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In vivo blood-gas and electrolyte sensors: Progress and challenges
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ln viva blood-gas and electrolyte sensors: progress and challenges Mark. E. Meyerhoff Ann Arbor, Ml, USA The management of critically ill patients requires the frequent and accurate measurement of blood gases and electrolytes in undiluted whole blood. To improve patient care, there is a growing desire to monitor these important parameters on a continuous basis via the use of intra-arterial chemical sensors. The progress and challenges relating to the development of suitable probes for making reliable analytical measurements within the confines of the human radial ‘artery are summarized.
Background The care of critically ill hospital patients [intensive care unit (ICU), coronary care unit (CCU), etc.] mandates the frequent measurement of a select group of critical care analytes in undiluted whole blood. Tests recognized as being essential for patient management include the blood gases (pH, POZ, PCO,) and electrolytes (Na+, K+, iCa2+). Normal range values for these analytes are summarized in Table 1. While historically these tests have been performed in centralized or satellite laboratories remote from the patient, long delays to obtain test results (so-called “turn-around time”) have created a demand for technologies that enable these measurements to be carried out more rapidly at or near the patient’s bedside. Although progress has been made recently via the developTABLE 1. Normal range values for blood-gas and electrolyte measurements Oxygen partial pressure (IQ) Carbon dioxide partial pressure
80-l 04 mmHg
(X02)
33-48
PH sodium (Na+) potassium (K+) ionized calcium (iCa”)
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mmHg 7.31-7.45 135-l 55 mmol/l 3.6-5.5 mmol/l 1 .14-l .31 mmol/l
ment of convenient to use, portable, blood-gaselectrolyte instruments [I], these instruments are only useful for discrete sample measurements, and are not capable of monitoring blood-gas and electrolyte values on a continuous basis. The availability of truly implantable blood-gas and electrolyte sensors (either electrochemical or optical) could provide clinicians with essentially “real time” diagnostic information and this could result in more timely therapeutic intervention and thus save lives. Twenty gauge arterial line catheters are placed in more than half of all ICU patients, typically for up to three days. These A-lines are implanted within the radial artery and are often equipped with pressure transducers to monitor blood pressure continuously (see Fig. la and b). At the same time, the A-line provides an open port to readily sample (via a syringe) the patient’s blood for periodic discrete sample blood tests. A steady stream of saline-heparin flows through the cannula to keep the implant site free of blood clots. The ultimate goal of researchers developing in viva sensors is to design blood gas and electrolyte probes that are small enough to slide down through the cannula of the A-line (see Fig. lb) so as to be in continuous contact with the flowing arterial blood (with mean linear flow velocities of 3.6 cm/s and mean volumetric flow-rates of 0.65 ml/s), enabling real-time measurements of blood-gas and electrolyte levels. To be useful clinically, in vivo blood-gas and electrolyte sensors must fulfill ull of the requirements outlined in Table 2. Meeting these requirements is an enormous challenge. For example, the I.D. of a 20 gauge A-line cannula is only 0.83 mm. Thus, in order to obtain accurate blood pressure signals, in vivo sensors must be considerably smaller, i.e., they must occupy < X)-60% of the total cross-sectional area of the cannuIa. While large size sensors (e.g., O.D. 1 mm) can be employed for preliminary development work, and as useful research tools for animal studies, ultimate application in humans requires the miniaturization of sensors to a size where multiple transducers, for example pH, PC02, P02, can all be combined into a single probe with a total O.D. of < 0.55 mm. In
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1993 Elsevier Science Publishers B.V. All rights reserved
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addition to the sensors, the same probe must also contain a thermocouple so that blood temperature can be monitored simultaneously. These readings are then used to compensate for any known temperature sensitivities of the sensors, and to provide temperature-corrected blood gas values (i.e., the blood pH and partial pressures of the gases if
a)
KJ-
salinelheparin flush solution
blood-gas/electrolyte monitor
A-line
catheter
b)
heparlnlsaline drip ‘;
TABLE 2. Requirements . . . . . .
analytical chemistry, vol.
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sensors
Small size (compatible with 20 gauge A-line) Long term calibration stability (one calibration insertion) Reliable values for 72 h under widely physiological conditions Blood compatibility (non-thrombogenic) Rapid response times Ability to sterilize
prior to varying
measured at 37°C in a conventional blood-gas instrument). The entire final probe must also be able to withstand appropriate sterilization (e.g., via ethylene oxide, gamma radiation, etc.) so that microbial contamination of the patient is not a concern. Beyond severe size restraints and the need to sterilize, signal stability is another major hurdle. Since it is most desirable that calibration of in viva sensors be performed only once prior to insertion, long term signal stability is very important, particularly when analytical accuracies of < + 5% are desired for the entire implant time (up to 72 h). While alternate in viva recalibration schemes involving periodic in vitro measurements on discrete samples of blood are possible (i.e., during the 72 h period), if the required frequency of such in vivo recalibrations is high (perhaps more than once a day), the advantages of the continuous in vivo measurement are partly negated. Analytical performance of implanted sensors, however, can also be influenced dramatically by the complexity of the in vivo environment, especially the propensity of blood clots to form on the surface of foreign materials (i.e., biocompatibility). In addition, the precise position of the sensor within the artery can influence output signals (so call “wall effects” [2]). These two issues relate directly to the innate difficulties of making reliable intra-arterial measurements even with the smallest and most stable analytical probes, and will be addressed in more detail in a separate section below.
Available sensor technology Fig. 1. (a) Schematic of arterial line catheter placement and associated components for continuous monitoring of blood pressure as well as blood-gases and electrolytes; (b) expanded view of A-line catheter with in vivo chemical sensor(s) inserted through cannula.
In vivo blood-gas and electrolyte sensors can be based on either electrochemical or optical measurements. Progress in the development of sensors based on these transduction modes are summarized below.
trends in analytical chemistry, vol. 7.2, no. 6, 1993
Electrochemical
sensors
The electrochemical probes are essentially miniaturized versions of the sensors already in use within large, commercial blood-gas-electrolyte instruments [3]. For example, PO;! levels can be determined with a catheter form of the classical Clark-style amperometric oxygen sensor (Fig. 2a). Various forms of such single parameters probes have been commercially available for some time with outer diameters on the order of 0.55 mm. Electrochemical reduction of oxygen (e.g., O2 + 2H+ + 4 e- + 2 OH-) occurs at a silver or platinum cathode, which is polarized at ca. -0.7 V vs. an Ag/AgCl reference/anode electrode. The resulting current flow (generally in the O-100 nA range) is directly proportional to blood PO*. Signal drift is a major concern with these devices, since the diffusion rate of oxygen through the outer polymer material can change with time. In addition, the surface area of the cathode can gradually increase as silver ions formed at the anode reference electrode migrate to the cathode and reduce, effectively creating a larger area cathode and thus a larger diffusion current for a given level of oxygen. Electrochemical pH measurements are generally made via miniaturized potentiometric ion-selective membrane probes, or metal/metal oxide based devices. The latter are subject to interferences from redox active compounds (e.g., ascorbic acid, uric acid, etc.), and are thus less attractive for direct blood pH measurements. The fragile nature and high electrical resistance of conventional glass
,_ _..._. Silver Cathode
(a)
Insulator Plug 04
Fig. 2. Sensing regions of amperometric oxygen (a) and tubular-style potentiometric pH (b) (or electrolyte) in viva sensors.
