- Email: [email protected]
Fast optochemical sensor for continuous monitoring of oxygen in breath-gas analysis
ELSEVIER
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Fast optochemical sensor for continuous monitoring of oxygen in breath-gas analysis C. Kolle a9*, W. Gruber a, W. Trettnak a, K. Biebernik a, C. Dolezal a, F. Reininger a, P. O’Leary b a Joanneurn
Research, Institute for Chemical and Optical Sensors, Steyrergasse 17, A-8010 b Institutefor Automation, University ofleoben, A-8700 koben, Austria
Graz, Austria
Abstract An application of an optochemical oxygen sensor to respiratory gas monitoring is presented. A new prototype module has been developed, which is suitable for the determination of oxygen concentration in mainstream breath-gas flow, directly at the patient’s mouth. The sensor is based on the measurement of the quenching of the intensity of photoluminescence of the dye platinum(U) -octaethylporphyrin-ketone by molecular oxygen. The dye has been immobilized in various polymers with and without plasticizer. The characteristics of the resulting oxygensensitive membranes have been determined with emphasis mainly on their oxygen response characteristics, such as calibration graphs, oxygen sensitivities, oxygen resolution and response times. The suitability of the optochemical oxygen sensors for clinical use in respiratory gasexchange measurements has been demonstrated in combination with a prototype instrument. Keywords:
Oxygen sensors; Optochemical sensors; Breath-gas analysis
1. Introduction
Fast real-time measurementsof O2 and CO2 in human breath are fundamental tasks in therapeutic and diagnostic medicine [ 11. For example, the shapeof curves obtained by recording O2 and COz in expired gas is modified by airway obstruction, and the analysis of the deformation provides an indirect measurementof lung disorders [ 21. For severaldiagnostic methods, the determination of carbon dioxide alone is not sufficient. Distribution disordersof gaseousdiffusion to or out of the blood system cannot be analysedby CO2 measurement alone becauseof the high diffusivity of carbon dioxide (about 20 timeshigher than for 0,) [ 31. Thus, the oxygen concentration in breath is an important additional parameter. Oxygen monitoring in pulmonary respiratory measurements involves a variety of clinical applications: adult and neonatal intensive care, anaesthesia,pulmonary function diagnostic (breath-by-breath oximetry) or stress/exercise testing (sports medicine). While in clinical monitoring applications (anaesthesia, intensive care) the measuring range
* Corresponding author. Phone: f43 316 876 226. Fax: f43 316 876 181. E-mail: [email protected]. 0925-4005/97/$17.00 0 1997 Elsevier Science S.A. All rights reserved PIISO925-4005 (97) 00040-3
covers 15 to 100 vol.% of oxygen, the oxygen concentration range in diagnostic medicine (15-21 vol.%) is essentially limited to breathing of ambient air. Generally mainstreammeasurements(the gasmixture is analysedin the primary flow directly at the patient’s mouth) and sidestream measurements(a small stream of gas is extracted from the primary flow) can be distinguished.Oxygen monitoring of gas mixtures in clinical applications is presently performed by instruments based on a variety of methods [ 1,4]. Conventional technical solutions may be classifiedasfollows: l Electrochemical sensors( polarographic or galvanic fuel cells) arecommonly employed in slow-responseapplications andthus restricted to the determination of the averageoxygen concentration. l Paramagnetic sensorsare fast, but this technology is utilized successfully in sidestreamconfigurations only. 0 Mass and Raman spectrometersare applied for sidestream measurements,with less suitability for widespread clinical applications becauseof their high price. l Zirconium oxide sensorsshow fast responsetimes but cannot be usedin the presenceof anaestheticagentsbecause of the high temperature ( > 650°C) inside the measurement
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cell. Thus, this principle is not frequently applied for medical purposes. Optical techniques for measuring oxygen in breath gas have also been suggested. For example, oxygen was measured based on its absorption band at 145 nm [5]. However, various substances such as water vapour or carbon dioxide were described to interfere. Another instrumentation has been presented, which is based on the measurement of oxygen concentration at the very narrow absorption line at 760 nm. A spectrally pure and precisely tuned laser is needed for this purpose, since the absorption of oxygen is very weak at this wavelength [ 61. Other approaches were based on optochemical oxygen measurements. The fast response of the luminescence (no data given) of fluoranthene using polyethylene as a matrix due to the change of oxygen partial pressure is