950
Anal. Chem. 1986, 58, 950-956
efforta of K. R. Loder and K. R. Hazel of Entropy in collecting field data and G. B. Howe of RTI for the GC analyses. We thank J. C. Suggs of USEPA for the statistical evaluation of field data. Registry NO.LCV, 603-48-5;p-SABA, 138-41-0;CO,630-08-0; cacotheline, 561-20-6.
LITERATURE CITED (1) Fed. Reglst. 1984. 4g (no. 29), 5326-5327. (2) Lambert, Jack L.; Wiens, Robert E. Anal. Chem. 1974,46,929-930. (3) Lambert, Jack L.; Chiang, Yuan C. Anal. Chem. 1983, 5 5 , 1828-1830. (4) Lambert, Jack L.; Chiang, Yuan C. Anal. Chem. 1984, 56, 808. (5) Ciuhandu, Gheorghe 2.Anal. Chem. 1957, 155, 321-327. (6) Levaggi, D. A.; Feidstein, M. Am. Ind. Hyg. Assoc. J . 1984, 25, 64-66. (7) Corsini, A,; Chan, A.; Mehdi, H. Talsnta 1984, 34, 33-38.
(6) Ferguson, Bruce €3.; Lester, Rlchard E.; Mitchell, W. J. ”A Study to Evaluate Carbon Monoxide and Hydrogen Sulfide Continuous Emission Monitors at an Oil Refinery”, EPA-600/4-82-054. Aug 1982. (9) “Traceabllky Protocol for Establishing True Concentrations of Q s e s Used for Calibration and Audits of Continuous Source Emission Monitors (Protocol No. )”I, June 1978, included wlthln “Quality Assurance Handbook for Air Pollution Measurement Systems. Volume 111. Stetionary Source Specific Methods”, EPA 6R0/4-77-027b, Aug 1977. (IO) Code of Federal Regulations, Title 40, Part 60, Appendix A, 1983 pp 593-61 2. (11) Youden, W. J. In “Statistical Manual of the Association of Officlal Analytical Chemists”; Washington, DC, 1975; pp 33-36. (12) Margeson, John H.;Knoll, Joseph E.; Midgett, M. Rodney; Oidaker, Guy E. 111; Reynolds, Wayne E. Anal. Chem. 1985. 5 7 , 1586-1590. (13) Code of Federal Regulations, Title 40. Part 60, Appendix A, 1983; pp 467-474.
RECEIVEDfor review July 8, 1985. Accepted November 27,
1985.
Development and Analytical Performance of Tubular Polymer Membrane Electrode Based Carbon Dioxide Catheters Walter N. Opdyckel and M. E. Meyerhoff*
Department of Chemistry, The University of Michigan, Ann Arbor, Michigan 48109
The development and analytical performance of a new potentlometric pC0, (partlai pressure of CO,) senslng catheter Is described. A unlque sensor geometry Is achleved by utlilzlng an Inner tubular polymer pH electrode In conjunction wlth an outer gas-permeabie silicone rubber tube. Once assembled, thls flexible catheter has an outer diameter of 1.1 mm. When filled wlth an approprlate reference soiutlon and Internal electrolyte, the resulting devlce responds rapldly (180 s for t,,,) and reproduclbly to pC0, changes In the normal and abnormal physiological concentration range (15-150 torr) with nearly Nemstlan behavlor (54-61 mV/decade change In pC0,). Long-term stability of the Inner pH electrode and consequently the gas sensor Is dependent on the Internal reference solutlon composilion. Favorable stabliHy (f2 mV over 45 h) is achleved wlth a phosphate buffered Internal reference solution. Contlnuous pC0, values obtalned with the new sensor during 6-h in vitro blood pump studles correlate well with conventional bioodgas Instruments. Preilminary results for sensors Implanted Intravascularly In a dog demonstrate their suitability for continuous In vivo monltorlng of PCOP.
