Anal. Chem. 1989, 61, 2566-2570
2566
In the second case, the ligand starts in the organic phase but has a reasonable solubility in the metal-ion-containing aqueous phase and the metal-ligand reaction takes place in the aqueous phase. Here, the rates of the mass transfer processes and chemical reaction cannot be treated as independent of one another. uM2and uR2 are not independent. A suitable impulse function for deconvolution cannot be obtained and the rate of a chemical reaction occurring in the presence of a comparably fast mass transfer cannot be obtained by deconvolution. In this mixed regime the chemical reaction rate is, in fact, difficult to measure in any way by solvent extraction alone (7). In contrast to this variable situation regarding the role of mass transfer with various classes of chemical reaction, the contribution of instrument band broadening will always be independent of both mass transfer and chemical reaction so that deconvolution with a suitable IRF (8,9)can be used to remove it, subject to the limitations of signal noise discussed above. This is fortunate since in current instrument design it is instrument band broadening, rather than mass transfer, which imposes a practical limit on the magnitude of chemical reaction rates that can be measured by rapid stir solvent extraction. Registry No. I-, 20461-54-5; Fe3+,20074-52-6.
LITERATURE CITED SchiU, G. In Ion Exchange and Solvent Extracbbn; Marinsky, J. A.; Marcus, Y.. Eds.; Marcel Dekker: New Yolk. 1974; Vol. 6. Chapter 1. Cantwell, F. F.; Carmichael, M. Anal. Chem. 1982, 5 4 , 697-702. De, A. K.; Khopkar, S. M; Chakners, R. A. Solvent Extracth of Meta/s; Van Nostrand: Toronto, 1970. Fossey, L.; Cantwell, F. F. Anal. Chem. 1982, 54, 1693-1697. H8ndbOok of Solvent Exfrectbn; Lo, T. C., Bakd, M. H., Hanson. C., Eds.; Wiley: New Yolk, 1983. Weber, W. P.; Gokel. G. W. Phase Transfer Cetalvsk: Swlnm-Verlag: New York, 1977. Danesi, P. R.; Cherizia, R. CRC Crn. Rev. Anal. Chem. 1980, 10, 1-12s. . Cantwell, F. F.; Freiser, H. Anal. Chem. 1988, 60, 226-230. Amankwa, L.; Cantwell, F. F. Anal. GI". 1989, 61. 1036-1040. Aprahamian, E., Jr.; Cantwell, F. F.; Freiser, H. LangmM 1985, 1 , 79-82. Hershey, A. V.; Bray, W. C. J . Am. Chem. Soc. 1936, 5 8 , 1760-1 772. Eigen, M.; Kustin, K. J . Am. Chem. Soc. 1962, 84. 1355-1361. "Stabilrty Constants" J . Chem. Soc. Spec. Pub/. 1964, No. 17.344. Laurence, 0. S.; Ellis, K. J. J . Chem. Soc., Dallon Trans. 1972, 2229-2233. Fudge, A. J.; Sykes, K. W. J . Chem. Soc. 1952, 119-124. Sykes, K. W. J . Chem. Soc. 1952, 124-129. Smith, R. M.;Martell, A. E. Crlthl Stabiltty Constants; Plenum: New York, 1976; Vol. 4.
. -
RECEIV~, for review April 5,1989. Accepted August 23,1989. This work was supported by the Natural Sciences and Engineering Research Council of Canada and by the University of Alberta.
Pulsed Amperometric Detection of Glucose in Biological Fluids at a Surface-Modified Gold Electrode Dilbir S. Bindra and George S. Wilson* Department of Chemistry, University of Kansas, Lawrence, Kansas 66045
A nonenzymatk glucose sensor that utlllzes permdectlve membranes to acMeve the selectlvlty requlred tor screening glucose In bWogkal HUMS has been described. Interference from endog.nousoxkkable substances such as amino ackls, urea, ascorbk acld, and uric acld, as well as the effect of chkride and protdns on g k o g e response, Is studkd by d n g flow Injection analysis. A set of membranesmade of Naflon pedhwhtedmemkaneand collagen, when arrangd In front of the worklng electrode (gold), result In slgnlfkant Improvement In the system selectlvlty. Even at physlologlcal pH, which Is far from belng the optimum pH for pulsed amperometrk detection of carbohydrates, the sensor shows a good llmR of detectlon (4-5 pg of glucose Injected).