membrane pH electrodes preclude their use as practical transducers for intra-arterial measurements. Modem polymer-membrane-type pH electrodes, prepared by doping PVC or other polymeric materials with hydrophobic amine compounds [4] do offer promise as useful in vivo probes. Fig. 2b illustrates one configuration for such devices. When the polymer tubing wall is doped with the appropriate amine (e.g., tridodecylamine), and a suitable buffer (0.25 M phosphate pH 6.8) is employed to fix the activity of protons within the internal electrolyte solution in contact with the internal Ag/AgCl reference electrode, the potential of the catheter probe vs. a #suitable external reference electrode can be re1ate.d to the pH of the sample. For in vivo potentiometric measurements, the external reference electrode (e.g., Ag/AgCl, saturated calomel, etc.) can actually be placed outside the artery, in one arm of the A-line (see Fig. 1). The continuous saline-heparin flush solution can serve as a useful electrolytic contact (i.e., salt bridge) to the blood, as it flows between the inner wall of the cannula and the outer wall of the in vivo probe. In vivo measurements of sodium, potassium, and ionized calcium (free calcium) can be made with a similar working (Fig. 2b) (and reference electrode arrangement; however, instead of doping the tubing walls of the device in Fig. 2b with lipophilic amines, highly selective ionophores for sodium, potassium or calcium are used as membrane-active species [5]. For example, in vivo potassium sensors are prepared with valinomycin (K+ selective antibiotic) as the active membrane component while ionized calcium levels can be monitored with alkylphosphates or other specially designed synthetic neutral organic ligands doped within the polymer films. These membrane active ionophores are very hydrophobic compounds and thus leach very slowly from the polymeric tubing into the surrounding blood (non-harmful levels). Although there also has been considerable effort aimed at adapting such polymer membrane ion-selective electrode technology to modern microelectric fabrication techniques to prepare multi-parameter implantable “solid-state” ion sensors (e.g., ion-selective field effect transistors (ISFETs) [6]), thus far these devices have failed to achieve the level of stability and selectivity required for accurate in vivo measurements. Electrochemical in vivo measurement of PC02 levels can be accomplished using miniaturized
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external reference electrode (in flush line)
e silicone rubber tubing \
potassium bicarbonate/ potassium chloride solution
PVC tubing
I
buffer
pH responsive tubing
/
Valinomycln-impregnated silicone tubing
Fig. 3. Dual potentiometric K+/PC02 sensing catheter described by Collison et al. (see ref. 8).
variations of the classical Severinghaus-style PCOZ sensor design in which the pH of thin layer of bicarbonate solution housed behind a gas-permeable membrane changes as a function of the PC02 level in the sample. In an early catheter configuration introduced by workers at General Electric Corp., a PdO-type pH electrode was used to detect the pH changes in the bicarbonate layer [7], More recently, polymeric pH electrodes (such as those in Fig. 2b) have been utilized successfully to fabricate catheter-type PC02 sensors for in vivo measurement (in dogs). In the concentric tube design shown in Fig. 3, the inner-tubular pH electrode detects pH changes in a thin layer of bicarbonate held within the outer silicone rubber tube [8]. When doped with an appropriate ionophore (e.g., valinomycin), the outer silicone rubber tube can serve both as an outher gas-permeable membrane for the PC02 measurement, and as an ionselective membrane for the simultaneous measurement of an electrolyte ion, in this case potassium. The potassium measurement is made by monitoring the membrane potential across the wall of the outer ionophore-doped silicone tube. (i.e., Ag/AgCl wire 1) vs. external reference electrode in contact with saline-heparin flush. It is important to recognize that all potentiometric devices respond to analyte concentrations in a logarithmic fashion in accordance with the Nernst equation: J%~II= K + (O.O59/zi) log [analyte] sample where f&l is the measured potential between the working ISE catheter and an external reference electrode, Kis the cell constant (sum of all constant potentials, including reference electrode poten-
tials), zi is the charge on the detected ion, and [analyte] is the concentration of the measured species (H+, K+, Na+, iCa2+ , PCO2). Thus, for pH, K+, Na+, PC02 measurements, changes of only 1 mV in the cell constant between the calibration step and the actual in vivo measurement will cause variations of + 4% with respect to the measured parameter, making reliable in vivo measurements with such devices quite difficult, unless the exquisite e.m.f. stability can be achieved. In general, instability of the measurement instrumentation is not the limiting factor in achieving accurate quantitative results over long time periods; rather it is changes in the membrane-sensing chemistries and in the compositions of the internal reference solutions (via osmotic effects) that can often yield the observed drift in potentiometric signals. Optical sensors
Although efforts to develop various electrochemical probes for intravascular measurements have been ongoing for more than 30 years, much of the most recent work in this area has focused on the use of modern fiber-optic chemical sensors [2,9,10]. These devices are prepared by immobilizing a wide variety of suitable indicator reagent dyes at the end of a narrow fiber and measuring either absorption, fluorescence, or phosphorescence resulting from the interaction of light with the immobilized indicators. These fiber-optic probes have certain inherent advantages over electrochemical sensors in that they: can be miniaturized more easily (using thin optical fibers of 100-150 pm diameter); have less noise (no transduction wires required); require no reference electrode; and have the potential to offer greater signal stability by employing multiple wavelength measurements to correct for changes in effective optical pathlengths, photobleaching of dyes, etc. The latter advantage is realized by referencing measurements of optical signals that originate at wavelengths that contain analytical information to wavelengths at which the spectroscopic signal does not change as a function of the analyte level in blood (e.g., isosbestic points in the indicator dye’s absorption or fluorescence spectrum) [ 111. Obviously, elaborate source-detection monitors which can be conveniently coupled to the various miniature optical fiber assemblies are also required for routine analytical measurements with these
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Light a
Light
Light out
Light in
b
Clad
Dye
Cladding
Fig. 4. Three configurations used to prepare in vivofiber-opticchemical design; and (c) transmission arrangement used by Optex Inc.
optical sensing devices. One inherent disadvantage in optical sensing schemes, when compared to electrochemical sensors, is their limited dynamic measurement range (e.g., optical pH sensor; often less than 2 orders of magnitude). While for general analytical measurements this would certainly be problematic, as shown in Table 1, the normal ranges for gases and electrolytes in blood are rather narrow. Thus, provided that the maximum optical change for a given sensor occurs over a concentration range that encompasses the normal physiological values (e.g., by selecting indicator dye with proper pK, for blood pH determinations), the measurement range limitation is not a major issue with respect to using optical sensors for in vivo monitoring. Three basic optical sensor configurations have been examined for use in prototype blood-gas catheters. These are illustrated in Fig. 4. In the so-called Lubbers design (Fig. 4a) [ 111, the indicator is immobilized on the distal end of a quartz or polymer fiber. Photons from an appropriate source (usually pulsed) are sent down the fiber and light coming back up through the same fiber can be monitored at the same (absorption) or longer wavelengths (fluorescence, phosphorescence). Microporous reflective materials (e.g., gold films, etc.) can be coated over the dye layer to enhance the intensity of the returning optical signal. In the Buckles configuration (Fig. 4b) [ll], the dye is immobilized on the outside walls of the optical fiber (after stripping away the fiber’s outer cladding material). Optical signals originating from the dye molecules confined to the outer walls of the
sensors; (a) Lubbers design; (b) Buckles
fiber occur as a result of the evanescent wave phenomena. Optex Inc. (Houston, TX, USA), has proposed the bent fiber approach (Fig. 4c) in which a small region of the fiber is cut away and replaced with appropriate indicator dye matrix. Unlike the other two configurations, the Optex approach functions in a manner equivalent to conventional spectrophotometry in that the transmitted light that passes through a fixed optical pathlength dye region/chamber is the signal detected. The dye chemistry used to measure the bloodgases varies considerably. Most optical measurements of PO;, are based on fluorescence quenching of the indicator [e.g., perylene dibutyric acid, Ru(I1) tri(dipyridine), etc.], in the presence of oxygen in accordance with the Stern-Volmer theory: (I&&
= 1 + kP0 2
where Zpo2 is the fluorescence intensity at a given PO2 level, I0 is the intensity at PO2 =I0, and k is a quenching constant for the particular dye. Thus ZJZpo2 is linearly related to PO2. Phosphorescence, or phosphorescence-lifetime measurements of immobilized metal ligand complexes can also be employed (i.e., the binding of oxygen decreases the excited state lifetimes). Sensors based on measurement of luminescent lifetimes have the inherent advantage of being insensitive to changes in optical pathlength and the amount of active dye present at the tip of the sensor without need to perform measuremore complex multiple-wavelength ments. Fluorescence or absorption pH measurements
262
can be made by immobilizing any number of classical indicator dyes that have pK, values at or near the normal pH of blood (pH 7.4) (e.g., 8-hydroxy1,2,6-pyrene trisulfonate among many others). One lingering issue relating to the use of indicators for accurate pH measurements is the effect of ionic strength on the pK, of the indicators [12]. Since optical sensors measure the concentrations of the protonated or unprotonated forms of the immobilized dyes as an indirect measure of the hydrogen ion activity, large variations in blood ionic strength via changes in blood sodium levels can, in principle, affect the accuracy of such measurements by varying the pK, of the immobilized species. Since reference values for blood pH are determined by conventional glass electrodes that detect proton activity directly, variations in the ionic strength will cause larger discrepancies between reference instrument values and the values determined by the optical pH sensor. It is important to note that these variations in ionic strength will also cause the pH of the blood itself to change as measured electrochemically (effect on pK, of bicarbonate, etc.); however, because the electrode measures this defined value directly, it will always be the reference method to which the pH value reported by the optical sensor will be compared. Optical sensors for carbon dioxide employ the fiber-optic pH sensors as internal transducers in an arrangement quite similar to electrochemical PC02 gas sensors described above. An outer hydrophobic gas-permeable film separates the sample from an internal layer of sample PCOZ and this is detected optically. Typically, the same dyes used for direct blood pH measurements can be employed to construct the pH sensor used for monitoring blood PCO, levels, since the variation in pH within the bicarbonate film falls within the same pH range as blood values. Fig. 5. illustrates the manufacturing complexity of putting three optical sensors together to form a single multiparameter pH, PCO;! and PO2 intravascular probe. In this patented design [ 131 (Abbott Laboratories), the PO2 and PC02 sensors are coated with a silicone rubber material which serves as the gas-permeable (and ion-impermeable) membranes for the two gas-sensing fibers. Since pH measurements require that sample hydrogen ions have access to the immobilized indicator dye, the pH sensor is positioned slightly forward of the gas sensors and is coated with a much different, more hydrophilic polymer layer
trends in analytical chemistw, vol. 12, no. 6, 7993
light
c
c z -
carbon dioxide
Adapted from Abbott Laboratories Patent U.S. Patent # 5,047,627
Fig. 5. Sensing region of optical electrolyte sensor based on thin polymer films doped with cation ionophores and chrome-pH ionophores.
