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INTRODUCTION
The invention of the transistor enabled the first radio telemetry capsules, which utilized simple circuits for in vivo telemetric studies of the gastro-intestinal tract. These units could only transmit from a single sensor channel, and were difficult to assemble due to the use of discrete components. The measurement parameters consisted of 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, with subsequent improvements in sensitivity and lifetime. Single-channel radiotelemetrycapsules have since been applied for the detection of disease and abnormalities in the GI tract where restricted access prevents the use of traditional endoscopy.
Most radio telemetry capsules utilize laboratory type sensors such as glass pH electrodes, resistance thermometers, or moving inductive coils as pressure transducers. 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 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,and has extended the concept of single-channel radio telemetry 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 multiparameterphysiological 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 system-level integration on a common integrated circuit platform comprising sensors, analog and digital signal processing, and signal transmission.
In this report, we present a novel analytical micro systemwhich incorporates a four-channel micro sensor array for real-time determination of temperature, pH, conductivity andoxygen. The sensors were fabricated using electron beam and photolithographic pattern integration, and were controlledby an application specific integrated circuit (ASIC), whichsampled the data with 10-bit resolution prior to communicationoff chip as a single interleaved data stream. An integrated radiotransmitter sends the signal to a local receiver (base station), prior to data acquisition on a computer. Real-time wireless datatransmission is presented from a modelin vitro experimentalsetup, for the first time.
Details of the sensors are provided in more detail later, butincluded: a silicon diode to measure the body core temperature, while also compensating for temperature induced signalchanges in the other sensors; an ion-selective field effect transistor, ISFET, to measure pH; a pair of direct contact goldelectrodes to measure conductivity; and a three-electrode electrochemical cell, to detect the level of dissolved oxygenin solution. All of these measurements will, in the future, beused to performin vivo physiological analysis of the GI-tract.
For example, temperature sensors will not only be used to mea-sure 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 withabnormal pH levels, particularly those associated with pancreatic disease and hypertension, inflammatory bowel disease, the activity of fermenting bacteria, the level of acid excretion, re-flux 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, as well as the role of oxygen in the formation of radicals causing cellular injury andpath physiological conditions (inflammation and gastric ulceration). The implementation of a generic oxygen sensor will alsoenable 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 wellas immune sensing protocols.
MICROELECTRONICPILLDESIGN AND FABRICATION
2.1.ISFET
This new line of pH meters and probes, based on ISFET (Ion Sensitive Field Effect Transistor) sensor technology, includes four pH meters and 10 pH probes. The pH meters are designed for ease-of-use and feature an interactive graphics LCD display with on-board Help and Auto-Read functions. All meters constantly monitor and display probe status and an estimation of its remaining life. The advanced meters have real-time clocks for time/date stamping, calibration alerts and high/low pH alarms.Titan Bench top pH meters operate on AC or battery power and offer a host of sophisticated features, including programmable user alarms and data logging. Argus Portable meters are rugged, waterproof and operate on a long-life rechargeable battery. Each meter is available in simple or advanced versions and is supported by a variety of probes covering almost every application. The portable Argususes an inductive (contact-less) battery charging system and IR data transfer eliminating the need for battery replacement or open contact points. This design ensures a completely watertight (IP67) rating.
Three new series of ISFET probes include the Red-Line general purpose series for routine applications, the Hot-Line series for testing to 105°C and in aggressive samples, and the Stream-Line series that are temperature and chemically resistant, and employ a flow-type reference junction to maximize performance in difficult samples.
2.2. pH value
pH is a measure of the acidity or basicity of an aqueous solution. Pure water is said to be neutral, with a pH close to 7.0 at 25 °C (77 °F). Solutions with a pH less than 7 are said to be acidic and solutions with a pH greater than 7 are basic or alkaline.pH measurements are in important in medicine, biology, chemistry, food science, environmental science, oceanography, civil engineering and many other applications.
In a solution pH approximates but is not equal to p[H], the negative logarithm (base 10) of the molarconcentration of dissolved hydronium ions (H3O+); a low pH indicates a high concentration of hydronium ions, while a high pH indicates a low concentration. Crudely, this negative of the logarithm matches the number of places behind the decimal point, so for example 0.1 molar hydrochloric acid should be near pH 1 and 0.0001 molar HCl should be near pH 4 (the base 10 logarithms of 0.1 and 0.0001 being −1, and −4, respectively). Pure (de-ionized) water is neutral, and can be considered either a very weak acid or a very weak base (center of the 0 to 14 pH scale), giving it a pH of 7 (at 25 °C (77 °F)), or 0.0000001 M H+.[1] For an aqueous solution to have a higher pH, a base must be dissolved in it, which binds away many of these rare hydrogen ions. Hydrogen ions in water can be written simply as H+ or as hydronium (H3O+) or higher species (e.g. H9O4+) to account forsolvation, but all describe the same entity. Most of the Earth's freshwater surface bodies are slightly acidic due to the abundance and absorption of carbon dioxide;[2] in fact, for millennia in the past most fresh water bodies have long existed at a slightly acidic pH level.
