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Microfabricated solid-state dissolved oxygen sensor
Sensors and Actuators B 83 (2002) 138±148
Microfabricated solid-state dissolved oxygen sensor Glen W. McLaughlina,*, Katie Bradena, Benjamin Francb, Gregory T.A. Kovacsa,b a
Department of Electrical Engineering, CISX-202, Stanford University, Stanford, CA 94305-4075, USA b Department of Medicine, Stanford University, Stanford, CA 94305, USA
Abstract This paper describes the design, fabrication and testing of a microfabricated oxygen concentration sensor consisting of a microfabricated thin-®lm electrode matrix overlaid with a solid-state proton conductive matrix (PCM) and encapsulated in a bio-inert polytetra¯uoroethelene (PTFE) ®lm. Through cyclic voltammetry (CV) and voltage step (VS) measurements, the device was shown to have a linear response with respect to dissolved oxygen concentration over a range from 0 to 601 mM (0±200 atm% O2). The diffusion constant of the PCM was measured to be D1 1:7 10 10 m2/s (2:7 10 11 m2/s, n 40), which translates to a response time constant of t 0:7 s (0.11 s, n 40). These results were consistent with those predicted by basic electrochemistry. # 2002 Published by Elsevier Science B.V. Keywords: Oxygen sensor; PTFE ®lm; Bio-inert device
1. Introduction Currently, medical oxygen sensors are large aqueousbased devices, that have inconsistent performance and limited shelf life [1]. Hand fabrication limits the ability of current devices to be cost effectively reduced in size, and leads to inconsistent working electrode size and separation. Microfabrication methods lower cost and size while improving device consistency. Use of a solid-state PCM extends shelf life by eliminating the need for re-hydration. Therefore, microfabricated sensors using a PCM encapsulated in a bio-inert polytetra¯uoroethelene (PTFE) ®lm are potentially smaller, more consistent in function, and longer-lasting than currently available sensors. 2. Design Each element of the test matrix consists of a working, reference, and counter electrode set. The surface areas of these are maintained at a 1:5:25 ratio. This ratio of surface areas was chosen such that the highest current density would be present at the working electrode, the site of primary chemistry. The current density is smaller at the counter electrode, because it has a substantially larger surface area. The reference electrode functions as a sense electrode and only the parasitic current of the measurement system ¯ows
* Corresponding author. E-mail address: [email protected] (G.W. McLaughlin).
0925-4005/02/$ ± see front matter # 2002 Published by Elsevier Science B.V. PII: S 0 9 2 5 - 4 0 0 5 ( 0 2 ) 0 0 0 2 1 - 7
through it. Thus, the current density at this electrode is signi®cantly smaller than at the working electrode. The dissolved oxygen sensor test matrix is constructed in a 4 4 array con®guration. Each array is individually addressable and consists of a set of electrode elements. Four different working electrode diameters: 10, 20, 40, and 80 mm are arrayed. The number of elements arrayed in each set also varies in four steps: 1, 2, 4, and 8. Thus, each column of the matrix has elements of the same working electrode diameter, while each row has arrays containing the same number of electrode elements (Fig. 1). 3. Fabrication The electrode test matrix was fabricated by methods described in [2]. Following fabrication, the completed matrix was ready for deposition of the PCM. The PCM is a mixture of Na®onTM per¯uorinated ion-exchange resin, 5 wt.% solution in a mixture of lower aliphatic alcohol and water (Aldrich, Milwaukee, WI, USA) [3] and polyvinylpyrrolidone (PVP-360) (Sigma, St. Louis, MO, USA) with 3 wt.% 2,6-bis(azidobenzylidene)-4-methylcyclohexanone (Aldrich) [4] at a ratio of 4:1. Na®onTM is similar to PTFE in that it is a hydrophobic ¯uoropolymer structure that is permeable to O2. Unlike PTFE, it contains a hydrophilic region, which acts as a proton transport site. This material belongs to the wide class of solid superacid catalysts and exhibits acid strength greater than that of 100% H2SO4. The superacidity of this material has been attributed to the electron-withdrawing effect of the per¯uorocarbon chain
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Fig. 1. Illustration of the test matrix of electrodes.
Fig. 2. Photograph of the completed Clark-type oxygen sensor test matrix. The matrix consists of 16 different type of electrode configurations each individually addressable. The working electrodes have been configured in four different diameters: 10, 20, 40, and 80 mm. The number of electrodes arrayed have also been configured in four different groups: 1, 2, 4, and 8.
acting on the sulfonic acid group. This mixture creates a ®lm with two important qualities: (1) it is a solid-state PCM, due to the Na®onTM component, which (2) because of PVP-360, adheres to the electrode surface over extended time. The surface of the device was treated with a mixture of Silane 174A (Sigma) and ethanol at a ratio of 1:50. A 10 ml of this mixture was pipetted onto the test matrix and distributed using compressed nitrogen. Next, 100 ml volume of the PCM was placed on the device surface via a micropipette. The device was then baked in a vacuum oven at a temperature of 100 8C for a duration of 24 h. The oven was maintained at a base pressure of 100 Torr with a constant ¯ow of nitrogen. The thickness of the PCM layer was x1 50 5 mm, estimated by dividing the known volume of the deposited mixture by the surface area coverage.
The ®nal fabrication step was the encapsulation of the device in a bio-inert, oxygen permeable, PTFE ®lm. The deposition conditions of the PTFE ®lm were similar to those presented in [2]. A photograph of the completed device is shown in Fig. 2. 4. Theory of operation Dissolved gas sensors operate by electro-chemically reducing oxygen dissolved in solution and measuring the resulting current. The advantage of the sensor described is that the electrodes are contained within a bio-inert oxygen permeable PTFE membrane. This membrane enables the device to be used in delicate biological media such as blood. Isolating
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Fig. 3. Illustration representation of the net reaction mechanism with an overlay of a diagrammatic representation of the sensor.
the electrodes from the media improves measurement accuracy because it prevents interference signals from other electroactive species [1,5]. Reduction of oxygen at the cathode is a two-step process [6]. A ¯ow diagram of the reactions within the device is presented in Fig. 3. The working electrode reactions are summarized by Eqs. (1.1) and (1.2).