reported [7]. A fast oxygen detector for analysing lung function using a fluorescent dye adsorbed on Porapack has been demonstrated to be as fast as 5 ms for 90% signal change [S] . With another fast-responding fluorescent sensor based on covalently immobilized pyrenebutyric acid on a porous glass support, a 90% response time smaller than 20 ms could be achieved [ 91. With an alternative method using [Ru(Ph,phen),]” immobilized in a silicone rubber matrix, a 95% response time smaller than 180 ms is obtained {lo]. In another approach, suitable for respiratory measurements with small laboratory animals, tetraphenyl-porphyrin is adsorbed on Porapack [ 111. In this set-up a 90% response time of 20 ms is reported. A further work with a measuring set-up in the sidestream reports a 90% response time of 31 ms [ 12]. Several luminescent oxygen-sensitive coatings designed for video luminescent barometry were studied with response times in the submillisecond range [ 131. However, it seems that the investigated systems have found no commercial application in breath-gas analysis to date. Up to now no fast-responding oxygen sensor for real-time monitoring of breath-gas in the mainstream is available. In accordance with available CO, modules for mainstream measurements, the required specifications for a mainstream oxygen module comprise a response time of less than 100 ms with a resolution of 0.1 vol.% of oxygen. The primary aim of this study was to investigate the possibilities offered by optical oxygen sensors in designing a prototype instrument for mainstream applications. The optical sensor is based on a new phosphorescent dye, which can be excited with a yellow light-emitting diode as the light source [ 14,151. Oxygen sensor membranes with the dye platinum( II) -octaethylporphyrin-ketone (PtOEPK) embedded in various polymers were investigated with respect to their oxygen response characteristics (calibration graphs, oxygen sensitivities, resolution and response times). Clinical tests (breath-by-breath analysis) were performed with a prototype instrument based on the measurement of luminescence intensity and in combination with a CO, sensor module from Hewlett-Packard (Andover, MA, USA).
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2. Experimental
2.1. Sensor membranes 2.1.1. Poly(viny1 chloride) membranes A solution of 1 g of poly (vinyl chloride) (PVC, high molecular weight; Sigma, Steinheim, Germany) in 10 ml tetrahydrofuran (THF; Merck, Darmstadt, Germany) was mixed with various amounts of solutions of 1 g of plasticizer in 10 ml THF. The following plasticizers were used: bis( 2ethylhexyl) sebacate (DOS), bis (2-ethylhexyl) phthalate (DOP) and tris( 2-ethylhexyl) trimellitate (TOTM) (all from Fluka, Buchs, Switzerland). 1 mg of platinum(II)octaethylporphyrin-ketone (PtO,EPK, Joanneum Research, Graz, Austria) was dissolved in 1 g of PVC or PVC/plasticizer solution. Membranes were cast on a transparent polyester support (175 km Mylar@, DuPont, Bad Homburg, Germany) by using a blade-coating technique and metal spacers with a thickness of 10 or 20 p,m, and left to dry at room temperature for at least 24 h.
2.1.2. Poly(styrene) membranes 1 mg of PtOEPK was dissolved in 1 g of a solution of 15 wt.% of poly( styrene) (average mol. wt. =280 000; Tg= 100°C; Sigma) in toluene. The dye polymer solution was cast onto Mylar@ with a wet-layer thickness of 20 km, and left to dry at room temperature for at least 24 h.
2.1.3. Sensor characterization Basic luminescence intensity measurements were carried out at 25’C and at ambient pressure (970 + 5 hPa) with a SPEX Fluorolog II fluorometer. The excitation and emission wavelengths were 592 and 758 nm, respectively. Dry gas mixtures were provided by a UTAH Medical Products PGM3 gas-mixing device with an accuracy of 0.25% absolute. Nitrogen served as the inert gas component.
2.2. Prototype instrument The prototype device for the determination of oxygen in breath gas consists of an airway adapter mounted on the primary flow tube and an electronic unit for the acquisition and processing of the signals coming from the sensor. To eliminate the condensation of water on the sensor surface, the sensor is heated and the temperature is regulated to 37°C. This airway adapter is connected via a shielded cable (approximately 2 m long) to the main module, which carries out the signal processing required to determine the fluorescence intensity from which the oxygen concentration can be calculated. The main electronics provide both analog (voltage) and digital (serial interface RS232) signals as output.
2.2. I. Airway-adapter construction The geometric dimensions of the measurement cell were chosen to enable easy handling and to facilitatemodifications during test of the optical and electronic components. The principle of the construction is shown in Fig. 1.
