Anal. Chem. 1989, 61, 2365-2372
other methods currently in use. The simplicity of ionophore synthesis and electrode fabrication should make this method attractive for future application in determining clinical CRP levels.
ACKNOWLEDGMENT The authors thank Ruth Silberman and Johanne Smid for their aasktance in ionophore synthesis, David Rice for assisting in the NMR study, and Paul Turner for his technical assistance.
LITERATURE CITED Kushner, Irving; Broder, Martin L.; Karp, David J. Clin. Invest. 1978, 61, 235-242. Gitlin. Jonathan D.: Gitlin. Joan I.: Gitlin. David, Arthrits Rheum. 1977, 20, 1491-1499. Pereira, De Sika J. A.: Elkon, Keith 6.;Hughes, G. R. v.; Dyck, Roland F.; Pepys, Mark B. Arthrnis Rheum. 1980, 23, 770-771.
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(4) Oewurz, Henry, Hosp. Practice 1982, (June), 67-81. (5) Robey, Frank A.; Liu, Teh-Yung J. Bkl. Chem. 1982, 258,3895-300. (6) Oliveira, Eduado B.; Gotschlich, Emil C.; Liu, Teh-Yung, J. Immunol. 1980, 124, 1396-1402. (7) Tanaka. Taeko; Robey, Frank A. J. Immunol. Methods 1983, 65, 333-341. (8) Pedersen, C. J. J. Am. Chem. SOC.1967, 89, 7017-7036. M. Y.; Rechnitz, G. A. Anal. Chem. 1984. 56. 801-806. Keating, (9) (10) Solsky, Robert L.; Rechnitz, G. A. Anal. Chlm. Acta 1881, 123, 135-141. (11) Keating, M. Y.; Rechnitz, G. A. Analyst 1983, (June), 766-766. (72) Petranek, J.; Ryba, 0. Anal. Chim. Acta 1974. 72, 375-380. (13) Osmond, Alexander P.; Freidenson, Bernard Gerwurz, Henry; Painter, Robert H.; Hofmann, Theo; Shelton, Emma R o c . Natl. Aced. Sci. U . S . A . 1977, 74, 739-743. (14) Volanakis, John E.; Clement, W. L.; Schrohenloher, Ralph E. J. Immunol. Methods 1978, 23, 285-295.
RECEIVED for review August 15, 1988. Revised manuscript received April 26, 1989. Accepted August 7, 1989.
Potentiometric Combination Ion/Carbon Dioxide Sensors for in Vitro and in Vivo Blood Measurements M. E. Collison,' G. V. Aebli, Jennifer Petty, and M. E. Meyerhoff* Department of Chemistry, University of Michigan, Ann Arbor, Michigan 48109
The development and analytical performance of a novel potentiometric combhation lon/pCO, sensor design for in vitro and In VIVO measurements are reported. The design is based on Incorporating an appropriate Ionophore wlthln the outer slilcone gas permeable membranes of both conventlonai macro and new catheter-type pC0, sensors. Simultaneous measurement of the potentials across the ion-selectlve/gas permeable membrane and the inner glass or polymer pH sensltlve membrane provldes the basis for continuous monltorlng of both Ionic and pC0, levels wlth the same devlce. A macro-sized K+/pC02 embodlment of the sensor Is constructed from a commercial Severinghaus COPsensor and is used to demonstrate the prlnciples and capablilties of the proposed design. A flexible, minlaturized (outer diameter = 1.2 mm) combination K+/pCO, catheter sensor is also described. The catheter-type sensor is fabricated by Inserting a tubular polymer membrane pH electrode Into an outer siilcone rubber tube doped with valinomycin. Continuous measurements of K+ and pC0, during 6-h blood pump studles uslng both the macro and catheter-type combination sensors correlate well wtth those of conventlonaibenchtop analyzers. I n addltkn, contlnuowr (4 h) Intravascular measurementswlth the combhation catheter sensor in dogs show good agreement wlth those of commercial blood analyzers ( R = 0.984 and 0.962 for pC0, and K', respectively).
In the modern health care setting, clinicians are increasingly dependent upon the reliable and rapid measurement of chemical variables for accurate diagnostic and therapeutic decisions (1). Consequently, there has been growing interest in the monitoring of blood gases and electrolytes continuously
Present address: Eli L i l l y Corp., Indianapolis,
IN.
0003-2700/89/0361-2365$01.50/0
using invasive chemical sensors (2-4). From the standpoint of patient safety (i.e. danger of infection and blood vessel damage), the invasiveness of implantable sensors will be a prime consideration in the clinical acceptance of these devices. Hence, a primary goal in the development of in vivo chemical sensors is the incorporation of multiple sensing elements per implant probe (3). In this report we describe the design and analytical performance of both macro in vitro and miniature catheter-type in vivo ion (K+)/pC02sensors. The proposed catheter device is a disposable, multiple-analyte sensor that is capable of continuous, simultaneous in vivo monitoring of a selected ion (e.g. K+) and carbon dioxide gas. Several approaches to the fabrication of multiple-analyte in vitro and in vivo sensors have been reported previously. Ion-selective field effect transistor (ISFET) devices have very small sensing areas, allowing a single miniaturized solid-state chip to be configured for simultaneous monitoring of several analytes (2, 5). While the ISFET concept has generated considerable interest, fundamental problems regarding sensor encapsulation, drift, and interference from electrically neutral acidic sample species (e.g. COJ must be resolved before these devices become practical for in vivo monitoring applications (3, 6,7). More recently, multiple-analyte fiber-optic sensors, particularly for pH, pCO,, and PO,, have been described (8). However, optical electrolyte sensors (for K+, Ca2+,Na+) that respond in a rapid and reversible fashion have not yet been developed, owing to the absence of appropriate indicators for these clinically important ions (9). Moreover, fundamental concerns regarding the effect of sample ionic strength on the accuracy of optical ion activity and pH measurements remain the subject of some debate in the literature (10). Miniature potentiometric combination ion/gas sensing devices have also been described previously. In 1976 General Electric Corporation developed a novel combination pH/pCO2 catheter sensor (11). Carbon dioxide sensing was achieved by using a metal oxide (Pd/PdO) pH electrode as the internal transducer in a classical Severinghaus-style gas sensing ar0 1989 American Chemical Society
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ANALYTICAL CHEMISTRY, VOL. 61, NO. 21, NOVEMBER 1. 1989
1
2
1 mm
1 dlam. vent Hole
I
Luer-lok InjeCIlon Adaptor
Assembly
Reference
*
I"I r na I Retersnce
.
