Micromachined sensing module for pO2, pCO2, and pH and its design optimization for practical use

Sensors and Actuators B 76 (2001) 565±572

Micromachined sensing module for pO2, pCO2, and pH and its design optimization for practical use Hiroaki Suzukia,*, Taishi Hirakawaa, Takuo Hoshib, Hidenori Toyookab a

Institute of Materials Science, University of Tsukuba, 1-1-1 Tennodai, Tsukuba, Ibaraki 305-8573, Japan Institute of Clinical Medicine, University of Tsukuba, 1-1-1 Tennodai, Tsukuba, Ibaraki 305-8575, Japan

b

Abstract The performance of a sensing module for pO2, pCO2, and pH was tested from a practical point of view and its design was modi®ed for the pO2 electrode and the pH combination electrode. Owing to the elongated lifetime of the on-chip liquid-junction reference electrode, pH measurement could be conducted repeatedly. The outputs from these electrodes taken in a 4.6 ml ¯ow channel on the chip and in a 100 ml beaker were compared. Although a substantial difference was observed between the measured currents of the pO2 electrode with a 0:2 mm  1:2 mm cathode, the difference was reduced markedly by using a 25 mm  25 mm cathode. A good coincidence was observed between the values of potential obtained with the pH electrode, while the data with the pCO2 electrode suggested that it needed a substantial modi®cation of design depending upon its applications. The sensing module was actually used to measure blood gas levels in whole human blood and the values were compared with those obtained by a commercial blood gas analyzer. The deviation of pH values was within 0.06 pH units and pO2 values were within an acceptable level. However, substantial deviations of pCO2 values were observed. # 2001 Elsevier Science B.V. All rights reserved. Keywords: Oxygen; Carbon dioxide; pH; Ag/AgCl; Liquid junction; Micromachining

1. Introduction The measurement of blood gases (pO2, pCO2, and pH) is one of the most critical targets of the micro total analysis system (mTAS) research [1]. Possible advantages of the microsystem will be that (1) sample volume can be reduced to several ml±sub ml, (2) measurement can be conducted on the spot, and (3) it gives us ¯exibility in choosing the style of the measurement (in-line, on-line, or off-line) along with the other advantages mentioned in the literature. In miniaturizing the system the electrochemical approach is promising because it can provide sensors of high performance owing to the accumulated knowledge and can realize very simple instrumentation compared with the other module of detection. Along with integrated sensor chips [2±4], microsensors have been incorporated in a microfabricated ¯ow channel [5±8]. Furthermore, some of them have actually been used in real blood-gas analysis [4,7]. As such, the development of microsensors and microsystems is steadily advancing. However, in reality, it is very dif®cult to realize a micro system on a tiny silicon and/or

* Corresponding author. Tel.: ‡81-298-53-5598; fax: ‡81-298-55-7440. E-mail address: [email protected] (H. Suzuki).

glass chip incorporating all necessary functions contained in a commercial blood analyzer. A good solution is to decompose the entire system into several modules with different functions such as sensing, pumping, and signal processing [9,10]. We have already reported a micromachined sensing module for pO2, pCO2, and pH [11]. The structure and fabrication are simple and compatible with micromachining processes, which makes the incorporation of the module in a more sophisticated system easier. In the development of micro potentiometric sensors and microsystems which incorporate them, an impediment has been the integration of a liquid-junction reference electrode, which our module has realized. Although its basic performance has already been evaluated and distinct responses have been obtained [11], a couple of points have remained as issues in upgrading its performance to a practical level. The ®rst point is the in¯uence of the oxygen consumption during the operation of the pO2 electrode. Any amperometric sensors including the pO2 electrode consume corresponding analytes by an electrochemical reaction. This generates a diffusion layer in the vicinity of the working electrode. An ill-designed electrode has a diffusion layer expanding far into the sample solution, thus, gives an output containing a substantial error [12]. The second point is the lifetime of the on-chip liquid-junction

0925-4005/01/$ ± see front matter # 2001 Elsevier Science B.V. All rights reserved. PII: S 0 9 2 5 - 4 0 0 5 ( 0 1 ) 0 0 6 3 4 - 7