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2.2.2. -
cable entry
pie-amplifier and heating electronics
slots for heating transistors
02 sensitive spot
airway adapter 60 mm
luminescence excitation LED
I’I
optical filters
L--
.A-
Fig. 1. Mechanical construction of the airway ments: (a) side view; (b) front view.
adapter for clinical
measure-
The airway adapter is made of black polyoxymethylene. The tube diameter fulfils the requirementsto match the commercially available COP-detection modules (Novametrix, Wallingford, CT, USA; Hewlett-Packard, Andover, MA, USA). Tb.e O,-sensitive membrane (10 mm diameters)is located at a window in the wall of the tube. The optical cell is made of aluminium (black, anodically oxidized) and definesthe geometry of the optical measurementhead while serving asa holder for the optical filters (coloured glass) and optoelectronic components (light-emitting diode, silicon photodiodes). The cell geometry hasbeen designedto maximize the signal yield while minimizing stray light. A robust aluminium metal housingguaranteesmechanicalstability and shielding of the primary electronics (preamplifier) during clinical test.
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System electronics
A modulated system has been chosen, where a fixed-frequency excitation signal is used. The applied synchronous detection techniquemeasuresonly the information (luminescence) which is at the samefrequency asthe excitation signal. Therefore this measurementmethod eliminates errors introduced by ambient light interference and electrical drifts. The block diagram of the implementedhardware arrangementis shown in Fig. 2. The front-end configuration comprisesa yellow LED with a suitable optical bandpassfilter (BG 39, Schott, Mainz, Germany) serving as excitation light source; the emitted fluorescenceis detected by a silicon photodiode. The radiation of the LED, which is scattered and reflected by the membrane, is simultaneously monitored via a second photodiode. This enablesa monitoring and correction of LED intensity fluctuations over time. The transimpedanceamplifiers perform the pre-conditioning of the detector signals.A heating-type thermostat,whichutilizes twopowertransistors, is implementedto eliminate temperaturevariations. Temperature sensingis accomplishedusing a small thermistor bead mounted inside the aluminium cell. The cell temperature is regulated to 37 + O.l’C. In order to reduce noise and interference, a shieldedcable is usedto connect the amplifier to the main module. The input stageof the main module consistsof a passive lowpassfilter and an active bandpassfilter for each channel. The essential part of the filter stage is a Sallen-Key bandpassfilter with positive feedback. The primary objective is to improve the signal-to-noise ratio by limiting the noise bandwidth and eliminate the d.c. component in order to match the signal amplitude to the input range of the following stage.A synchronousdemodulation systemis usedto obtain the intensity signal of the luminescenceand the excitation LED. This synchronous detection is performed via a suitable integrated circuit [ 161. It comprisesa programmablegain amplifier, the synchrondemodulator and a 20-bit C/A analog-to-digital converter (ADC) with an additional 1280-tap FIR filter integrated on one chip. A single-chip microprocessoris the central unit of the instrument and provides a digital serialoutput (RS232 and CAN) and an analog voltage output with an update rate of 30.52 Hz. The excitation square-wave signal (244 Hz) is generated by modulating a precision current source. A specialset-up of the prototype wasdesignedto compare different sensorsdue to their temporal responsecharacteristics and the noisebehaviour. This modified arrangementcomprisesa smallsamplechamberwith thin gas-tubefittings and replacesthe mainstreamairway adapter.Additional miniature solenoid valves are employed to generatesuddenchangesof oxygen concentrations.The response-timetestswere carried out at a temperatureof 37 & O.l”C with a flow rate of approximately 500 ml min- ‘. During the teststhe averageexcitation irradiance was set to 50 p,W cm-‘. For clinical tests (single-breath analysis in mainstream), a CO2 module (HP M1016A with transducer HP 1436012,
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,.-----.----......REFERENCECWANNEL ;
opicilr
filters \
Photodetector ,“I!
Adtier r~
7 r
; * i I
hmp%er F&i
> I
SIGNAL CiANNEL t; f
Fig. 2. Block diagram of-the electronic signal processing (for explanation, see text).
Hewlett-Packard, Andover, MA, USA) wasconnectedto the O2airway adapterof the prototype instrument. Both gasconcentrationswere recorded simultaneouslyin the samebreathgasstreamby using a personalcomputer. An additional small temperature sensor (BetaTHERM microCHIP thermistor probe diameter 0.018”, BetaTHERM Corporation, Shrewsbury, MA, USA) was located close to the sensingmembranedirectly at the primary flow. With this temperaturemeasurementthe temperaturebehaviour of the sensorduring breathing wasapproximated anda corresponding correction of the luminescence intensity signal was applied.
tests have been performed with the well-known poly(styrene) [ 14,151matrix.
