to transmit information in the form of digital pulses to
the external unit, and the optical receiver configured to receive
instructions from the external unit; the second sub-chip includes,
the interfacing circuits including initialization, sensor select,
and sensor calibration, the potentiostat and signal processing
circuits, and the transmitter configured to wirelessly transmit
digital pulses; and the third sub-chip includes the plurality of
sensor elements.
3. The analyte sensing device of claim 2, wherein the first
sub-chip, second sub-chip and third sub-chip are integrated via at
least one of through-Silicon-vias, partial-Silicon-vias and
interconnects.
4. The analyte sensing device of claim 1, wherein the interfacing
circuit includes a voltage control logic unit and the signal
processing circuit includes a potentiostat and an analog to digital
converter.
5. The analyte sensing device claim 1, wherein the plurality of
sensor elements are configured to monitor at least one of the
plurality of analytes.
6. The analyte sensing device of claim 5, further comprising
sensor-select and potentiostat circuits that sequentially address
the plurality of sensor elements.
7. The analyte sensing device of claim 1, further comprising a
digital-to-RF converter circuit and an antenna, wherein the
digital-to-RF converter circuit is configured to receive and
convert the digital pulses from the output of the signal processing
unit into a wireless radio frequency(RF) signal responsive to the
analytes, and wherein the antenna is configured to transmit the RF
signal to the external unit.
8. The analyte sensing device of claim 7, wherein the external unit
includes an external RF receiver configured to receive the RF
signal.
9. The analyte sensing device of claim 1, further comprising a
transducer configured to receive and convert the digital pulses at
the output of signal processing unit into a wireless ultrasound
signal responsive to the analytes.
10. The analyte sensing device of claim 9, wherein the external
unit includes an external ultrasound receiver configured to receive
the wireless ultrasound signal.
11. The analyte sensing device of claim 1 wherein the miniaturized
camera is used to implant the implantable sensor platform under the
skin using a needle based insertion device.
12. An analyte sensing device comprising: an external control unit
and an implantable sensor platform in wireless optical two-way
operable communication, wherein the implantable sensor platform can
pass though a 14 gauge or smaller bore needle, wherein the
implantable sensor platform comprises sub-chip #1, sub-chip #2 and
sub-chip #3 and a biocompatible coating surrounding at least a
portion of the sensor platform, wherein sub-chip #1 comprises: a
photovoltaic device that powers the sub-chip #1, sub-chip #2 and
sub-chip #3, a first optical receiver to receive instructions from
a mode select unit located in the external control unit via an
external unit optical transmitter, wherein the external unit
optical transmitter is located in the external control unit, and a
second optical receiver for providing information regarding light
intensity received from light-emitting diodes located in the
external control unit; wherein the first and second optical
receivers operate at wavelengths such that the first and second
optical receivers do not interfere with each other, a sensor
platform transmitter configured to transmit digital pulses,
relaying information selected from sensor output, calibration,
potentiostat check, or solar power level check received from a
driver located on subchip #2, one or more coatings providing
wavelength selection, transmission or reflection functions, wherein
sub-chip #2 comprises: a plurality of interfacing circuits selected
from initialization circuits, sensor select circuits, and sensor
calibration circuits, a potentiostat, a signal processing circuit,
a logic circuit, a demultiplexer, a multiplexer, and the driver to
enable transmission of feedback signals selected from a level of
radiation intensity received by the photovoltaic device, a
reference voltage of the potentiostat, or a sensor reading, and
wherein the driver on subchip #2 transmits a plurality of digital
pulses to the sensor platform transmitter located on subchip #1 and
transmits digital pulses using transmitter selected from wireless
RF and ultrasound, wherein sub-chip #3 comprises a plurality of
sensor elements operable for sensing of one or more analytes, and
wherein the plurality of sensor elements has one or more working
electrodes, a reference electrode and a counter electrode, wherein
the one or more working electrodes, reference electrode and counter
electrodes are in contact with the surface they are deposited on in
a way that they do not delaminate when exposed to body fluids,
wherein the plurality of sensor elements are in contact with the
potentiostat and other circuits located on subchip #2, wherein
sub-chip #1, sub-chip #2 and sub-chip #3 are electrically
interconnected using through-silicon-vias or partial-silicon-vias
and interconnects and integrated in a manner to operate in the
presence of body fluids; wherein the external control unit
comprises, in operable communication, an optical source comprising
light-emitting diodes and laser diodes, wherein the optical source
powers the photovoltaic device on sub-chip #1, a third receiver for
receiving digital pulses from the sensor platform transmitter on
subchip #1 of the implantable sensor platform, wherein the signal
processing circuit converts a sensor element output signal to
digital pulses, wherein the sensor platform transmitter converts
the digital pulses to RF and ultrasound pulses, wherein the
frequency of the digital pulses is determined by the sensor element
output signal which is controlled by an analyte level in the body
fluids, wherein the external unit optical transmitter transmits one
or more optical pulses to the first optical receiver, wherein the
one or more optical pulse relays instructions to sub-chip #2 for
the switching, multiplexing, demultiplexing and logic circuits of
sub-chip #2 to provide at least one function wherein the at least
one function includes an initialization function, a power check
function, a potentiostat circuit reconfiguration for analyte level
measurement function, an implantable sensor selection function, and
an implantable sensor calibration function, an integrated circuit
for processing and displaying an electrical pulse, wherein the
integrated circuit is in operable communication with an optical
receiver, a microcontroller comprising a program code; programmable
memory; a means to display the output and communicate with other
devices; a means of interfacing with the optical source, a receiver
located in the external control unit receiving digital pulses from
RF or ultrasound transducer serving as transmitter and an optical
transmitter located in the external control unit operating at
800-1000 nanometers, a power supply to power the external control
unit, one or more optical components providing wavelength
selection, transmission, or reflection functions, and a
miniaturized camera to align the implantable sensor platform with
the optical components of the external control unit, wherein the
miniaturized camera is used to assist in inserting the implantable
sensor platform under the skin using a 14 gauge or smaller
needle.
13. An implantable sensor platform delivery device for delivering
an implantable sensor platform into the body of living being,
comprising: an actuator, a plunger, and a needle cannula having a
hollow cannula bore that extends the length of the cannula, wherein
the hollow cannula bore is sized and shaped to moveably contain the
implantable sensor platform wherein the actuator is associated with
the plunger and wherein the plunger is configured to be moveably
located within the hollow cannula bore, such that when the
implantable sensor platform and the plunger is located within the
hollow cannula bore, movement of the actuator causes movement of
the plunger.
Description
CROSS REFERENCE TO RELATED APPLICATION
[0001] This application is a continuation application of co-pending
U.S. Non-Provisional patent application Ser. No. 11/862,866 filed
Sep. 27, 2007 and claims priority benefit of the filing date of
U.S. Provisional Application Ser. No. 60/827,104 filed Sep. 27,
2006, the contents of both of which are hereby incorporated by
reference in their entireties.
BACKGROUND
[0002] Careful metabolic monitoring and proper treatment can
improve control of metabolic diseases such as diabetes and obesity.
Knowing a patient's metabolism along with other physiological
parameters allows for correct dosing and delivery of medications
and nutrients. Improvements in metabolic measurement technology are
essential for better diagnostics and advances in treatment of
metabolic diseases and conditions. Treatment of metabolic diseases
and conditions ideally requires frequent and timely monitoring
which drives a need for monitors that are non-invasive, real-time,
portable, low cost, and accurate. Metabolic data are also useful in
assessing the physiological homeostatic conditions of patients and
healthy subjects in general.
[0003] Blood glucose concentration data is extremely useful for the
control of metabolic diseases such as diabetes and for monitoring
the overall metabolic condition of a human subject. An accurate,
real-time, noninvasive method for measurement of blood glucose
levels is of great interest in the diabetic community. Current
technologies involving the measurement of blood glucose by drawing
blood are invasive and often lead to poor patient compliance.
Measurement by probe involves frequent lancing and may result in
problems. An ideal non-invasive blood glucose sensor provides a
continuous signal and/or a signal on demand that can be used to
control devices, such as insulin pumps in closed loop feedback
applications.
[0004] In recent years, two different types of metabolic internal
units have been developed: non-invasive and minimally invasive.
Non-invasive optical internal units depend on light penetration
into the skin and spectroscopic measurement of metabolic levels;
however, lack of analyte specificity remains a problem for optical
internal units. Commercially available minimally invasive internal
units can function only for the short term (a few days) and require
frequent calibration via finger pricking. These commercially
available internal units are either incapable of continuous
monitoring of metabolic levels or are only suitable for use by
qualified medical personnel.
[0005] Therefore, there exists a need for a minimally invasive or
non-invasive metabolic internal unit suitable for use by the host
that allows continuous and/or on demand monitoring of metabolic
levels of specific analytes.
SUMMARY
[0006] An analyte sensing device comprises an external control unit
and an implantable sensor platform in wireless optical
communication, wherein the implantable sensor platform can pass
though a 14 gauge or smaller bore needle. This implantable sensor
platform comprises a variety of functional optoelectronic circuit
blocks for wireless powering, interactive communication,
programmable potentiostats interfacing with various electrochemical
sensors, mode-selection, signal processing, calibration, analog to
digital conversion, amplification, and optical transmission. The
outer surface of this miniaturized sensor platform is coated with
one or more biocompatible coatings, optionally capable of releasing
a variety of drugs and tissue response modifiers. The external
control unit comprises optical sources suitable for powering the
implantable sensor platform, along with transmitters and receivers
for transmitting and receiving optical commands to and from the
implantable sensor platform. These optical signals are then
converted to electrical pulses and processed by a microprocessor
located in the external unit. In addition, the external unit is
equipped with a miniaturized camera to assist in aligning the
various optical components of the external unit with that of the
implantable sensor platform.
DESCRIPTION OF THE FIGURES
[0007] FIG. 1. Schematic representation of an embodiment of an
implanted biosensor unit along with an embodiment of an external
user interface unit comprising mode select, monitoring and
calibration functions.
[0008] FIG. 2. Schematic representation of an embodiment of a
sensor platform shown as stack of three chips encased in a suitable
biocompatible coating. The sensor platform is compact enough for
implantation by needle and plunger.
[0009] FIG. 3. Schematic view of an embodiment of a three sub-chip
sensor platform along with its interface with the external control
unit.
[0010] FIG. 4. Embodiment of optical and optoelectronic components
housed within PDA unit and methodology to optically communicate
with the implanted sensor platform.
[0011] FIG. 5. Schematic of a programmable potentiostat interfacing
with two sensors whose signal is processed by the signal-processing
unit. The optoelectronic transmitter and receiver interface
communicating with the implanted chip and the modified PDA is also
shown.
[0012] FIG. 6. Schematic of a PDA communicating wirelessly with an
external unit that is located in the vicinity of the implantable
platform.
[0013] FIG. 7. Embodiment of a design of sensor-select circuit.
This circuit consists of a optical pulse receiving system, a timer,
set of D-Flip Flops, and a logic block. It interfaces calibration
and Mux (multiplexer) circuits.
[0014] FIG. 8. Figure A showing an ADC signal processor (MOSIS
fabricated chip) interfaced with a hybrid potentiostat. (B).
Measurements showing digital signal changing its pulse
characteristics as a function of glucose level. (C). Plot of
glucose level after converting the pulse frequency change.
