Anal. Chem. 1997, 69, 4032-4038
Miniaturized Ion-Selective Chip Electrode for Sensor Application Albrecht Uhlig,*,† Erno 1 Lindner,‡ Cristian Teutloff,† Uwe Schnakenberg,† and Rainer Hintsche†
Fraunhofer Institut fu¨ r Siliziumtechnologie; Fraunhoferstrasse 1, 25524 Itzehoe, Germany, and Institute for General and Analytical Chemistry, Technical University of Budapest, 1111 Budapest, Szent Gelle´ rt te´ r 4, Hungary
The performance of miniaturized potentiometric cells, with multilayer, planar ion-selective sensors in aqueous electrolyte solutions, human serum, urine, and whole blood, is presented. The basic steps of the fabrication with silicon technology are summarized. The effect of the contact surface between the internal reference system and the ion-sensitive membrane on the analytical characteristics of potassium- and calcium-sensitive sensors is studied. Silicone rubber-, high molecular weight PVC-, carboxylated PVC and aliphatic polyurethane (Tecoflex)based solvent polymeric membranes were dispensed into anisotropically etched wells on silicon wafers, and the resulted planar sensors were tested in terms of their ion sensitivity (slopes of the cell voltage-pK or pCa calibration curves), long-term stability, and reproducibility. For the assay of potassium in whole blood, the miniaturized potentiometric cell was built in a flow-through manifold. To achieve the required precision, the flow conditions were optimized and the sensors calibrated periodically. The results prove the feasibility of the new sensor design and satisfy the particularly difficult requirements for the analysis of biological samples. Ion-selective electrode (ISE) potentiometry is an elegant method of direct ion activity measurement in complex matrices.1 Various shapes of ion-selective electrodes and related sensors are widely used in blood gas analyzers for the measurement of clinically relevant ions and molecules.2-11 Ionophore cocktails for about 10 ions are commercially available, and new kinds are under †
investigation.12 However, in spite of the analytical and commercial success of polymeric membrane sensor-based analyzers, the in vivo acute and chronic application of such devices remains a serious challenge. Time-dependent ion activity measurements at the place of origin, i.e., single cells, on the surface of an organ, or even in the myocardial tissue, still attract considerable interest.6 In our previous paper, we reported on a valinomycin-based potassium-selective planar sensor fabricated on a silicon wafer with an integrated Ag/AgCl/p-HEMA (poly(hydroxyethyl methacrylate)) reference electrode, which was implanted into an 18 gauge catheter.13,14 This paper reports on the optimization of the planar potentiometric device. The transducer was fabricated with bulk silicon micromachining techniques using a double-sided wafer process combined with polymeric membrane coatings. Potentiometric devices having solid-state contact and using hydrogel as an inner liquid electrolyte solution were tested. Four different types of membrane matrix materials were used: high molecular weight PVC (PVC-HMW), carboxylated PVC (PVC-COOH), silicone rubber, and aliphatic polyurethane (Tecoflex).15,16 Three ionophores were utilized as model compounds: two with potassium and one with calcium selectivity. The sensors were tested in aqueous solutions, human serum, urine, and whole blood. The analytical characteristics of the sensors with different layer structures are compared on the basis of sensitivity (slope of the cell voltage-log ion activity functions), short- and long-term stability, reproducibility of the potential measurements, and response time. EXPERIMENTAL SECTION Chemicals. Chemicals were analytical grade (Merck, Fluka) or better. All solutions were made with deionized water. For slope and long-term stability measurements, appropriate amounts of KCl or CaCl2 were added to a 0.147 mol/L NaCl solution (Aldrich-Chemie, Steinheim, Germany). Moni-Trol I (normal) and Moni-Trol II (pathologic) were used for serum measurements (Baxter Deutschland GmbH, Ettlinger, Germany17). Whole blood was measured immediately after delivery using a lithium-heparinized syringe (Sarstedt, Nu¨mbrecht, Germany) as sample container.