INTRODUCTION The detection of glucose in biological fluids has long been essential in bioanalysis. As a result of development of the enzyme electrode by Clark (I) and by Updike ( Z), glucose detection based on the highly specific glucose oxidase catalyzed reaction has been the method of choice. Despite the specificity of the enzyme reaction, sensor response is influenced by the partial pressure of oxygen (a cosubstrate) in the medium and by the presence of electroactive interferents such as ascorbic
* To whom correspondence should be directed. 0003-2700/89/0361-2566$01.50/0
acid and uric acid. By the use of permselective membranes it has been possible to eliminate most of these difficulties but at some cost in sensor complexity (3). Glucose sensors based on direct oxidation of glucose at noble metal electrodes have been known for many years. Direct oxidation has been suggested as the basis for an implantable glucose sensor (4) or generally as a means for the detection of glucose in biological fluids. Glucose is not well-behaved electrochemically, and analysis based on the interpretation of current-voltage curves is complicated and time-consuming (5). Response is subject to electrode fouling by oxidation products (6) and to interference from amino acids, ascorbic acid, urea (7), and a variety of drugs such as acetaminophen. Moreover, chloride is known to have a significant inhibitory effect on glucose oxidation kinetics (5). In spite of the numerous difficulties mentioned above, direct oxidation of glucose can be made tractable by suitable conditioning or modification of the sensing electrode. For example, the application of multistep potential waveforms to the working electrode, which incorporates cleaning and activation steps along with detection, has made possible rapid, sensitive, and reproducible measurements. This technique, known as pulsed amperometric detection, has proven useful for the measurement of a wide range of analytes separated by liquid chromatography (8). A recent study (9) has shown that a Cu-based chemically modified electrode yields a stable response over a considerable period of time without recourse to a potential pulse sequence. 0 1989 American Chemical Society
ANALYTICAL CHEMISTRY, VOL. 61, NO. 22, NOVEMBER 15, 1989 2567
In this paper we describe a pulsed amperometric sensor that is able to reliably measure physiological glucose levels without any prior sample separation or cleanup. The interferences are limited to a large extent by using suitable membranes that not only selectively control the diffusion of certain species, but also protect the electrode from coming in direct contact with biological fluid. EXPERIMENTAL SECTION Apparatus. The flow injection system consisted of a Shimadzu Model LC-6A pump and SCL-6A controller, Waters Associates Model 710B autosampler, a Princeton Applied Research Model 400 electrochemical detector, and a flow-through thin-layer electrochemical cell consisting of single gold working electrode (MP1300), Pt counter electrode, and Ag/AgCl (saturated KC1) reference electrode. The detector output was processed by a Shimadzu CR 4A integrator, the peak area being used as the basis for analysis. The PARC Model 400 is designed to apply a repeating sequenceof three applied potentials to the electrochemical cell according to a specified timing sequence. Materials. All solutions were prepared from analytical grade chemicals with water from a Barnstead Nanopure I1 system. All buffers were refiltered through a 0.45-pm filter after preparation. L-ascorbic acid solutions were prepared just before use, as ascorbic acid is subject to oxidative decomposition in solution. Soluble Nafion (5% by weight in 90% lower aliphatic alcohols and 10% water) was obtained from Aldrich Chemical Company, Milwaukee, WI. The collagen and cellulose acetate (MW cutoff 12000-14000) membranes were supplied by YSI, Yellow Springs, OH, and Viscase Corporation, Chicago, IL, respectively. Normal control serum was obtained from Ortho Diagnostic Systems, Inc., Raritan, NJ. Normal human serum was also used and was obtained from Lawrence Memorial Hospital, Lawrence, KS. All measurements were performed in 0.1 M phosphate buffer (PB),pH = 7.4,unless specified otherwise. Samplesfor the g l u m recovery experiment were prepared by adding appropriate amounts of glucose (50-500 mg/dL) to control serum or human blood serum. These samples were diluted 1 5 with 0.1 M phosphate buffered saline (PBS), pH = 7.4 (100 mequiv/L NaCl), before injecting. Nafion Coating. A Nafion membrane was cast over the working electrode by