that enables rapid equilibration of protons with the dye. A novel multi-step photopolymerization methodology has been introduced recently [ 141 as a means of sequentially immobilizing indicator dyes and outer protective polymer coatings at the distal end of microfiber bundles to yield spatially resolved sensing sites that can be monitored simultaneously via a modern CCD imaging camera With further development, this in situ manufacturing process could ultimately reduce the complexity of preparing multiparameter in viva optical sensors. Reversible optical sensors for electrolyte ions have been described only recently [ 111. While not yet miniaturized for in viva measurements, the development of appropriate catheter configurations seems likely in the future. Fig. 6. illustrates the design and detection principles of one type of electrolyte sensor. Ionophores typically used for preparing potentiometric membrane electrodes (see above) are doped into very thin polymer films deposited at the distal end of optical fibers. In addition to the K+, Ca*+, or Na+ ionophore, a lipophilic pH indicator is also added to thin polymer film. When cations from the blood sample are selectively extracted into the bulk of the film by the ionophore, the pH indicator loses a proton to the sample to maintain charge neutrality in the
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mm
optical fiber light
mm
Polymer Film
I
I
I
I
2
4
6
8
Time (h) 0 R
= lonophore (e.g. VaIinomycin, etc.) = Chromo-pH ionophore:
X
R - and RH have different absorption spectra.
Fig. 6. Expanded view of three parameters optical blood-gas catheter described in patent granted to Abbott Laboratories (see ref. 13).
film. This results in a change in the optical absorption or fluorescence of the thin polymer layer. If the thickness of the films are kept < 10 pm, response times on the order of < 1 min can be achieved. The main limitation of this optical ionsensing configuration is that the pH of the sample also affects the overall extraction of ions into the film. Thus simultaneous measurement of the blood pH with a second pH “only” sensor would be required for accurate in viva measurements of electrolytes using such systems.
Analytical measurements environment
in the in vivo
Many of the sensors described above have been shown to function quite effectively as analytical devices in long term (72 h) benchtop laboratory studies (in buffered solutions or flowing blood). However, moving from the laboratory to the implantation in animal and human arteries has been far more difficult than initially envisioned. In the case of blood-gas sensors, the pattern of response illustrated in Fig. 7 is often observed [2]. With time, pH and PO2 values monitored by the implanted sensors become lower than corresponding in vitro measurements made on discrete samples with conventional blood-gas instruments. On the
q
in vitro blood gas measurement
Fig. 7. Typical accuracy vs. time pattern observed for many in viva blood-gas sensors after implantation onto the arteries of humans or animals (solid lines represent continuous output tracings from the implanted sensors).
other hand, PCO, values gradually increase with time. When the sensors are removed from the arteries and retested in vitro, very accurate results are often obtained. It should be noted that such behavior is less likely to occur when the test subjects are systematically heparinized to prevent thrombus formation on the surface of the sensor. Clinical application of such sensors, however, would require that they function effectively without systematic anticoagulation owing to the bleeding risks associated with such treatment [note: the saline-heparin flush illustrated in Fig. 1 does not equate to systematic anticoagulation; very low levels of heparin are present in such flush solutions, and this drip enters the blood downstream from the actual sensor tip (see Fig. 1b)]. In addition to the issue of in vivo biocompatibility, the patterns observed in Fig. 7 have also been correlated with the precise position of the sensor within the artery
PI. It is important to understand what can cause the response patterns simulated in Fig. 7. While initially many workers felt that the implanted sensors were failing to provide accurate values due to gradual protein fouling of the outer membranes, this turns out to be only a minor issue. Although protein adsorption to the polymer surfaces occurs, as shown in Fig. 8, this is only the first step in the much more complex process of thrombus forma-
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sensor
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Adsorption of Proteins * (Fibrin0 en, Von Wille%rand Factor, etc.)
Attachment of Platelets \
Thrombin 1 Clotting Factors
Thrombus (Fibrin with Entrapped Metabolic Cells)
Activated Platelets
Fig. 8. Sequence of events that can take place in vivo leading to formation of a metabolic thrombus on the surface of an implanted sensor (in non-systematically heparinized subjects). tion. As the clotting process proceeds, highly metabolic platelet cells adhere and are subsequently entrapped in the thrombus that can coat the surface of the sensor, along with other respiring blood cells (e.g., white cells, etc.). These living cells consume oxygen and liberate carbon dioxide. As a result, cell adhesion effectively forms an in situ biosensor; that is a sensor which has an immobilized layer of biologically active material in intimate contact with the transducer. The metabolic activity of the cells will produce local surface pH, PC02, and PO2 levels that are different than the bulk blood, in accordance with the pattern shown in Fig. 7. It has been suggested that the observed dependence of sensor accuracy on sensor placement/position within the artery can be traced to a similar in situ biological activity. That is, if the sensing site of an implanted probe ends up butted up against the inner wall of the artery, metabolic activity of the endothelia cells could cause local pH, PCOT, and PO2 values to differ from those in the bulk blood, particularly if placement of the sensor affects blood flow patterns at or near the inner wall, or total blood flow velocity within the artery. This is the so-called wall effect [2]. To reduce changes of in vivo thrombus formation (i.e., enhanced blood biocompatibility), spe-
vol. 72, no. 6, 1993
cially formulated polymers (e.g., hydrogels, etc.) are used to coat in vivo chemical sensors. Other approaches to reduce thrombus formation include covalent immobilization of heparin, or other anticoagulant species (e.g., prostaglandins, albumin, enzymes, etc.) to the outer surface of the sensors. Many of these common methods used to render implantable polymer devices less thrombogenic may, however, not always be compatible with the sensing chemistries employed within the probes. While success in eliminating or reducing clot formation has been suggested in a number of recent reports, the propensity to form thrombus on these implanted devices can never be eliminated completely. Indeed, the sensitivity of the entire human coagulation process to foreign implants will certainly vary from patient to patient, and thus even the most effective surface treatments may not work in every clinical situation. Efforts to overcome the wall effect problems have focused on engineering some type of mechanical positioning device to keep the implanted sensors in the center of the artery. In addition, wall effects may also be reduced by minimizing the distance the sensing probe extends beyond the end of the cannula. However, if this distance is not adequate, the sensors will be partially exposed to the flush solution, and this will also cause erroneous analytical results. Current status Despite the challenges of achieving size, stability, accuracy, biocompatibility, etc., a number of major biomedical companies are pursuing the development of implantable blood-gas and electrolyte sensors (see Table 3). One company, PuritanBennet, has developed and is currently marketing a three parameter optical blood-gas probe (PB 3300) which reportedly has yielded clinically acceptable analytical results for up to 72 h in human subjects [15]. Pfizer Inc., is preparing to introduce a hybrid-type three parameter blood-gas probe in which pH and PC02 are detected via optical fiber technology, and PO, is measured via an electrochemical catheter similar to that shown in Fig. 2a above. Other companies, however, have decided recently that reliable analytical measurements within the human radial artery are simply too difficult to achieve, opting to pursue a slight variation on the original in vivo sensor concept (see Table 2).