Valinomycin mpregnaled -lSilicone
I
Rubber Tubing
External Ret*,e"Ce Electrode
Potasslum
- Bicarbonate
I
Vsllnomycln Impregnated Slllcons Rubber Membrane
Flgure 2. Model combination K+lpC02 sensor fabricated from a commercial Severinghaus CO, sensor.
SOl"li0"
+ mm C
m e 1. Schematic dlaqam of the dual ion/pCO, catheter ccnfisured as a K+/pCO, sensor: (a)internal polymer membrane pH electrode module: (b) ionophore (valinomycin)impregnated silicone rubber tube module: (c) assembled K+/pCO, catheter sensor. rangement. The sensor's gas permeable membrane contained a hydrogen ion carrier, allowing it to also function as a hydrogen ion-selective membrane. Thus, by measurement of the potential between the inner PdO pH electrode and an external reference electrode, the pH of the blood sample could also be monitored. Unfortunately, this sensor's sensitivity to tip deformations, pronounced drift (due to inherent instabilities of the PdO pH electrode), and large temperature sensitivity precluded its routine use in vivo (12-14). Recently (15), we described the design and favorable performance of a novel disposable pC0, gas sensing catheter based on an internal tubular polymeric pH electrode that is housed inside a gas permeable silicone rubber tube. Since polymeric membrane pH electrodes are not sensitive to redox species (including O,), the resulting device is more stable than metal-oxide-electrode-based pC0, sensors. In addition, the unique tubular geometry allows the sensor dimensions to be reduced without a concomitant decrease in the pH-sensitive membrane area and an associated increase in membrane resistance. The main purpose of this work is to extend the functionality of the polymer-membrane-based pC0, catheter design by configuring it as a dual ion/pCO, sensor. The new combination sensor design, configured as a dual potassium ion/pCO, sensing catheter, is shown in Figure 1. This sensor differs fundamentally from the earlier polymer-membraneelectrode-based pC0, catheter in that the outer gas permeable silicone rubber tubing is impregnated with an appropriate ionophore (e.g. valinomycin for K+/pCO, sensor). In the case of valinomycin impregnation, the sample potassium ion activity can be monitored simply by measuring the potential that develops across the wall of the outer silicone rubber tubing (i.e. the potential between contact 1 of Figure 1 and a sample reference electrode). The sample pCOz is monitored by measuring the pH of the sensor's bicarbonate filling so-
lution via the internal polymer membrane pH electrode (i.e. the potential between contacts f and 2 of Figure 1). Lastly, the sensor can be adapted to monitor other electrolyte ions merely by impregnating the outer silicone rubber tubing with a different ionophore. The new combination sensor has, for the purposes of this communication, been configured as a dual K+/pCO, sensor because of the physiological importance of these two analytss. Accurate and timely measurement of a patient's blood pC0, plays a major role in establishing his or her acid-base status. Potassium ions play a crucial role in maintaining the transmembrane potentials necessary for normal cell functioning. Hypokalemia of the blood plasma has deleterious effects on cardiac rhythm, while hyperkalemia leads to cardiac arrest. A combination K+/pCO, in vivo sensor is of particular interest because of the relationship between pC0, and plasma POtassium concentration. Specifically, a rise in pC0, results in a transient fall in plasma potassium, whereas a fall in pCOl produces a transient rise in plasma potassium. While the magnitude of these changes in potassium levels would not be expected to have adverse effects in a normal patient, their occurrence in patients with preexisting abnormal plasma potassium levels and those on digitalis therapy may be clinically significant (16). The availability of an in vivo sensor capable of simultaneously monitoring both potassium and pC0, would help to assess the clinical significance of these changes. Prior to investigation of the proposed catheter design, the feasibility of the dual ion/gas potentiometric measurement scheme was first evaluated by using a commercial Severinghaus-style pC0, sensor as a model macro-size K+/pCO, in vitro sensing device (see Figure 2). This was accomplished by examining the effect of valinomycin impregnation of the gas permeable membrane on the sensor's pC0, response and by proving that simultaneous measurements of potassium can be made by monitoring the transmembrane potential of the outer ion-selective/gas permeable membrane. It will be shown that both the model macro and catheter Kf/pCOz sensors described herein provide an effective means of performing simultaneous in vitro and in vivo measurements of potassium and pCOz in whole blood.