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reference electrode. Because of its limited lifetime, pH measurement had to be ®nished in a very limited lifetime. In this study, these points were improved by modifying the design of the module. The performance of the improved sensing module and its application to blood gas analysis will be discussed in this report. 2. Experimental 2.1. Materials Glass wafers (No.7740, 300 , 500 mm thick) were purchased from Corning Japan, Tokyo. (1 0 0)-oriented silicon wafers (300 , 380 mm thick) were purchased from Shin-Etsu Chemical, Tokyo. A positive photoresist, S1400-31, and a negative photoresist, OMR-83, used for fabrication were purchased from Shipley Far East, Tokyo, and Tokyo Ohka Kogyo, Kawasaki, Japan, respectively. A photocurable polyimide prepolymer (Photoneece UR-3100) was a gift from Toray, Tokyo. A one-component RTV silicone, KE-42-TS, used for the gas-permeable membrane was purchased from Shin-Etsu Chemical. A photocurable hydrophilic resin, ENT-2000, was purchased from KANSAI PAINT, Osaka. A photocurable adhesive (BENEFIX PC) was purchased from Adell, Tokyo. Poly(2-hydroxyethyl methacrylate) (polyHEMA) was purchased from Aldrich Chem. The other reagents used in fabricating the device were of semiconductor grade and purchased from Kanto Chemicals, Tokyo and Wako Pure Chemicals Industries, Osaka. All reagents used to examine the device performance were purchased from Wako Pure Chemicals Industries. They were of analytical reagent grade and used without further puri®cation. Distilled water was used throughout the experiments. Arterial and venous blood samples were originally taken from patients to check blood gas levels during anesthesia. Remained samples were used in our experiment. The pO2, pCO2, and pH levels were measured again prior to the experiments using a commercial blood analyzer (CIBACORNING 860). 2.2. Structure and fabrication of the integrated module The module with microsensors for pO2, pCO2, and pH was fabricated by ordinary processes of photolithography and micromachining [11]. The planar dimensions of the chip were 19.2 mm wide and 17.0 mm long. Detecting electrode patterns were formed on a glass substrate (Fig. 1a). All metal layers were sputter-deposited. A 200 nm-thick gold layer with a 40 nm-thick chromium adhesive layer was used as a backbone layer, on which a silver cathode pattern, indicator electrode patterns, and Ag/AgCl patterns were formed. The thickness of these patterns was 300 nm. In delineating the active areas of the detecting electrodes and insulating leads, a 2.7 mm-thick polyimide layer was formed. For the pCO2 electrode and the pH combination electrode, a so-called

anodic iridium oxide ®lm (AIROF) was formed as a pHsensing element by cycling the potential of a thin-®lm iridium pattern 2000 times in 1.0 M LiClO4 solution between 0.9 and ‡1.0 V (versus Ag/AgCl) at 0.5 V s 1. The basic structure and dimensions of the Ag/AgCl element were the same for all sensors. A silver thin-®lm pattern was covered with the polyimide layer leaving six slits at the edges of the silver pattern and AgCl layers were grown from there. A ¯ow channel and micro containers for the electrolyte solutions for the respective electrodes were formed on a silicon substrate (Fig. 1b and c) by anisotropically etching it from both sides with the resulting depth of grooves of 190 mm. The volume of the micro container was 2.9 ml for the pO2 electrode, pCO2 electrode, and the liquid-junction reference electrode for the pH combination electrode, while that of the ¯ow channel was 4.6 ml when it was closed with a cover glass. Rectangular through-holes were formed at the crossings between the ¯ow channel and the micro containers for the pO2 and pCO2 electrodes, where gaspermeable membranes were formed. A through-hole with the same dimensions was formed for the pH indicator electrode. For the pO2 and pCO2 electrodes, two throughholes were formed at the ends of the micro containers to introduce electrolyte solutions (0.1 M KCl/0.1 M Tris-HCl buffer solution (pH 8.5) for the pO2 electrode and 20 mM NaHCO3/0.1 M KCl for the pCO2 electrode). The glass and silicon substrates were bonded with the photocurable adhesive (Fig. 1c). Eight modules were batch-fabricated on a 300 wafer and were diced into chips. Two portions were modi®ed compared with our previous model. First, the dimensions of the active area for the cathode of the pO2 electrode were changed from 1:2 mm  200 mm to 25 mm  25 mm. Second, the structure of the liquid-junction reference electrode was changed. Compared with the previous model, the width of the slit was narrowed from 100 to 50 mm and the ENT-2000 resin was applied to the slit areas to suppress the dissolution of AgCl more effectively (Fig. 1d). The structure of the liquidjunction was also modi®ed (Fig. 1e). A 600 mm square area was etched from the container side, and when the etching completed, an approximately 350 mm square through-hole was formed at the bottom of the ¯ow channel. The square cavity on the container side was connected with the container cavity with a groove of 100 mm in width and 500 mm in length. The square cavity and groove in the junction were used either with or without a polyHEMA restraining layer. Its complete ®lling was easily achieved by capillary action with a methanolic solution of 5% polyHEMA. Furthermore, the shape of the micro container was modi®ed keeping the internal volume the same as the previous model, and KCl was stored in the micro container in a solid form as much as possible. In order to do that, ®ne KCl powder ground with a mortar and a pestle was introduced from the through-hole near the pad in such a way that the powder ®lled the interior of the container cavity completely. To activate the electrode,