O2 2H 2e ! H2 O2
(1.1)
H2 O2 2H 2e ! 2H2 O
(1.2)
The reaction occurring at the anode, the counter electrode, converts the products from the cathode back into reactants. This is summarized by Eq. (1.3). 2H2 O ! O2 4H 4e
(1.3)
Unlike standard Clark cell con®gurations, there is no net consumption of either electrode. The dynamic performance of the device is predicted by solving the diffusion equation for an applied voltage step (VS), under the boundary conditions that the net concentration of oxygen at the surface of the electrode is zero at t 0, and the concentration of the oxygen at the interface of the device and environment is constant. The solution results in Eq. (1.4) [7]. ! 1 X C0 jp=x1 D1 t i t nFAD1 1 2e (1.4) x1 j1 This result is used to evaluate the performance of the device as well as to extract the diffusion constant for the PCM
which from this equation is predicted to be independent of electrode size and geometry. 5. Characterization Measurements of the electrode test matrix were made to determine the performance of the device using both cyclic voltammetry (CV) and chronoamperometric methods. CV measurements were used to determine the bias point of the device as well as to verify that no other competing reactions were present at the selected bias point. Chronoamperometric measurements were made to determine the dynamic response of the device. The combination of these measurements provides the experimental data needed to ®t theoretical models of the system. Cyclic voltammetry measurements of the devices were made over a voltage range of 0.9 to 0.9 V at a step increment of 50 mV/s starting at a 0 V bias. The measurements con®rmed that the device response followed the relationship predicted by the following Eq. (1.5) [9]. i i i 1 1 (1.5) e anf Z e 1 anf Z i0 il;c il;a where i0 is the exchange current, a the transfer coef®cient, n the number of electrons, f ratio of Faraday's constant to that of the ideal gas constant and temperature, Z the overpotential, and il,c and il,a are the diffusion limited cathodic and anodic currents, respectively. This equation predicts two
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Fig. 4. Cyclic voltammetry measurement of an 80 mm diameter working electrode with four elements arrayed. The amount of dissolved oxygen from that of atmospheric ranged from 5 to 208%. The measured current in region 2 increases linearly with that of the dissolved oxygen concentration.
regions: (1) the kinetic control region and (2) the diffusion control region. At low bias levels, the reaction at the working electrode is in the kinetic control region where the current has a strong dependence on the applied voltage. At larger bias points, the transport of oxygen to the working electrode shifts from operating under kinetic control to being diffusion limited. The point at which this happens is dependent on the ratio of the limiting current to that of the exchange current. The limiting current is proportional to the concentration and hence the transition between the kinetic (region 1) and the diffusion limited (region 2) regions occurs at an increased bias with increasing oxygen concentration. Measurements of a family of CV curves for a four element, 80 mm working electrode diameter array are presented in Fig. 4. Each element has a drawn surface area of 52,026.5 mm2, the spacing between the working electrodes is 80 mm. This ®gure shows the different possible regions of operation of the device and illustrates device response to a range of oxygen concentration from 5 to 208% with respect to atmospheric oxygen or 15± 625 mM. The ®gure shows that for low oxygen concentration, the kinetic control range of the reaction is from a bias level of 0 to approximately 0.6 V, while the diffusion control region ranges from approximately 0.6 to approximately 0.7 V. The transition between regions 1 and 2 can be set at the in¯ection point of the second derivative of the current/voltage relationship. The bottom curve shows the current beginning to rise above the ground reference level at the 0.7 V bias level. This is caused by an additional reaction occurring in parallel with the oxygen sensor. This side reaction introduces an additional current source into the signal path resulting in signal bias. The bias point chosen to make linearity
measurements of the electrodes was 0.65 V. This value was chosen, because it allows maximum dynamic range of the concentration level within the diffusion limited region as well as minimal signal bias from competing reactions. Fig. 5 shows a plot of CV measurements made at three different dissolved oxygen concentration levels with an overlay of the predicted values from Eq. (1.5). The measured data and that predicted by the theoretical model match well. The average R2 value of the data to the model was greater than 0.95. The region of physiological interest ranges from around 25 to 100 atm% O2 (75±300 mM). Oxygen concentration to current linearity measurements were made at 12 different levels of oxygen concentration on 10 of the 16 different sized electrode arrays. The current responses of the remaining six sets of electrodes were below the measurement capabilities of the instrumentation. Fig. 6 shows the linearity of eight element arrays with 10, 20, 40, and 80 mm working electrode diameters. The offset currents at the zero concentration level are primarily caused by two factors: (1) device capacitance and (2) proton conductive matrix resistance. Device capacitance creates an offset current related to the constant sweep rate of the measurement by I C dV/dt. The resistance of the proton conductive matrix also increases proportionately to the electrode area as well as possible contaminants contained within the ®lm. These factors combine to produce larger offset currents for larger electrode arrays. The current response of the electrode arrays with respect to dissolved oxygen concentration is linear as expected. The next measurement was to investigate the current-tooxygen-concentration coef®cient or the slope coef®cient of
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Fig. 5. Plot of three different dissolved oxygen concentrations of the 80 mm diameter working electrode with two elements arrayed and the predicted results from Eq. (1.5). The oxygen concentration region of physiological interest ranges from around 25 to 100 atm% O2. The measured data had on average an R2 value of fit to the theoretical model of greater than 0.95.
an LMS ®t of the data in Fig. 7, scaled with respect to electrode diameter. Fig. 7 shows the slope coef®cients of the same arrays measured in Fig. 6. An LMS ®t of the slope coef®cients with respect to the drawn area of the electrode has been performed to show that the coef®cients scale as predicted by Eq. (1.5). This suggests that the slope coef®cient of the response of each array should be proportional to
the electrode area, and hence to the square of the working electrode diameter. The slope coef®cients of other electrode arrays are presented in Table 1. The equations used to model the device predicted the current to scale linearly with respect to electrode area. This simple model assumes that vertical diffusion of dissolved oxygen dominates, thus making the contributions of lateral diffusion insigni®cant to the overall
Fig. 6. Cyclic voltammetry calibration measurements of eight electrode array sets with the working electrodes having diameters of 10, 20, 40, and 80 mm. Measured at a 0.65 V bias. Error bars for each point were determined by the standard deviation of a LMS linear fit.
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Fig. 7. Cyclic voltammetry scaling relationship between eight electrode array sets with working electrodes having diameters of 10, 20, 40, and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit. Each point of the plot was determined from 10 independent measurements.
