ANALYTICAL SCIENCES 2001, VOL.17 SUPPLEMENT 2001 © The Japan Society for Analytical Chemistry
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Microfabricated Devices for Real-time Measurement of In Vivo and In Vitro Biomolecules Ryoji KURITA,1 Katsuyoshi HAYASHI,2 Osamu NIWA,†,2 Keiichi TORIMITSU,3 Katsunobu YAMAMOTO,4 and Takeshi KATO,5 1 2† 3 4 5
NTT Advanced Technology, 3-1 Morinosato-Wakamiya, Atsugi, Kanagawa, 243-0124 Japan NTT Lifestyle & Environmental Technology Labs., 3-1 Morinosato-Wakamiya, Atsugi, Kanagawa, 243-0198 Japan (
[email protected]) NTT Basic Research Labs., 3-1 Morinosato-Wakamiya, Atsugi, Kanagawa, 243-0198 Japan BAS Japan, 1-36-4 Oshiage, Sumida-ku, Tokyo, 131 Japan, Yokohama City University, 22-2 Seto, Kanazawa-ku, Yokohama, Kanagawa, 236-0027 Japan
We used a micromachining technique to fabricate a miniaturized glucose and lactate sensor for in vivo monitoring with dual thin layer micro channels. The sensor response was fast and highly sensitive. Furthermore there was no chemical crosstalk because the sensor has a micro flow separator in the thin layer channel. We also employed this sensor structure in a very reliable glutamate sensor that performs in vitro monitoring by comparing two currents at a dual electrode modified/unmodified with glutamate oxidase. We were able to eliminate the baseline fluctuation and pumping noise, and observe the transient glutamate release from rat nerve cells. (Received on August 9, 2001; Accepted on September 13, 2001) Recently, microanalytical devices including capillary electrophoresis chips, microfluidic devices and integrated sensors have been developed to measure in vivo and in vitro biological molecules.1-5 With a view to measuring large molecules such as DNA and proteins, devices designed for high throughput analysis have been developed that can measure a number of samples in a limited time. These include chip based capillary array electrophoresis and DNA and proteome chips. In contrast, real-time measurement is very important when measuring small molecules such as neurotransmitters, hormones and metabolites. This is because the in vivo and in vitro concentrations of such molecules change very rapidly according to cell and organ activity. Therefore, the continuous monitoring of such molecules provides us with useful information for physiological studies. We have developed small volume on-line electrochemical biosensors for the continuous measurement of neurotransmitters such as L-glutamate,6-8 histamine9 and acetylcholine10. We have also reported the miniaturization of these sensors using micro-maching techniques. However, we must take various kinds of interference into consideration when attempting to use electrochemical sensors to detect biomolecules without a separation process. With in vitro measurement, the baseline current sometimes fluctuates as a result of electrostatic and pumping noise. We often observe a sharp capacitive current when we add a solution containing chemicals to stimulate the cells7. In contrast, multi-analyte detection that is unaffected by electroactive interferents is very important for in vivo analysis. Variations in the L-ascorbic acid concentration makes analyte detection difficult. Here, we report microfluidic devices integrated with multiple enzyme electrodes and micro-reactors
for the real-time measurement of in vivo and in vitro neurotransmitters and metabolites.
Experimental Sensor Fabrication Figure 1 shows the structure of our microfabricated dual channel sensor, which consists of two glass plates. We formed carbon film on the glass plates by chemical vapor deposition, and fabricated four carbon electrodes using photolithography and dry etching. We covered the part of the carbon electrode that did not operate as an electrode with photoresist. We fabricated a thin layer flow channel (1 mm wide and 20 µm deep) with THB photoresist (g) also using a photolithographic technique. We used this process to form not only the channel but also a micro flow separator (j) between two working electrodes. The surface area of each working electrode (g, f) was 0.2 mm2. We coated the two working electrodes with osmiumpoly (vinylpyridine) wired horseradish peroxidase (Os-gel-HRP) (BAS, Tokyo, Japan) using a small brush. We left the electrode to dry at room temperature for 1 hour, and then coated glucose oxidase (GOD) (Sigma, Tokyo, Japan) film on one of the Osgel-HRP modified electrodes also using a brush. The GOD solution contained 1 unit/µl GOD and 2%-bovine serum albumin (BSA), and 0.2%-glutaraldehyde, which we used to cross-link the film. We used the same procedure to coat the other working electrode with a solution containing 1unit/µl lactate oxidase (LOX) (Sigma), 2%-BSA and 0.2%-glutaraldehyde. We coated a reference electrode with silver paste. Finally, we pressed the glass plate with the electrodes and another glass plate (26 mm x 8 mm) together and fixed them in place with UV curable resin. For the in vitro measurement device, we attached a sampling and
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an outlet capillary to the device with UV curable resin. In contrast, we connected two Teflon tubes to the device that we fabricated for in vivo measurement. One end of the tube was conncted to a microdialysis (MD) probe and the other was used as a outlet tube. a
control the size and thickness of the channel with a relatively easy fabrication process. Working electrode 1
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Fig. 1 Structure of dual channel sensor. (a)(b) sampling or outlet capillary; (c) counter electrode; (d) reference electrode; (e) working electrode 1; (f) working electrode 2; (g) THB photoresist film (20 µm thick); (h) pads to potentiostats; (i) photoresist; (j) microseparator Measurement We used a dual syringe pump to introduce the solution into the device at several different flow rates. The four pads of the film electrodes were connected to two LC-4C potentiostats (BAS) and stored in a computer using an analog/digital converter DA-5 (BAS). All the experiments were performed at room temperature. For in vivo measurement, we perfused the MD probe with a buffer at a flow rate of 0.5-4.0 µl/min. The potentials of both working electrodes were held at -50 mV (vs Ag/AgCl). In contrast, the pointed end of the sampling capillary was located near the cells and the extracellular solution was drawn with the syringe pump. We measured the neurotransmitter release by stimulating the cells with 100 mM of KCl solution.