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Fig. 1. Structure of the integrated sensing module. (a) Glass substrate on which the detecting electrode patterns are formed. (b) Silicon substrate on which the container cavities for the electrolyte solutions are formed. The surfaces of (a) and (b) are bonded to complete the sensing module (c). The other side of the silicon substrate (b) can be seen. The gas-permeable membranes are formed at the bottom of the flow channel. (d) Thin-film Ag/AgCl element used in this module. The AgCl is formed in the black areas, (e) Structure of the liquid junction for the pH-combination electrode.

saturated KCl/AgCl solution was injected from the throughhole prior to use. 2.3. Other procedures When the performance of a single electrode was examined, only the electrode was cut out from the wafer. Corresponding electrolyte solution was introduced into the microcontainer by immersing the chip in the electrolyte solution in a plastic centrifuge tube, placing it in a chamber, and evacuating it. In evaluating the integrated module with the three electrodes, the electrolyte solutions were introduced into the corresponding micro containers using a microsyringe. Bubbles were removed by tapping one end of the chip on the table. Basic performance of the discrete electrodes was examined by immersing the sensitive area in an appropriate 100 ml solution ®lled in a 100 ml beaker. Here, one of the through-holes in the respective containers was closed with silicone rubber. The performance was also examined using the 4.6 ml ¯ow channel on the silicon

substrate. A cover glass was placed over the channel to minimize the error caused by the exchange of gasses between the atmosphere. The cathode of the pO2 electrode was poised at 0.8 V versus the internal Ag/AgCl anode using a Hokuto-Denko HA-151 potentiostat/galvanostat. The potential of the AIROF indicator electrode in the pCO2 electrode was measured using the internal thin-®lm Ag/AgCl reference electrode. In examining the performance of the pH indicator electrode, either the on-chip liquid-junction reference electrode or a macroscopic Ag/AgCl reference electrode (Horiba 2080A-06T with a ceramic-plug junction, internal ®lling solution: saturated with KCl and AgCl) was used. The opencircuit potential of the respective indicator electrodes was measured using a Hokuto-Denko HE-106 electrometer (input impedance: 1  1013 O). The outputs from these electrodes were recorded on a TOA Electronics PRR5011 strip-chart recorder. In preparing solutions with a predetermined pO2, pCO2, and pH values, pure oxygen and carbon dioxide gases were