[0015] FIG. 9. Photographs showing various components of subchips:
solar cells, laser as transmitters, signal processing chips (two
generation), three-electrode electrochemical sensors.
[0016] FIG. 10. (top) Hybrid approach to integrate sub-chips with
robust interconnects and bonding using vias and bumps. (bottom
left) Cross-sectional schematic illustration of bonding between a
via and bump for sub-chip #1 and sub-chip #2. Top view of
interconnect and power pad is also shown. (bottom right) Bonding
between sub-chip #2 and sub-chip #3 is shown using two approaches
for integration.
[0017] FIG. 11. Schematic of an embodiment of the three sub-chip
design (shown in FIGS. 1, 2 and 3). Here the top surface of
sub-chip #1 (4) is hermetically sealed using a transparent glass
window (111) sealed to the raised silicon walls (110) via anodic
bonding (112) between glass and silicon.
[0018] FIG. 12. Embodiment of a sensor platform having two
sub-chips on its top and bottom surfaces. Sub-chip #1 has three
pads on either side for power supply distribution (e.g., V.sub.dd,
V.sub.1, and C for common, shown in blue and are larger in size
than the via/bumps pads). The power is supplied to the sub-chip #2
using metalized vias labeled as V.sub.VDD etc. In addition, vias
are used to connect photodetector PD.sub.SS/PD.sub.M on sub-chip #1
to sub-chip #2 (having electronics such as sensor select, routing
logic/MUX, etc.). Note that the 1.55 .mu.m transmitter is located
on sub-chip #2 as this wavelength is transparent to Si platform and
chips.
[0019] FIG. 13. Circuit schematic showing various functions of
three sub-chips condensed into two sub-chips. In this version, the
modified PDA unit directly communicates with the implanted
unit.
[0020] FIG. 14. Circuit schematic showing various functions of
three sub-chips condensed into two sub-chips. In this version, the
modified PDA unit communicates with an external unit located in the
vicinity of the implanted sensor and the communication is via
Bluetooth.RTM. wireless technology.
[0021] FIG. 15. Advanced methodology to integrate two sub-chips
into one-wafer platform using wafer bonding technique. The
hermetical seal using glass-Si anodic bonding is also shown for the
solar cell/PDs chip.
[0022] FIG. 16. Schematic of implantable sensor platform
interfacing with a drug dispensing system. The drug delivery may
include a micro-electro-mechanical (MEM) components.
[0023] FIG. 17. Glucose sensor showing Ag reference electrode and
details of coatings on Pt working electrode. Schematic
representation of modified Clark amperometric glucose sensor, along
with various chemical, electrochemical and diffusion processes
associated with its operation. The glucose oxidase (GO.sub.x) layer
is coated with a semipermeable membrane to reduce the amount of
glucose entering the sensor. An HRP-modified hydrogel layer is then
applied to eliminate outer diffusion of H.sub.2O.sub.2. This is
followed by an outer composite hydrogel coating with embedded
microspheres at different stages of degradation and TRM
release.
[0024] FIG. 18. Schematic of the self assembly of semipermeable
membrane composed of humic acids and Fe.sup.3+ on the outer surface
of the electrochemical sensor. As the number of bilayers increases,
there is an increase in the tortuosity for glucose diffusion
towards the enzyme.
[0025] FIG. 19. One-second pulsed mode operation of sensor.
Calibration curve of current response vs. change in glucose
concentration. As the amount of glucose in the system is increased,
there is a corresponding rise in the current. The Figure on left
indicates the time required for the device to stabilize at each
glucose concentration. As the concentration of glucose increases,
the device reaches a stable reading faster.
[0026] FIG. 20. Current response as a function of increasing
semipermeable membrane thickness.
[0027] FIG. 21. Histological evaluation of subcutaneous tissue
samples taken from the vicinity of hydrogel composites containing
PLGA microspheres at 3 and 21 days post implantation. The
representative sections shown are 3 days after implantation (A
& B) and 21 days after implantation (C & D).
[0028] FIG. 22. schematic of a methodology to hermetically seal
sub-chips using one Si wafer as the carrier with provision to
placing sub-chips and interconnecting them. The hermetical seal
using glass-Si anodic bonding is shown on top as well as
bottom.
DETAILED DESCRIPTION
[0029] Disclosed herein is a device capable of monitoring the
metabolic levels of a plurality of analytes, in a continuous or
intermittent (e.g., on demand) operation. The device comprises an
external unit, which, for example, is worn around the wrist like a
wristwatch or carried like a Personal Digital Assistant (PDA) or a
cell phone, and a sensor platform that is suitable for implantation
under the skin or near the surface of another portion of a
patient's anatomy. The sensor may be implantable via a needle and
similarly removable via a needle, thus avoiding the need for
surgical implantation and removal.
[0030] The term "analyte" refers to a substance or chemical
constituent in a biological fluid (e.g., blood, interstitial fluid
or urine) that can be analyzed. In one embodiment, the analyte for
measurement by the devices and methods disclosed herein is
glucose.
[0031] "Biocompatibility" is the ability of a material to perform
with an appropriate host response in a specific application. The
terms "biocompatible membrane", "biocompatible layer," and the like
refer to a semipermeable membrane comprised of protective
biomaterials. In one embodiment, a biocompatible membrane is a few
microns thickness or more and is permeable to small-molecule
analytes oxygen and glucose, but is substantially impermeable to
biofouling agents (such as proteins) that could otherwise gain
proximity to and possibly damage the internal unit. This
"biocompatible membrane," or "biocompatible layer," may also
protect the sensor from damage and inconsistency in readings
resulting from inflammation and fibrous encapsulation. In some
embodiments, the biocompatible membrane comprises pores (e.g.,
typically from approximately 0.1 to approximately 1.0 micron).
[0032] An "electrochemical sensor" is a sensor configured to detect
the presence and/or measure the level of an analyte in a sample via
electrochemical oxidation and reduction reactions on the sensor.
These reactions are transduced to an electrical signal that can be
correlated to an amount, concentration, or level of an analyte in
the sample.
[0033] The sensor platform comprises one or more sensor elements. A
sensor element is a component of the sensor platform that is
capable of recognizing or reacting with an analyte whose presence
is to be detected by the sensor platform. Typically, the sensor
element produces a detectable signal after interacting with the
analyte to be sensed via an electrode in the sensor platform, for
example. Individual sensor elements within the sensor platform can
sense the same or different analytes. In this context, the sensor
platform can be adapted to measure multiple analytes
simultaneously. For example, multiple individual sensor elements
adapted to sense different analytes can be exposed to the external
environment at the same time. Alternatively, multiple individual
sensor elements adapted to sense different analytes can be exposed
to the external environment at different times. Other embodiments
include a sensor platform adapted to function as multi-analyte
sensor on a single chip (or, alternatively, on multiple chips). In
certain contexts, a signal from an individual analyte sensor
element within the plurality of analyte sensor elements that
contact and sense an analyte in a sensor platform are individually
interrogated and/or read. Alternatively, multiple analyte sensor
elements within a plurality of analyte sensor elements that contact
and sense an analyte in the sensor platform are interrogated and/or
read simultaneously and/or in combination.
[0034] In one embodiment, the sensor element utilizes an enzyme
(e.g., glucose oxidase (GOx)) that has been combined with a second
protein (e.g., albumin) in a fixed ratio (e.g., one that is
typically optimized for glucose oxidase stabilizing properties) and
then applied on the surface of an electrode to form a thin enzyme
constituent. In one embodiment, the sensor element comprises a GOx
and HSA (Human Serum Albumin) mixture. In this embodiment, the GOx
reacts with glucose present in the sensing environment (e.g., the
body of a mammal) and generates hydrogen peroxide, wherein the
hydrogen peroxide so generated is anodically detected at a working
electrode in the sensor platform.
[0035] An "electron transfer agent" is a compound that carries
electrons between an analyte and a working electrode, either
directly, or in cooperation with other electron transfer agents.
One example of an electron transfer agent is a redox mediator.
[0036] The measurement of analytes including glucose, lactate,
etc., is achieved using an external unit and an implantable
biosensor platform. The external unit can provide controls for
sensor unit selection and output display. In one embodiment, the
device integrates sensors with biocompatible coatings as well as
drug dispensing devices. In another embodiment, the device is
capable of additional wireless communication with health service
providers as appropriate.
[0037] The device of the present invention comprises a sensor
platform and an external unit that are in operable communication
through a set of transceivers. A transceiver comprises an optical
transmitter and an optical receiver. In one embodiment, the optical
transmitter is a light-emitting or laser diode. In another
embodiment, the optical receiver is a photodetector. The two
components of a transceiver are located in the external control
unit (the optical transmitter) and the implanted sensor platform
(the optical receiver), respectively or vise versa. The transceiver
orientation is defined by the direction of transmitted light. The
interactive coupling between two transceivers (transmitting in
opposing directions) establishes a feedback loop via other
circuits. Two transceivers with opposing location of transmitters
and receivers form a closed-loop, capable of wirelessly
transmitting and received commands, carrying out certain
instructions as well as transmitting certain information back and
forth among the two units. This interactive feedback loop enables
the remote operation of the sensor platform. In addition, with the
use of logic and routing circuits, the feedback loop provides
multiple functionalities including initialization, calibration, and
measurement of one or more analyte levels.
[0038] In one embodiment, the external unit comprises an
optoelectronic receiver suitable for receiving optical pulses from
the sensor platform, and converting these optical pulses to
electrical pulses. In addition, the external unit contains
integrated circuits suitable for processing and displaying the
analyte levers that are coded in terms of pulse characteristics. An
optical source located in the external unit powers a plurality of
photovoltaic cells, which in turn serve as the power source for the
implantable sensor platform.
[0039] In one embodiment, the sensor platform comprises a power
source; one or more electrochemical sensor elements suitable for
sensing of one or more analytes; one or more interfacing circuits
providing operating voltages and a reference voltage to the sensor
elements, wherein the interfacing circuit generates a signal
proportional to the amount of analyte present; (e.g., an ultrasound
transmitter) one or more signal processing circuits in operable
communication with the interfacing circuit, wherein the signal
processing circuit converts the analog sensor signal to digital
pulses, one or more electrical to optical converters in operable
communication with the signal processing circuit, wherein the
electrical to optical converter converts the digital pulse to
optical pulses; and a transmitter for transmitting the optical
pulses to the external unit. In one embodiment, the interfacing
circuit comprises a potentiostat. In one embodiment, the electrical
to optical converter is an infrared (IR) transmitter suitable for
the wireless relaying of analyte levels and power management
information to the external unit.
[0040] In one embodiment, the sensor platform comprises three
sub-chips. In one embodiment, sub-chip #1 comprises a wireless
photovoltaic powering solar cell array to power all components of
the sensor platform, a photodetector (PD.sub.M) to monitor the
power level, an infra red transmitter (TX.sub.D), and a
photodetector (PD.sub.SS), along with their associated circuits.