spreading a thin layer of Nafion solution on the clean and dry surface of the Kel-F block holding the electrode. Exactly 50 pL of the Nafon solution was placed over the electrode and was spread with a brush over the whole flow channel (2.5X 0.8 cm) including the gold electrode surface. The electrode was then left to dry at room temperature overnight. The thickness of the Nafion film was roughly estimated as 2-3 pm by using a density of 1.58gm/cm3for Nafion f i i (10). A precast collagen (100pm thick, dry state) or cellulose acetate (30 pm thick) membrane was used for the outer protective layer. The Teflon gasket (spacer) was replaced on top of the membranes before reassembling the cell. The three-step potential sequence for the detection of glucose at a gold electrode has been optimized elsewhere (11,12)and was adopted without any major modifications. In general, the choice of potentials is determined from the current-potential response curve of the analyte of interest. A sequence consisting of a detection potential (+lo0 mV, 0.499 s), an oxidative cleaning potential (+650 mV, 0.249s),and a cathodic reactivation potential (-800 mV, 0.166 s) was used for all experiments. As the time period for completion of one cycle is less than 1 s, the detection is essentially continuous as in direct current (dc) amperometry. RESULTS AND DISCUSSION Glucose Calibration Curves. The choice of 0.1 M PB, pH = 7.4, as the mobile phase was made to be consistent with physiological pH even though the optimal condition for measurement is pH = 12-14 (11). The protective membranes are, moreover, not stable at such pH extremes. The function of these membranes is discussed in the next section. The background current for a freshly polished gold electrode took around 3 h to reach a steady-state value after initial application of the pulse sequence. This coincides with the time it took for the sensor to produce a reproducible response
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00
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Time 100 (min) 150
200
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Flgue 1. Stabilizationof glucose response (A)and backgound curent (0)after initial application of the pulse: sample injected, 30 pL of 25 mg/dL glucose in 0.1 M PB, pH = 7.4.
IO
20
30
Time ( min ) Flgure 2. Multiple flow injection peaks for glucose: sample Injected, 30 pL of 60 mg/& glucose;mobk phase, 0.1 M PB, pH = 7.4 at 0.25 mL/min; working electrode, gold covered with Naflon and collagen.
to a given amount of glucose injected repeatedly (Figure 1). This phenomenon has been explained previously as being due to the continuous microscopic roughening of the electrode surface (until a constant area is obtained) by alternate formation and removal of surface oxide through electrochemical pulsing (13). However, reported times for background stabilization were only 5-10 min when 0.2 N NaOH was used as the mobile phase (13). Presumably the physiologic medium conditions do not favor formation and the subsequent reduction of the deposit formed. The glucose response was found to be linear within the physiological range. The detection limit was approximately 25 nmol in a 30-pL sample (4.5 pg of glucose) for SIN = 10. The reproducibility of the sensor response was evaluated by repetitive injection of the 30-pL sample. Ten injections were used to calculate a relative standard deviation of 1% . Multiple flow injection peaks obtained with a modified electrode are shown in Figure 2. A flow rate of 0.25 mL/min was selected, as it provided a good compromise between the sensitivity and the speed of analysis. The sensor response was found to be stable for a t least 1 week. Effect of Chloride. Figure 3 shows the effect of increasing chloride concentration in a glucose sample on the oxidation current for a bare gold electrode and a Ndion-coated gold electrode. About 80% of the response was lost for the bare gold electrode at physiological concentrations of chloride. Even though the variation in chloride ion concentration within
2568
ANALYTICAL CHEMISTRY, VOL. 61, NO. 22, NOVEMBER 15, 1989 120
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trode (0) and Naflon-coated gold etectrode (A):gkrcose concentration, 100 mg/dL; sample size, 30 pL.
BSA1 .o Concentration 2.0
(g/dL) 3.0
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ngWe 4. Effect of BSA on glucose response for bare gold electrode, and collagen on pH = 7.4 (01, (b) bare gold electrode, pH = 13 (A), Nafion-coated gold electrode, pH = 7.4 (0):glucose concentratlon, 100 mg/dL; sample size, 30 pL.