The clinical significance of blood gas measurements has long been recognized (1-3). Abnormal partial pressures of oxygen (PO,) and carbon dioxide (pC02) in blood can be indicative of respiratory or metabolic disorders. As a result, frequent blood gas analysis is crucial to the management of surgical and intensive care patients. Typically, these measurements are performed on discrete arterial blood samples using commercially available blood-gas analyzers (4-6).However, since serious changes in blood gas levels may occur in a matter of minutes or less, the lag time associated with discrete sample in vitro methods poses a danger to the patient’s welfare. For this reason, it would be desirable to develop a reliable and ’Present address: Diversey Wyandotte Corp., Wyandotte, MI. 0003-2700/86/0358-0950$01.50/0
relatively inexpensive device capable of continuously monitoring gas levels in vivo. In this report, we describe an essentially disposable electrochemical pCOz sensor that appears suitable for such purposes. Many methods already have been proposed for continuous monitoring of blood gases, and these techniques have been reviewed by several authors (7-11). While no method has yet received universal acceptance, considerable emphasis has been placed on inexpensive electrochemicaldevices that utilize outer gas-permeable membranes to achieve high selectivity. The Clark-type voltammetric PO, electrode has been successfully miniaturized for in vivo measurements (12, 13). However, miniaturization of the classical Severinghaus-type potentiometric pCO2 sensor (14,15)has met with less success, since this sensor relies on a fragile glass pH electrode as the internal sensing element. Miniature metal/metal oxide pH electrodes have been suggested as alternatives to glass for the fabrication of pCOz catheters (16-18);however, the sensitivity of such internal electrodes to redox species along with the delicate nature of the sensing tip region may explain why such devices have not found wide use despite some promising preliminary in vivo results (19-21). Recently (221, we reported on the development and analytical applications of new potentiometric gas sensors based on polymeric pH electrodes of the type described by Schulthess et al. (23). By use of this neutral carrier-type internal pH membrane electrode, small, inexpensive, and durable ammonia and carbon dioxide gas sensors may be fabricated whose response properties compare favorably with commercial glass membrane based sensors. The purpose of this paper is to extend this concept by describing the design and performance of a new pC02 gassensing catheter that is based on an internal tubular polymeric pH electrode. A schematic diagram of this sensor is shown in Figure 1. The unique feature of this device in comparison to other miniature pCOz sensors is the geometry of the sensing region. The polymeric pH-sensitive membrane is situated safely within the wall of the internal tubing rather than at the vulnerable tip. This protects the sensing region from Q I986 American Chemical Society
ANALYTICAL CHEMISTRY, VOL. 58, NO. 4, APRIL 1986
951
(22%) (chromatographic grade, Polysciences), dibutyl sebacate socket
==t
1
__
silicone rubber sealant
luer tip
f iiiing sol ut Ion
rilicone rubber tublng
Ag/AgCi electrodes
-
PVC tubing
reference solution ion-selective membrane
total length: 12cm
Flgure 1. Schematic diagram of the tubular polymer membrane
electrode-based pC0, catheter. damage during catheter placement or removal. This geometry also allows for sensor size reduction without a corresponding decrease in the pH-sensitive membrane area and a concomitant increase in electrode resistance. Further, this new design greatly reduces problems (variable time response, drift, etc.) normally associated with the junction between the bulk internal electrolyte and the thin-film sensing region found in conventional gas sensors (24, 25). In this report, we describe the development of this new pC0, sensor and the key factors that influence its performance. Specifically, we examine the effect of internal reference solution composition on the stability of the inner polymeric p H electrode as well as the effect of various filling solutions on the analytical performance of the final gas sensor (response times, slopes, etc.). In addition, to assess which parameters determine the response time of the new device, we directly compare the response times of the new sensor with analogous devices fabricated with capillary glass inner p H electrodes. The long-term stability of the sensor in flowing aqueous solution and blood is also evaluated as is its performance in a series of blood pump studies. Finally, data will be presented demonstrating the initial application of this device for in vivo pC0, monitoring in dogs. EXPERIMENTAL SECTION Apparatus. Most potentiometric measurements were made with a Coming Model 12 or a Fisher Model 825 M P pH/mV meter and recorded on a Fisher Recordall strip chart recorder. For the in vivo studies and some of the blood pump studies, the potentiometric measurementswere made with a custom-builthigh-input impedance differential amplifier with unity gain and 12 separate channels. The output of this system was monitored with a Fluke Model 2240C data logger. Reagents. All chemicals were reagent grade. All solutions and buffers were prepared with distilled-deionized water. All buffer concentrations refer to total ionic strength. The pH-sensitive polymer membrane paste consisted of trin-dodecylamine(20% w/w) (Eastman Kodak),poly(viny1chloride)
