Chemical Sensors for Bedside Monitoring of Critically Ill Patients Michael E. Collison Analytical Development Division Lilly Research Laboratories Lilly Corporate Center Indianapolis, IN 46285
Mark E. Meyerhoff Department of Chemistry University of Michigan Ann Arbor, Michigan 48109
"He will manage the cure best who has foreseen what is to happen from the present state of matters." The Book of Prognostics by Hippocrates (ca. 460-377 B.C.) Empedocles (ca. 500-430 B.C.), the "Father of Greek Chemistry," originated the theory of elements and compounds in which all substances derive from combinations of the four "elements": earth, air, fire, and water. Hippocrates adapted Empedocles' view of nature to a similar theory of human physiology (1), in which health existed when four humors, or body fluids (blood, phlegm, yellow bile, and black bile), were present in proper proportion to each other. Thus Hippocrates' call to assess "the present state of matters" may have been the first historical recognition of the need to monitor closely the physiological state of the human body in order to manage a more effective cure. 0003-2700/90/0362-425A/$02.50/0 © 1990 American Chemical Society
In the two and a half millennia since the Hippocratic doctrine, models of human physiology have grown in complexity, and the need for physiological monitoring has progressed beyond the balancing of the four humors. Indeed, modern clinicians are increasingly dependent on reliable measurements of key chemical variables in blood for proper clinical diagnosis and therapy. Until recently, physiological monitoring has been accomplished by analyzing discrete blood samples at centralized clinical chemistry laboratories remote from the patient's bedside.
REPORT This conventional approach, however, provides only "historical" measurement values because it takes precious time to transport the sample to the main lab, have the lab perform the appropriate test, and then return the test results to the clinician or nurse. In some hospitals, this so-called turnaround time can be 30 min or more. These inherent delays may prevent early detection of rapidly changing blood chemistries, particularly for critically ill, surgical, and other unstable patients. Moreover, for certain key analytes (e.g., blood gases), test values measured in remote laboratories may
not reflect the true physiological levels of the analyte because of errors induced by the required sample transport process (2). As the number of highly specialized and sophisticated critical care/surgical units at hospitals has grown, there has been an increase in the demand for measurement technologies that can monitor changes in a select list of critical care analytes (see Table I) at the bedside of seriously ill and surgical patients. Devices capable of continuously or semicontinuously monitoring blood gases, electrolytes, and certain metabolites can provide real-time information that can result in improved diagnosis and more timely therapeutic intervention. This growing demand for convenient critical care chemical testing technology challenges the analytical chemist to bring conventional analytical techniques from the clinical laboratory directly to the patient's bedside without compromising measurement accuracy and reliability. Although various complex technologies can be used to determine blood levels of the species listed in Table I, bedside analysis requires inexpensive and rugged testing instrumentation. In addition, the equipment should be simple to operate, ideally by nonlaboratory personnel with little or no training in analytical chemistry. We will discuss recent efforts to develop and imple-
ANALYTICAL CHEMISTRY, VOL. 62, NO. 7, APRIL 1, 1990 · 425 A
REPORT ment modern chemical sensing technologies (both electrochemical and optical) for measuring blood gases (and related parameters) and electrolytes. Enzyme-based biosensors suitable for bedside testing are not described here; their current state of development has
Table I.
been discussed extensively in recent reviews and monographs (3, 4). Blood gas and electrolyte measurements Po2 and Pco 2 · Blood gas analysis mandates simultaneous measurement of
Critical care analytes and their normal ranges in blood
Blood gases and related parameters
Metabolites
Electrolytes
Po2
80-104 Torr
Na+
Pco2
33-48 Torr
K+
pH
7.31-7.45
Ca2+
Hematocrit
40-54%
cr
Total hemoglobin Unsaturated hemoglobin
13-18g/100mL
135-155 mmol/L 3.6-5.5 mmol/L 1.14-1.31 mmol/L 98-109 mmol/L
Glucose Lactate Creatinine Urea
70-110 mg/100mL 3-7 mg/100mL 0.9-1.4 mg/100mL 8-26 mg/100 mL
95-100%
pH and the partial pressures of oxygen and carbon dioxide (Po2 and Pco 2 ) directly in undiluted whole blood samples. Almost all laboratory blood gas and electrolyte determinations are performed on commercial instruments that employ conventional Clark-style polarographic oxygen sensors for Prj2 measurements (Figure la). Such sensors operate in an amperometric mode and consist of a microplatinum cathode, a Ag/AgCl anode, an inert electrolyte, and an outer gas-permeable membrane. A negative voltage applied to the cathode forces the reduction of oxygen to water, hydrogen peroxide, or hydroxide ions. The outer oxygen-permeable membrane protects the platinum surface from protein fouling. This membrane also controls the rate of diffusion of oxygen from the sample to the surface of the platinum and thus the rate of electrochemical reaction (measured as a current flow). The flux of O2 to the cathode is directly related to sample Pç,0 levels. Measured currents are also dependent on membrane thickness, diffusivity of 0 2 through the membranes, and cathode size. Because these parameters can change with time, the output of such nonequilibrium devices tends to drift, necessitating frequent recalibration. This has hampered efforts to devise reliable catheter-type polarographic Po2 sensors that are suitable for continuous in vivo measurements over extended time periods (5). Newer optical Po2 sensors (Figure lb), already used in commercial extracorporeal bedside measurement systems, employ fluorescent dyes that are trapped or chemically immobilized in a thin layer adjacent to an optical fiber(s) (6, 7). In the presence of oxygen, the fluorescence intensity of the dye is quenched in an amount proportional to the Po2 level in the sample. Such sensors are equilibrium devices in that no oxygen is consumed during the measurement. In principle, optical oxygen sensors offer the advantage of potential long-term stability over conventional polarographic O2 sensors, particularly for continuous in vivo measurements. In practice, however, such improved stability can be difficult to achieve because of the low degree of quenching that is actually observed (20%), the instability of source and detector, and the photodegradation of the immobilized dyes.