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mixed with pure nitrogen gas in an appropriate ratio. The mixed gas was humidi®ed and was bubbled through the stirred solution. The pO2 and pCO2 values were calculated from their percentage in the mixed gas taking the barometric pressure at the time of the experiment and pH2O into consideration. The pH of the solution was measured accurately using a TOA Electronics HM-20S pH meter with a combination pH glass electrode previously calibrated in precision buffers of pH 4.01 and 6.86. In conducting the measurement in the ¯ow channel on the module, the entire module was placed on a hotplate controlled at 37.08C. In our previous study, the data were taken at 25.08C considering possible applications in general. Therefore, basic performance of the module was characterized at 25.08C and the temperature was raised to 37.08C for the experiment using whole blood samples. The precision of the temperature control was within 0.18C. In order to determine the pO2, pCO2, and pH values of whole blood samples, calibration solutions were prepared with 100 ml of 5 mM K2HPO4 solution, which had combinations of pO2 and pCO2 values of 149.0 and 74.6 or 0 and 149.0 mmHg, respectively. The current of the pO2 electrode at zero pO2 was also checked by adding excess Na2SO3 to the solution. The resulting pH values were carefully examined every time and were settled around 7.7 and 7.5 for each case. In conducting the measurement, the three electrodes on the module were calibrated with the above-mentioned standard solutions. One of the solutions was ®lled in the ¯ow channel and the outputs of the electrodes were recorded. After the ¯ow channel was rinsed with a phosphate buffered saline, the same procedure was repeated with another calibration solution and the blood samples. In carrying out the measurement along with the other experiments to characterize the performance of the module, the liquid junction of the on-chip reference electrode for the pH combination electrode was not ®lled with polyHEMA due to the reason mentioned later. 3. Results and discussion 3.1. Basic performance of the modified pO2 electrode By changing the dimensions of the cathode of the pO2 electrode, several aspects of its performance change. These include sensitivity, response time, and ¯ow dependence. In Fig. 2, calibration curves obtained with the 1:2 mm 200 mm cathode and the 25 mm  25 mm cathode at 258C are shown. Linear relationships were observed with both cathodes. Although the area of the smaller cathode was only 0.26% of that of the larger cathode, the current level did not decrease in proportion to the area. The area of the smaller cathode is in the range of the so-called microelectrode and the current will change in proportion to the length of the edge of the cathode rather than its area. Sensitivity was 3.3 nA/mmHg with the larger cathode and 0.06 nA/mmHg

Fig. 2. Calibration curves for the pO2 electrode. (a) With the 1:2 mm  0:2 mm cathode at 258C, (b) with the 25 mm  25 mm cathode at 258C, (c) with the 25 mm  25 mm cathode at 378C. The inset shows the expanded graphs of (b) and (c). The pO2 values were measured in a 50 mM KH2PO4±NaOH buffer solution (pH 7.0). Applied voltage: 0.8 V.

with the smaller cathode. For the pO2 electrode with the smaller cathode, the calibration curve was also taken at 378C. Although noticeable temperature dependence was observed, a good linear relationship was obtained again. The sensitivity was 0.17 nA/mmHg. The response was stable and distinct without any ¯uctuations or noise. The response time will change by making the cathode smaller, which was actually the case (Table 1). The pO2 electrode with the larger cathode gave faster response to both the increasing and decreasing step changes of pO2. As pointed out in our previous study, a residual current of several nA was observed. This is probably due to the in¯ux of oxygen from the portions in the micro container other than the sensitive area. A signi®cant improvement was observed in terms of the ¯ow dependence of the output current (Table 1). This is the result of the diffusion layer shrunk within the gas-permeable membrane. Table 1 Effect of decreasing the dimensions of the cathode of the pO2 electrodea Cathode area

0.2 mm  1.2 mm 25 mm  25 mm

Flow dependence (%)

90% response time (s) 0 ! Air saturated

Air saturated ! 0

15.3±19.6 1.4±2.4

23±45 33±51

23±50 43±69

a In determining the flow dependence, the output current was measured in a vigorously stirred solution (current: A) and a quiescent solution (current: B) and (B A)/A was calculated. In the former case, the diffusion layer was assumed to be completely disturbed. The 90% response time to decreasing pO2 was measured by adding Na2SO3, while the response time to the opposite direction was measured by moving the pO2 electrode chip to an air-saturated buffer solution. In conducting the experiment, three normally functioning chips were randomly selected.