(57%) (Sigma Chemical Co.), and sodium tetraphenylborate (1%) (Eastman Kodak). The membrane components were dissolved in about 1 mL of tetrahydrofuran (Fisher Scientific). The internal reference solution for the pH electrodes was usually 0.80 M phosphate buffer (pH 6.8) containing 0.10 M NaCl. For the electrode stability experiments, other buffers were used, principally 0.18 M and 0.35 M citrate-HC1 buffers (pH 4.75). A variety of filling solutions were tested in the CO, gas sensor. They included the commercial preparations Orion 95-02-02 and HNU ISE 10-22-01,as well as various combinations of sodium hydrogen carbonate with sodium chloride. The majority of experiments were performed with a 0.025 M NaHCO, solution in 0.125 M NaC1. Preparation of Polymer pH Electrodes and Gas Sensors. The polymer pH electrode assembly was prepared in the following manner. A length of Tygon tubing (Norton Plastics, 0.025 cm id., 0.076 cm 0.d.) was cut to a 12-cmlength. A 5-cm-long33-gauge syringe needle was inserted into the tube. About 2 cm from this end, a 1-cm-long slice was removed from the wall with a razor blade. The pH membrane was deposited in this gap by dropwise addition of the membrane components in THF. When the thickness of the pH membrane approximated that of the Tygon tube, the electrode was left to cure overnight. The electrode socket assembly was prepared from 34-gauge silver wire and a transistor socket (Radio Shack). Two pins on the socket were removed, and two silver wires (6 and 10 cm) were soldered to the remaining pins. The wires were anodized at 0.4 mA/cm2 in 0.1 M HC1 to produce two Ag/AgCl reference electrodes. The electrode tube was filled with the phosphate buffer using a 26-gauge syringe needle and then cut to a 10.4-cmlength. The electrode tube was fitted onto the 10-cm Ag/AgCl wire and sealed at each end with hot tweezers. At the transistor socket the exposed pins and wires were sealed in silicone rubber adhesive (G. C. Electronics), and the luer tip of a plastic syringe (Becton-Dickinson) was fitted over the silicone rubber, removing the excess adhesive. The COzgas sensor was completed with a 9.6-cm length of out& dimethyl silicone rubber tubing (0.09 cm i.d., 0.11 cm o.d., Patter Products) sealed on one end with a plug of silicone rubber adhesive. This tube was filled with an appropriate filling solution and then fitted over the pH and reference electrodes until the end plug touched the bottom of the pH electrode. The sensor was sealed at the luer tip with silicone adhesive and stored in a vial containing filling solution. By use of the above manual fabrication procedure, approximately two-thirds of all sensors prepared exhibited appropriate response to pC0,. The remaining sensors exhibited little or no response to COz apparently due to “leaks, short circuits, etc.” inherent to the manual fabircation of these very small devices. For the in vivo experiments the luer tip-transistor socket area was protected by an outer plastic ring containing epoxy cement. This reinforced the rather weak juncture for the rigors of in vivo experiments. Effects of Reference Solution Composition on the Stability of the Polymer pH Electrode. The stability of the pH electrodes was determined by daily measurement of the electrode potential in two buffered solutions. Typically, these solutions were 0.25 M phosphate buffers pH 6.0 and pH 7.5. The pH of these buffers was verified each day with a calibrated glass electrode. Each buffer contained 0.10 M NaCl so that the reference Ag/AgCl potentials were the same in both solutions. By use of these buffers, the cell potential and the electrode slope were determined daily for periods of 2-5 weeks. Two parameters were varied in this experiment. The first was the internal reference solution composition, and the second was the conditioning solution in which each electrode was stored. Both the citrate and phosphate solutions described above were used as both the reference solution and as conditioning solution. In some experiments, the hydrogen carbonate filling solution was also used as a conditioning solution. In another experiment, the effect of pCOz on the internal reference solution pH was determined. Electrodes containing the phosphate buffer were placed in a closed and thermostated (37 “C) cell containing 0.2 M citrate buffer (pH 4.8). Additions of NaHCO, were made until CO, saturation occurred.
952
ANALYTICAL CHEMISTRY, VOL. 58, NO. 4, APRIL 1986
a Orion
gas mixtures
b
r
ir'
I/
sampling .It.
aenror blocks
Figure 2. Blood pump apparatus used for (a) single and (b) multiple
sensor testing.