2.3. Sensors
The sensors were fabricated on two silicon chips located at the front end of the capsule. Chip 1 comprises the silicon diode temperature sensor, the pH ISFET sensor and a two electrode conductivity sensor. Chip 2 comprises the oxygen sensor and an optional nickel-chromium (NiCr) resistance thermometer. The silicon platform of Chip 1was based on a research product from EcoleSuperieureD’In-genieurs en Electro technique et Electroniquewith predefined n-channels in the p-type bulk silicon forming the basis for the diode and the ISFET. A total of 542 of such de-vices were batch fabricated onto a single 4-in wafer. In contrast, Chip 2was batch fabricated as a 9X9 array on a 380-m-thicksingle crystalline 3nSilicon wafer with <100>lattice orientation, precoated with 300 nm Si3N4, silicon nitride. One wafer yielded 80,5X5 mm2 sensors (the center of the wafer was used for alignment markers)
2.3.1.Sensor Chip 1
An array of 4X2 combined temperature and pH sensor platforms were cut from the wafer and attached on to a 100-m-thick glass cover slip using S1818 photo resist 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. 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. The left-hand-side (LHS) unit comprised the diode, while the right-hand-side (RHS) unit comprised the ISFET. The15X600 m (LXW) floating gate of the ISFET was precovered with a 50-nm-thick proton sensitive layer of Si3N4 for pH detection. Photo curable polyimide de-fined the 10-nL electrolyte chamber for the pH sensor (above the gate) and the open reservoir above the conductivity sensor
Photolithography
Photolithography (or "optical lithography") is a process used in microfabrication to selectively remove parts of a thin film or the bulk of a substrate. It uses light to transfer a geometric pattern from a photo mask to a light-sensitive chemical "photoresist", or simply "resist," on the substrate. A series of chemical treatments then either engraves the exposure pattern into, or enables deposition of a new material in the desired pattern upon, the material underneath the photo resist. In complex integrated circuits, for example a modern CMOS, a wafer will go through the photolithographic cycle up to 50 times.
Photolithography shares some fundamental principles with photography in that the pattern in the etching resist is created by exposing it to light, either directly (without using a mask) or with a projected image using an optical mask. This procedure is comparable to a high precision version of the method used to make printed circuit boards. Subsequent stages in the process have more in common with etching than to lithographic printing. It is used because it can create extremely small patterns (down to a few tens of nanometers in size), it affords exact control over the shape and size of the objects it creates, and because it can create patterns over an entire surface cost-effectively. Its main disadvantages are that it requires a flat substrate to start with, it is not very effective at creating shapes that are not flat, and it can require extremely clean operating conditions.
Sensor Chip 2
The level 1pattern (electric tracks,bonding pads, and electrodes) was defined in 0.m UV3resist by electron beam lithography. A layer of200 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-kΩ resistance).Level 3defined 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 was1.5x10-2mm2, whereas thecounter electrode made of gold had an area of1.1x10-1mm2.Level 4defined the microelectrode array of the working electrode, comprising 57 circular gold electrodes, each 10 m in diameter, with an interelectrode spacing of 25m and a combined area of4.5x10-2mm2. Such an array promotes electrode polarization and reduces response time by enhancing transportto the electrode surface. The whole wafer was covered with500 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 photo resist prior to dry etching with C2f6. 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 atLevel 5. Residual polyimide was removed in an O2barrel asher(2 min), prior to removal of the sacrificial titanium layer at Level6in a diluted HF solution (HF to RO water, 1:26) for 15 s. Theshort exposure to HF prevented damage to the PECVD Si3N4Layer.Thermally evaporated silver was oxidized to Ag|AgCl (50%of film thickness) by chronopotentiometry (120 nA, 300 s) atLevel 7in the presence of KCl, prior to injection of the internalreference electrolyte at Level 8. A 5X5mm2sheet of oxygen permeableteflon 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.
2.3.2.1. Plasma-enhanced chemical vapor deposition (PECVD)
It is a process used to deposit thin films from a gas state (vapor) to a solid state on a substrate. Chemical reactions are involved in the process, which occur after creation of a plasma of the reacting gases. The plasma is generally created by RF (AC) frequency or DC discharge between two electrodes, the space between which is filled with the reacting gases.