signal. This model is accurate for larger-sized electrode sets [8], but produces misleading scaling information in microsized electrodes [9,10]. To understand this relationship it is informative to examine the current density, mA/cm2, relationship of electrodes for varying sizes. Theory predicts that as the area of the electrode decreases, the signal contribution from vertical diffusion of the reactants will decrease proportionately to the total surface area. Theory for microelectrodes predicts that the signal contributed from the lateral diffusion of the reactants will decrease proportionately to the total perimeter of the electrode. The ratio of the perimeter to area of a circular electrode is proportional to the inverse of the radius. Thus, the current density of the electrode should increase with respect to decreasing electrode area. Data collected from the CV measurements have been normalized to consider the current density relationship with respect to electrode diameter. The current density is assumed to be the superposition of two components: (1) vertical diffusion and (2) lateral diffusion. The vertical diffusion term as described above is expected to be proportional to the area of the electrode, while the lateral term is expected to be
proportional to the perimeter of the electrode. This produces a current to electrode size relationship that can be explained by the following equation: I Ar 2 Br
(1.6)
where A is the constant for the vertical diffusion term and B is the constant for the lateral diffusion term. This equation can be divided by the area of the electrode to obtain the relationship of the current density with respect to the radius of the electrode. J A0
B0 r
(1.7)
The current density data collected from four different size element arrays is shown in Fig. 8. A east mean square ®t to Eq. (1.7) has been performed to determine the validity of the approximations. The estimated values of A and B were 0.46 (mA/(cm2 percent O2)) and 8.0 (mA/(cm2 percent O2)/mm), respectively (R2 > 0:95). Chronoamperometric or step response (ST) measurements were made on the electrode arrays to determine the dynamic
Table 1 Scaling of the change in working electrode currents with respect to a change in the percent O2 concentration dissolved in the solution Set
Eight Four Two One
Electrode diameter (mm) (pA/atm% O2) 10
20
40
80
7.4 0.2 (n 10) NA NA NA
21.8 1.8 (n 10) 8.2 0.6 (n 10) NA NA
98.5 2.6 (n 10) 32.2 1.8 (n 10) 17.6 0.6 (n 10) NA
336 118 47.2 14.0
10.6 (n 10) 1.8 (n 10) 0.2 (n 10) 0.03 (n 10)
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Fig. 8. Current density scaling properties of microelectrode arrays. The current density of the microelectrode arrays has a decreasing trend with respect to an increase in electrode diameter. This is caused by the surface to perimeter relationship. Each point of the plot was determined from four independently measured points, except for the 10 mm diameter electrode which was based off of two independently measured points. Each of the measured points was calculated from 10 independent measurements. Error bars were determined from the standard deviation of these collective measurements.
behavior of the devices. These measurements provided a means of extracting an estimate of the diffusion constant for oxygen within the Na®onTM:PVP-360 mixture. Measurements were made by applying a step voltage change to the working electrode of 0.65 V and measuring the current response at a sample rate of one sample/mS for a duration of 10 s. These measurements con®rmed that the electrode arrays follow the predicted relationship derived in Eq. (1.4). This equation also predicted that the step response would have a characteristic resembling an exponential decay with a maximum time constant: 2 p t D1 (1.8) x1 where x1 is the sum of the thickness of the proton conductive matrix and the PTFE ®lm, and D1 is the diffusion constant for oxygen within the proton conductive matrix. This relationship has been used as a method to extract the diffusion constant. A family of the characteristic ST response curves for one of the four different sized electrode sets measured is shown in Fig. 9. This ®gure illustrates the ST response measurements of an eight element, 20 mm working electrode diameter array. The family of curves range in dissolved oxygen concentration from 5 to 203 atm% O2. The lower current value curves correspond to the lower oxygen concentration. The current value measured by the system reaches a steady state approximately 5.0 s after the application of the step response. This value corresponded to about seven time constants or a predicted time constant of t 0:7 s (0.11 s, n 40). Fig. 10 shows a plot of the
step response measurements made at three different dissolved oxygen concentration levels with an overlay of the predicted values from Eq. (1.4). The measured data matches that predicted by the theoretical model (the average R2 0:85 0:02, n 40). The model is limited in this setting, because it does not account for the ®lter characteristics of the measurement system and the parasitic capacitance of the
Fig. 9. Step response measurements of an 20 mm diameter working electrode with eight elements arrayed. Each curve is a measurement taken with a different concentration of dissolved oxygen. The dissolved concentration ranged from 5 to 203%. The measured current increases linearly with that of the dissolved oxygen concentration.
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Fig. 10. Plot of three different dissolved oxygen concentrations of the 40 mm diameter working electrode with eight elements arrayed and the predicted results from Eq. (1.4). The measured data had on average an R2 value for the fit to the theoretical model of greater than 0.85.
device and cabling. These factors could easily account for the discrepancies between the theoretical model and the measured data. Step response linearity measurements were made on the same arrays of electrodes used for the CV measurements. These measurements were made 5.0 s after the application of the step input. Each of the four different sized electrode
arrays appeared to have consistent settling times. This is consistent with theory that predicts that the time constant is independent of electrode size. Fig. 11 shows the linearity of the ST response measurements of the eight element arrays with working electrode diameters of 10, 20, 40, and 80 mm. The ®gure shows that the current/oxygen concentration has a linear relationship as predicted.
Fig. 11. Voltage step scaling relationship between eight electrode array sets with working electrodes having diameters of 10, 20, 40 and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit.
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Fig. 12. Voltage step calibration measurements of eight electrode array sets with the working electrodes having diameters of 10, 20, 40, and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit. Each point of the plot was determined from 10 independent points.
The next measurement was to investigate the current-tooxygen-concentration coef®cient or the slope coef®cient of an LMS ®t of the data in Fig. 11, scaled with respect to electrode diameter. Fig. 12, shows the slope coef®cient for the same electrode arrays measured in Fig. 11. An LMS ®t of the slope coef®cients with respect to the square of the electrode diameter has been performed, demonstrating that they scale with the area of the working electrode as predicted by Eq. (1.4). The slope coef®cients of the measured electrode arrays are presented in Table 2. In order to extract the diffusion coef®cient of the above equation, the thickness of the proton conductive layer must to be known. The amount of PCM placed over the electrode set was 100 ml with 80 wt.% solvent. This produced a layer that had an estimated thickness of 50 5 mm that were based off of geometric measurements as well as the known variance of the ®lm dispensing apparatus. 2 p 2C0 ln i tm D1 t ln nFAD1 (1.9) x1 x1
This relationship shows that the slope of the resulting line with respect to time is proportional to the diffusion constant. An LMS ®t of the measured data produced the following estimate of the diffusion coef®cient for the Na®onTM:PVP360 mixture:
Table 2 Step response measurements of the scaling properties of a change in working electrode current with respect to a change in the percent O2 in solution
6. Discussion
Set
Eight
Electrode diameter (mm) (pA/atm% O2) 10
20
40
80
7.1 1.4 (n 10)
22.7 1.4 (n 10)
88.4 3.9 (n 10)
210 13.2 (n 10)
D1 1:7 10
10
m2 =s 2:7 10
11
m2 =s; n 40 (1.10)
Using this value in conjunction with Eq. (1.8), produced an estimate of the ®rst order time constant: t 0:7 s 0:11 s; n 40
(1.11)
This time constant estimation agreed to within 16% of the 40 measurements taken from four different electrode sets at different dissolved oxygen concentrations. From Eq. (1.4), it was predicted that the time constant would be, to a ®rst approximation, independent of dissolved oxygen concentrations.