Determination of Citric Acid (Citrate). To determine the amount of citric acid crossing the polymeric pH membrane, the following procedure was used. Two pH membranes were cast in glass rings (2.5 cm diameter) on a glass plate as previously described (26). These membranes were then pasted to Tygon tubing pieces having the same diameter (2.5 cm). One membrane assembly served as a test and the other as a blank. One side of each membrane (inner) was in contact with 500 pL of distilleddeionized water. The other interface (outer) was bathed in a 1.8 M citrate HC1 buffer (pH 4.8). For the blank, this latter solution was replaced with 1.0 M KCl. Every 7 days the 500-~Lwater samples were removed and evaporated on watch glasses. Fresh water was then added. After 4 weeks the experiment was stopped. The residue in the watch glasses was redissolved in 2 mL of water. Analysis of these samples was performed by an enzymatic method described in the Sigma product bulletin for aconitase (Product A5384). Response Characteristics of the COz Gas Sensors: Polymer vs. Glass Electrodes. Tubular microglass pH electrodes (0.08 cm 0.d.) were obtained from Ingold Electrodes, Inc. Carbon dioxide sensors based on these glass electrodes were prepared in a similar fashion to the PVC electrode-based systems. Extreme care was necessary due to the fragile nature of the glass electrodes. Response characteristics were determined by monitoring glass and PVC electrode-basedsensors simultaneously in thermostated cells (37 "C). Response times and reproducibility were examined by simultaneously switching both sensors between the two 0.15 M NaCl solutions saturated with gases containing 6.1% (43 torr) (the tank used in this particular study, labeled 5% COz in Nz, was actually found to contain 6.1% COz as determined by an independent method) and 10.0% (70 torr) COz in Nz (Air Products), respectively. Response ranges and slopes were determined by additions of NaHC03 to a 0.2 M citrate-HC1 buffer (pH 4.8). Flowing Blood Studies. Figure 2 shows two arrangements used for testing the COz sensors in flowing solutions and blood. For all experiments, the gas sensors were calibrated at 37 OC with 0.15 M NaCl solutions saturated with gas mixtures containing 5.0% (35 torr) or 10.0% (70 torr) COz in Nz. Experiments were carried out at 37 "C. Figure 2a shows the apparatus used to determine the sensors' long-term stability. Heparinized whole human blood reconstituted from outdated plasma and cells was obtained from the University of Michigan Hospital. An Orion C02gas sensor (95-02) was used as a reference sensor. The potentials of both sensors were monitored for periods of 3-20 h while the COz levels were adjusted
with gas mixtures of COz in N,. A more sophisticated apparatus is shown in Figure 2b. Heparinized canine blood was pumped (ExtracorporealBlood Pump 1812, Medtronic, Inc.) through the system at a rate of 2 or 3 L/min. A membrane oxygenator (SciMed Life Systems, Model I-2500-2A)was employed to change COz levels. Custom blend gases (Union Carbide) containing 5,10, and 12% COz and 13,20, and 40% Oz in a balance of Nz along with gas mixtures of 100% COz with 100% Oz were all used t o yield different COz levels. Blood samples were taken at approximately 20-min intervals and analyzed with a micro-13 blood gas analyzer (BGA) (Instrumentation Laboratories (IL)). The experiment lasted 6 h. In Vivo Experiments. The pCOz sensors developed were evaluated in vivo at Medtronic, Inc. (Minneapolis, MN), as part of their ongoing sensor program. A large mongrel dog (31 kg) was used in this experiment. The dog was anesthetized with injections of sodium pentabarbital (Nembutal). Carotid and femoral arteries served as sensor sites for four sensors. Fourteen-gauge catheter placement units were inserted into each exposed artery. The gas sensors were then inserted into the arteries through the placement units and sutured in place. Blood gas levels were monitored via the micro-13 blood gas analyzer and a Corning Model 165 blood gas instrument. Both instruments were maintained at 37 "C, while the temperature of the dog varied between 36.5 and 35.7 "C. No attempt was made to correct the in vitro pCOz values for temperature variations. Similarly, the catheter results were not corrected for these small temperature differences. The sampling site was the right brachial artery. Samples were taken and analyzed every 20 min for the duration of the experiment (6 h). Carbon dioxide levels were adjusted through the use of a mechanical respirator and different inhalation gases. A base-line level of COz was established when the dog breathed room air at 12 breaths/min. A low level of COz was produced by increasing respiration to 18 breaths/min and having the dog breathe 100% Oz. Similarly, a high COz level was achieved at 7 breaths/min and by breathing a gas mixture of 12% COz, 40% O,, and a balance of Nz. Sensors were calibrated before and after the experiments using the same procedure as for the blood pump studies to obtain the working slopes for the implanted sensors. However, absolute in vivo pCOz values were obtained by utilizing these slopes in conjunction with an initial in vivo pCOz value determined in vitro using the IL BGA. This initial value served as a means of obtaining K in the following Nernst equation for the sensor: Esensor
= K + s 1% (PC0Z)m VIVO
RESULTS AND DISCUSSION The reliability of the pCOz catheter proposed here should be directly related to the performance of the internal polymeric pH electrode used to fabricate the sensor. Therefore, we spent considerable effort studying and optimizing the stability of the miniature pH electrode prior to evaluating the final gas sensor design. Such studies along with the actual pC02 catheter results are presented below. Effect of Reference Solution Composition on the Stability of the Polymeric pH Electrode. All of the miniature tubular pH electrodes prepared displayed rapid and nearly Nernstian response to pH changes (e.g., slopes of 58-62 mV/pH unit) regardless of the composition of the buffer used as the internal reference solution. However, since the membrane potential of these polymer pH electrodes is governed by the following equation: E m = ( R T / F ) In (aH+(out)/aH+(in)) where (in) and (out) denote the internal and sample solutions, respectively, long-term stability and accurate pH measurements can only be achieved if the pH of the internal reference solution remains constant. The pH of the reference solution can be affected by acidic or basic species that can diffuse (in and/or out) through the polymeric p H membrane. Carbon dioxide is an example. Indeed, the COz permeability of pH-responsive polymer membranes has been exploited recently to design a new hy-
ANALYTICAL CHEMISTRY, VOL. 58, NO. 4, APRIL 1986
953
-90-
E,mV.