3.1. Oxygen calibration graphs The Stern-Volmer plots obtained with four different oxygen-sensitivemembranes(pure PVC membranesand PVC membranescontaining 25 vol.% of three different plasticizers) and oxygen concentrationsbetween 0 and 100vol.% are shown in Fig. 3. Only the calibration graph of the pure PVC membraneis linear and can be describedby the ideal SternVolmer equation:
IO 7= 1 +&.dw
(1)
3. Measurement results Three different plasticizers (DOS, DOP, TOTM) were identified for application in breath-gasanalysis. These not only fulfil the toxicity requirements,which are of essential importance for medical applications, but also show good physical and chemical properties. The characteristicsof the oxygen-sensitive membranes(calibration graph,oxygen sensitivity, oxygen resolution and responsetime) are affected by the type of polymer matrix, the type of plasticizer, the fraction of plasticizers in the polymer and the thicknessof the sensitive layer. The influence of matrix effects on the oxygen calibration curves hasbeensufficiently describedfor membranescontaining polystyrene, pure PVC and PVC with varying fractions of DOS [ 151. Since the various plasticizers may influence the oxygen responsecharacteristics,PVC sensor membranescontaining DOS, DOP and TOTM and pure PVC membraneswere comparedin this work. The time responseand the noise behaviour have not yet been investigated. Detailed measurementswith different polymer compositions were carried out. Preliminary clinical
where [O,] is the oxygen concentration in vol.%, KS, the Stern-Volmer quenchingconstant,I0 the luminescenceintensity in the absenceof oxygen andI the luminescenceintensity measuredat [O,] .
0
20
40
oxygen cmoentration
60
80
Go
[vol. %]
Fig. 3. Stem-Volmer plots obtained with oxygen-sensitive membranes consisting of pure PVC and of PVC with 25% plasticizer (TOTM, DOP and DOS) in dry oxygen/nitrogen mixtures aQWC.
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In all other cases(PVC with plasticizer), at least a twocomponent model has to be applied to describethe curve of the calibration graphssufficiently [ 171: IO -=
I
[
fl
l-f1
1+~svl[o,I
+ 1+~svz[o21
-I
1
(2)
where fi is the fractional contribution of KS”,, and (1 -fi) the fractional contribution of KSv2. To compare the sensitivities of the investigated sensor membranesover the full measurementrange, the tist derivative of the function Z/I, =f( 0,) is estimated by taking the tangents in known function values (measuredoxygen concentrations). This step is mathematically described as follows:
(3) The values obtained by these approximations (slope is negative) are shown on a logarithmic scalein Fig. 4. These graphs demonstratefor each sensormembranethe normalized luminescenceintensity change (I, = III,) for an increment of 1% in oxygen concentration over the full measurement range. As can be seen from Fig. 4, the PVC membrane with 25% DOS shows the highest sensitivity in the range O-10% oxygen, whereasthe pure PVC membrane showsthe highest sensitivity in the range 25-100% oxygen. Theseresults are not obvious from the graphsin Fig. 3, but are important for the choice of the most suitable sensor membranefor a given oxygen concentration range.
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level is strongly dependenton the oxygen level becauseof the non-linear Stem-Volmer relationship. Table 1 shows a comparison of the response-timecharacteristicsof different membraneswith the r.m.s. noise at one given oxygen concentration. Fig. 5 servesto illustrate the time responseof a selection of three different membranecompositions.
3.3. Clinical tests Before the breath sequenceof a patient was investigated, two calibration gases(one gasconsisting of 12% O2and 5% COZ in N2 and a secondof ambient air) were introduced into the measuringset-up. With these two calibration points, the oxygen concentration was calculated under the assumption of a linear Stern-Volmer equation (Eq. ( 1) ) in the range 1220.9 vol.% of oxygen. A comparison of the oxygen concentration measured with our prototype instrument and the carbon dioxide concentration measuredwith a commercially available insbument is shown in Fig. 6. The characteristic patterns are marked by alternating inspiratory and expiratory phasesduring breathing of a patient. While the capnogram (graphic presentation of COPconcentration) reflects the carbon dioxide exhalation, the oxygram (graphic presentationof O2concentration) corresponds to the oxygen uptake through the lungs. The capnogram B
180
e e g
170 HI
P 150
3.2. Response time and noise measurements The responsetimesfor 90% of the total signalchange( tgo) of various sensormembranes,which were recorded with the described special set-up, are summarized in Table 1. The measuredstandard deviation is in accordancewith the root mean square (r.m.s.) noise of the oxygen signal. This noise
140
?! $
130
6 0
120 ,9 time [s] (b)
- 180 dY t 170 2 160 E 150 ;
140 a 130 I 6 120 e time [s] Cc) - 180 fi f 170 I 160 B 150 \;;
oxygenconcentration [vol. %] Fig. 4. Oxygen sensitivities of oxygen-sensitive membranes consisting of pure PVC and of PVC with 25% plasticizer (TOTM, DOP and DOS) covering the full measurement range of O-100 vol.% oxygen (see also Fig. 3).