ANALYTICAL CHEMISTRY, VOL. 61, NO. 21, NOVEMBER 1, 1989
EXPERIMENTAL SECTION Instrumentation. Potentiometric measurements were made by using a Compaq portable computer (20-Mbytehard disk) with a Data Translation DT-2805-5716 analog/digital input/output board (Marlborough, MA) and a custom-built electrode interface module controlled by Labtech Notebook software (Laboratory Technologies Corp., Wilmington, MA). The 16-bit DT-2805-5716 A/D converter was operated to provide an input resolution of k0.04 mV. For most measurements, A/D sampling frequencies were 0.033 or 0.10 Hz. For sensor response time studies, 1-Hz sampling frequencies were used. High impedance electrode signals were interfaced to the data acquisition board through a custombuilt high input impedance interface consisting of eight differential, unity gain amplifier stages. Each stage employed two Analog Devices op-amps (AD 515, input impedance = lOI3 Q ) as voltage followers for indicator and reference electrode signals. The outputs of the voltage followers were fed to an AD 517 differential, unity gain amplifier. The output of each amplifier stage was further conditioned by a low-pass RC filter (99% rise time = 2.5 s) and a 60-Hz slot filter. Electrode resistance measurements were made with a Keithly 610-B electrometer calibrated against precision 10- and 1000-MQ resistors. Orion 95-02 and Instrumentation Laboratory 68652-10 commercial COP sensors (with internal glass membrane pH electrodes) were used to fabricate macro size model K+/pCOz sensors. Combination K+/pCOzsensor blood measurements (in vitro and in vivo) were compared to potassium measurements performed on a NOVA-6 blood electrolyte analyzer and to pCOz measurements performed on Radiometer ABL-2 or ABL-4 blood gas analyzers. Reagents. All chemicals were of reagent grade. All solutions and buffers were prepared with reverse-osmosis,deionized water. Buffer concentrations reported here refer to total ionic strength. The composition of the internal filling solution (IFS)(0.125 mol/L KCl and 0.025 mol/L NaHC03saturated with AgC1) used in both the macro and catheter combination sensors was similar to that of the optimal IFS used in the polymer-membrane-basedcatheter COz sensor (15). Combination catheter sensors used the same internal filling solution containing 0.0075% (w/w) Brij-35 surfactant (Sigma Chemical Co.) to facilitate insertion of the pH electrode module into the outer silicone rubber tubing (Figure 1). The surfactant was found to have no detrimental effect on the response, stability, selectivity, or lifetime of the inner pH electrode and the K+ selective valinomycin silicone rubber tube. The potassium ionophore, valinomycin, was obtained from Sigma (St.Louis, MO). The H+ ionophore tridodecylamine and lipophilic anion salt additive dimethyldioctadecylammonium bromide (DMDODABr) were obtained from Kodak (Rochester, NY). Potassium tetrakisb-chlorophenyl)borate (KTpClPB) and the plasticizers dioctyl sebacate (DOS) and o-nitrophenyl octyl ether (0-NPOE) were products of Fluka (Ronkonkoma, NY). The lipophilic cation/lipophilic anion salt additive (DMDODATpClPB) was prepared from DMDODABr and KTpClPB (17). General Sensor Calibration and Evaluation Procedures. In vitro measurements were conducted in thermostated doublewalled beakers at 25 or 37 OC. Combination sensor, commercial COz sensor, and conventional poly(viny1chloride) (PVC) membrane potassium electrode (18)signals were usually measured concurrently. For most in vitro studies, potassium signals were measured relative to a common double-junctionAg/AgCl reference electrode (Fisher) with 4.6 N sodium formate bridge electrolyte. This electrolyte was chosen because of its suitability for whole blood measurements (19). Combination sensors were calibrated simultaneously for COz and K+ responses by standard addition and two-point calibration methods. Standard addition calibrations were performed by making standard additions of 2.5 mol/L KCl, 0.25 mol/L NaHC03 solution to a 0.2 mol/L citrate-HC1, pH 4.8 buffer. Two-point span calibrations were accomplished with 0.14 mol/L NaCl, 0.025 mol/L NaHC03 calibrant solutions containing 1.00 and 10.0 mmol/L KC1 and equilibrated to two different COz/N2 gas mixtures (1.08% and 9.92% C02/Nz,Air Products, Tamaqua, PA). Silicone Rubber Impregnation Techniques. The procedure of Fogt et. a1 (20) was followed for valinomycin impregnations of silicone rubber membranes used in the macrosensor investigations. An impregnating solution composition of 10 mg of
2367
valinomycin in 2.5 mL of xylenes (Aldrich ChemicalCo.) was used. A round patch (0.13 mm thick, 2 cm in diameter) of methyl vinyl silicone rubber membrane (Sci Med Life Systems, Minneapolis, MN) was soaked in this solution for 1-2 min. Following impregnation, the membrane was suspended in a hood draft for 15 min to allow the swelling solvent to evaporate. Impregnation of silicone rubber tubing for catheters was optimized as described in the Results and Discussion section below. The final tubing impregnation procedure waa as follows: Silicone rubber tubing (0.94 mm i.d. X 1.22 mm 0.d.) was extruded by Patter Products (Beaverton, MI) using General Electric-871 silicone gum. Prior to impregnations, a 200-cm length of coiled tubing was cleansed by Soxhlet extraction with 95% ethanol for 30 min followed by toluene extraction for 15 min. Residual toluene was evaporated in a hood draft for 15 min, followed by 1 4 evaporation in a pump-evacuated chamber. The valinomycin impregnating solution consisted of 5 mg of valinomycin and 110 WLof dioctyl sebacate (DOS) dissolved in 1.25 mL of freshly distilled tetrahydrofuran (THF). The cleansed silicone rubber tubing was cut in 16.5-cm-length sections and impregnated by immersing one end in a 1.1-cm depth of impregnating solution for a totalof 15 min. Following impregnation, each tubing section was placed in a hood for 15 min to allow excess THF to evaporate. Excess DOS plasticizer remained on the surface of the impregnated region of the tubing and was used to estimate the length of valinomycin impregnation. Excess DOS plasticizer was then removed with compressed air and wiping with Kimwipes. Finally, residual THF was removed by 1-h evaporation in a pump-evacuated chamber. Macro Combination Sensor Assembly and Measurements. Macro combination potassium/carbon dioxide sensors (Figure 2) were constructed from commercial COz sensors by replacing the original gas permeable membranes with a valinomycin impregnated membrane, and the original IFS with the IFS described above. Carbon dioxide measurements were made by monitoring the potential of the internal glass pH electrode vs the internal Ag/AgCl reference. Potassium measurements were made by monitoring the transmembrane potential of the outer valinomycin impregnated silicone rubber membrane (i.e. between the IFS Ag/AgCl electrode and an external double-junction Ag/AgCl referernce electrode). Blood measurements with the model combination sensor were conducted with 1unit of packed red blood cells (RBCs,Red Cross Blood Services, Detroit, MI) diluted in 1005mL of normal saline plus 163 mL of deionized water plus 32 mL of 1mol/L NaHCO3 The blood suspension was anticoagulated with heparin and citrate, phosphate, dextran, adenine solution (CPDA). Before and after blood measurements, the model sensor was span-calibrated at 37 "C by using the calibrant solutions described above. Blood measurements were conducted for 6 h by immersing the model combination sensor and reference electrode in 100 mL of vigorously stirred RBC suspension thermostated at 37 OC. Blood pC02 levels were varied by equilibration of the blood with different gas mixtures (38- and 80-Torr pCOz in Nz) via continuous purging of the headspace above the blood. Changes in blood potassium levels were accomplished by making additions of 2.5 mol/L KCl and by dilution of the blood with normal saline. Throughout the blood measurements, periodic samples were analyzed for potassium on a NOVA-6 blood electrolyte analyzer and for pCOz on a Radiometer ABL-4 blood gas analyzer. Combination Catheter Sensor Fabrication and Measurements. Potassium-responsive catheter electrodes were constructed from valinomycin impregnated silicone rubber tubing that was sealed at the impregnated end with a 2-mm plug of The open silicone rubber adhesive (GC Electronics, Rockford, E). end of the tubing was inserted 1.3 cm into the tip of a in. long Luer-lok injection adapter (Medex, Inc., Hilliard, OH) containing a 1 mm diameter vent hole in its side. The adapter tip was then sealed by injection of silicone rubber adhesive around the tubing walls. After curing overnight, the valinomycin silicone rubber tubing modules (Val-SRTMs) were filled with surfactant containing IFS and stored at 4 "C in glass tubes filled with IFS. Internal Ag/AgCl reference electrodes were assembled from Teflon-coated silver wire (Medwire, Mount Vernon, NY, 0.007 in. o.d., tip coated with AgCl by anodizing for 60 min in 0.1 equiv/L HC1 at a current density of 0.4 mA/cm2) and miniature
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ANALYTICAL CHEMISTRY. VOL. 61, NO. 21, NOVEMBER 1. 1989
Flgure 3. Blood pump apparatus used lor flowingblood studles.