Fraunhofer Institut fu ¨ r Siliziumtechnologie. ‡ Technical University of Budapest. (1) Oesch, U.; Ammann, D.; Simon, W. Clin. Chem. 1986, 32, 1448-1459. (2) Lewenstam, A.; Maj-Zurawska, M.; Hulanicki, A. Electroanalysis 1991, 3, 727-734. (3) Smith, R. L.; Scott, D. C. IEEE Trans. Biomed. Eng. 1986, BME-33 (No. 2), 83-90. (4) Prohaska, O. J.; Kohl, F.; Goiser, P.; Olcaytug, F.; Urban, G.; Jachimowicz, A.; Pirker, K.; Chu, W.; Patil, M.; LaManna, J.; Vollmer, R. Transducers 1987, 812-815. (5) Malinkowa, E.; Oclejas, V.; Hower, R. W.; Brown, R. B.; Meyerhoff, M. E. Enhanced electrochemical performance of solid state ion sensors based on silicon rubber. Technical Digest Transducer ‘95, Eurosensors IX, Stockholm, Sweden, 25-29 Sept 1995; pp 851-854 (Chairman, I. Lundstro ¨m). (6) Lindner, E.; Cosofret, V. V.; Kusy, R. P.; Buck, R. P.; Rosatzin, T.; Schaller, U.; Simon, W.; Jeney, J.; Toth, K.; Pungor, E. Talanta 1993, 40 (7), 957967. (7) Janata, J.; McBride, P. T. J. Bioeng. 1978, 2, 459. (8) Reinhoudt, D. N.; Engbersen, J. F. J.; Brzozka, Z.; van den Vlekkert, H. H.; Honig, G. W. N.; Holterman, H. A. J., Verkerk, U. H. Anal. Chem. 1994, 66, 3618-3623. (9) Knoll, M.; Cammann, K.; Dumschat, C.; Eshold, J.; Sundermeier, C. Sens. Actuators 1994, B21, 71-76. (10) Bezegh, K.; Bezegh, A.; Janata, J.; Oesch, U.; Xu, A.; Simon, W. Anal. Chem. 1987, 59, 2846-2848. (11) Erickson, K. A.; Wilding, P. Clin. Chem. 1993, 39, 283-287.
(12) Selectophore, Ionophores for Ion-Selective Electrodes and Optrodes; Fluka: Buchs, Switzerland, 1991; p 31. (13) Hintsche, R.; Kruse, Ch.; Uhlig, A.; Paeschke, M.; Lisec, L.; Schnakenberg, U.; Wagner, B. Chemical microsensor systems for medical applications. Eurosensors VIII, Toulouse, France, Sept 1994; pp 376-379 (Chairman, G. Blasquez). (14) Uhlig, A.; Schnakenberg, U.; Lindner, E.; Dietrich, F.; Hintsche, R. Sens. Actuators 1996, B34 (1-3), 252-257. (15) Satchwill, T.; Harrison, D. J. J. Electroanal. Chem. 1986, 202, 75-81. (16) Cha, S. G.; Liu, D.; Meyerhoff, M. E.; Cantor, A. C.; Goldberg, H. D.; Brown, R. B. Anal. Chem. 1988, 63, 1666-1672. (17) Baxter M+D Moni-Trol, Kontrollserum zur Qualititssicherung in der klinischen Chemie.
4032 Analytical Chemistry, Vol. 69, No. 19, October 1, 1997
S0003-2700(96)00957-2 CCC: $14.00
© 1997 American Chemical Society
Membrane Cocktails. Four different membrane formulations were used. (a) carboxylated PVC (PVC-COOH)-based K+-selective membrane: 1 wt % BME-44,18 65.5 wt % bis(1-butylpentyl)decane1,10-diyl diglutarate, 0.5 wt % potassium tetrakis(4-chlorophenyl)borate, and 33 wt % carboxylated poly(vinyl chloride) (all Selectophore, Fluka Chemie AG, Neu-Ulm, Switzerland). (b) high molecular weight PVC (PVC-HMW)-based Ca2+selective membrane: 3.3 wt % ETH 1001, 63.7 wt % bis(1butylpentyl)decane-1,10-diyl diglutarate, 2.1 wt % potassium tetrakis(4-chlorophenyl)borate, and 30.9 wt % high molecular weight poly(vinyl chloride) (all Selectophore, Fluka). (c) silicone rubber (SR)-based K+-selective membrane: 2.5 wt % valinomycin, 83 wt % Silopren K 1000, and 14.5 wt % Silopren cross-linking agent K 11 (all Selectophore, Fluka). (d) aliphatic polyurethane-based K+-selective membrane: 49 wt % Tecoflex 85A (Thermedics, Inc., Woburn, MA), 49 wt % 2-nitrophenyl octyl ether, 1.2 wt % BME-44, and 0.8 wt % potassium tetrakis(4-chlorophenyl)borate) (Selectophore, Fluka). The membrane components were dissolved in tetrahydrofuran (a,b) or dichloromethane (c,d) (Selectophore, Fluka). For the preparation of the p-HEMA layer, the following cocktail was used: 57 wt % hydroxyethyl methacrylate, 38 wt % ethylene glycol, 1 wt % dimethoxyphenylacetophenone, 2.5 wt % poly(vinylpyrrolidone) K90, and 1.5 wt % tetra(ethylene glycol) dimethacrylate19 (all Fluka Chemie AG). The solution was polymerized using UV light radiation for 60 s under a 125 µm thick Mylar foil (DuPont) to avoid quenching. Fabrication of the Sensor. Figure 1 shows a SEM photograph of the cross section of the silicon sensor chip, without the IS membrane, with the “front side” (sample side) on the top. The pyramidal-shaped sensor structure with the lattice-like opening to the front of the sensor chip (see also Figure 2e) has the following functions: (i) guarantees a small contact surface to the sample together with a relatively large membrane reservoir, (ii) increases the contact surface (adhesion) between the ion-sensitive membrane and the internal reference system, and (iii) protects and stabilizes the membrane