2.3.3. Control Chip
The ASIC was a control unit that connected together the external components of the micro system. It was fabricated as a 22.5 mm2 silicon die using a 3-V, 2-poly, 3-metal 0.6-m CMOS process by Austria Microsystems (AMS) via the Euro-practice initiative. It is a novel mixed signal design that contains an analog signal conditioning module operating the sensors, an 10-bit analog-to-digital (ADC) and digital-to-analog(DAC) converters, and a digital data processing module. AnRCrelaxation oscillator (OSC) provides the clock signal.
The analog module was based on the AMS OP05B operational amplifier, which offered a combination of both a power-saving scheme (sleep mode) and a compact integrated circuitdesign. The temperature circuitry biased the diode at constantcurrent, so that a change in temperature would reflect a corresponding change in the diode voltage. The pH ISFET sensor was biased as a simple source and drain follower at constant current with the drainsource voltage changing with the threshold voltage and pH. The conductivity circuit operated at direct current measuring the resistance across the electrode pair as an inverse function of solution conductivity. An incorporated potentiostat circuit operated the amperometric oxygen sensor with a10-bit DAC controlling the working electrode potential with respect to the
Radio Transmitter
The radio transmitter was assembled prior to integration in thecapsule using discrete surface mount components on a single-sided printed circuit board (PCB). The footprint of the standardtransmitter measured 8X5X3mm including the integratedcoil (magnetic) antenna. It was designed to operate at a trans-mission frequency of 40.01 MHz at 20oC generating a signalof 10 kHz bandwidth. A second crystal stabilized transmitterwas also used. This second unit was similar to the free running standard transmitter, apart from having a larger footprint of10X5X3 mm, and a transmission frequency limited to 20.08MHz at 20oC, due to the crystal used. Pills incorporating thestandard transmitter were denotedType I, whereas the pills incorporating the crystal stabilized unit were denotedType II. Thetransmission range was measured as being 1 meter and the modulation scheme frequency shift keying (FSK), with a data rate of 1kbs-1.
1. Size of transmitter = 8 × 5 × 3 mm
2. Modulation Scheme = Frequency Shift Keying (FSK)
3. Data Transfer Rate = 1 kbps
4. Frequency = 40.01 MHz at 20 °C
5. Bandwidth of the signal generated 10 KHz
6. It consumes 6.8 mW power at 2.2 mA of current.
2.3.5. Capsule
The microelectronic pill consisted of a machined biocompatible (noncytotoxic), chemically resistant polyetherterketone (PEEK) capsule and a PCB chip carrier acting as a common platform for attachment of the sensors, ASIC, transmitter and the batteries (Fig.2.3). The fabricated sensorswere each attached by wire bonding to a custom made chip carrier made from a 10-pin, 0.5-mm pitch polyimide ribbon connector. The ribbon connector was, in turn, connected to an industrial standard 10-pin flat cable plug (FCP) socket attached to the PCB chip carrier of the microelectronic pill, to facilitate rapid replacement of the sensors whenrequired. The PCB chip carrier was made from two standard1.6-mm-thick fiber glass boards attached back to back by epoxyresin which maximized the distance between the two sensorchips. The sensor chips were connected to both sides of the PCBby separate FCP sockets, with sensorChip 1 facing the top face, withChip 2 facing down. Thus, the oxygen sensor onChip 2 hadto be connected to the top face by three 200-m copper leads soldered on to the board. The transmitter was integrated in thePCB which also incorporated the power supply rails, the connection points to the sensors, as well as the transmitter and theASIC and the supporting slots for the capsule in which the chipcarrier was located.
The ASIC was attached with double-sided copper conductingtape (Agar Scientific, U.K.) prior to wirebonding to the powersupply rails, the sensor inputs, and the transmitter (a processwhich entailed the connection of 64 bonding pads). The unitwas powered by two standard 1.55-V SR44 silver oxide(Ag2O)cells with a capacity of 175 mAh. The batteries were serial connected and attached to a custom made 3-pin, 1.27-mm pitch plugby electrical conducting epoxy. The connection to the matching socket on the PCB carrier pro-vided a three point power supply to the circuit comprising a negative supply rail (-1.55 V), virtual ground (0 V), and a positivesupply rail (1.55 V). The battery pack was easily replaced duringthe experimental procedures.