phone-jack connectors (Switchcraft No. TR-2A)) mounted in Luer-Ink syringe barrel tips. Potaeaium-responsive catheters were assembled by inserting a Ag/AgCI electrode into the Val-SRTM and sealing the vent hole with parafilm to prevent moistureshorting of the potentiometric circuit. Potassium catheter resistance measurements were made at r c " temperature in 0.125 mol/L KCI plus 0.025 mol/L NaHCOzsolution. Total membrane resistances were converted to specific resistances or resitivities ( 6 ) by using valinomycin impregnation lengths (ea. 2 0.2 cm) estimated from plasticizer deposits (see above) and from hromophenol blue dye impregnations. The internal pH electrode module of Figure 1was amembled in a fashion similar to that described for the original polymermembrane-based CO, catheter (15). Several details of assembly have heen modified in an effort to improve the 66% success rate reported for the fabrication of the original CO, catheter sensor: A 17.5-cm length of Tygon tubing (Cole-Parmer, 0.25 mm id., 0.76 mm 0.d.) was cut, and a pH responsive PVC membrane formed in its wall as previously described. The electrode tubing was then soaked overnight in 0.80 mol/L, pH 6.8 phosphate internal reference buffer. The internal Ag/AgCI electrode of Figure 1was prepared, as previously descrihed, from a 15.5-cm length of 34-gauge silver wire and then soldered to the inner conductor pin of a miniature phone jack connector (Switchcraft No. TR-2A). The shorter Ag/AgCI reference wire was made from a 6-cm length of Teflon-coated silver wire (as described above for the potassium catheter electrode) and soldered to the outer conductor pin of the phone jack mnnector. The two unused pins of the phone jack connector were removed and the Ag/AgCI wirepin joints reinforced and insulated with 1-em lenghts of in. i.d. heat-shrink tubing. The presoaked pH electrode tubing was then flushed and filled with fresh internal huffer hy using a 26-guage needle and syringe and the pH membrane end of the tubing sealed with a 2 mm long plug of silicone rubber adhesive. The 15.5-cm AgfAgCl wire of the electrode/phone-jack assembly was then inserted into the electrode tubing and the top of the pH electrode tuhe sealed with a low water-absorbing epoxy (0.5 g of Shell Epone 828 plus 0.19 g of Texam Jeffamine D230). After initial setting of the epoxy sealant, the internal pH electrode module was mounted in a trimmed 3-cm3Luer-lok syringe barrel, the tip of which was subsequently sealed by injection of silicone rubber adhesive. The assembled pH electrode modules were stored at 4 "C in glass tubes fded with IFS. Use of these assembly procedures enabled miniaturized pH electrode modules to he fabricated with nearly a 100% success rate. Combination K+/C02sensing catheters were constructed by inserting the internal pH electrode module into a valinomycin impregnated silicone rubber tuhe module containing potassium and bicarbonate IFS (see Figure 1). The vent hole of the ValSRTM was sealed with parafilm to prevent moisture-induced shorting of the potassium measurement circuit. Carbon dioxide measurements were made by monitoring the potential of the internal pH electrode (Le. between contacts 1and 2 of Figure l), and potassium measurements by monitoring the potential across
Flgure 4. Caiheter implant assembly used for in vlvo studies. 100 a)
*
;llp--,I 20 4 0
.-
'j,
n
2
4
I
12 1 4 1 b,
8
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4 6 TlME(h) Flgure 5. Results of blood measuremenis of CO, (a) and K+ (b) using the mods1 K+lpC02 sensor. In vitro values ( 0 )were determined wkh the ABL-4 BGA (CO,) and NOVA-6 (K+) instruments. Correlation of sensor CO, values (a) with in vitro values: slope 1.02 f 0.03, intercept = -1.10 Torr, r = 0,9909, n = 24. Correlation of Sensw K+ values (b) with in vivo values: sbpe = 0.962 0.008, intercept = 0.19 mmollL. r = 0.9993, n = 23. Equilibrations to C02 were achieved by purging hheadspace above the blood wkh various CO, 0
*
gas mixtures. the wall of the valinomycin impregnated silicone rubber tubing (i.e. between contact 1 of Figure 1 and an external reference electrode). By adoption of this modular design, catheter sensors can he easily disassembled (by removing the internal pH sensor) and their internal filling solutions replenished when necessary. The modular design also facilitates troubleshooting of malfunctioning sensors by allowing component electrodes to he easily interchanged. I n Vitro Blood Pump Characterizations of Catheter Sensors. The combination catheters were calibrated as described above. Blood stability measurements were conducted with the blood pump apparatus shown in Figure 3. A suspension of human red
ANALYTICAL CHEMISTRY, VOL.