g 120 0 52
53
54
55
56
-57 58 time [sl
59
60
61
62
Fig. 5. Time response characteristic of a PVC membrane (20 pm wet layer thickness) with (a) 50% DOS fraction, (b) 25% DOS fraction and (c) without plasticizer (37°C).
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Table 1 Response time and noise (r.m.s.)
measurements
of different
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(37°C)
Sensor polymer
Wet layer thickness WC
Response time tsw (ms) step 123-177 hPa O2
Noise r.m.s. ( =std.dev., at 177 hPa O2
PVC:DOS 50:50 PVC:DOS 75:25 PVC:DOS 75:25 PVC:TOTM 75:25 PVC:DOP 75:25 Polystyrene PVCDOS 90: 10 PVC without plasticizer
20 10 20 20 20 20 20 20
66 a7 100 140 200 800 1070 2400
0.83 0.54 0.28 0.26 0.27 0.21 0.28 0.31
35
Fig. 6. Characteristic
40
45
pattern of simultaneously
50
obtained
60 time (2.)
65
55
capnography
and oxygraphy
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80
75
during
hPa)
human breathing
of ambient
air
37,o 365
360 355 35a-l..,.I~~~~1.~‘,i.~‘~I’~“i’~“i~~”1~”’I”’~I 0 5 IO 15
Fig. 7. Detailed view of one breath cycle (in normal conditions record is a mirror image of the carbon dioxide curve).
the oxygen
should be a mirror image of the oxygram, Fig. 7 shows a single breath cycle (detail view of Fig. 6). From this curve it can be seen that the O2 recording seems to be slower in response. With the airway adapter we have observed temperature changes associated with breathing due to cooling of the sensor surface during inspiration, although the temperature is regulated. The ambient air flow across the sensor decreases the temperature of the oxygen-sensitive layer and thus causes changes in the luminescence intensity. The temperature controller is not able to compensate this temperature deviation immediately. Particularly in the case of hyperventilation, the
20 25 time (s)
34
35
40
45
Fig. 8. Comparison of (a) an uncompensated oxygen graph with (b) a realtime temperature-compensated oxygen curve. The lower graph is the monitored temperature at the sensor surface, which is taken for compensation.
luminescence intensity variations due to temperature effects can lead to significant errors in the measurement. The lower graph in Fig. 8 reflects the temperature change at the sensor surface during respiration. These temperature data were used to implement a compensation algorithm to obtain a stable inspiration baseline. In this approach a linear approximation is supposed to establish a model for the temperature behaviour of the sensor in the expected temperature range. Both the raw data recordings of breath-by-breath analysis and the mathematically treated data can be seen in Fig. 8.