The capsule was machined as two separate screw-fitting compartments. The PCB chip carrier was attached to the front section of the capsule. The sensor chips were exposed tothe ambient environment through access ports and were sealedby two sets of stainless steel clamps incorporating a 0.8-m-thick sheet of Viton fluoroelastomer seal. A 3-mm-diameter access channel in the center of each of the steel clamps (incl. the seal), exposed the sensing regions of the chips. The rear section of the capsule was attached to the front section by a 13-mmscrew connection incorporating a Viton rubber O-ring. The seals rendered the capsule water proof, as well as making it easy to maintain (e.g., during sensor and battery replacement). The complete prototype was 16X55 mm and weighted 13.5 g including the batteries. A smaller pill suitable for physiologicalin vivo trials (10X30 mm) is currently being developed from the prototype.
MATERIAL ANDMETHODS
3.1 Fabrication
Thermal evaporation of silver generates a dense metal layer, with characteristics closer to bulk metal compared to porouselectroplated silver. Although electroplating allow for a thickerlayer of silver to be deposited, the lifetime of a Ag|AgClreference electrode made from 500-nm-thick thermally evaporated silver was compared to a Ag|AgCl electrode made froma 5-m-thick electroplated layer. The results clearly demonstrated the potential of utilizing thermally evaporated silver inAg|AgCl electrodes to extend lifetime by more than 100%.However, a protective layer of 20 nm titanium was required toprevent oxidation of the silver in subsequent fabrication levels,and which had to be removed by immersion in a HF solution. Since HF also attacks, this procedure could not be used inChip 1 to avoid damage to the thin 50–nm layer ofdefining the pH sensitive membrane of the ISFET. In contrast,the 500-nm-thick PECVDdefining the microelectrodearray of the oxygen sensor, was tolerant to HF exposure.
The sensor lifetime was further extended through using athree-electrode electrochemical cell for the oxygen sensor, infavor of a two-electrode device. A two electrode unit utilizes thereference electrode as a combined counter and reference unit tochannel all the current from the reduction of oxygen. However, a three-electrode electrochemical cell bypasses the current flowfrom the working electrode by incorporating a separate counterelectrode subjecting the reference electrode only to the bias cur-rent of the input transistor stage of the operational amplifier, towhich the sensor is connected. Thus, the overall current channeled through the reference was reduced by at least three ordersof magnitude. This effect is important as it enables a reductionin the electrode area and improved long-term stability.
3.2. General Experimental Setup
All the devices were powered by batteries in order to demonstrate the concept of utilizing the microelectronic pill in re-mote locations (extending the range of applications fromin vivosensing to environmental or industrial monitoring). The pill wassubmerged in a 250-mL glass bottle located within a 2000-Mlbeaker to allow for a rapid change of pH and temperature of thesolution. A scanning receiver captured the wireless radio transmitted signal from themicroelectronic pill by using a coil antenna wrapped around the2000-mL polypropylene beaker in which the pill was located. A portable Pentium III computer controlled the data acquisition unit (National Instruments, Austin, TX) which digitally acquired analog data from the scanning receiver prior to recording it on the computer.
The solution volume used in all experiments was 250 mL.The beaker, pill, glass bottle, and antenna were located withina 25X25 cm container of polystyrene, reducing temperaturefluctuations from the ambient environment (as might be expected within the GI tract) and as required to maintain a stabletransmission frequency. The data was acquired using Lab View and processed using a MATLAB routine.
SensorCharacterization
The lifetime of the incorporated Ag|AgCl reference electrodes used in the pH and oxygen sensors was measured withan applied current of 1 pA immersed in a 1.0 M KCl electrolytesolution. The current reflects the bias input current of the operational amplifier in the analog sensor control circuitry to whichthe electrodes were connected.
The temperature sensor was calibrated with the pill sub-merged in reverse osmosis (RO) water at different temperatures. The average temperature distribution over 10 min was recordedfor each measurement, represented as 9.10C, 21.20C, 33.50C, and 47.90C. The system was allowed to temperatureequilibrate for 5 min prior to data acquisition. The controlreadings were performed with a thin wire K-type thermocouple (Radio Spares, U.K.). The signal from the temperature sensorwas investigated with respect to supply voltage potential, dueto the temperature circuitry being referenced to the negativesupply rail. Temperature compensated readings (normalized to230C) were recorded at a supply voltage potential of 3.123, 3.094, 3.071, and 2.983 mV using a direct communication link.Bench testing of the temperature sensor from 00C to 700C wasalso performed to investigate the linear response characteristicsof the temperature sensor.