blood cella was pumped through the system by a peristaltic pump (3.0 L/min; Sarns, Inc., Ann Arbor, MI) and thermostated at 37 "C by a Fisher Model 73 circulating water bath. Changes in pCOz were accomplished by equilibrating the blood to various gas mixtures with the use of a membrane oxygenator (Shiley, Inc., Model S-1OOA). Custom blend gas mixtures (Scott Specialty Gases) containing 1.4, 2.9, 5.67, 8.61, and 11.4% COz and 57.1, 35.7,21.4,10.7, and 2.9% 0,in a balance of N2were used to this end. The initial high potassium levels were varied by dilution of the blood with normal saline/bicarbonate solution. Blood samples were taken at 5- or 10-min intervals throughout the 6-h measurement period and analyzed for potassium on a NOVA-6 blood electrolyte analyzer and for pC02 on a Radiometer ABL-4 blood gas analyzer. All sensors were calibrated before and after blood measurements by using two-point span calibrant solutions. In Viuo Characterizatiom of K + / p C 0 2Catheters. In vivo measurements were performed on large mongrel dogs (ca. 15-20 kg) anesthetized with injections of sodium pentobarbital (Nembutal) and immobilized with injections of Pavulon (pancuronium bromide). AU drugs and KCl infusions were administered through the left cephalic vein. The right brachial artery was exposed for placement of a blood pressure sensor and for use as an arterial blood sampling site. The carotid and femoral arteries were exposed for use as catheter sensor implant sites. The catheter sensor implant assembly consisted of a 3.25 in. long, 14-gauge catheter guide (Beckton Dickinson), a Y-adapter, and an external Ag/AgCl reference electrode (see Figure 4). The external reference electrodes were housed in PVC tubes (0.51 mm i.d. X 1.52 mm 0.d.) containing 0.15 mol/L NaC1/AgN03 electrolyte gel (2% Natrosol250 HR gel) and were separated from the sample by a lo00 MW cutoff cellulose dialysis membrane (Spectrum, Los Angeles, CA) (21). Electrical contact was made between the reference electrode and the potassium-responsive tip of the catheter sensor by a normal saline solution containing heparin (2 or 5 units/mL) infused at 2 or 3.5 mL/h (Travenol AS2OS pump). Sensors were calibrated before the in vivo measurements by using the two-point calibration method described above (using 4.97% and 10.1% CO,; balance 0,calibrant gases). In vivo pC0, and K+ values were calculated from catheter sensor potentials by using eq 2 and 4 as derived from the standard Nernst relations shown in eq 1 and 3. (1) Eco, = Kco, + sco, 1% (PCOZ)in4VO
(4) Slope terms, SCO, and SK+, were determined from the two-point sensor precalibrations. Cell constants, Kco, and KK+,were calculated from precalibration slopes and initial in vivo pC0, and K+ values determined from in vitro analyses of an initial blood sample using the benchtop analyzers. During the 4-h in vivo measurements, blood samples were removed at 5-15-min intervala and analyzed for pC0, and K+ on Radiometer ABL-2 and NOVA-6 blood analyzers. The dogs' temperatures (esophogeal) typically varied in the range of 38 1 OC. No attempt was made to correct the instrumental or catheter sensor readings for these small temperature variations. The dogs were mechanically ventilated with 100% Op Blood CO, levels were varied by changing ventilation rates (ca. 4-35 breaths/min) and by introducing a dead space (ca. 325 cm3) in the airway of the respirator. Increases in blood potassium were induced by infusion of 20 mequiv of KCl over 20 min by using a Travenol infusion pump (Model AS2OS). Blood potassium was decreased by simultaneous injections of 20 g of dextrose and 20 units of insulin. Upon termination of in vivo measurements, the dogs were euthanized by bolus injections of KC1. Sensor implant arteries were immediately exposed, tied off, and excised at points above and below the implanted sensor. Each arterial wall was then cut open, and the sensor surfaces and inner arterial walls were examined for blood clots. Any clots found were quantitated by length and mass. The catheter sensors were recalibrated following the explant procedures. In a first in vivo experiment, systemic anticoagulation was
61,NO. 21, NOVEMBER 1, 1989 2369
Table I. Effect of Plasticizer and Salt Additives on the Resistivity of Potassium Catheter Electrodes -- - impreg.-.
nating soh
salt. mol/L DlastiDMDODAc i w PL resistivity: KTpClPB TpClPB DOS o-NPOE M%cm I
,
10
'2 3" 40
100 1.6 x 10-3 1.6 x 10-3
*
100
5b 6b
7.2 X
76 8b
7.2 x 10-3
38000 f 1500 6700 1100 39900 3600 6400 1040 39600 f 5400 17300 f 1300 4180 300 6630 210
100 100
*
"3.6 X 10" mol/L valinomycin in 1:l THF/xylene (1.25 mL total volume). b3.6X lo-$mol/L valinomycin in 1.25 mL of xylene. 'n = 3 potassium catheter electrodes.
achieved by bolus injections of heparin (56 uniia/kg of body weight every hour). A second experiment employed a continuous infusion of heparin. The activated coagulation time (ACT) test (22) was used to monitor heparin anticoagulant activity and to adjust infusion rates (every 30 min) according to established procedures (23).ACT tubes and reader (Hemochron-400)were obtained from Technidyne (Edison, NJ). The dog's base-line ACT was 153 f 13 s and was elevated to 550 s prior to implantation of catheter sensors. Elevated ACTS were maintained by continuous infusion of 250 units/mL heparin-saline solution at rates ranging from 52 to 104 units kg-I h-*. With this procedure, relatively constant anticoagulant activity was maintained throughout the experiment, with ACTS ranging from 322 to 591 s.