Results and discussion Accurate flow separation into two channels The chemical crosstalk becomes a serious problem when multiple enzyme electrodes are integrated in the same flow channel a small distance apart. In our device, we separate two working electrodes arranged in parallel with a separator (Fig1. (j)). This prevents the diffusion of hydrogen peroxide from one electrode to the other, even if the distance between two electrodes is short. However, the equal separation of the solution into two channels becomes more important when we want to measure both analytes without affecting the sensitivity. Figure 2 shows the response of 1 µM aq-ferr at our microfabricated device before modifying the enzyme and mediator at a flow rate of 1 µl/min. The potentials of both working electrodes were held at 550 mV (vs. Ag). The anodic currents started to increase 10 seconds after aq-ferr injection at both electrodes and were steady-state. The magnitude of the current observed at each electrode was about 700 pA. When the flow rate was changed from 0.5 to 4 µl/min, the difference in the current was 2-4%. This result shows that the aq-ferr solution separates equivalently into the two flow channels. This equal flow separation is achieved because our method can precisely
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Fig. 2 The response of 1µM aq-ferr at the microfabricated dual channel sensor before modifying the enzyme and mediator at a flow rate of 1 µl/min. Device for In vivo Glucose and Lactate measurement Figure 3 shows the responses of the integrated device for 10 mM glucose and 10 mM lactate at a flow rate of 2 µl/min. The dual electrode of the integrated device was modified with a bilayer of GOD/Os–gel–HRP or LOX/Os–gel–HRP described in experimental section. The dual electrode was potentiostated at 50 mV versus Ag. Each sample solution was introduced into the device three times. The magnitudes of the signals were almost the same when measuring the same sample, suggesting that measurement with our device is repoducible. The cathodic current at the working electrode modified with a bilayer of GOD/Os-gel-HRP started to increase after the MD probe was dipped in 10 mM of glucose solution and reached a steady state in a very short time. The cathodic current decreased rapidly and returned to the baseline when the MD probe was dipped in a buffer solution. This is due to the small inner volume of the device since a reduction in the flow cell volume of the on-line sensors effectively improves the temporal resolution and response time. A similar result can be obtained when measuring lactate solution with the other electrode in the device. The miniaturization of the device is also useful in that it allows us to attach the device directly to animals. This also reduces the volume of the system when connecting the device to the MD probe. (a)
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Fig. 3 Responses of 10 mM glucose and 10 mM lactate with the integrated sensor at a flow rate of 2 µl/min. (a) The current peak at a working electrode modified with a bilayer of GOD/Os-gelHRP. (b) The current peak at a working electrode modified with a bilayer of LacOx/Os-gel-HRP. We observed no chemical crosstalk, which was caused by the diffusion of hydrogen peroxide generated by the enzymatic reaction at one electrode to the neighboring electrode, when measuring glucose and lactate solutions, independently.