Table 1. Results of Glucose Recovery from Control Serum for Different Working Electrode Conditions conditions" bare gold (pH = 13) bare gold Nafion-coated gold collagen on Nafion-coated gold
--
70 recovery*
50 24 60 95 f 3 (n = 4)
:.
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3
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"All measurements made a t pH = 7.4, unless specified. Samples were prepared by adding 100 mg/dL glucose to control serum. Samples were diluted 1:5 before injecting. The recovery results were calculated by comparing the signals from glucosespiked serum samples (corrected for the serum blank) and glucose samples in 0.1 M PBS.
*
the normal physiological range (100-106 mequiv/L) had little effect on the glucose oxidation current, the overall loss in sensitivity was too significant to be overlooked. Nafion, in the form of a cation-exchangepolymer membrane, is effective in selectively excluding anions from the electrode surface. This property of Nafion has previously been exploited for elimination of a chloride interference effect on cupric ion-selective electrodes (14) and of ascorbic acid interference in the determination of dopamine (15). Nafion also has been successfully employed as a dialysis membrane to provide a protein- and interferent-free environment near the electrode for the determination of glucose in whole blood (IO). As seen in Figure 3, a Nafion-coated electrode preserves most of its response in the presence of chloride. The effect of chloride ion was far less pronounced when 0.2 N NaOH (pH = 13) was used as the mobile phase. This was primarily due to the greatly enhanced sensitivity for glucose at high pH as the absolute magnitude of the effect is the same. Effect of Proteins. The effect of proteins on glucose response was studied by adding bovine serum albumin (BSA) to the glucose standards. For a bare electrode the response was largely suppressed due to the surface poisoning by BSA adsorption. But unlike the case of dc amperometry, where protein adsorption tends to be irreversible, the electrode here regained its original activity within 10 min of protein injection as the pulsing slowly removed all of the surface adsorbed protein. At pH 13, though the poisoning was less severe (probably because denatured protein did not adsorb well to the electrode surface), it was still sufficient to affect the results in a glucose recovery experiment (Table I). The Nafioncoated electrode also did not offer complete protection against protein fouling, resulting in poor recovery of glucose from biological samples. It was,however, possible to put a collagen membrane in front of the Nafion to prohibit protein from
0.0 0
20
40
60
80
100
Concentration of added Glucose (mg/dL) Flgure 5. Glucose response curves for aqueous glucose samples (0) and glucose-spiked human blood serum samples ( 1 3 diluted) (A): sample size, 30 pL; working electrode, gold covered with Nafion and collagen.
reaching the inner Nafion layer. This bilayer membrane structure completely eliminated the inhibitory effect resulting from the protein adsorption while maintaining the desired permselective characteristics (Figures 4 and 5). A similiar bilayer coating, with a cellulose acetate film covering the Nafion layer, has been utilized previously by Wang and coworkers (16) for selective detection of cationic neurotransmitters in urine samples. Interference Studies. All interfering compounds were studied at their physiological minimum and maximum levels with glucose concentration being kept constant at 100 mg/dL. As expected,the glucose measurement was stronglyinfluenced by the presence of ascorbic acid, uric acid, amino acids, and acetaminophen. Under the conditions of measurement, the bare sensor was found to be far more sensitive toward ascorbic acid and uric acid than toward glucose. Therefore, even though both ascorbic acid (0.2-2 mg/dL) and uric acid (4.0-8.5mg/dL) are present in body fluids at concentrations that are significantly lower than that of glucose (60-110 mg/dL), the errors caused by their presence were fairly significant. For example, the response for a glucose sample (100 mg/dL) on a bare electorde was increased by 140% when ascorbic acid was added to the sample at its physiological maximum concentration (2 mg/dL). On the other hand, as both ascorbic acid and uric acid are present as anions at physiological pH, their transport to the
ANALYTICAL CHEMISTRY, VOL. 61, NO. 22, NOVEMBER 15, 1989 120.0
h
Table 11. Sensor Response for Various Carbohydrates
100.0
80.0
compound
response at equal concn bel)
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glucose galactose D-(-)-frUCtOSe lactose
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2569
physiological concn, mg/dL
response at physiological max (rel)
60-110