against mechanical defects. The dimension of the chip is 1 mm × 5 mm × 0.5 mm. The size of the active surface of the IS membrane on the front of the sensor is about 0.03 mm2. The total volume of the IS membrane is about 0.011 mm3. The wafer was fabricated with bulk silicon technology using a double-sided polished 4 in. wafer with (100) orientation. The process for sensor production starts with thermal oxidation to form an insulating SiO2 layer (500 nm). The first lithography defines the gridlike grooves (upside-down pyramidal-shaped cavities) on the “front side” (sample side) of the wafer. After patterning off the oxide by reactive ion etching (RIE), the photomask is removed. The grooves were etched anisotropically in aqueous tetramethylammonium hydroxide (TMAH) solution (25 wt %, 84 °C) up to the full depth of the upside-down pyramidal groove (Figure 2a). The fabrication is then continued on the “back side” (internal reference electrode side) of the wafer. The second lithography defines the Ti/Pt interconnections between the bonding pads and (18) Lindner, E.; To`th, K.; Jeney, J.; Horva`th, M.; Pungor, E.; Bitter, I.; AÄ gai, B.; To¨ke, L. Microchim. Acta 1990, I, 157-168. (19) van den Berg, A.; Koudelka-Hep, M.; van der Schoot, B. H.; de Rooij, N. F. DECHEMA Monogr. Band 1989, 126, 155-171.
Figure 1. SEM photograph of the sensor chip before IS membrane assembly.
Figure 2. Schematics of the process flow of sensor fabrication (a-e).
the membrane wells (the large pyramidal grooves on the back side of the silicon wafer). Lift-off is used for structuring. First, a HF dip (10% at 25 °C) is performed to produce an undercut of the photoresist mask. After evaporation of the metals, the photoresist is removed with dimethylformamide, leaving Ti/Pt structures embedded in the oxide (Figure 2b). The third lithography defines the “back-side” groove. The oxide structuring is performed by RIE. After resist stripping, the groove is anisotropically etched in the TMAH solution. The etching process Analytical Chemistry, Vol. 69, No. 19, October 1, 1997
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was performed in a time-controlled fashion until the tip of the frontside pyramidal groove was pierced. The merging apexes of the two pyramidal grooves formed holes throughout the wafer (Figure 2c). The side walls of the “back-side groove” were passivated by thermal oxidation. The silicon process is finished with coating the walls of the “back-side” grooves with a Ag/AgCl/Ag (30 nm, 600 nm, 10 nm) layer by sputtering and thermal evaporation. In this step, a specially designed silicon wafer is used as aperture blend.20 There were openings etched into the wafer, used as aperture blend, that matched the “back-side groove” of the sensor chip-wafer, both in position and in size. This aperture blend was positioned on top of the sensor chip-containing wafer during the Ag/AgCl/Ag coating. Thus, an additional structuring of the Ag/AgCl/Ag could be avoided. An overlap to the Ti/Pt metalization surrounding the back-side groove provides the electrical contact to the bonding pads (Figure 2d). The very thin Ag layer (10 nm) on top of the Ag/AgCl was used to protect the quality of the AgCl during extended storage. This protective layer could be removed before the final steps of the sensor fabrication (before the IS membrane cocktail was applied into the sensor wells). After dicing, the individual chips are contacted on the back side by thick wire bonding. Next, the chips are immersed into 0.1 M NaCl to furnish the internal reference system with the necessary chloride concentration and to transform the thin protective Ag layer into AgCl. After cleaning and drying, the different IS membrane cocktails are dispensed into the back-side grooves of the chips, using a microsyringe. This procedure is repeated several times (typically, 0.2 µL of membrane cocktail is dispersed four times) with small breaks to let the solvent evaporate. Due to the capillary forces, the membrane cocktail flows through the holes to the front side, forming a solidified drop. Finally, the back side of the single sensor chips was insulated using Dow Corning 3140 RTV silicone rubber (Figure 2e). The ready-to-use sensor chips were equilibrated for 24 h in Ringer solution before use. Experimental Equipment. Screening of the electrodes was carried out by determining the slope of the calibration curve (EMF versus log aK or log aCa) by adding small increments of a KCl or CaCl2 standard