The pH sensor of the microelectronic pill was calibrated instandard pH buffers [28] of pH 2, 4, 7, 9, and 13, which reflectedthe dynamic range of the sensor. The calibration was performedat room temperature (230C) over a period of 10 min, with the pill being washed in RO water between each step. A standardlab pH electrode was used as a reference to monitor the pH ofthe solutions. The pH channel of the pillwas allowed to equilibrate for 5 min prior to starting the data acquisition. Each measurement was performed twice. Bench testmeasurements from pH 1 to 13 were also performed using anidentical control circuit to the ASIC.
The oxygen sensor was bench tested with a standard laboratory potentiostat, over its dynamic range in phosphate buffered saline (PBS) usinga direct communication link at 23C. Cyclic voltammetry witha sweep potential from 0.1 to 0.45 V (versus Ag|AgCl) was per-formed in 1-mM ferroscene-monocarboxylic acid (FMCA) as amodel redox compound, to test the performance of the micro-electrode array. A three-point calibration routine was performedat oxygen concentrations of 0 mg L-1(PBS saturated with 2 M), 4 mg L-1(PBS titration with 2 M) and 8.2mg L (oxygen saturated PBS solution). The solution saturatedwith dissolved oxygen was equilibrated overnight prior to use.The dissolved oxygen was monitored using a standard Clarkelectrode .The reduction potential of water was assessed in oxygen depleted PBS, to avoid interference from oxygen, at the same time assessingthe lower potential limit that could be used for maximizing theefficiency of the sensor. The voltage was then fixed above thisreduction potential to assess the dynamic behavior of the sensorupon injection of saturated Na2SO4in oxygen saturated PBS.
3.4. Transmission
The pill’s transmission frequency was measured with respectto changes in temperature. TheType I pill (without crystal) wassubmerged in RO water at temperatures of 10C, 110C,230 Cand 490C, whereas theType II pill (with crystal) was submergedin temperatures of 20C, 250C, and 450C. The change in frequency was measured with the scanning receiver, and the resultsused to assess the advantage of crystals stabilized units at thecost of a larger physical size of the transmitter.
3.5. Dynamic Measurements
Dynamic pH measurements were performed with the pill sub-merged in a PBS solution at 230C. The pH was changed fromthe initial value of 7.3 by the titration of 0.1 M H2SO4and0.1 M NaOH, respectively. Subsequently, the pH was changedfrom pH 7.3 to pH 5.5 (after 5 min), pH 3.4 (after 8 min) topH 9.9 (after 14 min) and back to pH 7.7 (after 21 min). Astandard (bench-top) pH electrode monitored the pH of the solution. The solutions were sampled after the pH change, andmeasured outside the experimental system to prevent electronicnoise injection from the pH electrode. The temperature channelwas recorded simultaneously.
3.6. Sensor and Signal Drift
Long term static pH and temperature measurements were per-formed to assess signal drift and sensor lifetime in physiological electrolyte (0.9% saline) solutions. A temperature of 36.50Cwas achieved using a water bath, with the assay solutions continuously stirred and re-circulated using a peristaltic pump. Thesensors were transferred from solutions of pH 4 to pH 7, within2 h of commencing the experiment, and from pH 7 to pH 10.5,after 4 h. The total duration of the experiment was 6 h. Each experiment was repeated twice.
IMPORTANT OBSERVATIONS
The power consumption of the microelectronic pill with the transmitter, ASIC and the sensors connected was calculated to12.1 mW, corresponding to the measured current consumptionof 3.9 mA at 3.1-V supply voltage. The ASIC and sensors consumed 5.3 mW, corresponding to 1.7 mA of current, whereasthe free running radio transmitter (Type I) consumed 6.8 mW(corresponding to 2.2 mA of current) with the crystal stabilized unit (Type II) consuming 2.1 mA. Two SR44 Ag2O batteries used provided an operating time of more than 40 h for themicro system.
4.1. Temperature Channel Performance
The linear sensitivity was measured over a temperature range from 00C to 700C and found to be 15.4 mV0C-1. This amplified signal response was from the analog circuit, which waslater implemented in the ASIC. The sensor, once integrated in the pill, gave a linear regression of 11.9 bits-C with a resolution limited by the noise band of 0.40C. The diodewas forward biased with a constant current () with then-channel clamped to ground, while the p-channel was floating.Since the bias current supply circuit was clamped to the negative voltage rail, any change in the supply voltage potential would cause the temperature channel to drift. Thus, bench test measurements conducted on the temperature sensor revealed thatthe output signal changed by 1.45 mV per mV change in supply voltage expressed in millivolts, corresponding to a drift of-21mV h-1in the pill from a supply voltage change of-14.5mV-1.