-710.
-80.
E,mV.
-13 P
J
Time, days
Figure 3. Effect of storage solution composition on the potential of pH electrodes initially filled with citrate buffer. The potentials of two electrodes (vs. AgIAgCI) were monitored daily In a phosphate buffer (pH 6.0, 0.25 M) test solution. One electrode (0)was continuously stored in citrate-HCI buffer (pH 4.7, 0.18 M) for the duration of the experiment (except when tested in the phosphate buffer). The other electrode (a)was alternately stored in each buffer as indicated by the marked time scale (p, changed to phosphate: c, changed to cltrate).
drogen carbonate ion selective electrode (27). In this case, nonbuffered internal reference fills must be used. To eliminate the effects of CO, permeation (i.e., electrode drift), on pH measurements, we previously suggested the use of a high-ionic-strength citrate-HC1 buffer (pH 4.50) as a fill for the polymeric pH electrode. Such a buffer was also suggested by Anker et al. (28)for blood pH measurements with the same electrode. The citrate buffer was chosen because it best maintains a constant reference pH even in the presence of high sample pC0, levels. Unfortunately, when citrate buffers were used as reference solutions in the new tubular pH electrodes, it was found that the potentials of these electrodes changed dramatically when stored in non-citrate buffers (i.e., phosphate, pH 6.00). This is illustrated in Figure 3, which shows the absolute cell potentials (pH electrode vs. Ag/AgCl reference) with time for two electrodes containing citrate buffer reference solutions. As can be seen, the potential of the electrode that was stored in citrate buffer was fairly stable over the entire period tested, while the potential of the other electrode depended on the solution in which it was kept. Figure 4 expands on this point. Electrodes with inner phosphate reference solutions are far more stable when stored in various phosphate buffers than when stored in a citrate buffer. Further experiments (not shown) demonstrated that electrodes with phosphate fills were also stable when stored in hydrogen carbonate solutions, whereas those with citrate fills were not. Thus, it is the nature of the citrate buffers that causes the significant drift problems observed. Citrate buffer solutions (pH 4.7) contain at least 8% fully protonated citric acid. The solubility (lipophilicity)of the free citric acid (uncharged) in the polymeric pH membrane differentiates it from any of the other buffer components in either the citrate or phosphate buffers. Indeed, it has been reported that citric acid is readily extracted from aqueous solutions with organic solvents containing tridodecylamine (29). Therefore, it seems probable that the pH membrane will allow citric acid to diffuse through it provided that there is a difference in the chemical potential of citric acid across the membrane. Such diffusion of citric acid was demonstrated by enzymatically measuring the amount of citric acid appearing in a pure water solution as the other side of the membrane bathed in a citrate buffer (see Experimental Section). While actual rates of diffusion were not determined, this experiment clearly showed that citric acid was capable of permeating the membrane used to constructe the pH-sensitive electrode.
-120.
-160-
a
-200.
0
5
10
15
20
25
30
T i m e , days Figure 4. Effect of storage in citrate and phosphate buffers on pH electrodes inltlally filled with phosphate buffered reference solutions (0.25 M, pH 6.0). Potentials were determined in two phosphate buffers (pH 6.0 and 7.5, 0.25 M). Curve a is for an electrode stored In citrate-HCI buffer (pH 4.7, 0.18 M). Curves b and c represent the response of electrodes stored in 0.25 M phosphate buffer (pH 6 for b, pH 7.5 for c).
Strangely, this problem of citric acid diffusion and concomitant electrode drift was not noted in any of the earlier polymer pH electrode work (22). This was probably because the volumes of the internal solutions used in those efforts were far greater than the volumes used in our new miniature electrode (200 p L vs. 2 pL) while the area of the pH sensing membrane remained nearly the same. Thus, for a given rate of citric acid diffusion, the pH of the internal fill of the miniature electrodes will be affected to a greater extent in a given time period than the larger and more conventional polymeric pH electrodes. To avoid significant drift problems, all miniature polymeric pH electrodes were prepared with phosphate buffered internal reference solutions (pH 6.80). A rather high-ionic-strength buffer (0.8 M) was chosen to avoid problems associated with COz diffusion into this reference fill. Electrodes prepared in this manner did not respond to pC0, even at sample levels as high as 600 torr. Response Characteristics of pC0, Catheters. The response properties of the pC0, catheters prepared with the tubular pH electrode (Figure 1)varied somewhat depending on the internal filling solution used to prepare the sensors. Commerical HNU and Orion filing solutions yielded sensors with essentially the same response properties as those prepared with certain in-lab preparations, and their use was not pursued. Sensors prepared with the laboratory-prepared solutions (10,25, and 50 mM NaHC03 in 100,125, and 100 mM NaC1, respectively) did have significant differences in their response characteristics. For example, the response time for sensors made with the 10 mM bicarbonate fill was about twice as fast as that obtained with the 50 mM solution. The use of 10 mM bicarbonate also yielded sensors with lower detection limits (Le., 0.3 torr) as predicted by theory (30). Sensors employing
954
ANALYTICAL CHEMISTRY, VOL. 58, NO. 4, APRIL 1986
T i m e , hours
Figure 5. Long-term stability of the pC0, catheter in flowing solutions saturated with gas mixtures of 5 % (35 torr) and 10% (70 torr) CO, in N,. A 0.15 M NaCl solution was used for the first 20 h. Between hours 20 and 25, no recording was performed. A 0.15 M NaCi solution containing 1 % gelatin was used in the last half of the experiment. The apparatus shown in Figure 2a was used for this study.
