microelectronic pill

I. INTRODUCTION
THE invention of the transistor enabled the first radiotelemetry
capsules, which utilized simple circuits for
in vivo telemetric studies of the gastro-intestinal (GI) tract [1].
These units could only transmit from a single sensor channel,
and were difficult to assemble due to the use of discrete
components [2]. The measurement parameters consisted of
either temperature, pH or pressure, and the first attempts
of conducting real-time noninvasive physiological measurements
suffered from poor reliability, low sensitivity, and short
lifetimes of the devices. The first successful pH gut profiles
were achieved in 1972 [3], with subsequent improvements in
Manuscript received January 30, 2003; revised June 8, 2003. This work was
supported by the Scottish Higher Education Funding Council under Grant RDG
130. Asterisk indicates corresponding author.
*E. A. Johannessen is with the Department of Electronics and Electrical Engineering,
University of Glasgow, Rankine Building, Oakfield Avenue, Glasgow
G12 8LT, U.K. (e-mail: e.johannessen[at]elec.gla.ac.uk).
L. Wang, L. Cui, D. R. S. Cumming, and J. M. Cooper are with the Department
of Electronics and Electrical Engineering, University of Glasgow, Rankine
Building, Glasgow G12 8LT, U.K.
T. B. Tang, M. Ahmadian, A. F. Murray, and B.W. Flynn are with the School
of Engineering and Electronics, University of Edinburgh, King’s Buildings, Edinburgh
EH9 3JL, U.K.
A. Astaras and S. P. Beaumont are with the Institute for System Level Integration,
The Alba Centre, Alba Campus, Livingston EH54 7EG, U.K.
S. W. J. Reid is with the Department of Veterinary Clinical Studies, University
of Glasgow, Institute of Comparative Medicine,Veterinary School, Glasgow
G61 1QH, U.K. and also with the Department of Statistics and Modeling Science,
University of Strathclyde, Livingstone Tower, Glasgow G1 1XW, U.K..
P. S. Yam is with the Department of Veterinary Clinical Studies, University of
Glasgow, Institute of Comparative Medicine, Veterinary School, Glasgow G61
1QH, U.K.
Digital Object Identifier 10.1109/TBME.2003.820370
sensitivity and lifetime [4], [5]. Single-channel radiotelemetry
capsules have since been applied for the detection of disease
and abnormalities in the GI tract [6]–[8] where restricted access
prevents the use of traditional endoscopy [9].
Most radiotelemetry capsules utilize laboratory type sensors
such as glass pH electrodes, resistance thermometers [10], or
moving inductive coils as pressure transducers [11]. The relatively
large size of these sensors limits the functional complexity
of the pill for a given size of capsule. Adapting existing
semiconductor fabrication technologies to sensor development
[12]–[17] has enabled the production of highly functional units
for data collection, while the exploitation of integrated circuitry
for sensor control, signal conditioning, and wireless transmission
[18], [19] has extended the concept of single-channel radiotelemetry
to remote distributed sensing from microelectronic
pills.
Our current research on sensor integration and onboard data
processing has, therefore, focused on the development of microsystems
capable of performing simultaneous multiparameter
physiological analysis. The technology has a range of applications
in the detection of disease and abnormalities in medical
research. The overall aim has been to deliver enhanced functionality,
reduced size and power consumption, through systemlevel
integration on a common integrated circuit platform comprising
sensors, analog and digital signal processing, and signal
transmission.
In this paper, we present a novel analytical microsystem
which incorporates a four-channel microsensor array for
real-time determination of temperature, pH, conductivity and
oxygen. The sensors were fabricated using electron beam and
photolithographic pattern integration, and were controlled
by an application specific integrated circuit (ASIC), which
sampled the data with 10-bit resolution prior to communication
off chip as a single interleaved data stream. An integrated radio
transmitter sends the signal to a local receiver (base station),
prior to data acquisition on a computer. Real-time wireless data
transmission is presented from a model in vitro experimental
setup, for the first time.
Details of the sensors are provided in more detail later, but
included: a silicon diode [20] to measure the body core temperature,
while also compensating for temperature induced signal
changes in the other sensors; an ion-selective field effect transistor,
ISFET, [21] to measure pH; a pair of direct contact gold
electrodes to measure conductivity; and a three-electrode electrochemical
cell [22], to detect the level of dissolved oxygen
in solution. All of these measurements will, in the future, be
used to perform in vivo physiological analysis of the GI-tract.