Measurements were made with the oxygen sensors using both the CV and chronoamperometric methods. The model used made the assumption, that of the reactions involved in the device, the rate limiting step was the transport of oxygen to the working electrode, while the kinetics of the hydrogen as well as the hydrogen peroxide contained within the PCM were not limiting factors. Measurements have shown this to
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be a reasonable assumption at the dissolved oxygen concentration between 0 and 601 mM (0±200 atm% O2), for if the kinetics of these particles played a role as a limiting factor in the performance of the device one would expect to see a plateau in the measured current as the concentration of oxygen was increased, this was not observed (Figs. 6 and 11). Both of the measurements showed good linearity (R2 > 0:99 CV) and (R2 > 0:96 chronoamperometry) as measured from the 80 mm diameter eight electrode sets. Scaling the electrode size showed an increase in the current density as the diameter of the electrode was decreased, and the data matched theoretical prediction well (R2 > 0:95). Unfortunately, circuit parasitics of the measurement apparatus prevented the characterization of some of the ®ner geometry devices, which would have been useful in providing additional data points in the higher current density region of the curve. Using an improved measurement apparatus would help address this shortcoming. Several possible improvements can be made to the device for future studies. Better control of the Na®onTM:PVP-360 mixture would help reduce the variance in the ®lm thickness between different devices. One possible method would be to spin cast and develop the material using semiconductor processing equipment. The addition of the 2,6-bis(azidobenzylidene)-4-methylcyclohexanone to the mixture provides a means of crosslinking the ®lm when exposed to ultraviolet light to allow the PCM to be patterened. A second controlled pattering method would be to apply the Na®onTM:PVP-360 mixture using a controlled stencil placed on top of the electrode array in a similar process to that of silk screening, enabling a cost effective method of consistently producing patterned thickness controlled ®lms. Other performance improvements that could be made to the device would be the incorporation of on chip electronics to control and log the electrode measurements. This would minimize the parasitic effects of routing signals off of the sensor, allowing ®ner geometry devices to be used as well as increase the overall utility of the unit by alleviating the need for peripheral support equipment. 7. Conclusion A microfabricated, fast-responding dissolved gas oxygen sensor, which should meet biomedical and industrial needs, has been presented. The device's response is linear and proportional to electrode area. The sensor's performance is also consistent with the theoretical model, which provides useful design insights. Acknowledgements This work is funded by DARPA under contract N66600196-C-8631 and Medtronic Corporation (Minneapolis, MN, USA).
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References [1] L.C. Clark, Monitor and control of blood and tissure oxygen tension, Trans. Am. Soc. Artificial Intern. Organs 2 (1956) 144±156. [2] G.W. McLaughlin, A.F. Flannery, C.W. Storment, G.T.A. Kovacs, Integrated Clark-type oxygen sensor array using a pulsed plasma deposited PTFE-like film, in: Proceedings of Transducers 1999, Sendai, Japan, 7±10 June 1999, pp. 1176±1179. [3] H.-Q. Yan, Ch.-ch. Liu, Humidity effects on the stability of a solid polymer electrolyte oxygen sensor, Sens. Actuators B 10 (2) (1993) 133±136. [4] G. Jobst, G. Urban, A. Jachimowicz, F. Kohl, O. Tilado, Thin-film Clark-type oxygen sensor based on novel polymer membrane system for in vivo and biosensor applications, Biosens. Bioelectron. 8 (3/4) (1993) 123±128. [5] T. Togawa, T. Tamura, P.A. Oberg, Biomedical Transducers and Instruments, CRC Press, New York, 1997, pp. 294±301. [6] J. Maruyama, M. Inaba, Z. Ogumi, Rotating ring-disk electrode study on the cathodic oxygen reduction at NafionTM-coated gold electrodes, J. Electroanal. Chem. 458 (1998) 175±182. [7] G.W. McLaughlin, Microfluidic and biosensor applications of fluoropolymer films, Doctoral dissertation in electrical engineering, Stanford University, May 2001, pp. 145±196. [8] P.T. Kissinger, W.R. Heineman, Laboratory Techniques in Electroanalytical Chemistry, Marcel Dekker, New York, 1984, pp. 86±142. [9] A.J. Bard, L.R. Faulkner, Electrochemical Methods Fundamentals and Applications, Wiley, New York, 1980, pp. 1±277. [10] J. Wang, Analytical Electrochemistry, VCH Publishers, New York, 1994.
Biographies Glen W. Mclaughlin was born in San Francisco, CA, USA. He received the BSc degree in electrical engineering from Carnegie Mellon University, Pittsburgh, PA in 1992. He received the MSc and PhD degrees in electrical engineering from Stanford University, Stanford, CA in 1999 and 2001, respectively. He is currently the Chief Technical Officer of Novasonics Inc., a company developing advanced, hand-held ultrasound instruments. Katie Braden will receive the BSc and MSc degrees in electrical engineering from Stanford University in June 2002. Her experience includes design and fabrication of a thickness sensor for a ParyleneTM deposition system at the University of Illinois, and her current research in electrochemical sensors. In addition, she is President of the Stanford Society of Women Engineers. Benjamin Franc received the BSc degree in chemical engineering from Stanford University in 1996. He earned the MD degree from the University of Southern California, Los Angeles, CA in 2000. He has completed his internship in general surgery at Stanford University Hospital and is currently completing a residency in nuclear medicine at that institution. His experience includes work on the design and characterization of a bioartificial liver and tissue factor modulation. He is currently pursuing research in the areas of medical sensors and tissue engineering. Gregory T.A. Kovacs received the BSc degree in electrical engineering from the University of British Columbia, Vancouver, BC in 1984; the MSc degree in bioengineering from the University of California, Berkeley in 1985; the PhD degree in electrical engineering from Stanford University in1990 and the MD degree from Stanford University in 1992. His industry experience includes the design of electronic instrumentation, commercial and consumer product design, extensive patent law consulting, and the cofounding of several companies (most recently Cepheid Inc. (CPHD) in Sunnyvale, CA). In 1991, he joined Stanford University and is currently an
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Associate Professor of electrical engineering with an appointment in the Department of Medicine, by courtesy. He teaches courses in electronic circuits and micromachined transducers. He held the Robert N. Noyce Family Faculty Scholar Chair in 1992±1994, received an NSF Young Investigator Award in 1993, was a Terman Fellow in 1994±1997, was a University Fellow in 1996±1998, and was appointed to the Defense
Sciences Research Council in 1995 (he is currently Chairman). He is a Fellow National of the Explorers Club. His present research areas include sensors and actuators, medical instruments, biotechnology, and micro fabrication, all with emphasis on solving practical problems. He has numerous patents and technical publications including the text, ``micromachined transducers sourcebook'' (WCB/McGraw-Hill, 1998).