10 mv
I
1
+A:" + Figure 6. Comparison of response between (a)tubular glass and (b) polymer pH electrode-based gas-sensing catheters.
the 50 mM solution were, however, much more stable, especially at the higher C02levels found in physiological samples. A suitable compromise between the fast response and sensor stability at physiological C02 levels (pC0, = 20-80 torr) was reached by using the 25 mM hydrogen carbonate filling solution in all subsequent in vitro and in vivo studies. Sensors prepared with the 25 mM hydrogen carbonate solutions typically exhibited Nernstian response from 2 to 600 torr pCOz with a lower limit of detection of 1 torr. Slopes varied from sensor to sensor, but generally fell within the range of 54-61 mV/decade change in CO, concentration. The pC0, catheters were also quite reproducible with good long-term stability. Indeed, Figure 5 shows the absolute potential of a catheter over a 48-h period as GO2 levels are varied back and forth between 35 and 70 torr in flowing saline solution and then in flowing saline containing gelatin using the pump system shown in Figure 2a. Typically, useful sensors maintained their potentials within &2 mV for a given COz level over a t least a 40-h period. This good stability is clearly a result of our earlier efforts to minimize drift of the inner tubular pH electrode. Like all potentiometric gas sensors, the response time of the new pC0, catheter was dependent on analyte concentration. In the physiological pC0, range, the sensor required approximately 120 s to attain a stable equilibrium potential for step changes in GOz levels (Figure 6), while 95% of this was achieved in 580 s. This reequilibrium response sponse time appears to be limited by the diffusion of GOz into the inner polymeric pH electrode. Since this inner electrode is prepared with PVC tubing that is permeable to COz (29, equilibrium pH changes in the thin electrolyte film (between the outer silicone rubber tubing and inner PVC tubing) can only be achieved if the entire sensor reaches equilibrium with the sample solution. In contrast, when glass capillary electrodes (impermeable to CO,) with the same diameter as the
b. R
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Figure 7. Resuits of blood pump study for four catheter sensors. In vitro values (8)were determined with the IL micro-13 BGA. PVC tubings were used as inner pH detectors, much faster response to changes in C02 were observed. Figure 6 clearly illustrates this point. While it may be argued that the glass pH electrodes have slightly faster innate response times to pH changes than the polymeric electrodes, it is our view that such differences do not account for the full change in gassensor response times observed and that diffusion of analyte C02 into the polymer pH electrode is the limiting factor in time response for the new sensors. This time response limitation is, however, more than outweighed by the enormous practical advantages offered by using the polymer pH electrode in place of glass for in-vivo-type sensors. Blood Pump Studies. In order to determine the feasibility of using these tubular gas sensors for in vivo measurements, we first tested their performance in flowing blood. These studies were conducted with the apparatus shown in Figure 2. For each study, a reference method for pC02 determinations was employed. Correlation between pC0, values obtained with the new catheters and the reference methods is shown in Table I for several such experiments. It is apparent that the C02catheters track pC0, changes in blood quite well in the 15-175-torr range and can do so for extended periods (i.e., up to 20 h). This is further illustrated in Figure 7, which shows the continuous response of four pCOz catheters in relation to the discrete in vitro values obtained with the IL micro-13 blood gas analyzer (experiment 3 in Table I). Since clinical pC02 values rarely fall outisde the range of 15-80 torr (2),it is obvious that our sensors behave as required for biomedical applications. In Vivo Studies. In vivo experiments were performed in an anesthetized dog as described in the Experimental Section.
ANALYTICAL CHEMISTRY, VOL. 58, NO. 4, APRIL 1986
955
Table I. Blood Pump Studies: Correlation Data" no. of sensors
expt
ref method Orion COz sensor Orion COz sensor
20-95
IL micro-13
1
1
2a
2
1
2a
3
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3
26
0.6
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0.999
30-176
6
23
2.4
0.998
0.988
30-175
6
23
-0.4
0.921
0.970
30-175
6
23
0.6
0.976
0.991
30-175
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23
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0.938
0.992
. BGA
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L
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Flgure 8. Continuous in vivo response of two catheters placed in carotid arteries as compared to in vitro blood gas measurements (e) IL BOA, (W) Corning BGA).