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TABLE 3. Companies
developing
in vivo or on-line blood-gas-electrolyte
Puritan-Bennett (Carlsbad, CA, USA) Optex (Houston, TX, USA) CR Bard (Murray Hill, NJ, USA) Optical Sensors for Medicine (Eden Prairie, MN, USA) Abbott Laboratories (Bothell, WA, USA) Pfizer (New York, NY, USA) Otto Sensors (Cleveland, OH, USA) CDI-3M (Minneapolis, MN, USA) VIA Medical (San Diego, CA, USA)
Rather than threading very small sensors down through the A-line catheter, sensor arrays are being incorporated in the saline-flush solution line external to the patient (on-line systems). Periodically, samples of blood can be drawn up into the sensor chamber for measurements by reversing the pump direction used to infuse the patient with the flush solution. Several sensors can be placed in the chamber allowing the measurement of more parameters (blood-gases, electrolytes, and even glucose) [ 161. This approach partially overcomes the problems inherent in preparing very small sensors and at the same time reduces the signal stability requirements. Indeed, the flush solution can be used to continuously calibrate the sensors so that reliable measurements may be made even with sensors that have a significant drift. The downside of this approach is that the measurements are not truly continuous or real-time, and thus during calibration and sampling cycles rapid changes in the blood chemistry parameters may go undetected until the system is in the measurement mode. At this point, while implantable sensors for blood-gases and electrolytes would provide clinicians with invaluable real-time information that could enhance the quality of health care for critically ill patients, there are still lingering questions about the reliability of results obtained by these devices once they are placed within the confines of the radial artery. Although considerable research is in progress to solve the remaining issues that affect the accuracy of the various electrochemical and optical probes described above, it appears that the first generation of commercial multiparameter implantable sensors will likely serve more as trend monitors and will not obviate the need for routine blood-gas and electrolyte measurements on discrete blood samples. At the same time, the evolution of periodic on-line rather than continuous in vivo sensor technology could provide an attractive, temporary alternative until
systems
Optical intra-arterial blood gas (PB3300) Optical intra-arterial blood gas (BioSentry) Optical intra-arterial blood gas Optical intra-arterial blood gas Optical intra-arterial blood gas Hybrid electrochemical/optical intra-arterial blood gas Electrochemical intra-arterial blood gas On-line optical blood gas (CDI-System 2000) On-line electrochemical blood gas electrolytes, ‘glucose
in vivo sensors can truly meet all of the requirements outlined in Table 2. Acknowledgement The author is grateful to the National Institutes of Health (grant GM-28882) for supporting his research program on the development of new in vivo electrochemical gas and ion sensors. References M.E. Meyerhoff, Clin. Chem., 36 (1990) 1567. C.K. Mahutte, C.S.H. Sassoon, J.R. Muro, D.R. Hansman, T.P. Maxwell, W.W. Miller and M.Yafuso, J. Clin. Monit., 6 (1990) 147. G.S. Calabrese and L.M. O.‘Connell, Topics Cur-r Chem., 143 (1988) 51. P Anker, D. Ammann and W. Simon, Mikrochim. Acta., I (1983)1448. U. Oesch, D. Ammann and W. Simon, Clin. Chem., 32 (1986) 1448.
B.A. McKinely, B.A. Houtchens and J. Janata, ZonSelective Electrode Rev., 6 (1984) 1’73.
R.L. Coon, N.C.J. Lai and J.F? Karnpine, J. Appl. Phys., 40 ( 1976) 625.
M.E. Collison, G.V. Aebli, J. Petty and M.E. Meyerhoff, Anal. Chem., 61 (1989) 2365. B.A. Shapiro, Resp. Care, 37 (1992) 165. 10 W.W. Miller, M. Yafuso, C.F. Yan, H.K. Hui and S. Arick, Clin. Chem., 33 (1987) 1538. 11 O.S. Wolfbeis (Editor), Fiber Optic Chemical Sensors and Biosensors, Vols 1 and II; CRC Press, Boca Raton, FL, 199 1. 12 J. Janata, Anal. Chem., 59 (1987) 1351. 13 J.B. Yim, T.W. Hubbard, L.D. Melkerson, M.A. Sexton and B.M. Fieffen, U.S. Pcuent 5,047,627, September 10, 1991. 14 S.M. Barnard and D.R. Walt, Nature, 353 (1991) 338. 15 C.P Larson, Presentation at 14th International Symposium of the Electrolyte-Blood Gas Division of the American Association for Clinical Chemistry,
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Chatham, MA, May 1992.
16 D.K. Wong and W.S. Jordon, ht. J. Clin. Monit.
trends in analytical chemistry, vol. 12, no. 6, 1993
Mark E. Meyerhoff is at the Department of Chemisty, University of Michigan, Ann Arbor, MI, USA.
Camp., 9 (1992) 95.
Sensitive techniques for phospholipase determination in plants C. Thkvenot and J. Daussant Meudon, France
J. K&* and 0. Valentovii Prague, Czech Republic The breakdown of plant phospholipids appears to be mainly mediated by phospholipase D. Its role is not fully understood either in plant development or post-harvest metabolism. Thus, sensitive methods for the assessment of low activity levels are required for use in such studies. This article provides a comprehensive view of the available methods and presents the authors’ experiences with radiometric and biosensor assays.
Introduction The phospholipases catalyse the hydrolysis of diacylglycerophospholipids, which may be considered as derivatives of glycerophosphate two of whose hydroxyl groups are esterified by longchain fatty acids and whose phosphoryl group forms a phosphodiester bond with a polar moiety (Fig. 1). Phospholipids contain one of the following polar moieties: choline, ethanolamine, serine, glycerol or inositol. All four ester bonds in a phospholipid are susceptible to enzymatic cleavage. The phospholipases have specificity for the position on the glycerophosphate backbone rather than for the particular fatty acid or polar group. Phospholipase AZ, PLA;! (EC 3.1.1.4) which hydrolyses the acid moiety from the sn-2 position of a diacylglycerophospholipid is the best charac-
0165-9936/93/$06.00
D
terized phospholipase. Much work has been done on extracellular PLA, which is abundant in bee and snake venoms and in mammalian pancreas. It is among the enzymes with the smallest relative molecular mass and its sequence analysis and cloning have been reported. In contrast, PLAZ from other cell and tissue sources, are trace enzymes. It is more difficult to find data concerning phospholipase A,, PLA, (EC 3.1.1.3.2) which hydrolyses the acid moiety from the sn-1 position of a diacylglycerophospholipid. Enzymes with PLA, activity have been purified mainly from microorganisms and mammalian tissues. The enzymatic activity ascribed to simultaneous hydrolysis of acid moieties in sn-1 and sn-2 position of a diacylglycerophospholipid falls into the category of phospholipase B, PLB (EC 3.1.1 S). This concept, however, faced the criticism that the dual activity could reflect multiple enzyme activi-
(
0 \
cH2
II -o-c-
R1 pm1 &l
R2
-
; C -
/
0
-CH
/ 0 II CH2
-0-P-O I 0. /
z
Fig. 1. Sites of attack of phospholipases on a diacylglycerophospholipid. Rl and R2 represent longchain fatty acids; X represents a polar moiety; sn-1 , sn-2, m-3 refer to carbon positions on the glycerophosphate backbone.