Microfabricated solid-state dissolved oxygen sensor Glen W. McLaughlina,*, Katie Bradena, Benjamin Francb, Gregory T.A. Kovacsa,b a
Department of Electrical Engineering, CISX-202, Stanford University, Stanford, CA 94305-4075, USA b Department of Medicine, Stanford University, Stanford, CA 94305, USA
Abstract This paper describes the design, fabrication and testing of a microfabricated oxygen concentration sensor consisting of a microfabricated thin-®lm electrode matrix overlaid with a solid-state proton conductive matrix (PCM) and encapsulated in a bio-inert polytetra¯uoroethelene (PTFE) ®lm. Through cyclic voltammetry (CV) and voltage step (VS) measurements, the device was shown to have a linear response with respect to dissolved oxygen concentration over a range from 0 to 601 mM (0±200 atm% O2). The diffusion constant of the PCM was measured to be D1 1:7 10 10 m2/s (2:7 10 11 m2/s, n 40), which translates to a response time constant of t 0:7 s (0.11 s, n 40). These results were consistent with those predicted by basic electrochemistry. # 2002 Published by Elsevier Science B.V. Keywords: Oxygen sensor; PTFE ®lm; Bio-inert device
1. Introduction Currently, medical oxygen sensors are large aqueousbased devices, that have inconsistent performance and limited shelf life [1]. Hand fabrication limits the ability of current devices to be cost effectively reduced in size, and leads to inconsistent working electrode size and separation. Microfabrication methods lower cost and size while improving device consistency. Use of a solid-state PCM extends shelf life by eliminating the need for re-hydration. Therefore, microfabricated sensors using a PCM encapsulated in a bio-inert polytetra¯uoroethelene (PTFE) ®lm are potentially smaller, more consistent in function, and longer-lasting than currently available sensors. 2. Design Each element of the test matrix consists of a working, reference, and counter electrode set. The surface areas of these are maintained at a 1:5:25 ratio. This ratio of surface areas was chosen such that the highest current density would be present at the working electrode, the site of primary chemistry. The current density is smaller at the counter electrode, because it has a substantially larger surface area. The reference electrode functions as a sense electrode and only the parasitic current of the measurement system ¯ows
* Corresponding author. E-mail address: [email protected] (G.W. McLaughlin).
0925-4005/02/$ ± see front matter # 2002 Published by Elsevier Science B.V. PII: S 0 9 2 5 - 4 0 0 5 ( 0 2 ) 0 0 0 2 1 - 7
through it. Thus, the current density at this electrode is signi®cantly smaller than at the working electrode. The dissolved oxygen sensor test matrix is constructed in a 4 4 array con®guration. Each array is individually addressable and consists of a set of electrode elements. Four different working electrode diameters: 10, 20, 40, and 80 mm are arrayed. The number of elements arrayed in each set also varies in four steps: 1, 2, 4, and 8. Thus, each column of the matrix has elements of the same working electrode diameter, while each row has arrays containing the same number of electrode elements (Fig. 1). 3. Fabrication The electrode test matrix was fabricated by methods described in [2]. Following fabrication, the completed matrix was ready for deposition of the PCM. The PCM is a mixture of Na®onTM per¯uorinated ion-exchange resin, 5 wt.% solution in a mixture of lower aliphatic alcohol and water (Aldrich, Milwaukee, WI, USA) [3] and polyvinylpyrrolidone (PVP-360) (Sigma, St. Louis, MO, USA) with 3 wt.% 2,6-bis(azidobenzylidene)-4-methylcyclohexanone (Aldrich) [4] at a ratio of 4:1. Na®onTM is similar to PTFE in that it is a hydrophobic ¯uoropolymer structure that is permeable to O2. Unlike PTFE, it contains a hydrophilic region, which acts as a proton transport site. This material belongs to the wide class of solid superacid catalysts and exhibits acid strength greater than that of 100% H2SO4. The superacidity of this material has been attributed to the electron-withdrawing effect of the per¯uorocarbon chain
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Fig. 1. Illustration of the test matrix of electrodes.
Fig. 2. Photograph of the completed Clark-type oxygen sensor test matrix. The matrix consists of 16 different type of electrode configurations each individually addressable. The working electrodes have been configured in four different diameters: 10, 20, 40, and 80 mm. The number of electrodes arrayed have also been configured in four different groups: 1, 2, 4, and 8.
acting on the sulfonic acid group. This mixture creates a ®lm with two important qualities: (1) it is a solid-state PCM, due to the Na®onTM component, which (2) because of PVP-360, adheres to the electrode surface over extended time. The surface of the device was treated with a mixture of Silane 174A (Sigma) and ethanol at a ratio of 1:50. A 10 ml of this mixture was pipetted onto the test matrix and distributed using compressed nitrogen. Next, 100 ml volume of the PCM was placed on the device surface via a micropipette. The device was then baked in a vacuum oven at a temperature of 100 8C for a duration of 24 h. The oven was maintained at a base pressure of 100 Torr with a constant ¯ow of nitrogen. The thickness of the PCM layer was x1 50 5 mm, estimated by dividing the known volume of the deposited mixture by the surface area coverage.
The ®nal fabrication step was the encapsulation of the device in a bio-inert, oxygen permeable, PTFE ®lm. The deposition conditions of the PTFE ®lm were similar to those presented in [2]. A photograph of the completed device is shown in Fig. 2. 4. Theory of operation Dissolved gas sensors operate by electro-chemically reducing oxygen dissolved in solution and measuring the resulting current. The advantage of the sensor described is that the electrodes are contained within a bio-inert oxygen permeable PTFE membrane. This membrane enables the device to be used in delicate biological media such as blood. Isolating
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Fig. 3. Illustration representation of the net reaction mechanism with an overlay of a diagrammatic representation of the sensor.
the electrodes from the media improves measurement accuracy because it prevents interference signals from other electroactive species [1,5]. Reduction of oxygen at the cathode is a two-step process [6]. A ¯ow diagram of the reactions within the device is presented in Fig. 3. The working electrode reactions are summarized by Eqs. (1.1) and (1.2).