0018-9294/04$20.00 © 2004 IEEE
526 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 51, NO. 3, MARCH 2004
For example, temperature sensors will not only be used to measure
changes in the body core temperature, but may also identify
local changes associated with tissue inflammation and ulcers.
Likewise, the pH sensor may be used for the determination
of the presence of pathological conditions associated with
abnormal pH levels, particularly those associated with pancreatic
disease and hypertension, inflammatory bowel disease, the
activity of fermenting bacteria, the level of acid excretion, reflux
to the oesophagus, and the effect of GI specific drugs on
target organs. The conductivity sensor will be used to monitor
the contents of the GI tract by measuring water and salt absorption,
bile secretion and the breakdown of organic components
into charged colloids. Finally, the oxygen sensor will measure
the oxygen gradient from the proximal to the distal GI tract. This
will, in future enable a variety of syndromes to be investigated
including the growth of aerobic bacteria or bacterial infection
concomitant with low oxygen tension [23], as well as the role of
oxygen in the formation of radicals causing cellular injury and
pathophysiological conditions (inflammation and gastric ulceration).
The implementation of a generic oxygen sensor will also
enable the development of first generation enzyme linked amperometric
biosensors, thus greatly extending the range of future
applications to include, e.g., glucose and lactate sensing, as well
as immunosensing protocols.
II. MICROELECTRONIC PILL DESIGN AND FABRICATION
A. Sensors
The sensors were fabricated on two silicon chips located at the
front end of the capsule. Chip 1 [Fig. 1(a), ©, (e)] comprises
the silicon diode temperature sensor, the pH ISFET sensor and
a two electrode conductivity sensor. Chip 2 [Fig. 1(b), (d), (f)]
comprises the oxygen sensor and an optional nickel-chromium
(NiCr) resistance thermometer. The silicon platform of Chip 1
was based on a research product from Ecole Superieure D’Ingenieurs
en Electrotechnique et Electronique (ESIEE, France)
with predefined n-channels in the p-type bulk silicon forming
the basis for the diode and the ISFET. A total of 542 of such devices
were batch fabricated onto a single 4-in wafer. In contrast,
Chip 2 was batch fabricated as a 9 9 array on a 380- m-thick
single crystalline silicon wafer with lattice orientation,
precoated with 300 nm , silicon nitride, (Edinburgh
Microfabrication Facility, U.K.). One wafer yielded 80,
mm sensors (the center of the wafer was used for alignment
markers).
1) Sensor Chip 1: An array of 4 2 combined temperature
and pH sensor platforms were cut from the wafer and attached
on to a 100- m-thick glass cover slip using S1818 photoresist
(Microposit, U.K.) cured on a hotplate. The cover slip
acted as temporary carrier to assist handling of the device during
the first level of lithography (Level 1) when the electric connection
tracks, the electrodes and the bonding pads were defined.
The pattern was defined in S1818 resist by photolithography
prior to thermal evaporation of 200 nm gold (including an adhesion
layer of 15 nm titanium and 15 nm palladium). An additional
layer of gold (40 nm) was sputtered to improve the adhesion
of the electroplated silver used in the reference electrode
(see below). Liftoff in acetone detached the chip array from the
cover slip. Individual sensors were then diced prior to their re-attachment
in pairs on a 100- m-thick cover slip by epoxy resin
Fig. 1. The microelectronic sensors: (a) schematic diagram of Chip 1,
measuring 4:75  5 mm , comprising the pH (ISFET) sensor (1), the
5  10 mm dual electrode conductivity sensor (3) and the silicon
diode temperature sensor (4); (b) schematic diagram of Chip 2, measuring
5  5 mm , comprising the electrochemical oxygen sensor (2) and a NiCr
resistance thermometer (5). Once integrated in the pill, the area exposed
to the external environment is illustrated by the 3-mm-diameter circle; ©
photomicrograph of sensor Chip 1 and (d) sensor Chip 2. The bonding
pads (6), which provide electrical contact to the external electronic control
circuit, are shown; (e) close up of the pH sensor consisting of the integrated
3  10 mm AgjAgCl reference electrode (7), a 500-m-diameter and
50–m-deep, 10-nL, electrolyte chamber (8) defined in polyimide, and the
15  600 m floating gate (9) of the ISFET sensor; (f) the oxygen sensor
is likewise embedded in an electrolyte chamber (8). The three-electrode
electrochemical cell comprises the 1  10 mm counter electrode (10), a
microelectrode array of 57  10 m diameter (4:5  10 mm ) working
electrodes (11) defined in 500-nm-thick PECVD Si N , and an integrated
1:5  10 mm AgjAgCl reference electrode (12).