Sub-chip #1 preferably faces the portion of the external unit that
serves as the power source to power photovoltaic cells (e.g., super
luminescent LEDs or laser diodes). The PD.sub.SS puts sub-chip #1
in operable communication with sub-chip #2. For example, the
photodetector (PD.sub.SS) interfaces with a Mode Selector circuit
block on sub-chip #2, which in turn communicates with
Router/Logic/Mux circuits. Information regarding power levels is
ensures that the desired voltage and current levels are available
to operate all electronic and optoelectronic circuits of internal
implantable platform unit. This can prevent faulty internal unit
readings due to voltage-current levels below threshold. The
infrared transmitter TX.sub.D also serves to transmit information
to the external unit regarding the photovoltaic power level.
[0041] Sub-chip #1 can further comprise an eye-safe infrared (IR
.about.1.55 micrometer) InGaAsP--InP LED/laser source (TX.sub.D),
for example, bonded onto aSiO.sub.2 coated Si substrate in the
vicinity of the solar cell array. An 1.55 micrometer IR detector
(PD.sub.D), located in the external unit, detects the coded
internal unit signal. In an alternative embodiment, the external
unit further comprises a band-pass filter to reject radiation from
the powering LEDs that operate in a spectral regime, which affords
minimum absorption.
[0042] Sub-chip #2 comprises one or more interfacing circuits, one
or more signal processing circuits, and one or more electrical to
optical converters. According to one embodiment, sub-chip #2
comprises a Mode Selector and Router/MUX/Logic blocks, which
interface with programmable potentiostat and calibration circuits,
along with a signal processing analog-to-digital-converter (ADC)
interface and TX.sub.D driver electronics. For example, once a
sensor is selected, the programmable potentiostat provides
appropriate voltage values for working (V.sub.W), reference
(V.sub.REF), and counter (V.sub.C) electrodes of the selected
sensor, located on the Sub-Chip #3. The analog output of the
selected sensor is thus connected (via Router/Logic/MUX block) to
the potentiostat and ADC signal-processing unit. The digital output
from the ADC circuit is fed to the TX.sub.D driver, which in turn
is designed to interface with the infrared transmitter (TX.sub.D)
on sub-chip #1.
[0043] In one embodiment, the analog current developed in a glucose
sensor (e.g., due to the presence of glucose in the environment
adjacent to the implanted sensor platform) is converted into
voltage pulses of varying width by the ADC circuit. These pulses in
turn drive an infrared emitter (TX.sub.D). The emitter output is
received by an external photodetector (PD.sub.D) located in the
external unit, which can be worn on the wrist or located in a
modified PDA unit. Thus, the pulse duration or frequency carries
the glucose level information to the external unit, where it is
processed and displayed accordingly.
[0044] Sub-chip #3 comprises one or more electrochemical sensors,
for example, a glucose sensor, along with other micro-sensors
(e.g., oxygen, pH, insulin, and ion concentration). In one
embodiment, sub-chip #3 comprises an electrochemical sensor with
working platinum and auxiliary platinum electrodes in an
inter-digitated configuration, and a reference silver/silver
chloride electrode meandering between the two platinum electrodes.
Sub-chip #3 optionally comprises ionic sensors, in which
field-effect transistors with an electroactive gate material
coating are used.
[0045] FIG. 1 illustrates schematically an embodiment of the
functional blocks of both the external control unit (1) and the
sensor platform (2) subcutaneously implanted under the skin
(3).
[0046] The external control unit (1) comprises a microprocessor
(11), a software interface or program (12), a mode select
comprising various switches (13), and various electronic and
optoelectronic "Add-on Devices and Control Circuits" (14). In
addition, there is a display (15) and provision for interface with
"Other Devices" (16). The Add-on devices (14) include an optical
source ("719 nm Laser/LED (A)) or sources at wavelengths that are
not absorbed by the skin and subcutaneous tissue for powering solar
cells (41) located on Sub-Chip #1 (4) of the implanted sensor
platform (2). The Add-on devices (14) also includes a transmitter
(TX.sub.SS) (18), operating in the spectral range 800-980 nm, which
sends optical commands as coded pulses to the PD.sub.SS
photodetector (44), located on Sub-chip #1 (4) of the implanted
sensor platform (2). The Add-on devices (14) also includes a
photodetector (PD.sub.D) (19) operating at 1.55 .mu.m, which
receives information as coded optical pulses from the transmitter
(TX.sub.D) (45) located on Sub-chip #1 (4) of the implanted sensor
platform (2). An optical filter is optionally placed in front of
photodetector PD.sub.D (19) in order to allow transmission of
wavelengths of 1.55 .mu.m and reject away shorter wavelength
radiation.
[0047] FIG. 1 also shows an embodiment of the implantable sensor
platform (2) having a compact size of 0.5 mm width.times.5 mm
length.times.0.5 mm height. In this embodiment, the power source of
the internal unit comprises photovoltaic (PV) solar cells (41),
which are powered by high efficiency light-emitting diodes (LEDs)
(17) in the external unit. These PV cells (41), operating at
designed wavelength that allows transmission through the skin,
provide sufficient power output (voltage and current) needed by the
electronic and optoelectronic devices of the implantable sensor
platform unit (2).
[0048] In this embodiment, the implantable sensor platform (2)
comprises three sub-chips with the following functionality:
Sub-chip #1 (4) comprises the power source (41), the power level
monitor photodetector (PD.sub.M) (42), the optical command receiver
photodetector (PD.sub.SS) (44) along with its band-pass filter
(BPF) (43), and the transmitter (TX.sub.D) (45) for transmitting
the optical pulses to the external unit (1); Sub-chip #2 (5)
comprises the Mode Selector circuit block (51), which interprets
the optical commands from the transmitter (TX.sub.SS) (18) to the
photodetector (PD.sub.SS) (44) and communicates it via electrical
digital pulses to the Router/Logic/MUX circuit block (52). The
Router/Logic/MUX circuit block (52) interfaces with the
programmable Potentiostat (54), Calibration and Initialization
Circuits (53), signal processing circuits (Analog-to-Digital
Converter (ADC)) (55), and TX.sub.D Driver circuit (56). The The
Router/Logic/MUX circuit block (52) along with the programmable
Potentiostat (54) interfaces with various sensor elements located
on Sub-chip #3 (6); and Sub-chip #3 (6) comprises an number of
electrochemical sensors, whose share the same reference (61) and
counter (62) electrodes. Three working electrodes (63), (64), and
(65) are explicitly shown on Sub-chip #3 (6). In alternative
embodiments, the internal implantable sensor platform unit (2)
comprises two sub-chips or even one chip, if integration of
circuits is miniaturized further.
[0049] In operation, the implantable sensor platform unit (2)
receives the powering light (31) from the optical source (17)
through the skin (3). This powering light (31) is received by
photovoltaic (PV) cells (41) that provide power to all devices and
circuits in the implantable sensor platform (2) through bus lines
(23), shown in bold. This powering light (31) is also received by
the PD.sub.M photodetector (42), which provides information
regarding the input power level and hence the power produced by the
PV cells (41). The implantable sensor platform (2) also receives
through the PD.sub.SS photodetector (44) optical command and
control information (from the external unit via TX.sub.SS (18)) as
pulses in a certain frequency range (f.sub.1). These optical
commands enable selection of various function of the implantable
sensor platform such as initialization, sensor selection,
calibration and measurement of analyte levels, power level check,
etc. These functions are carried out by the Mode Selector (51) and
Router/Logic/MUX circuit (52) blocks. These two units provide
interface with all other electronic and optoelectronic and
electrochemical devices and circuits within the implantable sensor
platform unit (2). The PD.sub.SS photodetector (44) has a band pass
filter (43) or a coating that blocks the incoming powering light
(32) (from optical source (17)) and reflects it away (33). This
prevents undesirable interference of powering light (32) with the
optical pulses (34).
[0050] Once a function (e.g., initialization, sensor selection,
calibration or measurement) is selected (through the microprocessor
(11) and associated Software (12)), the Mode Select Unit (13) in
the external control unit (1) sends encoded electrical pulses,
which are transmitted optically by transmitter TX.sub.SS (18) to
the implantable sensor platform (2) where they are received by the
photodetector PD.sub.SS (44) and processed by the Mode Selector
(51) and Router/Logic/MUX (multiplexer) block (52). Upon execution
of a selected function, the result is transmitted to the TX.sub.D
driver (56), which in turn powers the TX.sub.D optical transmitter
(45). The TX.sub.D transmitter (45) relays the information via
optical pulses (35) of a different frequency range (f.sub.2)
through the skin (3) to the PD.sub.D photodetector (19) located in
the external unit (1). This signal is then processed by the Add-on
Devices & Control Circuits (14) of the external control unit
(1) in conjunction with the microprocessor (11) and the program
loaded in the Software Interface (12) unit. These steps constitute
a feedback loop to interactively implement a function. This loop is
repeated for every function including initialization, sensor
selection, calibration and measurement described below.
[0051] An exemplary initialization function operates as follows.
The initialization function checks if the solar cells are receiving
adequate optical power from the optical source (17). For this, the
Mode Selector (51) in conjunction with the Router/Logic/MUX block
(52) compares the output of the PD.sub.M photodetector (42) using a
comparator with a predetermined reference, available in the
Calibration Circuit unit (53). If the power level is adequate or
inadequate, a signal is transmitted using the TX.sub.D driver (56)
and TX.sub.D transmitter (45) to the external unit to take the
appropriate action (i.e. if power is adequate proceed with the next
function or if the power level is not appropriate, increase the
power level of the optical source (17) via the circuits in the
Add-on Devices & Control Circuits unit (14).
[0052] An exemplary sensor selection function operates as follows.
The sensor selection function, a command comprising an optical
pulse set is transmitted by TX.sub.SS (18) to the PD.sub.SS (44),
which selects one of the three working electrodes (63, 64, and 65)
shown in Sub-chip #3 (6). Once a sensor is selected, "sensor
calibration" is typically performed. Sensor calibration includes
configuring a programmable potentiostat (54) such that the voltage
between the working electrode (63 or 64 or 65) with respect to
reference electrode (61) is at the desired value dictated by the
electrochemical reaction involving the detection of a certain
analyte. Configuring of the potentiostat determines the appropriate
voltage or current levels, as well as the mode of operation
(continuous or pulsed for certain duration). This configuration is
achieved by Mode Selector (51) and Router/Logic/MUX (52) circuit
blocks in conjunction with the Potentiostat (54) and Calibration
Circuits (53). Once the optional calibration is accomplished, the
next function is sensor reading. This function is performed using
potentiostat (54) and the Signal Processor & ADC block (55).
The digital output of the Signal Processor & ADC block (55) is
then fed to the TX.sub.D driver (56), which in turn powers the
optical transmitter TX.sub.D (45). The analog electrochemical
current generated by the potentiostat-driven sensor [which includes
three electrodes: a working (63 or 64 or 65), a counter (62) and a
reference (61)] is read, amplified, and digitized by the Signal
Processor & ADC block (55). The Signal Processor & ADC
block (55) converts the magnitude of the electrochemical current
into digital pulses whose frequency is proportional to the analyte
level. The digital electrical pulses are converted into digital
optical pulses that are transmitted by TX.sub.D (45). These optical
pulses (35) pass through the skin (3) and are converted back to
electrical pulses by photodetector PD.sub.D (19). These electrical
pulses are decoded by the external control unit (1) using the
Add-on Devices and Control Circuits (14), and the analyte level is
displayed on Display (15).
[0053] Changing from one sensor to another is accomplished, for
example, by re-programming of the Router/Logic/MUX (52), which in
turn reconfigures appropriately the Programmable Potentiostat (54).