Figure 8 shows the results of the in vivo study for two reinforced sensors placed in the carotid arteries. In this experiment, the catheter pC02 values were calibrated in vivo to an initial value obtained with the BGA on a discrete sample (note: this value was used to set E,see Experimental Section). This method is preferred, since differences in conditions between the calibration medium and the in vivo environment of the sensors can cause emf offsets (4,6). From Figure 8, it is apparent that the catheters do follow the in vivo changes in pC0, within reasonable limits. Two other sensors placed in the femoral arteries also tracked the pC0, changes but exhibited significant positive drifts. On removal, it was observed that these two sensors were coated with a thick thrombus formation. The catheters placed in the carotid arteries also showed evidence of clot formation, but to a lesser extent. At this point, it is not clear whether the excessive clot formations caused the problems observed with the sensors placed in the femoral arteries, and further research aimed a t determining the exact effects of such clotting will be required. The effects of temperature and in vivo arterial pressure on the performance of the sensors were not examined in detail. While there was a finite difference in temperature between the dog and the calibrating medium, these differences were small and the errors associated with them were not deemed critical in demonstrating that the sensors functioned effectively in vivo. Further, while it is pomible that pulsating blood pressures could affect the performance of the sensors by
b,
time,
C T wsensors ~ failed; the response o f the four usable sensors is shown pinching off the thin bicarbonate internal electrolyte layer, there wm no evidence of such effeds in this preliminary study. Indeed, the noise levels of the sensors implanted were only slightly greater than those used for in vitro experiments. For future clinical applications, these sensors possess several noteworthy features. Since they are potentiometric devices, measurements are made under essentially zero-current conditions, and problems associated with mass transfer of the C02 (e.g., viscosity effect) should not cause errors in the pC02 values determined. In addition, like all Severinghaus-type CO, devices, these catheters are unaffected by anesthetic gases or other blood components. Most importantly, the approach described here enables fabrication of small and highly flexible gas sensors that are essentially disposable units. With further improvements in design and the use of appropriate anticoagulant coatings, we are confident that these new devices will eventually prove useful in a varieity of biomedical applications.
ACKNOWLEDGMENT We thank Eric J. Fogt and James Kelley of Medtronic, Inc., for performing the in vivo studies. In addition, we gratefully acknowledge the National Institutes of Health for supporting this work (Grant GM-28882-04). LITERATURE CITED (1) Spence, A. A. "Respiratory Monitorlng in Intensive Care"; Churchiii Llvlngstone: New York, 1980. (2) Shapiro, 8. A.; Harrison, R. A.; Walton, J. R. "Clinical Applications of Blood Gases"; Year Book Medical Publishers, Inc.: Chlcago, IL, 1982. (3) Durst, R., Ed. "Blood pH, Gases and Electrolytes"; NBS Special Publication No. 450, 1977. (4) Dowd, J.; Jenkins, L. C. Can. Anaesth. SOC.J. 1973, 20, 129-140. (5) Winckers, E. K. A.; Teunissen, A. J.; Van den Camp, R. A. M.; Maas. A. H. J.; Veefkind. A. H. J. Clln. Chem. Clin. Biochem. 1978, 76, 175- 185. (6) Bateman, N. T.; Musch, T. I.; Smith, C. A.; Dempsey, J. A. Resplr. PhySlol. 1980, 47, 217-226. (7) Blackburn, J. P. Br. J. Anaesth. 1978, 50, 51-62. (8) Severinghaus, J. W. Biotelemetry Patient Monit. 1979, 6, 9-15. (9) Hahn, C. E. W. J. Phys. Envlron. Sci. Instrum. 1981, 74, 783-797. (10) Beetham, R. Ann. Clln. Biochem. 1982, 79, 198-213. (11) Mindt, W.; Eberhardt. P. Med. Prog. Technol. 1982, 9, 105-111. (12) Parker, D.; Soutter, L. P. I n "Oxygen Measurements in Biology and Medlclne"; Payner, J. P., Hill, D. W., Eds.; Butterworths: London, 1975; Chapter 18. (13) Eberhard, P.; Fehlmann, W.; Mundt, W. Biotelemetry Patlent Monit. 1979, 6 , 16-31. (14) Lai, N. C.; Llu, C. C.; Brown, E. 0.; Neuman, M. R.; KO, W. H. Med. 8/01. Eng. 1975, 73, 876-881. (15) Parker, D.; Deipy, D.; Lewls, M. Med. 6lol. Eng. Cornput. 1978, 76, 599-600. (16) Van Kempem, L. H. J.; Kreuzer, F. Resplr. Physiol. 1975, 23, 37 1-379. (17) Markdahi-Bjarme, M.; Edwall, G. Med. 8/01.Eng. Cornput. 1981, 79, 447-456. (18) Macur, R. A.; LeBlanc, 0. H.; Grubb, W. T. US. Patent 3905889, 1975.