0 1993 Elsevier Science Publishers B.V. All rights reserved
257
ln viva blood-gas and electrolyte sensors: progress and challenges Mark. E. Meyerhoff Ann Arbor, Ml, USA The management of critically ill patients requires the frequent and accurate measurement of blood gases and electrolytes in undiluted whole blood. To improve patient care, there is a growing desire to monitor these important parameters on a continuous basis via the use of intra-arterial chemical sensors. The progress and challenges relating to the development of suitable probes for making reliable analytical measurements within the confines of the human radial ‘artery are summarized.
Background The care of critically ill hospital patients [intensive care unit (ICU), coronary care unit (CCU), etc.] mandates the frequent measurement of a select group of critical care analytes in undiluted whole blood. Tests recognized as being essential for patient management include the blood gases (pH, POZ, PCO,) and electrolytes (Na+, K+, iCa2+). Normal range values for these analytes are summarized in Table 1. While historically these tests have been performed in centralized or satellite laboratories remote from the patient, long delays to obtain test results (so-called “turn-around time”) have created a demand for technologies that enable these measurements to be carried out more rapidly at or near the patient’s bedside. Although progress has been made recently via the developTABLE 1. Normal range values for blood-gas and electrolyte measurements Oxygen partial pressure (IQ) Carbon dioxide partial pressure
80-l 04 mmHg
(X02)
33-48
PH sodium (Na+) potassium (K+) ionized calcium (iCa”)
0165.9936/93/$06.00
mmHg 7.31-7.45 135-l 55 mmol/l 3.6-5.5 mmol/l 1 .14-l .31 mmol/l
ment of convenient to use, portable, blood-gaselectrolyte instruments [I], these instruments are only useful for discrete sample measurements, and are not capable of monitoring blood-gas and electrolyte values on a continuous basis. The availability of truly implantable blood-gas and electrolyte sensors (either electrochemical or optical) could provide clinicians with essentially “real time” diagnostic information and this could result in more timely therapeutic intervention and thus save lives. Twenty gauge arterial line catheters are placed in more than half of all ICU patients, typically for up to three days. These A-lines are implanted within the radial artery and are often equipped with pressure transducers to monitor blood pressure continuously (see Fig. la and b). At the same time, the A-line provides an open port to readily sample (via a syringe) the patient’s blood for periodic discrete sample blood tests. A steady stream of saline-heparin flows through the cannula to keep the implant site free of blood clots. The ultimate goal of researchers developing in viva sensors is to design blood gas and electrolyte probes that are small enough to slide down through the cannula of the A-line (see Fig. lb) so as to be in continuous contact with the flowing arterial blood (with mean linear flow velocities of 3.6 cm/s and mean volumetric flow-rates of 0.65 ml/s), enabling real-time measurements of blood-gas and electrolyte levels. To be useful clinically, in vivo blood-gas and electrolyte sensors must fulfill ull of the requirements outlined in Table 2. Meeting these requirements is an enormous challenge. For example, the I.D. of a 20 gauge A-line cannula is only 0.83 mm. Thus, in order to obtain accurate blood pressure signals, in vivo sensors must be considerably smaller, i.e., they must occupy < X)-60% of the total cross-sectional area of the cannuIa. While large size sensors (e.g., O.D. 1 mm) can be employed for preliminary development work, and as useful research tools for animal studies, ultimate application in humans requires the miniaturization of sensors to a size where multiple transducers, for example pH, PC02, P02, can all be combined into a single probe with a total O.D. of < 0.55 mm. In
0
1993 Elsevier Science Publishers B.V. All rights reserved
258
trends in
addition to the sensors, the same probe must also contain a thermocouple so that blood temperature can be monitored simultaneously. These readings are then used to compensate for any known temperature sensitivities of the sensors, and to provide temperature-corrected blood gas values (i.e., the blood pH and partial pressures of the gases if
a)
KJ-
salinelheparin flush solution
blood-gas/electrolyte monitor
A-line
catheter
b)
heparlnlsaline drip ‘;
TABLE 2. Requirements . . . . . .
analytical chemistry, vol.
for in vivo ion-gas
12, no. 6, 7993
sensors
Small size (compatible with 20 gauge A-line) Long term calibration stability (one calibration insertion) Reliable values for 72 h under widely physiological conditions Blood compatibility (non-thrombogenic) Rapid response times Ability to sterilize
prior to varying
measured at 37°C in a conventional blood-gas instrument). The entire final probe must also be able to withstand appropriate sterilization (e.g., via ethylene oxide, gamma radiation, etc.) so that microbial contamination of the patient is not a concern. Beyond severe size restraints and the need to sterilize, signal stability is another major hurdle. Since it is most desirable that calibration of in viva sensors be performed only once prior to insertion, long term signal stability is very important, particularly when analytical accuracies of < + 5% are desired for the entire implant time (up to 72 h). While alternate in viva recalibration schemes involving periodic in vitro measurements on discrete samples of blood are possible (i.e., during the 72 h period), if the required frequency of such in vivo recalibrations is high (perhaps more than once a day), the advantages of the continuous in vivo measurement are partly negated. Analytical performance of implanted sensors, however, can also be influenced dramatically by the complexity of the in vivo environment, especially the propensity of blood clots to form on the surface of foreign materials (i.e., biocompatibility). In addition, the precise position of the sensor within the artery can influence output signals (so call “wall effects” [2]). These two issues relate directly to the innate difficulties of making reliable intra-arterial measurements even with the smallest and most stable analytical probes, and will be addressed in more detail in a separate section below.
Available sensor technology Fig. 1. (a) Schematic of arterial line catheter placement and associated components for continuous monitoring of blood pressure as well as blood-gases and electrolytes; (b) expanded view of A-line catheter with in vivo chemical sensor(s) inserted through cannula.
In vivo blood-gas and electrolyte sensors can be based on either electrochemical or optical measurements. Progress in the development of sensors based on these transduction modes are summarized below.
trends in analytical chemistry, vol. 7.2, no. 6, 1993
Electrochemical
sensors
The electrochemical probes are essentially miniaturized versions of the sensors already in use within large, commercial blood-gas-electrolyte instruments [3]. For example, PO;! levels can be determined with a catheter form of the classical Clark-style amperometric oxygen sensor (Fig. 2a). Various forms of such single parameters probes have been commercially available for some time with outer diameters on the order of 0.55 mm. Electrochemical reduction of oxygen (e.g., O2 + 2H+ + 4 e- + 2 OH-) occurs at a silver or platinum cathode, which is polarized at ca. -0.7 V vs. an Ag/AgCl reference/anode electrode. The resulting current flow (generally in the O-100 nA range) is directly proportional to blood PO*. Signal drift is a major concern with these devices, since the diffusion rate of oxygen through the outer polymer material can change with time. In addition, the surface area of the cathode can gradually increase as silver ions formed at the anode reference electrode migrate to the cathode and reduce, effectively creating a larger area cathode and thus a larger diffusion current for a given level of oxygen. Electrochemical pH measurements are generally made via miniaturized potentiometric ion-selective membrane probes, or metal/metal oxide based devices. The latter are subject to interferences from redox active compounds (e.g., ascorbic acid, uric acid, etc.), and are thus less attractive for direct blood pH measurements. The fragile nature and high electrical resistance of conventional glass
,_ _..._. Silver Cathode
(a)
Insulator Plug 04
Fig. 2. Sensing regions of amperometric oxygen (a) and tubular-style potentiometric pH (b) (or electrolyte) in viva sensors.