The Router/Logic/MUX (52) selects one of the desired working
electrodes (63, 64, or 65). All of these commands are executed from
instructions transmitted using the transmitter TX.sub.ss (18) and
its complementary photodetector (PD.sub.ss) (44) in the implantable
sensor platform (2). Their signals are processed by the Mode
Selector (51), which is interfaced to the Router/Logic/MUX (52).
Router/Logic/MUX (52) performs the reconfiguration and connection
to the calibration circuits. The results of calibration and
comparison are fed through Router/Logic/MUX (52) to the TX.sub.D
Driver (56) and Transmitter TX.sub.D (45) and relayed to the
external control unit to complete the instructional set and desired
function.
[0054] FIG. 2 illustrates an exemplary external control unit (1)
and implantable sensor platform (2) comprising three subchips (4,
5, and 6) that are coated with a biocompatible coating (68)
containing, for example, a number of tissue response modifying
agents. In one embodiment, sensor platform (2) is implanted
subcutaneously under the skin (3) with a hypodermic needle (69)
outfitted with a plunger (70). Implantation takes place, for
example, by lifting the skin up, inserting the needle containing
the sensor platform (2) along with its biocompatible coating (68)
and its plunger (70). After positioning the hypodermic needle (69)
to the proper depth, the plunger (70) is held fixed and the needle
(69) is removed, leaving the sensor platform with its biocompatible
coating at the desired location. Finally the plunger is also
removed. Care is exercised to ensure that the photovoltaic cells
(41) are facing up towards the skin (3). FIG. 2 also schematically
shows the subchip construction, interconnections and optical
powering and communication interface with the external unit (1).
Here, only the optical modules [laser/LED (17), TX.sub.SS (18) and
PD.sub.D (19)] of the external control unit (1) are shown. Some
interconnects are labeled (21). Working electrodes are shown with
specialized coatings (66 and 67), specific to the specific analyte
sensing.
[0055] FIG. 3 illustrates a three-dimensional (3D) schematic
representation of the 0.5.times.5 mm implantable wireless platform,
consisting of three sub-chips that are encapsulated within a
biocompatible coating (34), containing a variety of tissue response
modifiers (TRMs) to control, for example, tissue response and
induce neo-angiogenesis. Sub-chip #3 (6) consists of an
electrochemical sensor interfaced with sub-chip #2 (5) comprising a
potentiostat (54), ADC (55) and IR driver (56) (for glucose
transmitter TX-D (56)). Sub-chip #1 (4) includes the powering solar
cells and IR transmitters. The external command, control and
monitoring, modified-PDA unit is shown above the implantable sensor
capsule. This unit consists of powering-LEDs (17) and
photodetectors (19) for glucose level display (15) and power
control circuits (14 in FIG. 1), respectively. In one design
embodiment, the implantable device comprises three sub-chips, (each
of about. 0.5 mm wide, 5 mm long), which are stacked on top of each
other to achieve a compact configuration, suitable for
needle-assisted administration (see FIG. 2). Sub-chip #1 (top)
comprises the solar cells (41 in FIG. 2), photodetectors and
transmitters (44) (including PD 42 shown in FIG. 2). Sub-chip #3
(6, bottom chip) comprises the three-electrode (working, reference
and counter) electrochemical glucose sensor. Sub-chip #2 (5,
middle) comprises a potentiostat (54), which is operably linked by
vias (301 between subchip #1 and subchip #2, and 302 between
subchip #2 and subchip #3; see also FIG. 10) to the electrochemical
sensor (61-65 in FIG. 1) on subchip #3 and the solar cells on
subchip #1 (4), generating an amperometric current that is fed to
the signal-processing unit (55). This unit comprises a
transconductance amplifier and an analog-to-digital (ADC) converter
to convert amperometric current (i.e., glucose current) to voltage
pulses. These pulses are fed to a driver (56) which in turn feeds
an infra-red 1.55 .mu.m transmitter (TX) (45) (located on sub-chip
#1 (4)), using a low-power laser/LED. The optical pulses that carry
glucose level information (defined as TX.sub.D, D for display) are
received by a photodetector (PD.sub.D) (19/15) located in the
external unit. [Various components of sub-chips are shown in FIG.
9].
[0056] One feature of the device is an interactive feedback system
between the sensor platform and the external unit. This feedback
system provides the ability to, for example; verify adequate power
levels; account for other measurements such as blood pressure, body
temperature, as well as factors such as pH and oxygen level, which
assist to check biointernal unit calibration; and select a sensor
and retrieve the information.
[0057] For example, one feedback system provides information
regarding whether the LEDs (in the external unit) are powered in a
manner that ensures adequate optical input to the solar cells of
the sensor platform (for details see FIG. 13). The solar cells in
turn are supplying various units of the internal unit adequate
voltage, current and power levels. One scheme to ensure this is to
have a feedback system consisting of a photodetector (PD) monitor
measuring the level of optical power received from the LEDs in the
external unit(for details see FIG. 13). The PD output current is
proportional to the light intensity received. This current is
converted into a voltage, which compared in a comparator (COMP)
against a reference. This reference may be derived from the same
circuit, which supplies the reference electrode of the internal
unit. The difference between the reference voltage (V.sub.REF) and
the PD derived voltage is fed to the infrared transmitter (IR-M in
FIG. 3 and TX-M in FIG. 1). This signal is transmitted to the
external unit. The received PD input level is compared in the CTRL
unit, which in turn adjusts the power level of the LEDs.
[0058] The device can further comprise another feedback system
regarding information about other measurements by the internal
unit, such as blood pressure, oxygen and pH levels, which can
affect the calibration or accuracy of the internal unit.
[0059] In another embodiment, the device comprises a plurality of
sensors oberably (operable/fluid/electrical/optical) in
communication with the external unit. The operative connectivity
can be provided by any communication means, such as fluid,
electrical, optical, or a combination of at least one of the
foregoing. The plurality of sensors can be housed within a single
internal unit or multiple internal units. Once a user selects a
particular sensor using the sensor select switch in the external
unit, it activates an IR transmitter in the external unit. This
transmitter communicates and connects to the associated
photodetector (PD) shown schematically in FIG. 1 along with the
block marked VDD, Vref and Control. The communication code enables
the PD to interface with the potentiostat to set desired working
and reference voltages needed to operate the selected sensor. In
addition, the information is provided to the router switch, which
physically selects the sensor and connects the potentiostat output
to its terminals.
[0060] FIG. 4 shows a schematic representation of the optical
powering and communication interface located in the external
control unit (1). This interface also contains a miniaturized
camera (71), which assists in locating and aligning the implantable
sensor platform (2) with respect to the external control unit (1).
The optics elements within the external control unit (1) are housed
within the casing (76) and include three cube reflectors (73, 74,
and 75) equipped with the appropriate coatings to stir reflections
depending to the wavelength of the radiation, along with a number
of focusing lenses (72), transmitters, receivers (photodetectors)
and various other beam shaping components. The light (31) from the
optical powering source (17) is collimated by lens (72), steered by
two cube reflectors (73 and 74) to illuminate a broad area around
the sensor platform (2). This light also provides illumination for
the miniaturized camera (71). The 719 nm light penetrates the skin
adequately and provides enough contrast to the operator to locate
and align the control unit (1) with respect to sensor platform (2).
Optical commands from the TX.sub.SS transmitter (18) are directed
through a lens (72) and two cube reflectors (75 and 74) to
illuminate the similar broad area around the implanted sensor
platform (2). The optical signal from the TX.sub.D transmitter
(45), located in the implantable sensor platform (2) passes thought
the cube reflector (74) and deflected by the mirror of the cube
reflectors (75), collimated through lens (72) and detected by the
photodetector PD.sub.D (19). The insertion of long wavelength pass
filter (10) ensures that no interference from the 719 nm (17) and
800-950 nm (18) light sources occurs. Similarly, the band-pass
filter (43) on top of the PD.sub.SS (44) detector in the
implantable sensor platform (2) ensures no interference from the
719 nm (17) and the 1.55 .mu.m (45) sources, respectively.
[0061] FIG. 5 illustrates the selection of one sensor (80 or 81)
from an implantable sensor platform that comprises more than one
sensor (80, 81). It also demonstrates the operation of a
programmable potentiostat consisting of operational amplifiers (87)
and associated resistor network (82) assisting in the determination
of the appropriate voltage between the Reference (R, shown in 80 or
81) and the Working (W, shown in 80, 81) electrodes of the selected
sensor. A set of coded optical (850-980 nm) pulses (34) (see FIG.
2) transmitted by the TX.sub.SS (18), located in the modified-PDA
unit (101) are received by the photodetector PD.sub.SS (44). The
pulses enable selection of a function (such as potentiostat
reference check, sensor selection, calibration or reading, or
checking solar input power level etc.). These are enabled with the
help of sensor select block 51, switches (shown as 82, 83, 84; and
realized by transistors in integrated circuit chip), circuits
lumped in block 85 and Mux 86. The outcome of this function is
communicated back to the modified-PDA using the 1.55 .mu.m optical
pulse transmitter (45, located at bottom right corner). Using
logic/router block 85 (consisting of reference comparator and
calibration circuits), we can provide the information to MUX
(Multiplexer) circuit (86), which in turn connects, to the driver
(56) of TX.sub.D. The use of different wavelengths and pulse
frequencies for the two transceivers (one for sensor select and the
other for rest of the functions) will further ensure minimal
interference. Here, the sensors (80,81) interface with the current
mirror 88, and signal processor 89, and Mux 86, and the driver 56,
and the transmitter TX.sub.D (45). The current mirror (88)
processes the sensor current and the ADC (89) converts into
electrical pulses of different frequencies depending on the
sensor-produced current level. The driver (56) makes it suitable to
drive the output light emitting device (45, TX.sub.D). In the case
of ADC, we show an operational amplifier (501), a comparator (502),
One shot (monostable multivibrator, 503), and a Divider by 2
circuit (Div 2, 504). There are many variations to implement
this.
[0062] In one embodiment, two sensors (S1 and S2) are connected to
a programmable potentiostat whose reference voltages will be
selected on the sensor under test. The signal processing units can
be miniaturized by reducing the design rules (fineness of
microelectronic features) from 0.35 .mu.m to 0.12 .mu.m and
below.
[0063] In one embodiment, two sensors (S1 and S2) are connected to
a programmable potentiostat whose reference voltages will be
selected on the sensor under test. The control lines and associated
switches (to be replaced by FETs) in the FIGS. 5 and 6 are used to
perform select function (including sensor selection S1 or S2 (see
FIG. 7)) or CP1 and CP2. The signal processing units can be
miniaturized by reducing the design rules (fineness of
microelectronic features) from 0.35 .mu.m to 0.12 .mu.m and
below.