Figure 1. Gas sensor designs suitable for whole blood measurements. (a) Clark-style amperometric Po2 sensor, (b) optical P02 sensor based on fluorescence quenching, (c) optical and potentiometric PCOz sensors, and (d) differential potentiometric PC02 sensing arrangement using dual polymeric H+-selective sensors.
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Both optical and potentiometric Pco 2 sensors are based on detection of equilibrium pH values in a small volume of electrolyte held behind an outer gas-permeable membrane (Figure lc). When the partial pressure of CO2 is
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REPORT equal on both sides of the membrane, the pH of the inner electrolyte is gov erned by the dissociation of carbonic acid formed from diffusing C 0 2 in ac cordance with the well-known Henderson-Haselbach equation: pH(thin layer) = l o g (aHC0JK\
· Pc02)
(1) where K\ is the first dissociation con stant for carbonic acid. In practice, the activity of bicarbonate in the inner electrolyte is held constant by using a high concentration of sodium bicar bonate (e.g., 50 mmol/L) such that pHfthin layer) = K ' -
log P C 0 2
(2)
K' takes into account both the fixed amount of biocarbonate and the disso ciation constant for carbonic acid. In traditional blood gas instruments, the pH of the inner bicarbonate solution is monitored by a standard glass mem brane combination electrode that pro vides a wide dynamic Pco2 measuring range. Optical Pco2 sensors for extra corporeal and in vivo sensing employ optical pH probes as inner transducers. Because these devices use pH-sensitive indicator dyes (6, 7), the dynamic mea suring range of the resultant Pco 2 probes is more restricted but still quite adequate for most biomedical situa tions. With the advent of solvent/polymer ic membrane electrodes for measure ment of specific ionic species, including protons, a novel differential detection scheme can also be used to measure Pco 2 levels reliably in whole blood (Fig ure Id). T h e analyte C 0 2 diffuses through the H + -selective polymeric membranes of two "pH sensors" pre pared by doping the membranes with neutral trialkylamine species (e.g., tridodecylamine). A small volume of bicarbonate/NaCl solution is used as the inner solution behind one pH mem brane, and the inner solution of the sec ond electrode is composed of a strong buffer. If both polymeric membranes respond to the proton activity of the sample identically, then at equilibrium this sample pH will cancel in the differ ential measurement mode, yielding:
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0.059 log iV K a m p l f l
(3)
The cell constant, K, is dependent on the pH of the strong buffer used as the inner reference fill of pH electrode E2 (Figure Id). This differential P C o 2 measurement is an attractive approach for systems that require disposable/ planar-sensing elements. pH and electrolytes. Potentiometric ion-selective membrane electrodes are the dominant methodology em ployed for determining pH and electro
428 A · ANALYTICAL CHEMISTRY, VOL. 62, NO. 7, APRIL 1, 1990
lytes in whole blood. Although pH and Na + -selective glass membrane elec trodes are still used in many laboratory instruments, newer bedside testing in strumentation, including some proto type in vivo sensors, rely on modern solvent/polymeric membrane electrode technology to measure H + , Na + , and other important ionic species such as Ca 2+ and K + (8). Solvent/polymeric membranes function as ion-selective transducers only when an appropriate ionophore is incorporated within the polymeric film. The ionophore serves as a reversible and reusable binding re agent that selectively extracts the ana lyte ion into the organic membrane phase, thereby creating a charge sepa ration or phase boundary potential at the membrane sample/interface. In the differential C 0 2 sensor de scribed above, the alkylamine com pounds doped into the polymer mem brane are the proton-selective ionophores. Indeed, the potential E