membrane pH electrodes preclude their use as practical transducers for intra-arterial measurements. Modem polymer-membrane-type pH electrodes, prepared by doping PVC or other polymeric materials with hydrophobic amine compounds [4] do offer promise as useful in vivo probes. Fig. 2b illustrates one configuration for such devices. When the polymer tubing wall is doped with the appropriate amine (e.g., tridodecylamine), and a suitable buffer (0.25 M phosphate pH 6.8) is employed to fix the activity of protons within the internal electrolyte solution in contact with the internal Ag/AgCl reference electrode, the potential of the catheter probe vs. a #suitable external reference electrode can be re1ate.d to the pH of the sample. For in vivo potentiometric measurements, the external reference electrode (e.g., Ag/AgCl, saturated calomel, etc.) can actually be placed outside the artery, in one arm of the A-line (see Fig. 1). The continuous saline-heparin flush solution can serve as a useful electrolytic contact (i.e., salt bridge) to the blood, as it flows between the inner wall of the cannula and the outer wall of the in vivo probe. In vivo measurements of sodium, potassium, and ionized calcium (free calcium) can be made with a similar working (Fig. 2b) (and reference electrode arrangement; however, instead of doping the tubing walls of the device in Fig. 2b with lipophilic amines, highly selective ionophores for sodium, potassium or calcium are used as membrane-active species [5]. For example, in vivo potassium sensors are prepared with valinomycin (K+ selective antibiotic) as the active membrane component while ionized calcium levels can be monitored with alkylphosphates or other specially designed synthetic neutral organic ligands doped within the polymer films. These membrane active ionophores are very hydrophobic compounds and thus leach very slowly from the polymeric tubing into the surrounding blood (non-harmful levels). Although there also has been considerable effort aimed at adapting such polymer membrane ion-selective electrode technology to modern microelectric fabrication techniques to prepare multi-parameter implantable “solid-state” ion sensors (e.g., ion-selective field effect transistors (ISFETs) [6]), thus far these devices have failed to achieve the level of stability and selectivity required for accurate in vivo measurements. Electrochemical in vivo measurement of PC02 levels can be accomplished using miniaturized
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trends in analytical chemistry, vol. 72, no. 6, 1993
external reference electrode (in flush line)
e silicone rubber tubing \
potassium bicarbonate/ potassium chloride solution
PVC tubing
I
buffer
pH responsive tubing
/
Valinomycln-impregnated silicone tubing
Fig. 3. Dual potentiometric K+/PC02 sensing catheter described by Collison et al. (see ref. 8).
variations of the classical Severinghaus-style PCOZ sensor design in which the pH of thin layer of bicarbonate solution housed behind a gas-permeable membrane changes as a function of the PC02 level in the sample. In an early catheter configuration introduced by workers at General Electric Corp., a PdO-type pH electrode was used to detect the pH changes in the bicarbonate layer [7], More recently, polymeric pH electrodes (such as those in Fig. 2b) have been utilized successfully to fabricate catheter-type PC02 sensors for in vivo measurement (in dogs). In the concentric tube design shown in Fig. 3, the inner-tubular pH electrode detects pH changes in a thin layer of bicarbonate held within the outer silicone rubber tube [8]. When doped with an appropriate ionophore (e.g., valinomycin), the outer silicone rubber tube can serve both as an outher gas-permeable membrane for the PC02 measurement, and as an ionselective membrane for the simultaneous measurement of an electrolyte ion, in this case potassium. The potassium measurement is made by monitoring the membrane potential across the wall of the outer ionophore-doped silicone tube. (i.e., Ag/AgCl wire 1) vs. external reference electrode in contact with saline-heparin flush. It is important to recognize that all potentiometric devices respond to analyte concentrations in a logarithmic fashion in accordance with the Nernst equation: J%~II= K + (O.O59/zi) log [analyte] sample where f&l is the measured potential between the working ISE catheter and an external reference electrode, Kis the cell constant (sum of all constant potentials, including reference electrode poten-
tials), zi is the charge on the detected ion, and [analyte] is the concentration of the measured species (H+, K+, Na+, iCa2+ , PCO2). Thus, for pH, K+, Na+, PC02 measurements, changes of only 1 mV in the cell constant between the calibration step and the actual in vivo measurement will cause variations of + 4% with respect to the measured parameter, making reliable in vivo measurements with such devices quite difficult, unless the exquisite e.m.f. stability can be achieved. In general, instability of the measurement instrumentation is not the limiting factor in achieving accurate quantitative results over long time periods; rather it is changes in the membrane-sensing chemistries and in the compositions of the internal reference solutions (via osmotic effects) that can often yield the observed drift in potentiometric signals. Optical sensors
Although efforts to develop various electrochemical probes for intravascular measurements have been ongoing for more than 30 years, much of the most recent work in this area has focused on the use of modern fiber-optic chemical sensors [2,9,10]. These devices are prepared by immobilizing a wide variety of suitable indicator reagent dyes at the end of a narrow fiber and measuring either absorption, fluorescence, or phosphorescence resulting from the interaction of light with the immobilized indicators. These fiber-optic probes have certain inherent advantages over electrochemical sensors in that they: can be miniaturized more easily (using thin optical fibers of 100-150 pm diameter); have less noise (no transduction wires required); require no reference electrode; and have the potential to offer greater signal stability by employing multiple wavelength measurements to correct for changes in effective optical pathlengths, photobleaching of dyes, etc. The latter advantage is realized by referencing measurements of optical signals that originate at wavelengths that contain analytical information to wavelengths at which the spectroscopic signal does not change as a function of the analyte level in blood (e.g., isosbestic points in the indicator dye’s absorption or fluorescence spectrum) [ 111. Obviously, elaborate source-detection monitors which can be conveniently coupled to the various miniature optical fiber assemblies are also required for routine analytical measurements with these
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261
vol. 7.2, no. 6, 1993
Light a
Light
Light out
Light in
b
Clad
Dye
Cladding
Fig. 4. Three configurations used to prepare in vivofiber-opticchemical design; and (c) transmission arrangement used by Optex Inc.
optical sensing devices. One inherent disadvantage in optical sensing schemes, when compared to electrochemical sensors, is their limited dynamic measurement range (e.g., optical pH sensor; often less than 2 orders of magnitude). While for general analytical measurements this would certainly be problematic, as shown in Table 1, the normal ranges for gases and electrolytes in blood are rather narrow. Thus, provided that the maximum optical change for a given sensor occurs over a concentration range that encompasses the normal physiological values (e.g., by selecting indicator dye with proper pK, for blood pH determinations), the measurement range limitation is not a major issue with respect to using optical sensors for in vivo monitoring. Three basic optical sensor configurations have been examined for use in prototype blood-gas catheters. These are illustrated in Fig. 4. In the so-called Lubbers design (Fig. 4a) [ 111, the indicator is immobilized on the distal end of a quartz or polymer fiber. Photons from an appropriate source (usually pulsed) are sent down the fiber and light coming back up through the same fiber can be monitored at the same (absorption) or longer wavelengths (fluorescence, phosphorescence). Microporous reflective materials (e.g., gold films, etc.) can be coated over the dye layer to enhance the intensity of the returning optical signal. In the Buckles configuration (Fig. 4b) [ll], the dye is immobilized on the outside walls of the optical fiber (after stripping away the fiber’s outer cladding material). Optical signals originating from the dye molecules confined to the outer walls of the
sensors; (a) Lubbers design; (b) Buckles
fiber occur as a result of the evanescent wave phenomena. Optex Inc. (Houston, TX, USA), has proposed the bent fiber approach (Fig. 4c) in which a small region of the fiber is cut away and replaced with appropriate indicator dye matrix. Unlike the other two configurations, the Optex approach functions in a manner equivalent to conventional spectrophotometry in that the transmitted light that passes through a fixed optical pathlength dye region/chamber is the signal detected. The dye chemistry used to measure the bloodgases varies considerably. Most optical measurements of PO;, are based on fluorescence quenching of the indicator [e.g., perylene dibutyric acid, Ru(I1) tri(dipyridine), etc.], in the presence of oxygen in accordance with the Stern-Volmer theory: (I&&
= 1 + kP0 2
where Zpo2 is the fluorescence intensity at a given PO2 level, I0 is the intensity at PO2 =I0, and k is a quenching constant for the particular dye. Thus ZJZpo2 is linearly related to PO2. Phosphorescence, or phosphorescence-lifetime measurements of immobilized metal ligand complexes can also be employed (i.e., the binding of oxygen decreases the excited state lifetimes). Sensors based on measurement of luminescent lifetimes have the inherent advantage of being insensitive to changes in optical pathlength and the amount of active dye present at the tip of the sensor without need to perform measuremore complex multiple-wavelength ments. Fluorescence or absorption pH measurements
262
can be made by immobilizing any number of classical indicator dyes that have pK, values at or near the normal pH of blood (pH 7.4) (e.g., 8-hydroxy1,2,6-pyrene trisulfonate among many others). One lingering issue relating to the use of indicators for accurate pH measurements is the effect of ionic strength on the pK, of the indicators [12]. Since optical sensors measure the concentrations of the protonated or unprotonated forms of the immobilized dyes as an indirect measure of the hydrogen ion activity, large variations in blood ionic strength via changes in blood sodium levels can, in principle, affect the accuracy of such measurements by varying the pK, of the immobilized species. Since reference values for blood pH are determined by conventional glass electrodes that detect proton activity directly, variations in the ionic strength will cause larger discrepancies between reference instrument values and the values determined by the optical pH sensor. It is important to note that these variations in ionic strength will also cause the pH of the blood itself to change as measured electrochemically (effect on pK, of bicarbonate, etc.); however, because the electrode measures this defined value directly, it will always be the reference method to which the pH value reported by the optical sensor will be compared. Optical sensors for carbon dioxide employ the fiber-optic pH sensors as internal transducers in an arrangement quite similar to electrochemical PC02 gas sensors described above. An outer hydrophobic gas-permeable film separates the sample from an internal layer of sample PCOZ and this is detected optically. Typically, the same dyes used for direct blood pH measurements can be employed to construct the pH sensor used for monitoring blood PCO, levels, since the variation in pH within the bicarbonate film falls within the same pH range as blood values. Fig. 5. illustrates the manufacturing complexity of putting three optical sensors together to form a single multiparameter pH, PCO;! and PO2 intravascular probe. In this patented design [ 131 (Abbott Laboratories), the PO2 and PC02 sensors are coated with a silicone rubber material which serves as the gas-permeable (and ion-impermeable) membranes for the two gas-sensing fibers. Since pH measurements require that sample hydrogen ions have access to the immobilized indicator dye, the pH sensor is positioned slightly forward of the gas sensors and is coated with a much different, more hydrophilic polymer layer
trends in analytical chemistw, vol. 12, no. 6, 7993
light
c
c z -
carbon dioxide
Adapted from Abbott Laboratories Patent U.S. Patent # 5,047,627
Fig. 5. Sensing region of optical electrolyte sensor based on thin polymer films doped with cation ionophores and chrome-pH ionophores.