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RECEIVED for review September 13,1985. Accepted November 25, 1985.
Homogeneous Enzyme-Linked Competitive Binding Assay for the Rapid Determination of Folate in Vitamin Tablets Leonidas G. Bachas and Mark E. Meyerhoff* Department of Chemistry, The University of Michigan, Ann Arbor, Michigan 48109
A rapid and sensltlve homogeneous enzyme-llnked competltive blndlng assay for folate that exhlblts unique dosereponse characteristics Is descrlbed. The method utlllzes a hlghly substituted glucose 6-phosphate dehydrogenase-folate conjugate In conjunctlon wlth folate blndlng proteln. I n the absence of folate, the enzyme conjugate Is inhiblted up to 70 %, I n presence of folate, actlvlty Is regalned In an amount proportlonal to the folate concentration. Such reversal occurs over an extremely narrow concentratlon range and thls range can be flne tuned to deslred values by varylng the concentratlon of folate blndlng proteln and/or the enzyme-folate conjugate. Theoretical models suggest that thls unusual behavior can be attributed to the favorable equilibrium constants between folate bindlng proteln and conjugate. The new assay Is shown to be preclse, accurate, selectlve, and useful for the measurement of folate In vitamin preparatlons. Speculation as to the advantageous use of thls novel assay in a yes/no type quality control situation is also presented.
The use of enzyme labels rather than radiotracers in the development of competitive binding assays for important biological molecules is a rapidly emerging area of bioanalytical research (1-3). The technique exploits the inherent chemical amplification of an enzyme-catalyzed reaction along with the selective molecular binding characteristics of certain proteins to achieve assays with high specificity and low detection limits. When antibodies are utilized as the binding reagent, the resulting methods are termed enzyme immunoassays (EIAs). Such methods overcome many of the problems associated with the more classical radioimmunoassay procedures (RIA) and are now being employed routinely in clinical chemistry laboratories (4). Enzyme immunoassays may be either heterogeneous (separation required) or homogeneous (no separation step). The homogeneous type methods are preferred since they often result in much faster assays. Ngo and Lenhoff (5) have reviewed the host of homogeneous EIA arrangements that have been proposed over the past decade. The most widely used approach is still the EMIT type assay pioneered by Ruben-
stein et al. (6) for the detection of haptenic species (low molecular weight drugs, hormones, etc.). In this case, the binding of hapten-enzyme conjugates by antihapten antibodies causes a modulation of the enzyme activity in solution. Usually, the modulation involves inhibition of enzyme activity and this inhibition is reversed in an amount proportional to the concentration of analyte in the sample (7, 8). Glucose 6-phosphate dehydrogenase (GGPDH) has been the enzyme label most often employed and a wide assortment of simple drug assays based on this system are commercially available from Syva Corp. (Palo Alto, CA). To date, antibodies have been the binders exclusively used in the EMIT type of assays. However, it is known that antibodies can exhibit considerable binding recognition toward the bridging group, which couples the hapten molecules to the enzyme label. This is particularly true when the same bridge is used to prepare the protein-hapten conjugate employed to elicit antibody production (9-11). Thus, the antibodies that form have often greater affinity for the enzyme labeled hapten than the unlabeled hapten and this can seriously diminish the detection capabilities (detection limits, steepness of dose-response curve, ED50 value, etc.) of the EMIT assays. While the effects of such bridging group recognition on the analytical properties of heterogeneous EIAs have been studied experimentally (9-1 I ) and, more recently, theoretically (12),such effects on EMIT type of assays have received only limited attention (13-15). We recently reported on the use of endogenous binding proteins rather than antibodies in the development of heterogeneous enzyme-linked competitive binding methods for folate, cyanocobalamin, and thyroxime (16,17). Since binding proteins do not recognize the bridging group, it was our belief that assays with improved analytical characteristics would result. Indeed, the folate assay (16) exhibited excellent detection limits, high sensitivity (Le., steepness of the descending portion of the dose-response curve), and a rather unique biphasic (hooked) dose-response curve. These characteristics were attributed to the binding properties of the immobilized folate binding protein. The purpose of this report is to extend this concept by describing our findings when folate binding protein (FBP) is utilized in solution (nonimmobilized) along with highly substituted GGPDH-folate conjugates in an EMIT
0003-2700/88/0358-0956$01,50/00 1986 American Chemical Society