4.2.pH Channel Performance
The linear characteristics from pH 1 to 13 corresponded toa sensitivity of-41.7mVpH-1unit at 230C, which is in agreement with literature values although the responsewas lower than the Nernstian characteristics found in standard glass pH electrodes (-59.2mVunit). The pH ISFETsensor operated in a constant current mode (), withthe drain voltage clamped to the positive supply rail, and thesource voltage floating with the gate potential. The Ag|AgClreference electrode, representing the potential in which thefloating gate was referred to, was connected to ground. Thesensor performance, once integrated in the pill [Fig. 3(a)], corresponded to 14.85 bitspH-1which gave a resolution of 0.07pH per datapoint. The calibrated response from the pH sensorconformed to a linear regression, although the sensor exhibited a larger responsivityin alkaline solutions. The sensor lifetime of 20 h was limitedby the Ag|AgCl reference electrode made from electroplatedsilver. The pH sensor exhibited a signal drift of -6mVh-1(0.14 pH), of which -2.5 mVh-1 was eliminated to be due to the dissolution of AgCl from the reference electrode. The temperature sensitivity of the pH-sensor was measured as 16.8mVC-1. Changing the Ph of the solution at 400C from ph 6.8 to pH 2.3 and pH 11.6 demonstrated that the two channels were completely independent of each other and that there was no signal interference from the temperature channel.
4.3.Oxygen sensor performance
The electrodes were first characterized using the model redox compound FMCA, showing that the oxygen sensor behaved with classic microelectrode characteristics. The reduction potential of water was subsequently measured at -800 mV (versus the integrated Ag|AgCl) by recording the steady-state current in oxygen-depleted PBS, thereby excluding any interfering species.
In order to calibrate the sensor, a three point calibration was performed (at saturated oxygen, and with oxygen removed by the injection of Na2SO3 to a final concentration of 1 M). the steady state signal from the oxygen saturated solution was recorded at a constant working electrode potential of -700 mV(versus Ag|AgCl).which was below the reduction potential for water. This generated a full-scale signal of 65 nA corresponding 8.2 mg O2L-1. The injection of Na2S03 into the PBS after 90 s provided the zero point calibration. This fall in the reduction current provided corroborative evidence that dissolve oxygen was being recorded, by returning the signal back to the base line level once all available oxygen was consumed. A third, intermediate point was generated through the addition of 0.01 M Na2SO3. The resulting calibration graph form a linear regression expressed in nanoamperes. The sensitivity of the sensor was 7.9 nA mg-1 O2,with the resolution of 0.4mg L-1 limited by noise or background drift. The lifetime of the integrated Ag|AgCl reference electrode, made from thermal evaporated silver, was found to be to 45 h,with an averagevoltage drift of -1.3 mVh-1 due to he dissolution of the AgCl during operation. Both measurements of FMCA and oxygen redox behavior indicated a stable Ag|AgCl reference.
Conductivity sensor performance
The prototype circuit exhibited a logarithmic performance from 0.05 to 10 ms cm-1 which conformed to a first-order regression analysis expressed in millivolts. The sensor saturated at conductivities above 10 ms cm-1 due to the capacitive effect of the electric double layer, a phenomena commonly observed in conductimetric sensor systems.
4.5.Control chip
The background noise from the ASIC corresponded to a constant level of 3-Mv peak-to-peak, which is equivalent to one least significant bit (LSB) of the ADC. Since the second LSB were required to provide an adequate noise margin, the 10-bit ADC was anticipated to have an effective resolution of 8 bits.
4.6.Transmission frequency
Frequency stabilized units were essential to prevent the transmission drifting out of range, particularly if the pill was subject to a temperature change during operation. The standard type 1 transmitter exhibited a negative linear frequency change from 39.17 MHz at 100C to 38.98 MHz at 490C, corresponding to -4 kHz0C-1 expressed in hertz. The narrow signal bandwidth of 10 kHz gave a temperature tolerance of only -+1.30C before the signal is lost. In contrast, the type 2 transmitter exhibited a positive linear frequency change from 20.07 MHz at 20C to 20.11 MHz at 400C, corresponding to 0.9 Khz0C-1. Considering the identical signal bandwidth of 10 KHz, the temperature tolerance was increased to -+5.50C. The transmitter’s signal magnitude was not affected with the pill immersed in the different electrolyte solutions or RO water, compared to the pill surrounded by air only. Tests were also conducted with the pill immersed in the large polypropylene beaker filled with 2000 mL of PBS without the signal quality being compromised. The electromagnetic noise baseline was measured to 78 dB of S/N in the 20 MHz band of the crystal stabilized transmitter.