RESULTS AND DISCUSSION Model Combination Sensor. Before conducting investigations with the proposed catheter design of Figure 1, we studied the response of a model combination ion/gas sensor constructed from commercial Severinghaus-style CO, sensors. One of the goals of the model sensor investigations was to study the effect of valinomycin impregnation of the outer gas permeable membrane on the response of the CO, sensor. Sensors assembled with valinomycin impregnated membranes and with nonimpregnated membranes showed identical, Nernstian responses to C02. Likewise, valinomycin impregnation had no significant effect on sensor response times. From Figure 2 it is seen that the two measurement circuits of the combination sensor share a common conductor (i.e. the internal filling solution Ag/AgCl electrode). The model sensor was also used to investigate the independence of these two measurement circuits (data not shown). In a first set of experiments, different (constant) background levels of potassium (0, 1,and 6 mmol/L) were shown to have no significant effect on sensor's CO, calibration response. Likewise, different (constant) background pC02 levels (0,44, and 84 Torr) had no measurable effect on the sensor's K+ calibration response. Finally, the model combination sensor responded rapidly, independently, and in a Nernstian fashion to both analytes when simultaneous standard additions of potassium and bicarbonate were made to an acidic sample buffer. Together, these results demonstrated the independence and theoretical predictability of the two analytical signals obtained from the combination ion/gas sensor configuration. The results of continuous COz and K+ measurements using the macro in vitro sensor in whole blood are shown in Figure 5. The CO, and K+ signals obtained with the combination sensor showed nearly perfect correlation with periodic blood pC0, and K+ values determined on the bench-top blood analyzers during the entire 6-h measurement period (Figure 5 legend). These results demonstrate the combination sensor's suitability for continuous, simultaneous monitoring of blood K+ and pC02 levels. CombinationK+/pC02Catheter Sensor. Two methods
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of impregnating silicone rubber with valinomycin have been reported. Fogt and co-workers used an impregnating solution composition of 3.3 mmol/L valinomycin in xylene swelling solvent and a total impregnation time of 2 min (20). Marsoner et al. employed THF, freon, or octane as swelling solvents, with a valinomycin concentration of 25 mmol/L and a total impregnation time of 5 min (24). Initial attempts to prepare potassium catheter electrodes by either technique yielded potassium electrodes with large static sensitivities and high membrane resistivities (6 ca. 20000-40000 M k m ) independent of swelling solvent and impregnation times employed. Conventional static insensitive PVC K+ electrodes exhibit resistivities in the range of 100-1OOOMQ-cm,which may serve as target resistivities for producing static insensitive K+ silicone rubber catheter electrodes. Attempts to reduce the resistivity of valinomycin impregnated silicone rubber tubing focused on the addition of lipophilic salts (KTpClPB and DMDODA-TpClPB) and plasticizers (DOS (e = 3.89) and o-NPOE (e = 24.2)) to the valinomycin impregnating solution. Mobile, lipophilic anionic sites have been shown to reduce the resistivities of ion-selective membranes (17,25). Also, ionophore mobilities in polymeric ion-selective membranes are strongly dependent on the plasticizer content of the polymer and on the dielectric constant of the plasticizer (26,27). Because silicone rubber is a nonplasticized, cross-linked network polymer, it was reasoned that addition of plasticizer may reduce tubing resistivities by increasing the mobilities of the ionophore charge carriers and by increasing the dielectric constant of the silicone rubber matrix (dimethyl silicone rubber e = 2.5-3.0). From Table I, it is seen that addition of plasticizer caused a 6-10-fold reduction in tubing resistivities, while lipophilic salts had no beneficial effect on resistivities relative to plasticizer additives. Thus, it may be concluded that addition of plasticizer to the silicone rubber matrix is more beneficial than increasing the concentration of charged sites in the membrane. We believe that the benefits of the added plasticizer derive from the enhanced mobility of the ionophore in the presence of the plasticizer. Moreover, the similar effects of both DOS and o-NPOE suggest that the dielectric effects of the plasticizers are not as important as their plasticization capabilities. The lower dielectric DOS plasticizer was used in all subsequent impregnations because it affords a wider K+ response range than that of the higher dielectric o-NPOE (27). Finally, it should be noted that none of these additives reduces tubing resistivities to the range (100-lo00 M h m ) required for static insensitive potassium-responsive electrodes. The effect of silicone rubber source on tubing resistivities was also studied by comparing tubing extruded from Rhone-Poulenc (RP) (RS-1927) and General Electric (GE871) silicone gums supplied and extruded by Patter Products (Beaverton, MI). Direct comparisons of R P and GE tubing resistivities were made by impregnating several extrusion lots from each tubing type with 3.6 mmol/L valinomycin and 100 pL of DOS dissolved in 1.25 mL of THF. Resistivities of RP tubing lots ranged from 6OOO to 17 OOO Mil-cm. By contrast, potassium electrodes prepared from GE tubing were static insensitive, with resistivities ranging from 140 to 800 MQ-cm (well within the target resistivity range). These pronounced differences in tubing resistivities point to some property of the silicone rubber acting as a major determinant of tubing resistivities. Such relationships between the physical and chemical properties and the resistivity of silicone rubbers have been cited recently (28-32). All subaequent potassium catheter electrodes were prepared with low-resistivity, static insensitive GE tubing as described in the Experimental Section. Catheter electrodes prepared in this fashion showed greater than 99% of the theoretical Nernstian sensitivity (slope = 58.99 f 0.08 mV/decade; n =