solution, under stirring, to 10 mL of background electrolyte of a 0.14 M NaCl + 10-3 M KCl. Eight electrodes were tested at the same time using a home-made eight-channel potentiometric measuring unit in combination with the Origin 3.7 software. An Orion Model 90-02 Ag/AgCl double-junction reference electrode with 0.14 M NaCl outer filling solution was used as reference electrode. The long-term stability measurements were performed in flowing aqueous electrolytes, serum, urine, and whole blood. The sensors were placed into a Plexiglas flowthrough cell, 20 mm in length and 1.4 mm i.d. In the flow-through manifold, the cell voltages were measured against a Model 8-702 flow-through reference electrode (Microelectrodes, Inc., Londonderry, NH). The cell was equipped with a computer-controlled electrolyte management system consisting of a pump and two valves. The PC-controlled system allowed the control of pumping velocity, valve positions, and sampling rate via an AD/DA module. The automation and the experiment protocol ensured reproducible experimental conditions. In the long-term measurements, the sensors were flushed alternately with aqueous isotonic Ringer (20) Schnakenberg, U.; Lisec, T.; Hintsche, R.; Kuna, I.; Uhlig, A.; Wagner, B. Sens. Actuators 1996, B34 (1-3), 476-480.
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Table 1. Slope of the Calibration Curves of Planar Microelectrodes Fabricated with Different Membrane Formulations and Tested in the 10-4-10-1 mol/L Concentration Range with Constant 0.14 M NaCl Background matrix
ionophore
slope
polyurethane silicone PVC-COOH PVC-HMW
BME-44 valinomycin BME-44 ETH-1001
56 ( 4 mV/pK+ 56 ( 3 mV/pK+ 50 ( 4 mV/pK+ 26 ( 3 mV/pCa2+
solutions (single-point calibration) and the sample solution (aqueous electrolyte, human serum, urine, whole blood). The output signal of the sensors was measured using a home-made potentiometric measuring unit.14 The potential data were collected under stopped-flow conditions at two different concentrations (c1 and c2). The values recorded by the data acquisition system (Ec1 and Ec2) were the mean values of the last 100 data points, collected within a 200 s time period in the respective sample. In the case of serum, urine, and whole blood measurements, the mean values were calculated from 50 data points, collected in the last 100 s of a 200 s measuring cycle. The potassium concentrations of the whole blood samples were deduced from the calibration curve of the sensors in standard serum samples. RESULTS To demonstrate the performance of the new sensor design, the silicon-made transducer was combined with four types of membrane matrices, commonly used for the fabrication of ionselective macroelectrodes: PVC-HMW, PVC-COOH, silicone rubber, and polyurethane.15,16,21 Each of the commercially available ionophores was reported to have its best performance in one of the membrane matrices.12 In this work two potassium- and one calcium-selective ionophores were used as model compounds. Valinomycin was selected as potassium-selective ionophore because it works well in each of the membrane matrices. The other potassium-selective ionophore (BME-44)18 has higher lipophility compared to valinomycin. The calcium-selective ionophore (ETH 1001) was selected due to its special medical relevance and to prove that the microelectronically fabricated sensors show the usual ion carrier selectivity dependencies. Comparison of the Slope ∆E/∆pM and the Response Time Values of Different Membrane Formulations and Sensor Structures. The slope values ∆E/∆pM were measured in the concentration range 10-4-10-1 mol/L after soaking the prepared sensors in a Ringer solution for 3 h. The mean values of five sensors for each membrane formulation are summarized in Table 1. The tested sensors show a near-Nernstian response in the investigated range. The source of the large standard deviations in the slope values is not clear. Membranes of the same composition mounted into a conventional ISE body always had theoretical slopes. The most crucial factors determining the practical value of ion sensors for medical applications are their stability, reproducibility, and response time. The stability and reproducibility of planar sensors are significantly influenced by the osmolality fluctuations of the sample. Fortunately, the osmolality of normal human serum is fairly constant. However, (21) Pick, J.; To´th, K.; Pungor, E.; Vasak, M.; Simon, W. Anal. Chim. Acta 1973, 64, 477.