4.7.Dual channel wireless signal transmission
Dual channel wireless signal transmission was recorded from both the pH and temperature channels at 230c, with the pill immersed in a PBS solution of changing pH. The calibration graphs for the temperature and pH channel were used to convert the digital units from the MATLAB calculated routine to the corresponding temperature and pH values.
The signal from the pH channel exhibited an initial offset of 0.2 pH above the real value at pH 7.3. In practice, the pH sensor was found to exhibit a positive pH offset as the solution became more acidic, and a negative pH offset as the solution became more alkaline. The temperature channel was unaffected by the pH change, confirming the absence of crosstalk between the two channels.
DISCUSSION
All of the components of the sensors and the capsule, exposedto the local environment, had to be able to resist the corrosiveenvironment in the digestive tract, and at the same time be non-toxic (biocompatible) to the organism. If toxic materials wereused (such as in batteries and the Ag|AgCl reference electrode), care would need to be taken to prevent leakage from the microsystem and into the surrounding environment.
5.1. Sensor Performance
The temperature circuit was sensitive to the supply voltage. The n-channel of the silicon diode was clamped to ground, whereas the bias current supply circuit was clamped to the negative supply rail. Thus, an increase of 7.25 mV h(froma total 14.5 mV hfrom the positive and negative supply rail) would reduce the bias current by 0.5%, resulting in adiode voltage change ofmV h[20]. A potential dividercircuit clamped between ground and the positive supply railwas used to create an offset signal prior to the amplificationstage. The change in offset signal corresponds tomVh, resulting in a total signal change ofmV hprior toamplification with a gain of 6.06, resulting in a total change ofmV h. The theoretical calculation conforms to within40% of the experimental result, which can be explained byreal circuit device tolerances (such as supply voltage effect onthe operational amplifiers) which deviates from the theoreticalpredictions.
The pH channel recordings from the pill (Fig. 4.1) deviatedfrom the true value measured with the glass pH electrode, bytransmitting pHresponsivity below the calibrated value. Inacidic solutions, this resulted in a pH response slightly abovethe true value, whereas the response in alkaline solutions wasbelow the true value. In neutral solutions, the pH channel exhibited an offset of 0.2 units above the real value. The resultsof the long-term measurements conducted in Fig. 4.2 suggestedthat the recorded values would match the real pH of the solution if left to equilibrate for 2 h. Thus, the combined effect ofcalibration offset and short equilibration time to a changing pH, could explain the signal offset between the measured and realpH presented in Fig. 6. The discrepancy between the real andrecorded value was possibly due to an inherent memory effect inthe pH sensitivemembrane, where the magnitude inresponse to a changing pH depended on the previous pH value. Thedifference between the initial pH measurement and the solution value of pH 4 and 7 (Fig.4.2) was comparable to the offsetmagnitudes seen in Fig. 4.1.
Considering Fig. 4.2, the offset recorded for pH 10.5 was dueto additional factors, such as drift in the reference electrode andsupply voltage. The potential divider circuit, which clamped thedrain potential of the ISFET, was connected between groundand the positive supply rail. Thus, a corresponding change inthe positive supply rail ofmV hwould result in adrain voltage change ofmV hfrom the potential divider circuit. The additional drift from the Ag|AgCl reference ofmV hbalanced the remaining drift ofmV recorded. The additional discrepancy found at pH 10.5 was most likely a result of long-term signal drift from the inter-action of proton reactive sites in the bulk of themembrane, with the drift becoming more predominant in alkaline solutions and at higher temperatures.
Bench testing of the oxygen sensor proved satisfactoryoperation of the electrochemical cell with a low noise (1%of full signal magnitude) and rapid response time of 10 s.However, signal resolution was limited to the standard error ofmg L. The signal discrepancy was caused bycontamination or deposits on the working electrode surface, which reduced the sensitivity, and by ambient temperature variation, changing the amount of dissolved oxygen by 2%C. Cleaning the surface in abarrel asher restored thefunction. However, signal drift was also caused by electrolytepenetration of the interface between the PECVDlayerand the underlying gold working electrode comprising themicroelectrode array. This represented a more serious problem,since it effectively increased the combined surface area of theworking electrode resulting in an increase in signal magnitudeat a constant dissolved oxygen level.
The conductivity sensor is currently being redesigned to ex-tend the dynamic range. The sensor will be an interdigitatedgold planar electrode using to prevent the absorption of organic compounds onto its surface.
Methods of digital signal processing will be considered afterdata acquisition to improve the performance from each sensorwith respect to signal drift. In contrast, analog signal algorithms (artificial neural networks) will be used in the sensor electronicsto cancel out the memory effect of the pH sensor, and the reduction in sensitivity caused by contamination of the sensor surface.