[0064] In one embodiment, the device is run in self-calibration
mode. The potentiostat reference voltage [CP1 or CP2 (Potential
Reference Check 1 or 2) by comparing the voltage difference between
the reference and corresponding working electrode], power level
and/or voltage out put of the solar cells that are powering the
entire chip and transceivers can be checked when the device is in
self-calibration mode. Built-in comparators and logic are used to
achieve the self-calibration functions. In operation, a comparator
receives a specific signal and compares it with the reference
(voltage) and depending on the difference generates a decision,
which is then executed via the logic circuits. For example, the
circuit block labeled "Ref. Comparator and Calibration Circuits"
(85) along with the multiplexer (Mux) (86) enable utilization of
the 1.55 .mu.m transmitter to report back to the PDA the desired
information. (FIG. 5)
[0065] FIG. 6 shows an alternate scheme wherein a PDA unit is
interfaced via wireless technology (e.g., Bluetooth.RTM.) to an
auxiliary external unit (91), which communicates with the implanted
sensor. The signal (from PDA or modified PDA (101)) is received by
unit 91, this signal is converted into optical pulses by an optical
transmitter TX.sub.SS (18). This signal is received by
photodetector PD.sub.SS (44) in the implanted unit. The information
from the implanted sensor is received by photodetector PDM (42) now
located in the auxiliary external unit (91) which in turn transmits
it back to the PDA (101). This type of arrangement is envisioned in
circumstances where a PDA unit is not a dedicated unit for
monitoring analytes.
[0066] FIG. 7 shows an embodiment of a design of sensor select
circuit. The circuit has a photodetector (shown as diode on the
extreme left) PD.sub.SS (44). This photodetector receives optical
pulses (800-950 nm) from the transmitter TX.sub.SS (18) located in
the external unit (see FIG. 2). The signal is amplified by circuit
shown in block (703) which consists of an inverting amplifier (two
transistors) and a tri-state buffer (702). Block (704) is the timer
that generates clock pulse based on our logic block (705)
operation. Depending on the pulse code, one of the six outputs
(labeled as 707, 708, and 709) gets activated. The activation
depends on block (706, consisting of D-Flip Flops) and Logic Block
(705). For example, if the pulse code is set for selection of
Sensor #1 left most output SN1, shown as part of (707), is
selected. This means a high voltage is available, which in turn
will activate switches 82, 83, and 84 (see FIG. 5). This selection
also activates the Router/Logic/Mux (FIG. 1, FIG. 5) circuit in a
certain way to enable the activation of TX.sub.D (45) through
driver (56). The transmitter TX.sub.D, operating at 1.55 .mu.m,
sends a signal showing the selection of Sensor 1. Details of logic
(705) and MUX (86 in FIG. 5) circuits are not shown here.
[0067] When Mode select (13, FIG. 1) sends a code for calibration
check of potentiostat, CP1 and CP2 (shown as 709) are activated.
Similarly, mode select (13) can select calibration control to check
power level of solar cells (41) using CSP-I and CSP-0 (708). In
FIG. 5, the solar power check is written as CSC (in block 51).
[0068] Circuit designers may implement this concept in a variety of
ways.
[0069] FIG. 8 shows photographs of an embodiment of a working
glucose sensor and the data obtained therefrom. FIG. 8(A) shows an
ADC signal processor (MOSIS fabricated chip) interfaced with a
hybrid potentiostat. FIG. 8(B) shows the digital pulses on an
oscilloscope along with the potential difference (680 mV shown on a
meter) between the working and reference electrode of a functional
glucose sensor. The electrochemically detected glucose
concentration is encoded in terms of digital pulse frequency by the
chips in FIG. 8(A). FIG. 8(C) illustrates a plot of glucose
concentration as a function of digital pulsed frequency, generated
by the signal processing chips shown in FIG. 8(A).
[0070] FIG. 9 shows first and second generation microelectronic
sensor component embodiments. The dashed-line square and circle
indicate the first and second generation components, respectively.
Selected components (1-5) are further magnified in panels (B-F)
along with their actual dimensions. (B) shows a 1.sup.st generation
(MOSIS--fabricated) integrated circuit that contains an ADC signal
processor unit and two RF antennas. (C) shows a planar
electrochemical sensor with three electrodes (working (Pt), counter
(Pt) and reference (Ag/AgCl, whitish-appearing, right-most
electrode)). (D) shows an InGaAsP semiconductor laser stack (8 in
number) emitting infra red (IR) radiation of 1.55 .mu.m. The sensor
platform may be equipped with one of these lasers. (E) shows an
individual, Si-based solar cell device. Typically 5 or 6 panels may
be connected in series to power up all components of the
implantable sensor platform. (F) shows a second generation
(MOSIS--fabricated) integrated circuit containing the potentiostat
and ADC signal processors, without the RF antennas. Stacking of all
components from the second generation will result in about a
0.3.times.5 mm implantable platform with provision for 0.1 mm thick
biocompatible coating on each side. The overall 0.5.times.5 mm
sensor can be implanted through a 16 gauge hypodermic needle.
[0071] FIG. 10 shows an exemplary layout of three sub-chips using a
hybrid approach to integrate sub-chips with robust interconnects
1003 between subchip #1 and subchip #2 (labeled as 112G for ground
interconnect between subchip #1 and subchip #2, 112P for power or
voltage supply) and interconnects 1004 between subchip #2 and
subchip #3 (e.g. labels I23W, I23R and I23C; here W for working
electrode, R for reference electrode and C for counter electrodes
connections between sensors on subchip #3 and potentiostat on
subchip #2). The subchips are connected using vias (1001) and bumps
(1002). The vias are V12P, V12S, V12G (numeral 12 refer to subchips
#1 and 2, and the letter P, S, and G refer to function such as
power, signal and ground). The bumps 1002 shown on subchip #3 mate
with the vias on the subchip above this (that is subchip #2). Bumps
are designated as B12P (the bump on subchip #2 locks with the via
on subchip #1; here P refers to power or voltage connection) and
B12G (G for ground connection). Cross-sectional schematic
illustration of bonding between a via and bump for sub-chips #1 and
#2 (bottom left). Top view of interconnect and power pad are also
shown. Bonding between sub-chip #2 and sub-chip #3 is shown using
two approaches used in the integration (bottom right). In FIG. 2
the word `vias` (301 and 302) was used for the complete connection
between two subchips including via, interconnect and the bump.
[0072] FIG. 11 shows another embodiment of a three sub-chip design
(shown in FIGS. 1, 2 and 3) Here the top surface of Sub-chip #1 (4)
is hermetically sealed using a transparent glass window (111)
sealed to raised silicon walls (110) via anodic bonding (112)
between the glass and the silicon. Similar hermetic seals are
employed between the bottom surface of Sub-Chip #1 (4) and the top
surface of Sub-Chip #2 (5), as well as the bottom surface of
Sub-Chip #2 (5) and the back surface of Sub-Chip #3 (6). The front
surface of Sub-Chip #3 (6) hosts various electrodes that are
designed to operate in body fluids. All interconnects among
sub-chips are vias filled with non-corrosive metals that provide
hermetic seals for microelectronic and optoelectronic components of
the implantable sensor platform.
[0073] FIG. 12 shows an embodiment where the three sub-chip design
has been implemented as two sub-chips. This design is possible
using finer design rules (i.e., 0.35 micron versus 0.5 micron chip
processing) for the potentiostat and signal processing components
(usually located on sub-chip #2, FIG. 11), which results in saved
space on a Si chip. As a result, functions of sub-chip #2 and
sub-chip #3 could be integrated as one sub-chip (referred to as
(120) the new-sub-chip #2); thus a sensor platform having two
sub-chips can be realized. An exemplary embodiment is illustrated
in FIG. 12. New-Sub-chip #2 (120), according to this embodiment,
comprises a .about.0.25.times.4 mm MOSIS chip (122) (like part 5 of
FIG. 1, or electrical circuit of FIG. 5) bonded face down on the
carrier 0.3.times.5 mm platform (120) that hosts along its
perimeter the electrochemical sensor electrodes (61,62,63,64). The
bonding pads (123) are connected to the bump and or vias (shown as
1002 and 124). Two working electrodes (63,64) are placed at either
side along the width in a meander-form to improve adhesion of the
glucose oxidase coating. Lengthwise, a Pt counter (62) and an
Ag/AgCl reference electrode (61) are also placed to function in
conjunction with either working electrode. In addition, a 1.55
.mu.m InGaAsP transmitter TX.sub.D (45) (and TX.sub.M which is for
monitoring and could be one unit via MUX (18) circuits) is placed
on the upper right corner, adjacent to the working electrode 2
(64). The wafer carrier is a high resistivity (>20,000 Ohm/cm)
Si (100) in which bumps, vias and interconnects are patterned and
metallized prior to growing the organic layers for the
electrochemical sensor. The bumps and vias are also shown. Both the
MOSIS and the 1.55 .mu.m InGaAsP transmitter chip are placed face
down and their pads are connected to the pads on the carrier
chip.
[0074] The subchip #1 (4) is shown as part new subchip #1 (121). In
this embodiment, new-sub-chip #1 has 3 pads on either side for
power supply distribution (e.g., V.sub.dd, V.sub.1 for voltages,
and C for common). The power is supplied to the new-sub-chip #2
using metalized vias labeled as V.sub.VDD (124). In addition, vias
are used to connect photodetectors PD.sub.SS (44) and PD.sub.M on
new-sub-chip #1 to new-sub-chip #2 (having electronics such as
sensor select, routing logic/MUX, etc.). Note that the 1.55 .mu.m
transmitter is located on new-sub-chip #2 as this wavelength is
transparent to Si platform and chips. Solar cells (41) are not
shown individually as in FIG. 2.
[0075] FIG. 13 shows embodiments of various circuits that may be
integrated into two sub-chips: (subchip #1 (121 of FIG. 12 or 4 of
FIG. 1) and subchip #2 (120) of FIG. 12. It also shows some
circuits (14 called Add-on Devices & Control Circuits in FIG.
1) that may be housed in the modified PDA unit (101). Here, the
modified PDA is communicating directly with the implanted unit.
[0076] In some embodiments, it is advantageous to employ an
auxiliary external unit (91) [of FIG. 6] that communicates with a
PDA (101) on one hand and the implanted sensor platform on the
other. This is shown in FIG. 14. In this embodiment, the PDA unit
is communicating with an auxiliary external unit (91) via wireless
technology (e.g., Bluetooth.RTM.). The external unit communicates
with the implanted unit using optical communication.
[0077] FIG. 15 shows a methodology to integrate two sub-chips into
one wafer platform using wafer bonding methodology. This advanced
integration does not require vias and bumps as discussed above. In
this embodiment, the photodetectors [PD.sub.M (44) and PD.sub.SS
(42)] and solar cells (41) are realized on top part of the Si chip
(.about.20-40 microns in thickness) having electrical resistivity
in the 1-10 Ohm-cm. This wafer is bonded [using silicon dioxide
layer (151) or other wafer fusion techniques known in the
literature] to a high resistivity (.about..about.10,000 Ohm-cm)
wafer. This wafer or the bottom side has sensors [comprising of
various electrodes (61), (62), (63), and (64)] processed along with
signal processing chip (5), which is mounted in a recessed part.
The recessed part is created by deep RIE (reactive ion etching) or
other technique. The top part of the wafer can be hermetically
sealed using anodic bonding using glass 111 and bonded seals
112.