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3.2. Durability of the novel thin-film Ag/AgCl element and lifetime of the novel liquid-junction reference electrode Durability of the novel thin-®lm Ag/AgCl element with 50 mm-wide slits was checked in a solution saturated with KCl and AgCl. In conducting this experiment, 80% of the silver layer had been converted into AgCl. It has been suggested that the durability of the element would be improved by covering the slits with a hydrophilic polymer layer [13]. Therefore, the ENT-200 resin was ®lled in the slit areas. The effect was evident (Fig. 3). By just making the slits narrower, the lifetime was approximately doubled (Fig. 3b). Furthermore, coating the slits with the ENT2000 resin made the lifetime longer. The longest one functioned nearly 20 h (Fig. 3c). However, compared with our similar thin-®lm Ag/AgCl element with only one slit of the same width at the center of the element, lifetime was shorter [13]. This suggests that it is more advantageous to reduce the number of slits and form the AgCl layer more deeply into the silver layer. A critical point to elongate the lifetime of the liquidjunction reference electrode is the structure of the liquid junction. An ill-designed junction cannot suppress the effusion of internal KCl even if it is stored in a solid form [11]. Junctions with and without the restraining material were compared. When polyHEMA was ®lled in the cavity of the junction, a substantial liquid-junction potential exceeding 10 mV was observed. Furthermore, it depended upon the amount of polyHEMA and increased to 30±40 mV when a substantial portion of the junction was ®lled with it. This is clearly because the effusion of internal KCl was blocked and the liquid-junction potential was not eliminated effectively. Therefore, the same experiment was repeated without the restraining material. To our surprise, not only the liquidjunction potential was substantially eliminated but also the reference electrode could maintain a stable potential within 1 mV for approximately 18 h (Fig. 3d), in marked contrast

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to the previous liquid-junction reference electrode which functioned for approximately 1 h [11]. The result indicates that the solid KCl precipitated around the edge of the junction served to restrain the effusion of the internal electrolyte. 3.3. Basic performance of the pH and pCO2 electrodes The basic performance of the pH and pCO2 electrodes at 258C has already been reported [11]. The response pro®les and calibration curves at 378C were also examined. However, no substantial difference was observed except for a change in the slope of the calibration curves. When the electrodes are used for practical purposes, the stability of the potential of both the indicator electrode and the reference electrode will be required. As already mentioned, the drift of the potential of the liquid-junction reference electrode was within 1 mV during its lifetime. When the thin-®lm Ag/ AgCl element is used in a solution containing NaHCO3, deterioration of the AgCl layer is enhanced [14]. However, within the lifetime of the entire sensing module, no substantial drift of potential was observed. The stability of the AIROF used as the indicator electrode for the pH and pCO2 electrodes was also checked (Fig. 4). The drift of the potential was ‡0.38 mV/h on average in the acid region and ‡0.56 mV/h in the neutral region. However, the potential became unstable in the alkaline region, which agrees with the data given in the literature [15]. As long as the module is used in a solution of pH around 7.4 for the pH measurements or in the internal electrolyte solution for the pCO2 electrode with pH around 8.7, no substantial errors will be produced. 3.4. Effect of sample volume on the output of respective electrodes It is possible that an amperometric sensor like the pO2 electrode gives a different output to the same amount of

Fig. 3. Variation of the potential of the thin-film Ag/AgCl element ((a)±(c)) and the completed on-chip liquid junction reference electrode (d). (a) Slit width: 100 mm, (b) slit width: 50 mm, (c) slit width: 50 mm with the ENT-2000 protecting layer. The curves (a)±(c) were taken in a saturated KCl±AgCl solution and the curve (d) was taken in a 50 mM KH2PO4±NaOH buffer solution (pH 7.0) at 258C. About 80% of the silver layer had been converted into AgCl. A macroscopic commercial Ag/AgCl reference electrode (Horiba 2080A-06T) was used as a potential standard.

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Fig. 4. Variation of the potential of the AIROF indicator electrode in solutions of various pH. The data were taken in (a) 50 mM citrate±NaOH buffer (pH 5.1), (b) 50 mM KH2PO4±NaOH buffer (pH 7.0), (c) 50 mM borate±NaOH buffer (pH 10.1) at 258C. A macroscopic commercial Ag/AgCl reference electrode (Horiba 2080A-06T) was used as a potential standard.