6 K+ catheter electrodes; T = 25 "C) to potassium ions extending from 0.25 to 63 mmol/L (normal serum potassium range = 3.3-5.5 mmol/L). The reproducibility of catheter electrode equilibrium potentials was better than f0.2 mV (*0.8% K+ activity), exceeding that required for reliable whole blood measurement of K+ (33). Selectivity coefficients (WJ;) were determined by the separate solutions method (34)using 0.1 mol/L solutions of the chloride salts of Na+, H+, Li+, h@+, and Ca2+ interferent ions. For 2-day-old Val/SR catheter electrodes (n = 5), all selectivity coefficients = -4.020 f 0.008, = -4.40 f 0.12, mLi+ = -4.230 f 0.006, l@%,v = -4.692 f 0.006, mt,,,2= + -4.616 f 0.009) were identical with those of a conventional PVC membrane electrode and were suitable for direct whole-blood measurements of potassium ions (33). These selectivity coefficients remained constant when monitored over a 30-day period, excepting those of H+ ions, which increased approximately 2 orders of magnitude from log Kf$t,H+= -4.4 to -2.5. During the same period, PVC membrane electrode proton response remained constant. Two likely mechanisms of increased proton-exchange capacity of the silicone rubber matrix include the use of fumed silica (SiOz) filler particles in most silicone rubber formulations and moisture-induced hydrolytic depolymerization of the silicone rubber polymer (35),both of which may act as sources of SiOproton-exchange sites in the silicone rubber membrane. Further investigations of the pH response of silicone-rubber-based electrodes are not warranted since the increases in pH response to not compromise the accuracy of potassium determinations at physiological pH values (33). However, the increased pH response does necessitate buffering of calibrant solutions in the physiological pH range in order to obtain accurate potassium calibrations. Shelf lives of the catheter potassium electrodes were determined by periodic span calibrations over a 30-day period, during which time all electrodes retained greater than 95% Nemstian potassium sensitivity (SK+) and cell constants (KK+ of eq 4) changed by less than 5 mV. Finally, the Val/SR catheter electrodes were h t e d for accuracy of blood potassium determinations by 5 h of continuous measurements in fresh whole dog blood. Near-perfect correlation between catheter and NOVA instrumental potassium values was obtained. Combination K+/pC02catheter sensors were assembled and calibrated as described in the Experimental Section. The sensors responded independently, and in a Nernstian fashion (Sco, = 55.1 f 0.3 mV/decade; SK+ = 57.2 f 0.4 mV/decade; T = 25 "C), to simultaneous changes in K+ activity and pC02. Potassium response times, as defined by IUPAC (36),were less than 10 s, while 95% equilibration to pC0, changes was attained in less than 2 min at physiological pCOz levels. Within-day calibration reproducibility waa assessed by generating several calibration curves over a period of 6 h. All combination catheter sensors showed better than f0.5-mV reproducibilities (standard deviations) in equilibrium potentials throughout the test period. The temperature sensitivities of miniature PVC-based pH electrodes, PVC-based pC02 catheters, Val/SR potassium catheters, and combination K+/pC02catheter sensors were determined by performing aqueous two-point calibrations a t 25 and 45 "C. The results of these measurements were compared with temperature coefficients reported for other catheter sensor designs. The miniature polymer-membrane-based pH electrodes had temperature sensitivities (-0.33 mV/"C) that were 3 times smaller than those reported for Pd/PdO miniature pH electrodes (-1 mV/OC, -3.7% H+activity/"C) (37). Likewise, polymer-membrane-based pC0, catheters had temperature sensitivities (-0.54 mV/"C) 4 times smaller than those reported for PdO-based pC0, catheters (-2.2 mV/OC, -8.0% pCO,/"C) (14). Thus, a lower temperature sensitivity
mt,~+
(mt,Na+
ANALYTICAL CHEMISTRY, VOL. 61, NO. 21, NOVEMBER 1, 1989
1
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100
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Figure 6. Results of blood pump study for two Combination catheter s e n m . In vitro values (0)determined with the ABL-4 BOA (CO,) and NOVA-6 (K') instruments. Correlation of sensor COP values (a) wlth in vitro values: slope = 1.000 f 0.007, intercept = -0.71 f 0.33 Torr, r = 0.9986, n = 62. Correlation of sensor K values (b) with in vitro values: slope = 1.042 0.008, intercept = -0.22 f 0.05 mmol/L, r = 0.9964, n = 62.
*
is another apparent advantage of employing polymer-membrane-based pH electrodes as internal sensing elements in pCOz catheters. Temperature sensitivities of the siliconerubber-based potassium catheters (-0.47 mV/"C, -1.7% K+ activity/OC) were comparable to those reported for PVC-based potassium catheter electrodes (38).As expected, temperature coefficients of the combination catheter sensors were equal to those observed for the component K+ and pC02 catheters. Sensors designed for in vivo monitoring applications must meet stringent calibration stability requirements. From the Nernst equation (e.g. eq 1 and 3)) two types of calibration instability are apparent: slope or sensitivity drift, and cell constant (K") drift. Cell constant drift can be corrected for by periodic recalibration to a blood sample determined on a bench-top instrument. Ideally, blood sample recalibrations would be required no more than once every 8 h. A target Kmu drift rate may then be set a t less than f l mV drift (f4% analyte concentration) per 8 h, or ic0.125 mV/h. Sensitivity drift, on the other hand, cannot be easily corrected for after sensor implantation. Therefore, a conservative target sensitivity drift rate may be set a t less than f l % change in sensitivity per 12 h, or *0.083% /h. The long-term stability of the combination catheter sensors in aqueous solutions was assessed by continuous monitoring of three sensors as they were cycled between 1.0 mmol/L K+/37-Torr pC0, and 10.0 mmol/L K+/91-Torr pC02 span calibrants (37 "C) for a 45-h period. All sensors were calibrated before and after the stability measurements so that drifts in calibration parameters could be assessed. The K," and sensitivity drift rates for the combination catheters were all less than the target drift rates: KE&+ = -0.022 f 0.022 mV/h; K ~ ~ ~=c -0.029 o, f 0.008 mV/h; slope$' = 0.027 f 0.019%/h; slope?$ = 0.079 f 0.017% /h. Furthermore, the catheter sensor K,u drift rates were actually superior to those of commercial Orion COZ and
' i 10
2
4
TIME (h)
Figure 7. Continuous response of combination catheter sensors toward COP(a)and K+ (b) during in vivo measurements with continuous heparin infusion. In vitro values (0)were determined with the ABL-2 BGA (CO,) and NOVA-6 (K') instruments. Correlation of three catheter sensors' CO, values (a) with in vitro values: slope = 1.07 f 0.02, intercept = -1.8 f 0.9 Torr, r = 0.984, n = 93. Correlation of two catheter sensors' K+ values (b) with in vitro values: slope = 0.96 f 0.04; intercept = 0.19 f 0.16 mmol/L, r = 0.962, n = 55 (excluding