5.2. Micro system
The temperature tolerance of the radio-transmitter excludeddynamic temperature measurements for both theType I andTypeIIpill within the range of the temperature sensor. In order to increase the data collection rate, the potential for signal transmission at the European Industrial, Scientific and Medical (ISM) Standard (433.9 MHz) will be explored.
The combined power consumption from the microsystem washigher than the theoretical predictions made during the designprocess, which considered the implementation of sleep modeand the use of an on-off keying (OOK) transmitter. In the sleepmode, the sensors and analog circuitry were powered up prior todata sampling by the digital processing unit, and then turned off. A constant power mode was used in the experiments, since boththe oxygen and pH sensors required time to stabilize (15 s) after being switched on. Thus, the current consumption from theASIC and sensors was measured to 54% above the predicted 1.1mA from the system implementing sleep mode. The FSK typeradio transmitters reduced the load from the ASIC by drawingpower directly from the batteries rather than the chip, althoughit consumed on average twice the amount of current than a com-parable OOK type transmitter (calculated to 1 mA). The simulated power consumption of the free runningType I transmitter (2.45 mA) was 12% above the measured data due to the reduced power consumption at the measured frequency of 39.08 MHz at25C in contrast to the calculated frequency of 40.01 MHz usedin the model. The measured current consumption of the crystal stabilizedType II unit was comparable to the modeled data. Thecapacity of the enlarged SR44cells used as power supplyunits could meet the increased current demand, in contrast to theSR26cells originally proposed.
5.3.Observations on receiver computer
1. 2 SR44 Ag2O batteries are used.
2. Operating Time > 40 hours.
3. Power Consumption = 12.1 Mw
4. Corresponding current consumption = 3.9mA
5. Supply Voltage = 3.1 V
5.4.Range
1. Temperature from 0 to 70 ° C
2. pH from 1 to 13
3. Dissolved Oxygen up to 8.2 mg per liter
4. Conductivity above 0.05 mScm-1
5. Full scale dynamic Range analogue signal = 2.8 V
5.5.Accuracy
1. pH channel is around 0.2 unit above the real value
2. Oxygen Sensor is ±0.4 mgL.
3. Temperature & Conductivity is within ±1%.
Advantages
1. It is being beneficially used for disease detection & abnormalities in human body. There fore it is also called as MAGIC PILL FOR HEALTH CARE.
2. Adaptable for use in corrosive & quiescent environmentt
3. It can be used in industries in evaluation of water quality, Pollution Detection, fermentation process control & inspection of pipelines.
4. Micro Electronic Pill utilizes a PROGRAMMABLE STANDBY MODE, So Power consumption is very less.
5. It has very small size, hence it is very easy for practical usage’
6. High sensitivity, Good reliability & Life times.
7. Very long life of the cells(40 hours), Less Power, Current & Voltage requirement (12.1 mW, 3.9 mA, 3.1 V)
8. Less transmission length & hence has zero noise interference
6.2.Disadvantages
1. It cannot perform ultrasound & impedance tomography.Tomography is imaging by sections or sectioning, through the use of any kind of penetrating wave.
2. Cannot detect radiation abnormalities.
3. Cannot perform radiation treatment associated with cancer & chronic inflammation.
4. Micro Electronic Pills are expensive & are not available in many countries.
5. Still its size is not digestible to small babies.
CONCLUSION
We have developed an integrated sensor array system whichhas been incorporated in a mobile remote analytical microelectronic pill, designed to perform real-timein situ measurements of the GI tract, providing the firstin vitro wireless transmitted multichannel recordings of analytical parameters. Further work will focus on developing photopatternable gel electrolytes andoxygen and cationselective membranes. The microelectronic pill will be miniaturized for medical and veterinary applications by incorporating the transmitter on silicon and reducing powerconsumption by improving the data compression algorithm andutilizing a programmable standby power mode.
The generic nature of the microelectronic pill makes itadaptable for use in corrosive environments related to environ-mental and industrial applications, such as the evaluation ofwater quality, pollution detection, fermentation process controland the inspection of pipelines. The integration of radiationsensors and the application of indirect imaging technologiessuch as ultrasound and impedance tomography, will improvethe detection of tissue abnormalities and radiation treatmentassociated with cancer and chronic inflammation.
In the future, one objective will be to produce a device, analogous to a micro total analysis system (TAS) or lab on a chip sensorwhich is not only capable of collecting and processing data, but which can transmit it from a remote location. The overall concept will be to produce an array of sensor devicesdistributed throughout the body or the environment, capable oftransmitting high-quality information in real-time.