ANALYTICAL SCIENCES 2001, VOL.17 SUPPLEMENT This shows that the hydrogen peroxide diffusion is suppressed by the separator fabricated between the glucose and lactate electrodes. For in vivo measurement, electroactive molecules such as L-ascorbic acid are major interferents. The effect of the acid can be removed by using an integrated ascorbate oxidase immobilized micro-reactor upstream of the glucose and lactate electrodes. Calibration curve Figure 4 shows calibration curves for glucose and lactate solutions measured with an integrated device at a flow rate of 2 µl/min. The cathodic current was proportional to the glucose or lactate concentration over a wide concentration range. The detection limits for both glucose and lactate were 2.3 µM (S/N = 3). The wide linear range and low detection limit is sufficient for monitoring the glucose and lactate concentrations in various media such blood or cerebrospinal fluid. In fact, variations in the glucose and lactate levels in rat brain were continuously monitored with the device and they could be used for biological sample measurements of at least one day’s duration. 10
i439 bilayer films. (B) shows the current at the working electrode modified with BSA/Os-gel-HRP without GluOx. We observed a transient cathodic current peak after a slight decrease in the current in trace (A). In contrast, we observed only a slight decrease in the current without a cathodic peak in trace (B). These results clearly indicate that the transient peak is due to the glutamate release from the cultured nerve cells. Figure 5 (a-b) shows the trace we obtained by deducting the current shown as (b) from that shown as (a) in Fig. 5. It clearly shows that the electrostatic noise and pumping noise were greatly reduced, and only the glutamate concentration change can be observed. The above results prove that we can continuously measure biochemicals with high selectivity and reliability in a differential measurement mode using our dual channel microfabricated biosensor.
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Fig. 4 Calibration curves for glucose (circles) and lactate (squares) with the integrated sesnor. Differential measurement of in vitro glutamate using the device Trace level analysis is required for measuring neurotransmitter release from cells since the transmitter concentration and sample volume are very low for such measurements. However, various kinds of noise are present when we attempt to detect bio-chemicals with high sensitivity. These include the electrostatic noise or baseline noise caused by pumping induced pressure fluctuations or the capacitive current observed when another solution is added to stimulate a sample such as cultured cells. To avoid the above noises, we fabricated the device integrated with two electrodes in a dual channel for the differential electrochemical detection of glutamate. The structure of our dual channel glutamate sensor is similar to that of the glucose and lactate sensor described in the previous section. We coated the two working electrodes with Os-gel-HRP, and then with glutamate oxidase (GluOx:Yamasa Shyoyu, Choshi, Japan) film on one of the Os-gel-HRP modified electrodes. The GluOx solution contained 0.2 units/µl GluOx and 2%-BSA, and 0.2%-glutaraldehyde, which we used to crosslink the film. We then used the same method to coat the other working electrode with 2%-BSA and 0.2%-glutaraldehyde solution without GluOx. Figure 5 shows the variations in glutamate concentration from rat cortex neurons cultured for two weeks stimulated by 100 mM KCl solutions obtained using our microfabricated glutamate sensor. (A) shows the current at the working electrode modified with BSA-GluOx/Os-gel-HRP
30 sec Fig. 5 Variations in the glutamate concentration from cultured rat cortex neurons stimulated by KCl obtained using the dual microfabricated glutamate sensor. (A) shows the current at the working electrode modified with BSA-GluOx/Os-gel-HRP bilayer films. (B) shows the current at the working electrode modified with BSA/Os-gel-HRP without containing GluOx. (AB) shows the trace when trace (B) was deducted from trace (A). Acknowledgments The authors thank Drs. Saburo Imamura and Hisao Tabei for encouraging this project. The authors also thank Ms. Wakako Tanaka for drawing the sensor photomask and Ms. Yuriko Furukawa for preparing the cell culture. References 1. D. J. Harrison, K. Fluri, K. Seiler, Z. Fan, C. S. Effenhauser, A. Manz, Science, 1993, 261, 895 2. A. J. Woolley, D. Hadley, P. Landre, A. J. deMolleo, R. A. Mathies, M. A. Northrup, Anal. Chem., 1996, 68, 4081 3. L. C. Waters, S. C. Jacobson, N. Kroutchinina, J. Khandurina, R. S. Foote, J. M. Ramsey, Anal. Chem., 1998, 70, 158 4. K. Sato, M. Tokeshi, T. Odake, H. Kimura, T. Ooi, M. Nakao, T. Kitamori, Anal. Chem., 2000, 72, 1144 5. S. P. A. Fodor. J. L. Read, M. C. Pirrung, L. Strryer, A. T. Lu, D. Solas, Science, 1991, 251, 767 6. O. Niwa, T. Horiuchi, R. Kurita, H. Tabei, K. Torimitsu, Anal. Sci., 1998, 14, 947 7. O. Niwa, R. Kurita, T. Horiuchi, K. Torimitsu, Electroanalysis, 1999, 11, 356 8. O. Niwa, R. Kurita, Z. Liu, T. Horiuchi, K. Torimitsu, Anal. Chem., 2000, 72, 949 9. O. Niwa, R. Kurita, K. Hayashi, T. Horiuchi, K. Torimitsu, K. Maeyama, K. Tanizawa, Sens. and Actu. B, 2000, 67, 43 10. O. Niwa, T. Horiuchi, R. Kurita, K. Torimitsu, Anal. Chem., 1998, 70, 1126