Figure 3. Potential-time plots of the new planar K+ and Ca2+ sensors with solid-state contact and different IS membranes. The transients were recorded in 1, 3, 5, 7, and 9 mmol/L KCl (a, c, and d) and CaCl2 (b) solutions with 0.147 M NaCl background.
the best results were collected using a measurement protocol where the sensors were subjected only to temporary changes in the osmolality. Besides the optimized measurement protocol, we assume that the encouraging results are due to the relatively thick sensing membrane in our new sensor design. The reproducibility tests and the response time measurements were carried out in the biologically important concentration range, at millimolar level, between 1 × 10-3 and 9 × 10-3 M. Figure 3a-d gives the potential-time plots of all four membrane formulations by varying the concentration in steps of 0.002 mol/ L. The response time of the silicone rubber-based planar sensor is comparable to that of a conventional, large ISE with massive internal filling solution. The t90% values of the PVC-based sensors are in the range of 10-15 s, whereas the polyurethane-based sensor seems to be somewhat slower. However, under flowthrough conditions, no significant response time differences were detected. The potential-time transients recorded after injecting an incremental amount of standard into a vigorously stirred beaker sometimes show a slow drift (Figure 3). However, the signal of the same sensor in the flow-through manifold was very stable (Figure 8). The apparent differences in the response times of the different sensor formulations are most likely due to the variations in the experimental conditions (cell geometry and flow conditions). The real response time of the sensors could be determined at much higher flow rates, but these are not realistic for small sample volumes like those in whole blood or human serum analysis.22 The selectivity coefficients for Na+ were
determined from the calibration plots determined at constant Na+ interference level. The selectivity coefficients for all interferences of biological relevance were determined with valinomycin-based membrane. These values were found to be in the same range as those reported in the literature for macroelectrodes.12 Determination of Long-Term Stability. The long-term stability of the sensors was tested in flow-through mode. The K+ concentration was varied between c1 ) 0.001 mol/L and c2 ) 0.004 mol/L, while the Ca2+ concentration was varied between c1 ) 0.0025 mol/L and c2 ) 0.005 mol/L. The background NaCl concentration was 0.147 mol/L in both cases. In Figure 4, a segment of the results of a long-term experiment (potential-time plot) is given for the silicone rubber-based sensor. The sensor follows the changes of potassium concentration with high reproducibility. As can be seen in the figure, both the absolute potential values and the sensor sensitivity, ∆E ) Ec2 - Ec1, were stable in the time frame of the experiment. It must be noted that the described sensors need considerable time to reach appropriate potential stability after the first contact with aqueous electrolyte. In the first 3 h of electrolyte contact, the sensor signal shows a substantial drift (≈10 mV/h). However, this drift decreases quickly with conditioning time. After 5 h of continuous electrolyte contact, 0.2 mV/h drift can be generally achieved for all types of sensors. After complete equilibration, (22) Lindner, E.; Toth, K.; Pungor, E.; Berube, T. R.; Buck, R. P. Anal. Chem. 1987, 59 2213.
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Figure 6. Schema of a sensor design variant with p-HEMA as internal electrolyte contact (for legend, see Figure 2).
Figure 4. Potential-time transients of a valinomycin/silicone rubber membrane-based sensor under flow-through conditions. The concentration of the KCl solutions was alternated between 1 and 4 mmol/ L. Background electrolyte, 0.147 mol/L NaCl.