errant values that resulted from blood contamhation of a catheter reference electrode (c)). conventional PVC K+ macrosensors (n = 1): KL&+ = 0.051 mV/h; @:t802 = 0.10 mV/h; slopegy' = 0.026%/h; slope%? = 0.050%/h. Thus, the combination K+/pC02catheter design is capable of accurate monitoring of both analyks for extended periods without recalibration. Blood Pump Studies. To determine the feasibility of using the combination K+/pC02 catheter sensors for in vivo measurements, we first tested the performance of these devices in flowing blood. These studies were conducted with the b l d pump apparatus of Figure 3 for measurement periods of 6 h (see Experimental Section). The results of these measurements are shown in Figure 6. The combination catheter sensors' COz and K+ traces gave nearly perfect correlation to bench-top instrument readings over the entire 6-h test period (Figure 6 legend). In Vivo Studies. In vivo experiments were performed in anesthetized dogs. Catheter K+ and pCOz signals were calibrated in vivo by using precalibration slopes and an initial discrete blood sample, analyzed on bench-top instruments, to calculate the sensors' in vivo cell constants. In a first experiment, the dog was anticoagulated by hourly injections of heparin. Two combination catheter sensors were successfully implanted during this experiment (right femoral and left carotid arteries), and continuous in vivo measurements were conducted for 4.5 h. During this fiit experiment catheter signals were in reasonably good agreement with instrumental values for the first 2.5 h of the experiment. After this time, however, large discrepancies between catheter traces and instrumental values were observed. After termination of the experiment, large (2-3 cm long, 0.011 f 0.002 g), solid blood clots were found covering the sensor surfaces and blocking
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the implant arteries. These heavy clot formations are the most likely cause of the discrepancies between catheter signals and instrumental readings later in the experiment. Such discrepancies would result from changes in the local chemical environment of the sensor caused by the physical presence of blood clots and by the respiration of cells entrapped within the bloot clots (e.g. COz evolution). Obviously, such local changes would not be reflected in bench-top analyses of discrete blood samples. The possibility of drift in the sensors’ K+ and C 0 2 calibration responses was checked by recalibrating the sensors following removal of the blood clots. No such drift was observed. In order to assess the in vivo performance of the combination catheter sensors in the absence of blood clots, a second experiment was performed using continuous infusion of heparin (see Experimental Section). Three combination catheter sensors were successfully implanted during this second experiment. However, one sensor’s K+ signal showed large fluctuations that were attributed to shorting of the K+ measurement circuit by body fluids contacting the top of the sensor. The continuous traces from the remaining two sensors are shown in Figure 7. Note that the deviations of one catheter’s K+ trace in Figure 7b from 3.5 to 4 h were caused by back-diffusion of blood through the catheter implant guide to the reference electrode assembly. Subsequent flushing of the reference electrode arm and catheter guide with normal saline restored agreement between the catheter and instrumental K+ values. Excluding the errant K+ values, the COz and K+ catheter traces showed nearly perfect correlation with bench-top instrumental readings throughout the 4.25-h measurement period (Figure 7 legend). Indeed, for the entire 4.25-h in vivo measurement period, the mean difference between catheter and instrumental pCOBvalues (n = 93) was 1.2 f 3.3 mmHg, and the mean difference between catheter and instrumental K+ values (n = 55) was 0.04 f 0.33 mmol/L. However, a slight increase in the discrepancy between catheter and instrumental K+ values was noted with time. In the first half hour of in vivo measurements the average K+ discrepancy was -0.09 f 0.25 mmol/L (n = 12), while in the final half hour of in vivo measurements the average discrepancy was 0.28 f 0.20 mmol/L (n = 14). In spite of the slightly increased discrepancy, the catheter K+ signals still retain acceptable accuracy even after 4.25 h of continuous in vivo measurements. Following these measurements, the sensors were explanted and examined for clot formations. The efficacy of the continuous heparin infusion protocol was evidenced by the complete lack of measurable clots on the sensor and arterial wall surfaces. Thus, in the absence of significant blood clot formations, the combination catheter sensor is capable of continuous in vivo monitoring of K+ and pCO,, with accuracies comparable to those of conventional bench-top blood gas and electrolyte analyzers. Comparison of the results from these two in vivo experiments demonstrates the detrimental effect of blood clots on the accuracy of the in vivo catheter sensors. Since these effects can be eliminated by maintaining adequate systemic heparinization during in vivo measurements, it is likely that the proposed combination catheter design will be immediately applicable to the many clinical procedures that require systemic anticoagulation (e.g. heart surgery, hemodialysis, etc.). To make the sensors more generally applicable for routine clincial use, reductions in the thrombogenicity of the sensor’s surfaces will be required. To this end we attempted to adsorb heparin to valinomycin impregnated silicone rubber membranes according to the procedure of Grode et al. (39). Unfortunately, this adsorption procedure destroyed the K+ response of the valinomycin impregnated silicone rubber membranes. Thus, further efforts to improve the blood compat-
ibility of the combination catheter sensor must focus on alternate means of immobilizing anticoagulant and/or antiplatelet agents, which do not impair the ionic response of the ionophore impregnated silicone rubber tubing. With such improvements, the combination catheter sensor described here should prove useful in a wide variety of biomedical applications.
ACKNOWLEDGMENT We thank Mallinckrodt Sensor Systems for assisting us in performing the blood pump studies reported here. We also thank Dr.Paul Knight and the staff of the Anesthesiology Laboratory of the University of Michigan Medical School for their help in performing the in vivo studies.
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RECEIVED for review June 8,1989. Accepted August 14,1989. We gratefully acknowledge The National Institutes of Health (GM No. 28882) for support of these studies.