O2 2H 2e ! H2 O2
(1.1)
H2 O2 2H 2e ! 2H2 O
(1.2)
The reaction occurring at the anode, the counter electrode, converts the products from the cathode back into reactants. This is summarized by Eq. (1.3). 2H2 O ! O2 4H 4e
(1.3)
Unlike standard Clark cell con®gurations, there is no net consumption of either electrode. The dynamic performance of the device is predicted by solving the diffusion equation for an applied voltage step (VS), under the boundary conditions that the net concentration of oxygen at the surface of the electrode is zero at t 0, and the concentration of the oxygen at the interface of the device and environment is constant. The solution results in Eq. (1.4) [7]. ! 1 X C0 jp=x1 D1 t i t nFAD1 1 2e (1.4) x1 j1 This result is used to evaluate the performance of the device as well as to extract the diffusion constant for the PCM
which from this equation is predicted to be independent of electrode size and geometry. 5. Characterization Measurements of the electrode test matrix were made to determine the performance of the device using both cyclic voltammetry (CV) and chronoamperometric methods. CV measurements were used to determine the bias point of the device as well as to verify that no other competing reactions were present at the selected bias point. Chronoamperometric measurements were made to determine the dynamic response of the device. The combination of these measurements provides the experimental data needed to ®t theoretical models of the system. Cyclic voltammetry measurements of the devices were made over a voltage range of 0.9 to 0.9 V at a step increment of 50 mV/s starting at a 0 V bias. The measurements con®rmed that the device response followed the relationship predicted by the following Eq. (1.5) [9]. i i i 1 1 (1.5) e anf Z e 1 anf Z i0 il;c il;a where i0 is the exchange current, a the transfer coef®cient, n the number of electrons, f ratio of Faraday's constant to that of the ideal gas constant and temperature, Z the overpotential, and il,c and il,a are the diffusion limited cathodic and anodic currents, respectively. This equation predicts two
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Fig. 4. Cyclic voltammetry measurement of an 80 mm diameter working electrode with four elements arrayed. The amount of dissolved oxygen from that of atmospheric ranged from 5 to 208%. The measured current in region 2 increases linearly with that of the dissolved oxygen concentration.
regions: (1) the kinetic control region and (2) the diffusion control region. At low bias levels, the reaction at the working electrode is in the kinetic control region where the current has a strong dependence on the applied voltage. At larger bias points, the transport of oxygen to the working electrode shifts from operating under kinetic control to being diffusion limited. The point at which this happens is dependent on the ratio of the limiting current to that of the exchange current. The limiting current is proportional to the concentration and hence the transition between the kinetic (region 1) and the diffusion limited (region 2) regions occurs at an increased bias with increasing oxygen concentration. Measurements of a family of CV curves for a four element, 80 mm working electrode diameter array are presented in Fig. 4. Each element has a drawn surface area of 52,026.5 mm2, the spacing between the working electrodes is 80 mm. This ®gure shows the different possible regions of operation of the device and illustrates device response to a range of oxygen concentration from 5 to 208% with respect to atmospheric oxygen or 15± 625 mM. The ®gure shows that for low oxygen concentration, the kinetic control range of the reaction is from a bias level of 0 to approximately 0.6 V, while the diffusion control region ranges from approximately 0.6 to approximately 0.7 V. The transition between regions 1 and 2 can be set at the in¯ection point of the second derivative of the current/voltage relationship. The bottom curve shows the current beginning to rise above the ground reference level at the 0.7 V bias level. This is caused by an additional reaction occurring in parallel with the oxygen sensor. This side reaction introduces an additional current source into the signal path resulting in signal bias. The bias point chosen to make linearity
measurements of the electrodes was 0.65 V. This value was chosen, because it allows maximum dynamic range of the concentration level within the diffusion limited region as well as minimal signal bias from competing reactions. Fig. 5 shows a plot of CV measurements made at three different dissolved oxygen concentration levels with an overlay of the predicted values from Eq. (1.5). The measured data and that predicted by the theoretical model match well. The average R2 value of the data to the model was greater than 0.95. The region of physiological interest ranges from around 25 to 100 atm% O2 (75±300 mM). Oxygen concentration to current linearity measurements were made at 12 different levels of oxygen concentration on 10 of the 16 different sized electrode arrays. The current responses of the remaining six sets of electrodes were below the measurement capabilities of the instrumentation. Fig. 6 shows the linearity of eight element arrays with 10, 20, 40, and 80 mm working electrode diameters. The offset currents at the zero concentration level are primarily caused by two factors: (1) device capacitance and (2) proton conductive matrix resistance. Device capacitance creates an offset current related to the constant sweep rate of the measurement by I C dV/dt. The resistance of the proton conductive matrix also increases proportionately to the electrode area as well as possible contaminants contained within the ®lm. These factors combine to produce larger offset currents for larger electrode arrays. The current response of the electrode arrays with respect to dissolved oxygen concentration is linear as expected. The next measurement was to investigate the current-tooxygen-concentration coef®cient or the slope coef®cient of
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Fig. 5. Plot of three different dissolved oxygen concentrations of the 80 mm diameter working electrode with two elements arrayed and the predicted results from Eq. (1.5). The oxygen concentration region of physiological interest ranges from around 25 to 100 atm% O2. The measured data had on average an R2 value of fit to the theoretical model of greater than 0.95.
an LMS ®t of the data in Fig. 7, scaled with respect to electrode diameter. Fig. 7 shows the slope coef®cients of the same arrays measured in Fig. 6. An LMS ®t of the slope coef®cients with respect to the drawn area of the electrode has been performed to show that the coef®cients scale as predicted by Eq. (1.5). This suggests that the slope coef®cient of the response of each array should be proportional to
the electrode area, and hence to the square of the working electrode diameter. The slope coef®cients of other electrode arrays are presented in Table 1. The equations used to model the device predicted the current to scale linearly with respect to electrode area. This simple model assumes that vertical diffusion of dissolved oxygen dominates, thus making the contributions of lateral diffusion insigni®cant to the overall
Fig. 6. Cyclic voltammetry calibration measurements of eight electrode array sets with the working electrodes having diameters of 10, 20, 40, and 80 mm. Measured at a 0.65 V bias. Error bars for each point were determined by the standard deviation of a LMS linear fit.
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Fig. 7. Cyclic voltammetry scaling relationship between eight electrode array sets with working electrodes having diameters of 10, 20, 40, and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit. Each point of the plot was determined from 10 independent measurements.