[0078] FIG. 16 shows a methodology where the internal unit
implantable platform (160) is integrated with drug delivery devices
(shown as three modules below the command Control Display Wrist
Unit (101D)). These units are also implanted separately or
integrated along with the internal unit. Based on the implanted
sensor reading of glucose or insulin, the control unit 101D (e.g.
modified PDA unit) sends the signal to the insulin dispenser to
dispense desired amount of insulin. Here, we have three sub-units
or sub-modules identified as subchip #1 (powering module similar to
part 4 of FIG. 1), Dispenser subchip #2 (161), and Dispenser
subchip #3 (162). Subchip #1 provides the power to unit (161) and
unit (162). Dispenser subchip (161) consists of sensor, signal
processing [as carried out by subchip #2 (5) and subchip #3 (6) in
FIG. 1], and control circuits (163) to actuate MEM
(microelectrochemical) actuators (164, 165). MEM actuators activate
Valves (166, 167) located on the Dispenser subchip #3 (162) which
connect to the insulin drug reservoir (168). The insulin is
dispensed via 173 and 174 if needed at two locations. The implanted
biosensor in turn provides information regarding the glucose levels
(169) or insulin levels. This information is fed to the processor
(170) of the command unit (101D) where it is compared with expected
insulin levels per software protocol and executed by Level
Comparator unit (171). This in turn activates Drug Delivery Module
(172) which relays the information to control electronic circuits
(163). The command unit (101D) has optical power source (17) for
the solar cells (41) on subchip #1 (4).
[0079] FIG. 17 illustrates glucose sensor with an Ag reference
electrode and details of coatings on Pt working electrode. This is
a schematic representation of modified Clark amperometric glucose
sensor, along with various chemical, electrochemical and diffusion
processes associated with its operation. The working Pt electrode
(63 or 64 if more than one) is coated with electropolymerized
o-phenylddiamine. PPD layer (171) (e.g., prevents permeation of
ascorbic acid, acetaminophen etc.) is coated with glucose oxidase
(GO.sub.x) layer (172), which in turn is coated with a
semipermeable humic acid membrane (173) to reduce the amount of
glucose entering the sensor. An HRP-modified hydrogel layer (174)
is then applied to eliminate the outer diffusion of H.sub.2O.sub.2.
This is followed by an outer composite hydrogel coating with
embedded microspheres at different stages of degradation and TRM
release (175). The tissue is represented by (176).
[0080] In one embodiment, an enzyme-based glucose sensor operates
on the principle of detection of hydrogen peroxide (H.sub.2O.sub.2)
formed by glucose oxidation. Glucose oxidase (GO.sub.x) acts as a
catalyst, which turns glucose into gluconic acid (Reaction 1) and
produces H.sub.2O.sub.2. H.sub.2O.sub.2 is electrochemically
oxidized (Reaction 2) under an applied potential of 0.7 V and the
current measured is related to the glucose concentration (see
diagram in FIG. 17). The semipermeable membrane, depicted in FIG.
17, in addition to assisting in prevention of biofouling, is
designed to regulate glucose diffusion. It is well known that for
an enzyme-based implantable sensor to work at its optimum
efficiency in tissue, the ratio of oxygen, which regenerates the
enzyme, to the permeating glucose should remain constant.
Typically, the physiological levels for glucose and oxygen in
subcutaneous tissue are 5.6 and 0.1 mM, respectively. Thus, in the
absence of a diffusion-limiting barrier for glucose, the kinetics
of H.sub.2O.sub.2 production may be oxygen-limited due to the
significantly larger amount of glucose compared to oxygen. At high
glucose concentrations, this oxygen limit can lead to reduced
glucose sensitivity. Therefore, a method to achieve accurate
monitoring of glucose over the entire physiological concentration
range with high sensitivity and short response time an outer
membrane with tunable permeability properties is needed. Such a
membrane has been developed through layer-by-layer grown
polyelectrolytes and/or multi-valent cations.
[0081] The GOx and/or carrier protein concentration may vary. For
example, the GOx concentration is about 50 mg/ml (approximately
10,000 U/ml) to about 700 mg/ml (about 150,000 U/ml). In
particular, the GOx concentration is about 115 mg/ml (approximately
22,000 U/ml). In such embodiments, the HSA concentration is about
0.5%-30% (w/v), depending on the GOx concentration. In particular,
the HSA concentration is about 1-10% w/v, and most particularly is
about 5% w/v. In alternative embodiments, collagen or BSA (Bovine
Serum Albumin) or other structural proteins used in these contexts
can be used instead of or in addition to HSA. Although GOx is
discussed as an enzyme in the sensor element, other proteins and/or
enzymes may also be used or may be used in place of GOx, including,
but not limited to glucose dehydrogenase or hexokinase, hexose
oxidase, lactate oxidase, and the like. Other proteins and/or
enzymes may also be used, as will be evident to those skilled in
the art. Moreover, although HSA is employed in the example
embodiment, other structural proteins, such as BSA, collagens or
the like, can be used instead of or in addition to HSA.
[0082] For embodiments employing enzymes other than GOx,
concentrations other than those discussed herein may be utilized.
The concentration may be varied not only depending on the
particular enzyme being employed, but also depending on the desired
properties of the resulting protein matrix. For example, a certain
concentration may be utilized if the protein matrix is to be used
in a diagnostic capacity while a different concentration may be
utilized if certain structural properties are desired. Those
skilled in the art will understand that the concentration utilized
may be varied through routine experimentation to determine which
concentration (and of which enzyme or protein) may yield the
desired result.
[0083] FIG. 18 shows the self-assembly of semipermeable membrane
composed of humic acids and Fe.sup.3+ on the outer surface of the
electrochemical sensor. As the number of bilayers increases, there
is an increase in the tortuosity for glucose diffusion towards the
enzyme. The glucose oxidase (GO.sub.x) layer is coated with a
semipermeable humic acid membrane to reduce the amount of glucose
entering the sensor.
[0084] For testing purposes, a miniaturized sensor (shown in FIG.
9) made of platinum evaporated on a high resistivity Si wafer has
been developed. To ensure Pt bonding to the Si wafer, a number of
coatings have been employed to eliminate delamination problems once
implantation takes place. In particular, Au/Ti/Pt/Ti/Ag coatings
may be employed. Silver is removed from the working and counter
electrodes. Ag is converted to AgCl to form a reference electrode.
In the case of the working electrode, a film of poly
(o-phenylenediamine) is electropolymerized on the working
electrode, following which the sensing enzyme, i.e., glucose
oxidase, is immobilized on top. Finally a semipermeable membrane
composed of humic acids and Fe.sup.3+ ions is grown on the device
through electrostatic self-assembly. The sensor is tested via
amperometry in phosphate buffered saline (PBS) solution maintained
at 37.degree. C. Glucose is added to the solution following which a
pulse of 0.7V is applied to the device every 5-10 minutes until a
constant current reading is obtained. Once the device has
stabilized, glucose concentration is incremented and the process
repeated to generate a calibration curve as shown in FIG. 16.
[0085] FIG. 19 illustrates pulsed mode operation of the sensor. In
pulsed mode operation, voltages are applied to various sensor
electrodes for a certain duration. Calibration curve of current
response vs. change in glucose concentration is shown for different
glucose concentrations. As the amount of glucose in the system is
increased, there is a corresponding rise in the current. The Figure
on left indicates the time required for the device to stabilize at
each glucose concentration. As the concentration of glucose
increases, the device reaches a stable reading faster.
[0086] FIG. 20 shows that by varying the number of dip-cycles (2, 5
and 10), the current is reduced by almost 10 fold, while
maintaining current linearity. The thickness of the film depends on
the number of dip-cycles. The ability of these membranes to act as
an efficient barrier for glucose permeation has thus been
demonstrated.
[0087] FIG. 21 shows histological evaluations of subcutaneous
tissue samples taken from the vicinity of hydrogel composites
containing PLGA microspheres at 3 and 21 days post implantation.
The representative sections shown are 3 days after implantation (A
& B) and 21 days after implantation (C & D). There seems to
be no inflammation showing the effectiveness of the coatings.
[0088] FIG. 22 shows a schematic cross-section of an embodiment of
implantable sensor platform with a methodology to hermetically seal
sub-chips using one high resistivity Si wafer as the carrier (220)
which has one recessed region (221) on one side. In this region
solar cells (41) are bonded on metallic pads (224) deposited on an
oxide layer (223). Photodetectors (44) and (42) are also placed on
this side. The carrier wafer (220) body is recessed (222) on the
other side as well. Thinned sub-chips 2 (5) and subchip #3 (6) are
placed in the recessed region (222).
[0089] The sub-chips can be electrically interconnected using
interconnects like (226) which run in vias like (225) as shown or
other standard interconnect techniques may be used. The hermetical
seal provided by the glass-Si anodic bonding (112) is shown on top
as well as bottom surfaces of the carrier (220). This is an
alternate approach for three sub-chip integration as shown in FIGS.
11 and 15.
[0090] A variety of optional items may be included in the sensor
platform. One optional item is a temperature probe. One exemplary
temperature probe comprises two probe leads connected to each other
through a temperature-dependent element that is formed using a
material with a temperature-dependent characteristic. An example of
a suitable temperature-dependent characteristic is the resistance
of the temperature-dependent element. The two probe leads comprise,
for example, a metal, an alloy, a semimetal, such as graphite, a
degenerate or highly doped semiconductor, or a small-band gap
semiconductor. Examples of suitable materials include gold, silver,
ruthenium oxide, titanium nitride, titanium dioxide, indium doped
tin oxide, tin doped indium oxide, or graphite. The
temperature-dependent element can further comprise a fine trace
(e.g., a conductive trace that has a smaller cross-section than
that of the probe leads) of the same conductive material as the
probe leads, or another material such as a carbon ink, a carbon
fiber, or platinum, which has a temperature-dependent
characteristic, such as resistance, that provides a
temperature-dependent signal when a voltage source is attached to
the two probe leads of the temperature probe. The
temperature-dependent characteristic of the temperature-dependent
element can either increase or decrease with temperature.
[0091] The sensor platform comprises components manufactured from
biocompatible materials, such as materials that are corrosion
resistant, INCLUDING Pt, SiO.sub.2 coatings, and glass thin films.
In addition, corrosion resistant materials that are harmless to
tissues in biologic environments, such as silicon and heavily
boron-doped silicon can be used in the manufacture of the
components of the internal unit. Another method by which the
corrosion resistance of the internal unit can be improved is
through coating of the internal unit with titanium, iridium,
Parylene (a biocompatible polymer), or various other common and/or
proprietary thick and thin films.
[0092] The sensor platform optionally comprises a biocompatible
coating. The bioactive polymers are generally biocompatible, that
is, physiologically tolerated, and do not cause substantial adverse
local or systemic responses. While synthetic polymers such as
poly(tetrafluoroethylene), silicones, poly(acrylate),
poly(methacrylate), hydrogels, and derivatives thereof are most
commonly used, natural polymers such as proteins and carbohydrates
are also suitable. The bioactive polymer layer functions to protect
the implant, preserve its function, minimize protein adsorption
onto the implant, and serve as a site for the delivery of the
tissue response modifying agents and drugs as well as other drugs
and factors.
[0093] In one embodiment, the bioactive polymer layer comprises a
hydrogel. Hydrogels are formed from the polymerization of
hydrophilic and hydrophobic monomers to form gels and are
described, for example, in U.S. Pat. No. 4,983,181 and No.