that enables rapid equilibration of protons with the dye. A novel multi-step photopolymerization methodology has been introduced recently [ 141 as a means of sequentially immobilizing indicator dyes and outer protective polymer coatings at the distal end of microfiber bundles to yield spatially resolved sensing sites that can be monitored simultaneously via a modern CCD imaging camera With further development, this in situ manufacturing process could ultimately reduce the complexity of preparing multiparameter in viva optical sensors. Reversible optical sensors for electrolyte ions have been described only recently [ 111. While not yet miniaturized for in viva measurements, the development of appropriate catheter configurations seems likely in the future. Fig. 6. illustrates the design and detection principles of one type of electrolyte sensor. Ionophores typically used for preparing potentiometric membrane electrodes (see above) are doped into very thin polymer films deposited at the distal end of optical fibers. In addition to the K+, Ca*+, or Na+ ionophore, a lipophilic pH indicator is also added to thin polymer film. When cations from the blood sample are selectively extracted into the bulk of the film by the ionophore, the pH indicator loses a proton to the sample to maintain charge neutrality in the
263
trends in analytical chemistry, vol. 72, no. 6, 1993
mm
optical fiber light
mm
Polymer Film
I
I
I
I
2
4
6
8
Time (h) 0 R
= lonophore (e.g. VaIinomycin, etc.) = Chromo-pH ionophore:
X
R - and RH have different absorption spectra.
Fig. 6. Expanded view of three parameters optical blood-gas catheter described in patent granted to Abbott Laboratories (see ref. 13).
film. This results in a change in the optical absorption or fluorescence of the thin polymer layer. If the thickness of the films are kept < 10 pm, response times on the order of < 1 min can be achieved. The main limitation of this optical ionsensing configuration is that the pH of the sample also affects the overall extraction of ions into the film. Thus simultaneous measurement of the blood pH with a second pH “only” sensor would be required for accurate in viva measurements of electrolytes using such systems.
Analytical measurements environment
in the in vivo
Many of the sensors described above have been shown to function quite effectively as analytical devices in long term (72 h) benchtop laboratory studies (in buffered solutions or flowing blood). However, moving from the laboratory to the implantation in animal and human arteries has been far more difficult than initially envisioned. In the case of blood-gas sensors, the pattern of response illustrated in Fig. 7 is often observed [2]. With time, pH and PO2 values monitored by the implanted sensors become lower than corresponding in vitro measurements made on discrete samples with conventional blood-gas instruments. On the
q
in vitro blood gas measurement
Fig. 7. Typical accuracy vs. time pattern observed for many in viva blood-gas sensors after implantation onto the arteries of humans or animals (solid lines represent continuous output tracings from the implanted sensors).
other hand, PCO, values gradually increase with time. When the sensors are removed from the arteries and retested in vitro, very accurate results are often obtained. It should be noted that such behavior is less likely to occur when the test subjects are systematically heparinized to prevent thrombus formation on the surface of the sensor. Clinical application of such sensors, however, would require that they function effectively without systematic anticoagulation owing to the bleeding risks associated with such treatment [note: the saline-heparin flush illustrated in Fig. 1 does not equate to systematic anticoagulation; very low levels of heparin are present in such flush solutions, and this drip enters the blood downstream from the actual sensor tip (see Fig. 1b)]. In addition to the issue of in vivo biocompatibility, the patterns observed in Fig. 7 have also been correlated with the precise position of the sensor within the artery
PI. It is important to understand what can cause the response patterns simulated in Fig. 7. While initially many workers felt that the implanted sensors were failing to provide accurate values due to gradual protein fouling of the outer membranes, this turns out to be only a minor issue. Although protein adsorption to the polymer surfaces occurs, as shown in Fig. 8, this is only the first step in the much more complex process of thrombus forma-
264
sensor
trends in analytical chemistry
Adsorption of Proteins * (Fibrin0 en, Von Wille%rand Factor, etc.)
Attachment of Platelets \
Thrombin 1 Clotting Factors
Thrombus (Fibrin with Entrapped Metabolic Cells)
Activated Platelets
Fig. 8. Sequence of events that can take place in vivo leading to formation of a metabolic thrombus on the surface of an implanted sensor (in non-systematically heparinized subjects). tion. As the clotting process proceeds, highly metabolic platelet cells adhere and are subsequently entrapped in the thrombus that can coat the surface of the sensor, along with other respiring blood cells (e.g., white cells, etc.). These living cells consume oxygen and liberate carbon dioxide. As a result, cell adhesion effectively forms an in situ biosensor; that is a sensor which has an immobilized layer of biologically active material in intimate contact with the transducer. The metabolic activity of the cells will produce local surface pH, PC02, and PO2 levels that are different than the bulk blood, in accordance with the pattern shown in Fig. 7. It has been suggested that the observed dependence of sensor accuracy on sensor placement/position within the artery can be traced to a similar in situ biological activity. That is, if the sensing site of an implanted probe ends up butted up against the inner wall of the artery, metabolic activity of the endothelia cells could cause local pH, PCOT, and PO2 values to differ from those in the bulk blood, particularly if placement of the sensor affects blood flow patterns at or near the inner wall, or total blood flow velocity within the artery. This is the so-called wall effect [2]. To reduce changes of in vivo thrombus formation (i.e., enhanced blood biocompatibility), spe-
vol. 72, no. 6, 1993
cially formulated polymers (e.g., hydrogels, etc.) are used to coat in vivo chemical sensors. Other approaches to reduce thrombus formation include covalent immobilization of heparin, or other anticoagulant species (e.g., prostaglandins, albumin, enzymes, etc.) to the outer surface of the sensors. Many of these common methods used to render implantable polymer devices less thrombogenic may, however, not always be compatible with the sensing chemistries employed within the probes. While success in eliminating or reducing clot formation has been suggested in a number of recent reports, the propensity to form thrombus on these implanted devices can never be eliminated completely. Indeed, the sensitivity of the entire human coagulation process to foreign implants will certainly vary from patient to patient, and thus even the most effective surface treatments may not work in every clinical situation. Efforts to overcome the wall effect problems have focused on engineering some type of mechanical positioning device to keep the implanted sensors in the center of the artery. In addition, wall effects may also be reduced by minimizing the distance the sensing probe extends beyond the end of the cannula. However, if this distance is not adequate, the sensors will be partially exposed to the flush solution, and this will also cause erroneous analytical results. Current status Despite the challenges of achieving size, stability, accuracy, biocompatibility, etc., a number of major biomedical companies are pursuing the development of implantable blood-gas and electrolyte sensors (see Table 3). One company, PuritanBennet, has developed and is currently marketing a three parameter optical blood-gas probe (PB 3300) which reportedly has yielded clinically acceptable analytical results for up to 72 h in human subjects [15]. Pfizer Inc., is preparing to introduce a hybrid-type three parameter blood-gas probe in which pH and PC02 are detected via optical fiber technology, and PO, is measured via an electrochemical catheter similar to that shown in Fig. 2a above. Other companies, however, have decided recently that reliable analytical measurements within the human radial artery are simply too difficult to achieve, opting to pursue a slight variation on the original in vivo sensor concept (see Table 2).