Figure 5. Stability and reproducibility of the potential difference between two standard solutions (∆E ) Ec2 - Ec1) during 12 h. A segment of the experiment is shown in Figure 4.
the potential values show a high reproducibility (see Figures 4, 7, and 8). The drift has no effect on the sensor sensitivity. In Figure 5, the calculated potential differences (∆E values) are plotted over the measuring time, using the data given in Figure 4. The ∆E values for consecutive concentration changes are stable immediately after starting the measurement, with nearly no drift. This allows reliable analysis with the new device if the sensors are calibrated frequently. The long-term stabilities of the sensors’ sensitivity were evaluated in a 100 h experiment by determining the potential difference in two standard solutions (∆E ) Ec2 Ec1). Within the time interval studied, the ∆E values of the different membrane formulations and sensor structures had a standard deviation of 2-4% (from (0.6 to 1.2 mV for ∆E ≈ 30 mV), without any systematic decrease. By optimizing the flowthrough chamber and mathematically filtering the sensor signal, the reliability of the ∆E values of the sensors might be further improved. The sensor performance in the long-term experiment supports the advantages of the new design. Due to the large contact surface between the IS membrane and the internal reference electrode, in the time frame of the experiment, none of the membrane peeled off or produced low-resistance pathways below the membrane. The lattice-like structure, with small surface area contact with the sample, prevented a loss in sensor sensitivity due to the leaching out of membrane ingredients (ionophore, plasticizer, additives) into the sample. 4036 Analytical Chemistry, Vol. 69, No. 19, October 1, 1997
To improve the stability of the device, an electrolyte-saturated hydrogel layer (p-HEMA) was placed between the inner reference electrode and some of the silicone rubber-based membranes. The p-HEMA was photopolymerized according to the procedure given in ref 19. A schema of the sensor structure is given in Figure 6. Using the hydrogel layer, the charge transfer processes at both sides of the IS membrane are thermodynamically well defined. The improvement with respect of the sensor stability was smaller than expected on the basis of literature data. This might be due to the difficulties in controlling the deposition of the coating layers (Ag/AgCl/p-HEMA/IS membrane). A direct contact between the IS membrane and the Ag/AgCl internal reference electrode could not be prevented with absolute certainty. Due to these results and the difficulties in the reliable fabrication, all further experiments were done using sensors having solid-state contact. The practical value of sensor devices is greatly influenced by the time necessary to reach the optimal sensor performance within certain acceptable range (start-up time) as well as by the lifetime (shelf and use lifetimes). All these characteristics are primarily influenced by the storage of the sensors. After the first complete equilibration period, our sensors were kept dry (in air) between the experiments (for a maximum of 3 weeks). The rewetted sensors had the same slope immediately after the electrolyte contact as given for new sensors. The reproducibilities of the slope values were similar (∆E values determined in two solutions of different concentration). A drift of 0.2 mV/h could be reached after 30 min of solution contact. Thus, no special storage is needed for a fast start-up time and the sensor’s lifetime is not at risk due to wet storage. Performance in Human Serum. In measurements of biological samples, the loss of membrane ingredients into the sample of high lipophilicity is a primary concern. Our sensors were designed to minimize this problem. The contact surface between the IS membrane and the sample is small compared to the membrane volume. On the other hand, the small IS membrane surface is more vulnerable to materials like proteins, assumed to deposit at the membrane surface. The signals of both PVC-COOH and PVC-HMW membranebased sensors were not sufficiently stable in the biological samples. Immediately after serum contact, we observed a large drift of the sensor signal without reaching a stable value. This might be caused by the interaction of the small IS membrane surface with the sample (e.g., protein deposition) due to insufficient biocompatibility of the PVC membranes.4 Thus, our further experiments were restricted to the silicone rubber- and polyurethane membrane-based sensors. In a continuous experiment over 7 h, the sensors’ responses were tested using two different standard serum samples (S1(normal) and S2(toxic)) and three calibration solutions (C1, C2, and C3) with 1, 4, and 11 mmol/L K+ levels in 0.147 M NaCl. A segment of the potential-time recording of the
Figure 7. Segment of the experimental protocol (potential-time transients recorded in flow-through mode) used in serum measurements. (A) Ag/AgCl/silicone rubber/valinomycin-based sensor. (B) Ag/AgCl/polyurethane/BME-44-based sensor. The chronological order of samples: C1, C2, C1, C3, C1, S1, C1, S1, C1, S1, C1, C2, C1, C3, C1, S2, C1, S2, C1, S2, C1 (C, aqueous calibration solution; S, serum).