[Fig. 1©]. The left-hand-side (LHS) unit comprised the diode,
while the right-hand-side (RHS) unit comprised the ISFET. The
m (L W) floating gate of the ISFET was precovered
with a 50-nm-thick proton sensitive layer of for pH
detection [24].
Photocurable polyimide (Arch Chemicals n.v., Belgium) defined
the 10-nL electrolyte chamber for the pH sensor (above
the gate) and the open reservoir above the conductivity sensor
(Level 2).
The silver chloride reference electrode mm was
fabricated during Levels 3 to 5, inclusive. The glass cover slip,
to which the chips were attached, was cut down to the size of
the mm footprint (still acting as a supporting base)
prior to attachment on a custom-made chip carrier used for electroplating.
Silver (5 m) was deposited on the gold electrode
JOHANNESSEN et al.: MULTICHANNEL SENSORS FOR REMOTE BIOMEDICAL MEASUREMENTS IN A MICROSYSTEMS FORMAT 527
defined at by chronopotentiometry ( 300 nA, 600 s) after removing
residual polyimide in an barrel asher (Electrotech,
U.K.) for 2 min. The electroplating solution consisted of 0.2 M
, 3MKI and 0.5M . Changing the electrolyte
solution to 0.1 M KCl at Level 4 allowed for the electroplated
silver to be oxidized to AgCl by chronopoteniometry (300 nA,
300 s). The chip was then removed from the chip carrier prior to
injection of the internal 1 M KCl reference electrolyte required
for the Ag AgCl reference electrode (Level 5). The electrolyte
was retained in a 0.2% gel matrix of calcium alginate [25].
The chip was finally clamped by a 1-mm-thick stainless-steel
clamp separated by a 0.8- m-thick sheet of Viton fluoroelastomer
(James Walker, U.K.). The rubber sheet provided a uniform
pressure distribution in addition to forming a seal between
the sensors and capsule.
2) Sensor Chip 2: The level 1 pattern (electric tracks,
bonding pads, and electrodes) was defined in 0.9 m UV3
resist (Shipley, U.K.) by electron beam lithography. A layer of
200 nm gold (including an adhesion layer of 15 nm titanium
and 15 nm palladium) was deposited by thermal evaporation.
The fabrication process was repeated (Level 2) to define the
5- m-wide and 11-mm-long NiCr resistance thermometer
made from a 100-nm-thick layer of NiCr (30- resistance).
Level 3 defined the 500-nm-thick layer of thermal evaporated
silver used to fabricate the reference electrode. An additional
sacrificial layer of titanium (20 nm) protected the silver from
oxidation in subsequent fabrication levels. The surface area
of the reference electrode was mm , whereas the
counter electrode made of gold had an area of mm .
Level 4 defined the microelectrode array of the working electrode,
comprising 57 circular gold electrodes, each 10 m in
diameter, with an interelectrode spacing of 25 m and a combined
area of mm . Such an array promotes electrode
polarization and reduces response time by enhancing transport
to the electrode surface [26]. The whole wafer was covered with
500 nm plasma-enhanced chemical vapor deposited (PECVD)
. The pads, counter, reference, and the microelectrode
array of the working electrode was exposed using an etching
mask of S1818 photoresist prior to dry etching with C . The
chips were then diced from the wafer and attached to separate
100- m-thick cover slips by epoxy resin to assist handling. The
electrolyte chamber was defined in 50- m-thick polyimide at
Level 5. Residual polyimide was removed in an barrel asher
(2 min), prior to removal of the sacrificial titanium layer at Level
6 in a diluted HF solution (HF to RO water, 1:26) for 15 s. The
short exposure to HF prevented damage to the PECVD
layer.
Thermally evaporated silver was oxidized to Ag AgCl (50%
of film thickness) by chronopotentiometry (120 nA, 300 s) at
Level 7 in the presence of KCl, prior to injection of the internal
reference electrolyte at Level 8. A mm sheet of oxygen
permeable teflon was cut out from a 12.5- m-thick film and attached
to the chip at Level 9 with epoxy resin prior to immobilization
by the aid of a stainless steel clamp.