signal. This model is accurate for larger-sized electrode sets [8], but produces misleading scaling information in microsized electrodes [9,10]. To understand this relationship it is informative to examine the current density, mA/cm2, relationship of electrodes for varying sizes. Theory predicts that as the area of the electrode decreases, the signal contribution from vertical diffusion of the reactants will decrease proportionately to the total surface area. Theory for microelectrodes predicts that the signal contributed from the lateral diffusion of the reactants will decrease proportionately to the total perimeter of the electrode. The ratio of the perimeter to area of a circular electrode is proportional to the inverse of the radius. Thus, the current density of the electrode should increase with respect to decreasing electrode area. Data collected from the CV measurements have been normalized to consider the current density relationship with respect to electrode diameter. The current density is assumed to be the superposition of two components: (1) vertical diffusion and (2) lateral diffusion. The vertical diffusion term as described above is expected to be proportional to the area of the electrode, while the lateral term is expected to be
proportional to the perimeter of the electrode. This produces a current to electrode size relationship that can be explained by the following equation: I Ar 2 Br
(1.6)
where A is the constant for the vertical diffusion term and B is the constant for the lateral diffusion term. This equation can be divided by the area of the electrode to obtain the relationship of the current density with respect to the radius of the electrode. J A0
B0 r
(1.7)
The current density data collected from four different size element arrays is shown in Fig. 8. A east mean square ®t to Eq. (1.7) has been performed to determine the validity of the approximations. The estimated values of A and B were 0.46 (mA/(cm2 percent O2)) and 8.0 (mA/(cm2 percent O2)/mm), respectively (R2 > 0:95). Chronoamperometric or step response (ST) measurements were made on the electrode arrays to determine the dynamic
Table 1 Scaling of the change in working electrode currents with respect to a change in the percent O2 concentration dissolved in the solution Set
Eight Four Two One
Electrode diameter (mm) (pA/atm% O2) 10
20
40
80
7.4 0.2 (n 10) NA NA NA
21.8 1.8 (n 10) 8.2 0.6 (n 10) NA NA
98.5 2.6 (n 10) 32.2 1.8 (n 10) 17.6 0.6 (n 10) NA
336 118 47.2 14.0
10.6 (n 10) 1.8 (n 10) 0.2 (n 10) 0.03 (n 10)
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Fig. 8. Current density scaling properties of microelectrode arrays. The current density of the microelectrode arrays has a decreasing trend with respect to an increase in electrode diameter. This is caused by the surface to perimeter relationship. Each point of the plot was determined from four independently measured points, except for the 10 mm diameter electrode which was based off of two independently measured points. Each of the measured points was calculated from 10 independent measurements. Error bars were determined from the standard deviation of these collective measurements.
behavior of the devices. These measurements provided a means of extracting an estimate of the diffusion constant for oxygen within the Na®onTM:PVP-360 mixture. Measurements were made by applying a step voltage change to the working electrode of 0.65 V and measuring the current response at a sample rate of one sample/mS for a duration of 10 s. These measurements con®rmed that the electrode arrays follow the predicted relationship derived in Eq. (1.4). This equation also predicted that the step response would have a characteristic resembling an exponential decay with a maximum time constant: 2 p t D1 (1.8) x1 where x1 is the sum of the thickness of the proton conductive matrix and the PTFE ®lm, and D1 is the diffusion constant for oxygen within the proton conductive matrix. This relationship has been used as a method to extract the diffusion constant. A family of the characteristic ST response curves for one of the four different sized electrode sets measured is shown in Fig. 9. This ®gure illustrates the ST response measurements of an eight element, 20 mm working electrode diameter array. The family of curves range in dissolved oxygen concentration from 5 to 203 atm% O2. The lower current value curves correspond to the lower oxygen concentration. The current value measured by the system reaches a steady state approximately 5.0 s after the application of the step response. This value corresponded to about seven time constants or a predicted time constant of t 0:7 s (0.11 s, n 40). Fig. 10 shows a plot of the
step response measurements made at three different dissolved oxygen concentration levels with an overlay of the predicted values from Eq. (1.4). The measured data matches that predicted by the theoretical model (the average R2 0:85 0:02, n 40). The model is limited in this setting, because it does not account for the ®lter characteristics of the measurement system and the parasitic capacitance of the
Fig. 9. Step response measurements of an 20 mm diameter working electrode with eight elements arrayed. Each curve is a measurement taken with a different concentration of dissolved oxygen. The dissolved concentration ranged from 5 to 203%. The measured current increases linearly with that of the dissolved oxygen concentration.
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Fig. 10. Plot of three different dissolved oxygen concentrations of the 40 mm diameter working electrode with eight elements arrayed and the predicted results from Eq. (1.4). The measured data had on average an R2 value for the fit to the theoretical model of greater than 0.85.
device and cabling. These factors could easily account for the discrepancies between the theoretical model and the measured data. Step response linearity measurements were made on the same arrays of electrodes used for the CV measurements. These measurements were made 5.0 s after the application of the step input. Each of the four different sized electrode
arrays appeared to have consistent settling times. This is consistent with theory that predicts that the time constant is independent of electrode size. Fig. 11 shows the linearity of the ST response measurements of the eight element arrays with working electrode diameters of 10, 20, 40, and 80 mm. The ®gure shows that the current/oxygen concentration has a linear relationship as predicted.
Fig. 11. Voltage step scaling relationship between eight electrode array sets with working electrodes having diameters of 10, 20, 40 and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit.
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Fig. 12. Voltage step calibration measurements of eight electrode array sets with the working electrodes having diameters of 10, 20, 40, and 80 mm. Error bars for each point were determined by the standard deviation of a LMS linear fit. Each point of the plot was determined from 10 independent points.
The next measurement was to investigate the current-tooxygen-concentration coef®cient or the slope coef®cient of an LMS ®t of the data in Fig. 11, scaled with respect to electrode diameter. Fig. 12, shows the slope coef®cient for the same electrode arrays measured in Fig. 11. An LMS ®t of the slope coef®cients with respect to the square of the electrode diameter has been performed, demonstrating that they scale with the area of the working electrode as predicted by Eq. (1.4). The slope coef®cients of the measured electrode arrays are presented in Table 2. In order to extract the diffusion coef®cient of the above equation, the thickness of the proton conductive layer must to be known. The amount of PCM placed over the electrode set was 100 ml with 80 wt.% solvent. This produced a layer that had an estimated thickness of 50 5 mm that were based off of geometric measurements as well as the known variance of the ®lm dispensing apparatus. 2 p 2C0 ln i tm D1 t ln nFAD1 (1.9) x1 x1
This relationship shows that the slope of the resulting line with respect to time is proportional to the diffusion constant. An LMS ®t of the measured data produced the following estimate of the diffusion coef®cient for the Na®onTM:PVP360 mixture:
Table 2 Step response measurements of the scaling properties of a change in working electrode current with respect to a change in the percent O2 in solution
6. Discussion
Set
Eight
Electrode diameter (mm) (pA/atm% O2) 10
20
40
80
7.1 1.4 (n 10)
22.7 1.4 (n 10)
88.4 3.9 (n 10)
210 13.2 (n 10)
D1 1:7 10
10
m2 =s 2:7 10
11
m2 =s; n 40 (1.10)
Using this value in conjunction with Eq. (1.8), produced an estimate of the ®rst order time constant: t 0:7 s 0:11 s; n 40
(1.11)
This time constant estimation agreed to within 16% of the 40 measurements taken from four different electrode sets at different dissolved oxygen concentrations. From Eq. (1.4), it was predicted that the time constant would be, to a ®rst approximation, independent of dissolved oxygen concentrations.