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trends in analytical chemistry, vol. 12, no. 6, 1993
TABLE 3. Companies
developing
in vivo or on-line blood-gas-electrolyte
Puritan-Bennett (Carlsbad, CA, USA) Optex (Houston, TX, USA) CR Bard (Murray Hill, NJ, USA) Optical Sensors for Medicine (Eden Prairie, MN, USA) Abbott Laboratories (Bothell, WA, USA) Pfizer (New York, NY, USA) Otto Sensors (Cleveland, OH, USA) CDI-3M (Minneapolis, MN, USA) VIA Medical (San Diego, CA, USA)
Rather than threading very small sensors down through the A-line catheter, sensor arrays are being incorporated in the saline-flush solution line external to the patient (on-line systems). Periodically, samples of blood can be drawn up into the sensor chamber for measurements by reversing the pump direction used to infuse the patient with the flush solution. Several sensors can be placed in the chamber allowing the measurement of more parameters (blood-gases, electrolytes, and even glucose) [ 161. This approach partially overcomes the problems inherent in preparing very small sensors and at the same time reduces the signal stability requirements. Indeed, the flush solution can be used to continuously calibrate the sensors so that reliable measurements may be made even with sensors that have a significant drift. The downside of this approach is that the measurements are not truly continuous or real-time, and thus during calibration and sampling cycles rapid changes in the blood chemistry parameters may go undetected until the system is in the measurement mode. At this point, while implantable sensors for blood-gases and electrolytes would provide clinicians with invaluable real-time information that could enhance the quality of health care for critically ill patients, there are still lingering questions about the reliability of results obtained by these devices once they are placed within the confines of the radial artery. Although considerable research is in progress to solve the remaining issues that affect the accuracy of the various electrochemical and optical probes described above, it appears that the first generation of commercial multiparameter implantable sensors will likely serve more as trend monitors and will not obviate the need for routine blood-gas and electrolyte measurements on discrete blood samples. At the same time, the evolution of periodic on-line rather than continuous in vivo sensor technology could provide an attractive, temporary alternative until
systems
Optical intra-arterial blood gas (PB3300) Optical intra-arterial blood gas (BioSentry) Optical intra-arterial blood gas Optical intra-arterial blood gas Optical intra-arterial blood gas Hybrid electrochemical/optical intra-arterial blood gas Electrochemical intra-arterial blood gas On-line optical blood gas (CDI-System 2000) On-line electrochemical blood gas electrolytes, ‘glucose
in vivo sensors can truly meet all of the requirements outlined in Table 2. Acknowledgement The author is grateful to the National Institutes of Health (grant GM-28882) for supporting his research program on the development of new in vivo electrochemical gas and ion sensors. References M.E. Meyerhoff, Clin. Chem., 36 (1990) 1567. C.K. Mahutte, C.S.H. Sassoon, J.R. Muro, D.R. Hansman, T.P. Maxwell, W.W. Miller and M.Yafuso, J. Clin. Monit., 6 (1990) 147. G.S. Calabrese and L.M. O.‘Connell, Topics Cur-r Chem., 143 (1988) 51. P Anker, D. Ammann and W. Simon, Mikrochim. Acta., I (1983)1448. U. Oesch, D. Ammann and W. Simon, Clin. Chem., 32 (1986) 1448.
B.A. McKinely, B.A. Houtchens and J. Janata, ZonSelective Electrode Rev., 6 (1984) 1’73.
R.L. Coon, N.C.J. Lai and J.F? Karnpine, J. Appl. Phys., 40 ( 1976) 625.
M.E. Collison, G.V. Aebli, J. Petty and M.E. Meyerhoff, Anal. Chem., 61 (1989) 2365. B.A. Shapiro, Resp. Care, 37 (1992) 165. 10 W.W. Miller, M. Yafuso, C.F. Yan, H.K. Hui and S. Arick, Clin. Chem., 33 (1987) 1538. 11 O.S. Wolfbeis (Editor), Fiber Optic Chemical Sensors and Biosensors, Vols 1 and II; CRC Press, Boca Raton, FL, 199 1. 12 J. Janata, Anal. Chem., 59 (1987) 1351. 13 J.B. Yim, T.W. Hubbard, L.D. Melkerson, M.A. Sexton and B.M. Fieffen, U.S. Pcuent 5,047,627, September 10, 1991. 14 S.M. Barnard and D.R. Walt, Nature, 353 (1991) 338. 15 C.P Larson, Presentation at 14th International Symposium of the Electrolyte-Blood Gas Division of the American Association for Clinical Chemistry,
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Chatham, MA, May 1992.
16 D.K. Wong and W.S. Jordon, ht. J. Clin. Monit.
trends in analytical chemistry, vol. 12, no. 6, 1993
Mark E. Meyerhoff is at the Department of Chemisty, University of Michigan, Ann Arbor, MI, USA.
Camp., 9 (1992) 95.
Sensitive techniques for phospholipase determination in plants C. Thkvenot and J. Daussant Meudon, France
J. K&* and 0. Valentovii Prague, Czech Republic The breakdown of plant phospholipids appears to be mainly mediated by phospholipase D. Its role is not fully understood either in plant development or post-harvest metabolism. Thus, sensitive methods for the assessment of low activity levels are required for use in such studies. This article provides a comprehensive view of the available methods and presents the authors’ experiences with radiometric and biosensor assays.
Introduction The phospholipases catalyse the hydrolysis of diacylglycerophospholipids, which may be considered as derivatives of glycerophosphate two of whose hydroxyl groups are esterified by longchain fatty acids and whose phosphoryl group forms a phosphodiester bond with a polar moiety (Fig. 1). Phospholipids contain one of the following polar moieties: choline, ethanolamine, serine, glycerol or inositol. All four ester bonds in a phospholipid are susceptible to enzymatic cleavage. The phospholipases have specificity for the position on the glycerophosphate backbone rather than for the particular fatty acid or polar group. Phospholipase AZ, PLA;! (EC 3.1.1.4) which hydrolyses the acid moiety from the sn-2 position of a diacylglycerophospholipid is the best charac-
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D
terized phospholipase. Much work has been done on extracellular PLA, which is abundant in bee and snake venoms and in mammalian pancreas. It is among the enzymes with the smallest relative molecular mass and its sequence analysis and cloning have been reported. In contrast, PLAZ from other cell and tissue sources, are trace enzymes. It is more difficult to find data concerning phospholipase A,, PLA, (EC 3.1.1.3.2) which hydrolyses the acid moiety from the sn-1 position of a diacylglycerophospholipid. Enzymes with PLA, activity have been purified mainly from microorganisms and mammalian tissues. The enzymatic activity ascribed to simultaneous hydrolysis of acid moieties in sn-1 and sn-2 position of a diacylglycerophospholipid falls into the category of phospholipase B, PLB (EC 3.1.1 S). This concept, however, faced the criticism that the dual activity could reflect multiple enzyme activi-
(
0 \
cH2
II -o-c-
R1 pm1 &l
R2
-
; C -
/
0
-CH
/ 0 II CH2
-0-P-O I 0. /
z
Fig. 1. Sites of attack of phospholipases on a diacylglycerophospholipid. Rl and R2 represent longchain fatty acids; X represents a polar moiety; sn-1 , sn-2, m-3 refer to carbon positions on the glycerophosphate backbone.
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