4,994,081, which are incorporated by reference herein. Hydrogels
consist largely of water, and may be crosslinked by either chemical
or physical methods. Chemical crosslinking is exemplified by the
free-radical induced crosslinking of dienes such as ethylene glycol
dimethacrylate (EGDMA), and the like. Physical crosslinks are
formed by copolymerizing a hydrophobic co-monomer with the
water-soluble monomer, and then by contacting the copolymerized gel
with water. Physical association of the hydrophobic regions of the
gel results in the formation of physical crosslinks. Control of the
ratio of hydrophilic to hydrophobic monomers allows control of the
final properties of the gel. Physical crosslinks can also be formed
by freeze/thaw methods, for example freeze/thawing a poly(vinyl
alcohol) (PVA) hydrogel. Highly water-swollen hydrogels are
bioactive, and have minimal impact on the diffusion rates of small
molecules. Hydrogels are also intrinsically mobile, and therefore
have minimal deleterious effects on associated peptide tissue
response modifiers.
[0094] Hydrogels may be formed by the polymerization of monomers
such as 2-hydroxyethyl methacrylate, 2-hydroxyethyl methacrylate,
fluorinated acrylates, acrylic acid, and methacrylic acid, and
combinations thereof. Suitable hydrogels include copolymers of
2-hydroxyethyl methacrylate, wherein the co-monomers are selected
to improve mechanical strength, stability to hydrolysis, or other
mechanical or chemical characteristics. Copolymerization with
various acidic monomers can decrease the buffer capacity of the
gel, and thus modulate the release of the tissue response modifier.
Suitable co-monomers include, but are not limited to,
3-hydroxypropyl methacrylate, N-vinyl pyrrolidinone, 2-hydroxyethyl
acrylate, glycerol methacrylate, n-isopropyl acrylamide,
N,N-dimethylacrylamide, glycidyl methacrylate, and combinations
thereof. Suitable hydrogels are terpolymers of 2-hydroxyethyl
methacrylate (HEMA), N-vinyl pyrrolidinone (NVP), and
2-N-ethylperflourooctanesulfanamido ethyl acrylate (FOSA) with
added EGDMA to provide controlled crosslinking. HEMA is
hydrophilic, and swells in the presence of water. The hydroxyl
groups of HEMA also provide potential sites for the covalent
attachment of tissue response modifiers, slow release delivery
systems, and the like. Acrylic acid, methacrylic acid, and other
functionalized vinyl monomers can also be employed to provide these
attachment sites. NVP is amphiphilic, wherein the backbone ring
provides hydrophobicity and the polar amide group provides
hydrophilicity. Poly(vinyl pyrrolidinone) is water soluble,
physiologically inactive, and forms complexes with a number of
small molecules such as iodine and chlorhexidine. Use of NVP
improves the toughness of polymerized HEMA, and provides for the
enhanced solubility of the other monomers under bulk polymerization
conditions.
[0095] An example of a bioactive layer generated by self-assembly
is the formation of NAFION.TM./Fe.sup.3+ multilayer films.
NAFION.TM. is a perfluorinated electrolyte having sulfonic acid
functionalities that has been previously used as a semipermeable
membrane for electrochemical sensors. However, the strong
ion-exchange properties of NAFION.TM. lead to calcification in
vitro and in vivo. The sulfonate (R--SO.sub.3) groups present in
the hydrophilic domains of the membrane act as nucleating sites for
deposition of calcium phosphate. These crystals tend to inhibit
metabolite transport through the membrane, and also cause the
membrane to become brittle and eventually crack. Electrostatic
assembly of NAFION.TM. and Fe.sup.3+ from dilute solutions of
ferric citrate at a pH about 2 to 6 can be used to prevent calcium
deposition.
[0096] A natural bioactive coating is a mussel adhesive protein
(MAP). Self-assembly of biological materials such as mussel
adhesive proteins allows the incorporation of materials, which
improve implant biocompatibility. MAP produced by the blue seal
mussel (Mytilus edulis) generally comprises 75 to 85 repeating
decameric units having the primary sequence of KPSY-Hyp-Hyp-T-DOPA,
wherein Hyp is hydroxyproline and DOPA is
3,4-dihydroxyphenylalanine. DOPA is a strong metal chelating agent,
particularly with Ca.sup.2+ and Fe.sup.3+, and the strong
self-aggregation of DOPA in the presence of cations results in
supra-molecular self-assembly. Accordingly, a substrate comprising
metal chelating groups, for example free amine groups, is
sequentially immersed first in a solution comprising metal ions
(i.e., Ca.sup.2+ and/or Fe.sup.3+) (followed by optional washing in
fresh solvent); and second, in a solution comprising the
poly(ligand) (i.e., the MAP protein) (followed by optional washing
in fresh solvent). The thickness of the membrane will be directly
proportional to the number of sequential immersion cycles. The
assembly of the membrane may be monitored with Variable Angle
Spectroscopic Ellipsometry (VASE), UV-VIS and Quartz Crystal
Microbalance. The strong chelation between Ca.sup.2+ and DOPA in
the MAP membrane results in a substantial decrease in porosity,
allowing the permeation of small molecules such as glucose and
oxygen, while excluding permeation of larger molecules.
Additionally, the introduction of small amount of crosslinking, via
the Michael addition from neighboring lysine repeats by slight
increase of pH above 8.5, which may be used to further fine-tune
the permeability of such assemblies to levels.
[0097] Humic acids may also be polymerized, or self-assembled into
a biocompatible layer. Humic acids or "humic substances" are
heterogeneous, high-molecular weight organic acids having a large
proportion of DOPA, and are resistant to microbial degradation. The
known ability of humic acids to donate and accept electrons from a
variety of metals and organic molecules explains their capability
to shuttle electrons between the humic-reducing microorganisms and
the Fe(III)-Fe(II) oxide. It has been suggested that humic acids
participate in a biological electron transfer as a result of the
electron accepting ability of quinone moieties when reduced to
hydroquinones and vice-versa. This renders the Fe.sup.3+/humic acid
assembled membranes an attractive vehicle for the attachment to
various kind of biocompatible layer.
[0098] Other components may also be incorporated into the bioactive
polymer layer, such as poly(ethylene oxide) (PEG), to minimize
protein adsorption. Poly(ethylene oxide) is most readily
incorporated into the hydrogel, for example, by co-polymerization
of a vinyl monomer having poly(ethylene oxide) side chains, for
example poly(ethylene glycol) methacrylate (which is commercially
available from Aldrich Chemical Co.), or a divinyl-terminated
poly(ethylene glycol) macromonomer. Copolymerization of HEMA and
poly(ethylene glycol) methacrylate in the presence of AIBN yields a
more flexible, unhydrated copolymer. The optimal molecular weight
and content of poly(ethylene oxide) for each application can be
determined by protein adsorption studies.
[0099] To provide further chemical functionality on the bioactive
polymer layer, particularly a hydrogel layer, either polyvinyl
alcohol or polyethylene imine may be employed as macromolecular
surfactants. Where hydroxyl functionalities are available, the
coupling is promoted by tresylation. Poly(ethylene oxide) may also
be grafted to hydroxyl groups on the surface of the polymer layer
by tresylation coupling with Jeffamine, an amine-terminated
poly(ethylene oxide) commercially available from Huntsman.
[0100] In one embodiment, the biocompatible layer comprises a
biocompatible membrane, which is permeable to analytes, such as
oxygen and glucose, but is impermeable to, for example, white blood
cells and macrophages to prevent these cells from contacting other
components of the internal unit. The biocompatible membrane can
comprise polymers including, but not limited to, polypropylene,
polysulphone, polytetrafluoroethylene (PTFE), and poly(ethylene
terephthalate) (PET). The biocompatible layer should be biostable
for long periods of time (e.g., several years).
[0101] The internal unit can also comprise a mass
transport-limiting layer to act as a diffusion-limiting barrier to
reduce the rate of mass transport of the analyte, for example,
glucose or lactate, into the internal unit. By limiting the
diffusion of the analyte, the steady state concentration of the
analyte in the proximity of the working electrode (which is
proportional to the concentration of the analyte in the body or
sample fluid) can be reduced. This extends the upper range of
analyte concentrations that can still be accurately measured and
can also expand the range in which the current increases
approximately linearly with the level of the analyte.
[0102] In some embodiments, the mass transport limiting layer can
also limit the flow of oxygen into the internal unit. This can
improve the stability of internal units that are used in situations
where variation in the partial pressure of oxygen causes
non-linearity in internal unit response. In these embodiments, the
mass transport limiting layer restricts oxygen transport by at
least 40%, specifically at least 60%, and more specifically at
least 80%, than the membrane restricts transport of the analyte. In
these embodiments, the mass transport limiting layer comprises a
film that is less permeable to oxygen, for example, by having
density closer to that of the crystalline polymer, such as
polyesters including polyethylene terephthalate.
[0103] FIG. 16 shows a methodology where the internal unit platform
is integrated with drug delivery devices also implanted or
integrated along with the internal unit.
[0104] In one embodiment, the drug delivery device delivers a
tissue response modifier. "Tissue response modifiers" as used
herein are factors that control the response of tissue adjacent to
the site of implantation. One facet of this response can be broadly
divided into a two-step process, inflammation and wound healing. An
uncontrolled inflammatory response (acute or chronic) results in
extensive tissue destruction and ultimately tissue fibrosis. Wound
healing includes regeneration of the injured tissue, repair
(fibrosis), and in-growth of new blood vessels (neovascularization
and angiogenesis). For fibrosis, the body utilizes collagen from
activated fibroblasts to "patch and fill" the unregenerated areas
resulting from trauma and inflammation.
[0105] Fibrosis can lead to "encapsulation" or "entombment" of the
sensor in fibrotic tissue and this can lead to loss of analyte
supply and loss of functionality of the sensor. In-growth of new
blood vessels is critical to the ultimate outcome of wound healing.
A number of other responses are also included within this category,
for example fibroblast formation and function, leukocyte
activation, leukocyte adherence, lymphocyte activation, lymphocyte
adherence, macrophage activation, macrophage adherence, thrombosis,
cell migration, cell proliferation including uncontrolled growth,
neoplasia, and cell injury and death. Adverse tissue responses to
implantation may also arise through genetic disorders, immune
diseases, infectious disease, environmental exposure to toxins,
nutritional diseases, and diseases of infancy and childhood.
[0106] Tissue response modifiers are therefore a broad category of
organic and inorganic, synthetic and natural materials, and
derivatives thereof which affect the above responses to tissue
injury upon implantation. Such materials include but are not
limited to synthetic organic compounds (drugs), peptides,
polypeptides, proteins, lipids, sugars, carbohydrates, certain RNA
and DNA molecules, and fatty acids, as well metabolites and
derivatives of each. Tissue response modifiers may also take the
form of, or be available from genetic material, viruses,
prokaryotic or eukaryotic cells. The tissue response modifiers can
be in various forms, such as unchanged molecules, components of
molecular complexes, or pharmacologically acceptable salts or
simple derivatives such as esters, ethers, and amides. Tissue
response modifiers may be derived from viral, microbial, fungal,
plant, insect, fish, and other vertebrate sources.