Figure 8. Segment of the experimental protocol (potential-time transients recorded in flow-through mode) used in whole blood measurement. (A) Ag/AgCl/silicone rubber/valinomycin-based sensor. (B) Ag/AgCl/polyurethane/BME-44 based-sensor.
sensors is given in Figure 7. Both sensor types show reproducible equilibrium potential values. The response time of the polyurethane-based sensor was found to be shorter than that of the silicone rubber membrane-based electrodes. Furthermore, the reproducibility of ∆ES1 or S2 ) E(S1 or 2) - E(C1) was also better with the polyurethane membrane-based electrodes, as evaluated for three sensors of each membrane type. Reproducibilities of (0.07 and (0.03 mV were measured using the silicone rubberand the polyurethane membrane-based sensors at ∆E ≈ 30 mV potential change, respectively. Thus, both types of membrane provide highly reliable data. Based on its better biocompatibility, the polyurethane membrane is preferred.23 Performance in Whole Blood and Urine. In distinct medical applications, the continuous monitoring or frequent determination of electrolyte concentrations would be preferable. In order to test the sensors’ long-term performance after repeated exposure to whole blood or urine, a flow-through experiment was (23) Lindner, E.; Cosofred, V. V.; Ufer, S.; Buck, R. P.; Kao, W. J.; Neuman, M. R.; Anderson, J. M. J. Biomed. Mater. Res. 1994, 28, 591-601.
designed modeling a catheter application.14 The sensors were flushed alternately with calibration solution of 1 × 10-3 M KCl + 0.147 M NaCl and the blood or urine sample. Duration of the experiments was typically 4 h. Figure 8 gives a segment of the potential-time plot of a whole blood measurement. Both the silicone rubber- and the polyurethane membrane-based sensors showed fast and stable response. The reproducibility of ∆Eb ) E(blood) - E(cal) was evaluated for three sensors of each type. It was found to be the same for both types of electrodes ((0.06 mV for ∆E ) 30 mV), and thus it was comparable to the data gained in aqueous electrolyte solutions and human serum samples. In addition, in the time frame of our experiments, ∆Eb did not show any tendency indicating sensor fouling. This is probably due to the continuous cleaning/calibration steps between blood measurements. In the electrolyte management, used for blood experiments, the sensor is only in contact with whole blood 25% of experimental time. After the blood contact, no considerable change in slope, offset voltage, and response time could be determined, as deduced from subsequent testing of the sensors Analytical Chemistry, Vol. 69, No. 19, October 1, 1997
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in aqueous solution. The performances of our sensors in real blood samples were also compared with reference values, determined in a commercial clinical laboratory using atomic absorption spectroscopy (AAS). A blood sample of 4.2 mmol/L (AAS) potassium concentration was assayed as 4.1 mmol/L using the polyurethane membranebased sensor. The analysis of another sample with nominal potassium concentration of 3.6 mmol/L (AAS) resulted cK+ ) 3.5 mmol/L using the silicone rubber-based membrane sensor. Considering the standard deviation of both methods, we find the correspondence of the values encouraging. Both the polyurethane membrane- and silicone rubber-based sensors were additionally tested in urine in a flow-through experiment. The sensor signal is stable. No increased drift was observed. CONCLUSIONS This paper demonstrates the performance of a miniaturized potentiometric chip device in combination with four different types of membrane and three different ionophores. For all membranes used, the sensors show near-Nernstian slope, good resolution,
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sufficient lifetime, and excellent reliability. Compared with conventional ion-selective macroelectrodes, the introduced miniaturized sensor shows a larger drift, mainly in the start-up period. On the other hand, the slope of the sensor is very stable. Thus, the disadvantage of stronger drift could be compensated by more frequent calibration. Among the tested transducer/membrane combinations, those with polyurethane and silicone rubber membranes showed the best performance. The combination of the new sensor device with an appropriate solution management system resulted in excellent signal stability and reproducibility in both serum and whole blood samples and ensured a close correlation of the concentration values determined by the potentiometric device and independent methods. The results are promising with respect to in vivo chronic applications. The performance of the new sensors in aqueous solutions makes them attractive also for process control and environmental applications. AC960957D X
Abstract published in Advance ACS Abstracts, July 15, 1997.