Measurements were made with the oxygen sensors using both the CV and chronoamperometric methods. The model used made the assumption, that of the reactions involved in the device, the rate limiting step was the transport of oxygen to the working electrode, while the kinetics of the hydrogen as well as the hydrogen peroxide contained within the PCM were not limiting factors. Measurements have shown this to
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be a reasonable assumption at the dissolved oxygen concentration between 0 and 601 mM (0±200 atm% O2), for if the kinetics of these particles played a role as a limiting factor in the performance of the device one would expect to see a plateau in the measured current as the concentration of oxygen was increased, this was not observed (Figs. 6 and 11). Both of the measurements showed good linearity (R2 > 0:99 CV) and (R2 > 0:96 chronoamperometry) as measured from the 80 mm diameter eight electrode sets. Scaling the electrode size showed an increase in the current density as the diameter of the electrode was decreased, and the data matched theoretical prediction well (R2 > 0:95). Unfortunately, circuit parasitics of the measurement apparatus prevented the characterization of some of the ®ner geometry devices, which would have been useful in providing additional data points in the higher current density region of the curve. Using an improved measurement apparatus would help address this shortcoming. Several possible improvements can be made to the device for future studies. Better control of the Na®onTM:PVP-360 mixture would help reduce the variance in the ®lm thickness between different devices. One possible method would be to spin cast and develop the material using semiconductor processing equipment. The addition of the 2,6-bis(azidobenzylidene)-4-methylcyclohexanone to the mixture provides a means of crosslinking the ®lm when exposed to ultraviolet light to allow the PCM to be patterened. A second controlled pattering method would be to apply the Na®onTM:PVP-360 mixture using a controlled stencil placed on top of the electrode array in a similar process to that of silk screening, enabling a cost effective method of consistently producing patterned thickness controlled ®lms. Other performance improvements that could be made to the device would be the incorporation of on chip electronics to control and log the electrode measurements. This would minimize the parasitic effects of routing signals off of the sensor, allowing ®ner geometry devices to be used as well as increase the overall utility of the unit by alleviating the need for peripheral support equipment. 7. Conclusion A microfabricated, fast-responding dissolved gas oxygen sensor, which should meet biomedical and industrial needs, has been presented. The device's response is linear and proportional to electrode area. The sensor's performance is also consistent with the theoretical model, which provides useful design insights. Acknowledgements This work is funded by DARPA under contract N66600196-C-8631 and Medtronic Corporation (Minneapolis, MN, USA).
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References [1] L.C. Clark, Monitor and control of blood and tissure oxygen tension, Trans. Am. Soc. Artificial Intern. Organs 2 (1956) 144±156. [2] G.W. McLaughlin, A.F. Flannery, C.W. Storment, G.T.A. Kovacs, Integrated Clark-type oxygen sensor array using a pulsed plasma deposited PTFE-like film, in: Proceedings of Transducers 1999, Sendai, Japan, 7±10 June 1999, pp. 1176±1179. [3] H.-Q. Yan, Ch.-ch. Liu, Humidity effects on the stability of a solid polymer electrolyte oxygen sensor, Sens. Actuators B 10 (2) (1993) 133±136. [4] G. Jobst, G. Urban, A. Jachimowicz, F. Kohl, O. Tilado, Thin-film Clark-type oxygen sensor based on novel polymer membrane system for in vivo and biosensor applications, Biosens. Bioelectron. 8 (3/4) (1993) 123±128. [5] T. Togawa, T. Tamura, P.A. Oberg, Biomedical Transducers and Instruments, CRC Press, New York, 1997, pp. 294±301. [6] J. Maruyama, M. Inaba, Z. Ogumi, Rotating ring-disk electrode study on the cathodic oxygen reduction at NafionTM-coated gold electrodes, J. Electroanal. Chem. 458 (1998) 175±182. [7] G.W. McLaughlin, Microfluidic and biosensor applications of fluoropolymer films, Doctoral dissertation in electrical engineering, Stanford University, May 2001, pp. 145±196. [8] P.T. Kissinger, W.R. Heineman, Laboratory Techniques in Electroanalytical Chemistry, Marcel Dekker, New York, 1984, pp. 86±142. [9] A.J. Bard, L.R. Faulkner, Electrochemical Methods Fundamentals and Applications, Wiley, New York, 1980, pp. 1±277. [10] J. Wang, Analytical Electrochemistry, VCH Publishers, New York, 1994.
Biographies Glen W. Mclaughlin was born in San Francisco, CA, USA. He received the BSc degree in electrical engineering from Carnegie Mellon University, Pittsburgh, PA in 1992. He received the MSc and PhD degrees in electrical engineering from Stanford University, Stanford, CA in 1999 and 2001, respectively. He is currently the Chief Technical Officer of Novasonics Inc., a company developing advanced, hand-held ultrasound instruments. Katie Braden will receive the BSc and MSc degrees in electrical engineering from Stanford University in June 2002. Her experience includes design and fabrication of a thickness sensor for a ParyleneTM deposition system at the University of Illinois, and her current research in electrochemical sensors. In addition, she is President of the Stanford Society of Women Engineers. Benjamin Franc received the BSc degree in chemical engineering from Stanford University in 1996. He earned the MD degree from the University of Southern California, Los Angeles, CA in 2000. He has completed his internship in general surgery at Stanford University Hospital and is currently completing a residency in nuclear medicine at that institution. His experience includes work on the design and characterization of a bioartificial liver and tissue factor modulation. He is currently pursuing research in the areas of medical sensors and tissue engineering. Gregory T.A. Kovacs received the BSc degree in electrical engineering from the University of British Columbia, Vancouver, BC in 1984; the MSc degree in bioengineering from the University of California, Berkeley in 1985; the PhD degree in electrical engineering from Stanford University in1990 and the MD degree from Stanford University in 1992. His industry experience includes the design of electronic instrumentation, commercial and consumer product design, extensive patent law consulting, and the cofounding of several companies (most recently Cepheid Inc. (CPHD) in Sunnyvale, CA). In 1991, he joined Stanford University and is currently an
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Associate Professor of electrical engineering with an appointment in the Department of Medicine, by courtesy. He teaches courses in electronic circuits and micromachined transducers. He held the Robert N. Noyce Family Faculty Scholar Chair in 1992±1994, received an NSF Young Investigator Award in 1993, was a Terman Fellow in 1994±1997, was a University Fellow in 1996±1998, and was appointed to the Defense
Sciences Research Council in 1995 (he is currently Chairman). He is a Fellow National of the Explorers Club. His present research areas include sensors and actuators, medical instruments, biotechnology, and micro fabrication, all with emphasis on solving practical problems. He has numerous patents and technical publications including the text, ``micromachined transducers sourcebook'' (WCB/McGraw-Hill, 1998).