[0107] Exemplary tissue response modifiers include, but are not
limited to, anti-inflammatory agents such as steroidal drugs, for
example corticosteroids such as Dexamethasone
(9-alpha-fluoro-16-alpha-methylprednisolone), a potent, broad
spectrum steroidal anti-inflammatory and anti-fibrotic drug with
known efficacy in a diabetic rat model, methyl prednisone,
triamcoline (fluoroxyprednilisone), hydrocortisone
(17-hydroxycorticosterone); and non-steroidal drugs, for example
Ketoprofin (2-(3-benzophenyl)propionic acid), cyclosporin, Naproxin
((+)-6-methoxy-alpha-methyl-2-naphthalene acetic acid), and
Ibuprofin (4-isobutyl-alpha-methylphenyl acetic acid).
[0108] Other exemplary tissue response modifiers include
neovascularization agents such as cytokines. Cytokines are growth
factors such as transforming growth factor alpha (TGFA), epidermal
growth factor (EGF), vascular endothelial growth factor (VEGF), and
anti-transforming growth factor beta (TGFB). TGFA suppresses
collagen synthesis and stimulates angiogenesis. It has been shown
that epidermal growth factor tethered to a solid substrate retains
significant mobility and an active conformation. VEGF stimulates
angiogenesis, and is advantageous because it selectively promotes
proliferation of endothelial cells and not fibroblasts or collagen
synthesis, in contrast to other angiogenic factors. In addition to
promoting would healing, the improved blood flow resulting from the
presence of neovascularization agents should also improve the
accuracy of sensor measurements.
[0109] Another type of tissue response modifier is a neutralizing
antibody including, for example, anti-transforming growth factor
beta antibody (anti-TGFB); anti-TGFB receptor antibody; and
anti-fibroblast antibody (anti-CD44). Anti-TGFB antibody has been
shown to inhibit fibroblast proliferation, and hence inhibit
fibrosis. Because of the importance of TGFB in fibrosis, anti-TGFB
receptor antibodies inhibit fibrosis by blocking TGFB activation of
fibroblasts. Recent studies have demonstrated that anti-CD44
antibody induces programmed cell death (apoptosis) in fibroblasts
in vitro. Thus, use of anti-CD44 antibody represents a novel
approach to inhibition of fibroblast formation, and therefore
fibrosis. Other anti-proliferative agents include Mitomicyin C,
which inhibits fibroblast proliferation under certain
circumstances, such as after vascularization has occurred.
[0110] Adhesive ligands ("binding motifs") may also be used as
tissue response modifiers, wherein the adhesive ligands are
incorporated into the polymer layer to stimulate direct attachment
of endothelial cells to implant surfaces. Such attachment promotes
neovascularization at the implant/tissue interface. Where the
surface density of binding motifs has an effect on the cellular
response, variation in the density of the binding motifs allows
control of the response. Exemplary adhesive ligands include but are
not limited to the arginine-glycine-aspartic acid (RGD) motif, and
arginine-glutamic acid-aspartic acid-valine (REDV) motif, a
fibronectin polypeptide. The REDV ligand has been shown to
selectively bind to human endothelial cells, but not to bind to
smooth muscle cells, fibroblasts or blood platelets when used in an
appropriate amount. Sensors detecting body temperature, blood
gases, ionic concentrations can be incorporated in the implantable
sensor platform. The analyte sensing device of Claim 1, wherein the
sensor element comprises a body temperature sensor, a blood
pressure sensor, a pH sensor, an oxygen sensor, a glucose sensor, a
lactate sensor, or a combination comprising one or more of the
foregoing sensors.
[0111] In operation, the device can use any mechanism (e.g.,
enzymatic or non-enzymatic) by which a particular analyte can be
quantitated.
[0112] The devices and methods disclosed herein can be applied to
determine the metabolic levels of many analytes present in
biological fluids, including, but not limited to, glucose, amino
acids, and lactate. Suitable analytes include analytes that are
substrates for oxidase enzymes.
[0113] Some analytes, such as oxygen, can be directly
electrooxidized or electroreduced on the working electrode. For
other analytes, such as glucose and lactate, an electron transfer
agent and/or a catalyst can facilitate the electrooxidation or
electroreduction of the analyte. Catalysts can also be used for
those analytes, such as oxygen, that can be directly
electrooxidized or electroreduced on the working electrode. For
example, some embodiments can quantitate metabolic glucose levels
by using a membrane comprising glucose oxidase (see FIG. 17) that
catalyzes the conversion of glucose and molecular oxygen to
gluconate and hydrogen peroxide:
Glucose+O.sub.2.fwdarw.Gluconate+H.sub.2O.sub.2. Because for each
glucose molecule converted to gluconate, there is a proportional
change in the co-reactant O.sub.2 and the product H.sub.2O.sub.2,
one can monitor the current change in either the co-reactant or the
product to determine glucose concentration.
[0114] In one embodiment, the sensor element comprises an
electrochemical pH sensor. Since a large number of biological
processes are pH-dependent, there is a great need for outfitting
miniaturized biosensors with a pH sensing element. The need for
maintaining biocompatibility limits the use of traditional
materials for the fabrication of pH-sensors (i.e., electrically
semiconducting oxides such as MoO.sub.2,.sup.43 IrO.sub.2,.sup.44
or RuO.sub.2.sup.45) due to their toxicity. Biocompatible polymers
that contain nitrogen or oxygen moieties amenable to protonation
have been used to develop potentiometric pH biosensors. Polyphenol,
polyaniline, poly(1,2-diaminobenzene), poly(4,4'-diaminoddiphenyl
ether), etc. have been employed in fabricating pH sensors. The
electrochemistry of these polymers however, is greatly affected by
redox reagents (such as H.sub.2O.sub.2) and based upon prior
experience with poly(o-phenyl diamine), positioning such a sensing
element in the vicinity of a glucose sensor (which produces
H.sub.2O.sub.2) could affect the measurements. More recently,
linear polyethylenimine (L-PEI) and linear polypropylenimine
(L-PPI) modified Pt electrodes have been successfully used for the
development of miniaturized electrochemical pH sensors with a
linear pH range from 4-9. The non-semiconducting nature of both
L-PEI and L-PPI polymers render them ideal for operation within a
redox-prone environment and their biocompatibility and long-term
stability (when operated in a three-electrode configuration)
renders them ideal for the development of our miniaturized pH
sensors.
[0115] In one embodiment, ethylenediamine (EDA) or
1,3-diaminopropane (1,3-DAP) is be electropolymerized onto flat
substrate by means of cyclic voltametry in solutions composed of
10.sup.-2 M N-lithiotrifluromethane-sulfonimide (LiTFSI) in pure
EDA or 1,3-DAP, by biasing the working electrode at 3V with respect
to a standard reference electrode, for a duration ranging from 5-60
min..sup.4 This will result in the electrodeposition of either
L-PEI or L-PPI on the biased electrode, leaving the rest of the
electrodes intact. Planar electrodes will be defined
microlithographically. These electrodes will be grown by
evaporating first a thin layer of Ti to improve adhesion on the
SiO.sub.2-covered Si wafer, followed by the deposition of a thick
layer or Pt and an optional second thin layer of Ti to enable the
adhesion of a SiN overlayer. This SiN layer is used to protect the
edges of the microlithographically defined Pt electrodes from
delaminating in an aqueous environment. Following SiN patterning,
the remaining Ti is stripped off by immersion of the wafer in a
titanium etchant (i.e., H.sub.2SO.sub.4/H.sub.2O-1/1, 80.degree.
C.) to expose the underlying Pt layer.
[0116] Electroplating Ag on top of one of the patterned Pt
electrode will be used to selectively grow the reference Ag/AgCl
electrode. This can be accomplished by electroplating in a solution
comprising KCN, K.sub.2CO.sub.3, and KAg(CN).sub.2. Subsequent
electrochemical oxidation of Ag to AgCl will take place at a
constant current of 40 .mu.A (at .about.0.5 V) in 0.1 M HCl for
approximately 10-30 minutes Since only the reference electrode is
connected to the voltage source, no deposition occurs on the other
electrodes, which remain clean for the subsequent electrodeposition
of L-PEI or L-PPI (described above). The use of an auxiliary Pt
electrode can improve device reliability and long-term
operation.
[0117] The fabrication of such a pH sensor is simple and
straight-forward. The thickness of the electropolymerized L-PEI or
L-PPI are reported to influence the sensor response. In the case
where there is interference of H.sub.2O.sub.2 with the pH sensor,
this should to be quantified and included in the multi-parameter
sensor response characteristics.
[0118] In one embodiment, the sensor element comprises an
electrochemical oxygen sensor. Variations in the partial pressure
of O.sub.2 in the blood is expected to have a significant effect on
the glucose sensor response. This is because of the dual role of
O.sub.2 in GO.sub.x enzymatic catalysis to form H.sub.2O.sub.2 and
its subsequent oxidation to regenerate O.sub.2. Providing an
independent assessment of O.sub.2 concentration could improve our
level of confidence in sensor accuracy and reliability. Design
simplicity, stability and good current linearity over the range of
oxygen from 0 to 99.5% v/v have rendered electrochemical-based
Clark sensors as the preferred method for O.sub.2 sensing. A number
of planar miniaturized versions of it have already been
developed,.sup.4 and variations in these are outlined below.
[0119] Planar electrodes will be defined microlithographically, as
described earlier. The Pt working electrode may be covered with a
biocompatible diffusion limiting membrane to control O.sub.2
permeability. Fine tuning the thickness of this membrane aids in
minimizing response time and maintaining sensitivity..sup.4
Layer-by-layer (LBL) growth of Nafion/Fe.sup.3+ thin films allow
for adjusting permeability of a variety of species. By adjusting
the pH, the conformation of film growth could be tailored so as to
acquire films of desired thickness. A pH of 4.5, for example,
induces surface spreading of Nafion onto the substrate, thus
ensuring a film growth consisting of surface spread and tightly
meshed polymeric chains that exhibit high tortuosity to permeation.
Moreover, the presence of Fe.sup.3+ groups prevents the potential
calcification of these films due to interactions of the negative
SO.sup.3- groups of Nafion and the physiologically present
Ca.sup.2+ions. This may be helpful to prevent in vivo degradation
of these devices. The precise localization of such films may be
performed using the well-established technique of micro-contact
printing along with LBL assembly. Polyacrylamide hydrogels will be
employed for the construction of these stamps, defined by
crosslinking them onto lithographically etched masters. The applied
force of the hydrogel on the substrate and time of contact will be
adjusted accordingly.
[0120] The fabrication of this sensor is straightforward, although
it requires considerable skill in terms of integrating it with the
other two electrochemical sensors on the same chip. Depending on
feature size, stamp micropositioning is critical. Four-degree of
movement (x, y, z and tilt) stages along with corresponding
controllers may be helpful in micropositioning.
[0121] In one embodiment, glucose sensor response is determined as
a function of temperature, pH and oxygen. As outlined above, the
interdependence of temperature, pH and oxygen content, together
with various glucose levels and film-specific construction
parameters create a multi-dimensional problem. A system to
integrate all variables into a single calibration platform would be
useful.
[0122] Standardizing all basic elements of the sensors and keeping
the number of independent variables to a minimum is an objective
after individual sensing functionality and longevity are
established. This will be followed by conducting a series of
calibrations.
TABLE-US-00001 TABLE I Typical Glucose sensor voltages Sensor
V.sub.REF V.sub.Working Type CO.sub.2 Sensor specific Sensor
specific Electrochem Ionic Sensor specific Sensor specific
Electrochem Glucose 0.7 V 1.2 V Electrochem
* * * * *