analyte depending upon the sample volume, because the expansion of the diffusion layer following the consumption of the analyte by the electrochemical reaction will not be neglected in a minute volume. On the other hand, detection of potentiometric sensors including the pCO2 electrode and the pH combination electrode is based on an equilibrium on the solution/membrane interface and does not consume the analyte. Therefore, it is expected that the same reading is obtained irrespective of sample volume. However, the output potential of the pCO2 electrode might be affected because it requires the transport of carbon dioxide from the sample solution to the internal electrolyte layer. The transport can also occur between the sample solution and the internal electrolyte solutions for the other electrodes. To examine these effects, the output of the electrodes was measured in a buffer solution of 100 ml and 4.6 ml. One of the electrodes was ®rst stabilized in 100 ml of the solution and the output was measured. Then, the module was placed on a hotplate maintained at the same temperature and the solution in the beaker was ®lled in the ¯ow channel on the silicon substrate. In order to minimize errors caused by the escape of the analyte gases, a cover glass was placed on the silicon substrate. The on-chip reference electrode was used for the pH measurement in both cases. Table 2 summarizes the result. The difference in potential was very small for the pH combination electrode, which con®rmed that the electrode gave accurate data even with a very small amount of sample solution. Unnegligible deviation was observed between the values measured with the pCO2 electrode. This is considered to be the result of the transport of carbon dioxide from the ¯ow channel to the microcontainers. Concerning the pO2 electrode, values were compared between the 0:2 mm  1:2 mm cathode and the 25 mm  25 mm cathode. A signi®cant difference exceeding 30% was observed when the 1:2 mm  200 mm cathode was used. The deviation was markedly reduced with the

Table 2 Effect of sample volume on the output of respective sensorsa Sample volume

100 ml

4.6 ml

pH electrode pCO2 electrode pO2 electrode 0.2 mm  1.2 mm cathode 25 mm  25 mm cathode

177.3 mV 398.5 mV

175.8 mV 383.5 mV

680 nA 15.0 nA

434 nA 14.5 nA

a Sample solution was 50 mM KH2PO4±NaOH. pO2: 246 mmHg, pCO2: 246 mmHg, pH: 6.37. The pH of the buffer solution was originally 7.00 but was lowered to the above value by bubbling carbon dioxide.

25 mm  25 mm cathode. It is considered that the transport of oxygen between the ¯ow channel and the micro container had little in¯uence on the output of the pO2 electrode. The pO2 in the electrolyte solution is usually higher than that in sample solution and oxygen is constantly supplied from the through-holes to introduce the electrolyte solution. Therefore, because expansion of the diffusion layer is approximately comparable to the size of the cavity of the sensitive area for the 25 mm  25 mm cathode [16], errors caused by the transport of oxygen will not be substantial. 3.5. Application to blood gas analysis The accuracy of measurement by the integrated module was tested using whole human blood samples. The pO2, pCO2, and pH values of the samples had been checked with the commercial blood analyzer prior to the experiment. Table 3 shows the values obtained by the module and the blood analyzer. In measuring with the module, the calibration solution or the blood sample was ®lled in the ¯ow channel on the silicon substrate and was covered with a cover glass. Also, the on-chip liquid-junction reference electrode was used for the pH measurement. Coincidence

H. Suzuki et al. / Sensors and Actuators B 76 (2001) 565±572 Table 3 Comparison of the pO2, pCO2, and pH values obtained by the sensing module (a) and a commercial blood analyzer (CIBA-CORNING 860) (b)a Sample

1 2 3 4 5 6 7 a b

pO2 (mmHg)

pCO2 (mmHg)

pH

a

b

a

b

a

b

135.5 150.0 45.0 126.5 161.8 ±b ±

120.4 150.2 45.2 120.2 158.2 ± ±

32.7 51.1 57.4 62.1 15.9 10.3 5.8

39.7 36.5 46.4 35.7 33.7 47.8 35.1

7.49 7.48 7.41 7.41 7.50 7.36 7.48

7.47 7.49 7.42 7.47 7.50 7.39 7.44

The experiment was conducted at 378C. Data not taken due to malfunctioning of the pO2 electrode: ±.

of the values depended upon the electrodes and was excellent for pH. Deviation of the values was 0.06 even for the worst case. Coincidence of the pO2 values was also satisfactory except for a couple of cases. However, compared with these, deviations could not be neglected with the pCO2 values. This is again considered to be the result of the transport of carbon dioxide. Note that the pCO2 values obtained by the module can take values both higher and lower than those obtained by the analyzer, because the transport can also occur during the calibration procedure. By modifying the structure of the sensitive area, the accuracy of the measurement will be improved. 3.6. Future directions Concerning the durability of the thin-®lm Ag/AgCl element, the number of slits correlated directly with the durability of the element. If we convert the same proportion of silver layer into AgCl, fewer slits would be preferable. We used the slit con®guration considering the decrease in conductivity. However, once the AgCl layer is grown into the silver layer, the geometry of the interface between silver and AgCl determines the conductivity and the slit only serves as a liquid junction. Therefore, if the Ag/AgCl element is used in the pH combination electrode or the pCO2 electrode under a potentiometric mode, arrays of pinholes will be better than the slit con®guration. Note that this discussion only applies to the potentiometric sensors. When the Ag/AgCl element is used as the anode for the pO2 electrode, the conductivity of the element can become a limiting factor. As the dimensions of the cathode decrease, the in¯uence of conductivity decreases. The length or number of the slits should be optimized in relation to the generated current in order to measure the pO2 range which is of interest in the application. Another critical point is the design of the micro containers. As evidenced by the obtained data, the structure of the sensitive area should be modi®ed to make the response of the pO2 and pCO2 electrodes faster. It would also decrease the residual current of the pO2 electrode and the error caused by the transport of carbon dioxide between the ¯ow channel and the micro containers.

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4. Conclusions The integrated sensing module was examined for the measurement of samples of a minute volume. For the pO2 electrode, accuracy was markedly improved by using a 25 mm  25 mm cathode. The elongated lifetime of the liquid-junction Ag/AgCl reference electrode made pH measurements easier. The pO2, pCO2, and pH values of whole blood samples obtained with the module were compared with those obtained by a blood gas analyzer. As expected, good coincidence was observed for pO2 and pH. However, noticeable discrepancies were observed for pCO2. This is probably due to the transport of carbon dioxide between the sample solution and the internal electrolyte solutions. To upgrade the performance further, a substantial modi®cation might be necessary including the design of the micro container, which will be our next issue. Acknowledgements The authors are grateful for the ®nancial support of the Research for the Future Program of JSPS, Tateisi Science and Technology Foundation, and the Mikiya Science and Technology Foundation.

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[13] H. Suzuki, H. Shiroishi, S. Sasaki, I. Karube, Microfabricated liquid junction Ag/AgCl reference electrode and its application to a onechip potentiometric sensor, Anal. Chem. 71 (1999) 5069±5075. [14] H. Suzuki, H. Arakawa, S. Sasaki, I. Karube, Micromachined Severinghaus-type carbon dioxide electrode, Anal. Chem. 71 (1999) 1737±1743. [15] M.L. Hitchman, S. Ramanathan, Evaluation of iridium oxide electrodes formed by potential cycling as pH probes, Analyst 113 (1988) 35±39. [16] I. Fatt, An ultramicro oxygen electrode, J. Appl. Physiol. 19 (1964) 326±329.

Biographies Hiroaki Suzuki received his BE and ME degrees in applied physics and his PhD degree in bioelectronics and biotechnology from The University of Tokyo in 1981, 1983, and 1993, respectively. In 1983, he joined Fujitsu Laboratories Ltd., Japan. In 1996, he moved to the Institute of Materials Science, University of Tsukuba, Japan, where he became associate professor of materials science. His current research interests include micromachining, micro bio/chemical sensors, and mTAS.

Taishi Hirakawa received his BE and ME degrees in materials science from University of Tsukuba. He had studied micro bio/chemical sensors in the laboratory of Associate Prof. Suzuki. Takuo Hoshi received his MD degree from Toyama Medical and Pharmaceutical University in 1993. In 1993, he started anestesiology resident at Kitasato University Hospital. In 1994, he moved to Tsukuba University Hospital and continued his practice. In 1999, he finished anesthesiology resident course at Tsukuba University Hospital and he entered graduated school of medicine, University of Tsukuba to research about anesthesiology and intensive care. Hidenori Toyooka received his MD and PhD degrees in Medicine from Tokyo University in 1967 and 1980, respectively. From 1974 he took residency in Department of Anesthesiology, Yale-New Haven Hospital. In 1978, he returned to Tokyo University Hospital as a lecturer in Department of Emergency Medicine. From 1986 he worked in Tokyo Medical and Dental University as an associate professor in Department of Anesthesiology and Critical Care Medicine. In 1996, he moved to the Institute of Clinical Medicine, University of Tsukuba, as a professor in Anesthesiology. His current research interests include respiratory physiology in Critical Care Medicine.