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4. Discussion A number of oxygen-sensitive membranes have been investigated with respect to their utility in respiration-gas measurements. The results reflect that there are several influencing parameters which must be considered for the selection of a breath-gas sensor. The calibration graph obtained with an oxygen-sensitive membrane is influenced by the type of polymer used, the type of plasticizer and the concentration of the plasticizer. Addition of plasticizer to PVC leads to an increased oxygen sensitivity of the membrane. But whereas the calibration graph was linear with pure PVC, it became clearly non-linear in the presence of the plasticizer, thus complicating the calibration of the sensor. The sensor needs a goodresolution over the whole concentration range of interest to meet the requirements of a reliable measurement. Whereas PVC membranes with 25% DOS showed very high sensitivity to oxygen below 10 vol.%, it was found that pure PVC gave the best results (and consequently highest resolution), in comparison to plasticized PVC membranes, at O2 concentrations above approximately 25%. Preliminary results have shown that the storing of the membranes at 25°C over 1 month did not change the sensor characteristics significantly. Current investigations regarding the behaviour at 37’C are in progress. Another feature which is of utmost importance for breathgas analysis is the response time of the sensor. In this case the added plasticizer improves the response characteristics. A problem associated with the used set-up is the fact that it is very difficult to quantify the true response time of the oxygen-sensitive membranes. The measured data represent a sum of the response characteristics of the pneumatic set-up, the speed of gas-concentration change and the limited bandwidth of the electronic detection unit. Several parameters influencing the response time can be deduced from Table 1: the type of polymer and plasticizer, the fraction of plasticizers in the polymer and the thickness of the sensitive layer. These factors are also inseparably linked with the required resolution (limited by the noise) of the measuring system (Table 1). The obtained noise depends to a high extent on the excitation light intensity. A further increase of this intensity may cause increasing photobleaching of the dye. This photobleaching is a limiting factor of the stability and calibration requirements of the optochemical oxygen sensor at the moment. With the frequently used system the maximum error due to photobleaching of the dye (PtOEPK PVC:DOS 75:25) within 1 h of continuous measurement is smaller than 0.1 vol.% oxygen. Thus, to meet the requirements of ameasuring stability of &- 0.1 vol.% 02, a recalibration is necessary every hour. An obvious reduction of excitation light intensity again leads to high demands on the quality of signal processing to meet the resolution requirements of an efficient fast oxygen analyser. As a result of all these facts, it can be concluded that it is a difficult task to select an adequate combination of sensor composition and instrumentation set-up.
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The results obtained from the clinical tests are very promising with respect to the agreement between oxygen and carbon dioxide measurements. The rapid rise of CO2 concentration at the beginning of the expiration (Fig. 7)) corresponding to expiration air, seems to be time shifted to the decrease in O2 concentration. This is due to the specific geometric set-up of the detectors. The steeper slope of the carbon dioxide response characteristic is true. However, these initial experiments have been made with relatively slow-responding sensors (poly( styrene) membranes). Further tests with the faster oxygen-sensitive membranes (Table 1) have not yet been performed. Using these membranes, we expect an exact conformity of the time response of COZ and O2 measurements. The experiments only cover the requirements for lungfunction tests in the measurement range 15-21% OZ. For example, for intensive care or anaesthesia, measurements in the full range up to 100% oxygen are necessary and possible cross-sensitivities to anaesthetic agents have to be investigated. Corresponding work relating the choice of a suitable sensor membrane in combination with an optimized electronic set-up for a required measurement range is in progress. With the additional temperature-compensation method suppression of an essential temperature influence could be obtained. However, this method is not yet mature. One of the keys to this problem certainly is to tune the temperature response time of the oxygen sensor membrane with the response time of the utilized temperature sensor in the gas stream. Another approach to compensate temperature effects could be the combination of the optochemical oxygen sensor with a luminescence-based temperature sensor [ 181.
5. Conclusions
and possible improvements
A prototype instrument for respiration gas exchange based on an optochemical oxygen sensor has been developed and verified experimentally. It can be emphasized that this detection principle is a viable alternative to other oxygen-measurement techniques in medical gas-exchange measurements, with the optical sensors offering some essential features: 0 The measurement can be carried out directly in the primary flow. l The sensor is small and cheap. Sterilization problems, which are evident with any medical application, can be prevented by changing the sensor spot after each measurement. 0 The oxygen sensor works in the concentration range O100% oxygen, opening applications in many fields of clinical gas analysis. l In comparison with other methods (e.g., paramagnetic), the sensor is not affected by humidity, nor is it poisoned by CO,. l The sensors require no moving parts. Consequently, they are more stable and cheaper to construct than paramagnetic instruments. It is possible to apply the sensors in vibrating and harsh environments.
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In the future we intend to intensify’our investigations in the selectionof a proper polymer or polymer/plasticizer combination, especially with respect to sensorstability due the photobleaching of the dye. To overcome temperatureproblems,the temperaturecharacteristicsof the sensormembranes have to be investigated in order to find a reliable temperature compensation.Moreover, and looking further ahead,we may expect that there areconsiderableadvantagesin replacingthe intensity-measuring principle with a luminescencelifetimemeasuringtechnique [ 19-2 11.
Acknowledgements The authors wish to thank ProfessorK. Harnoncourt and his colleagues (II. Medizinische Abteilung, LKH Graz) for their assistanceduring the clinical tests.
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[6] Fast Oxygen Gas Sensor, data sheet, Oxygraf, Palo Alto, CA, USA, 1994. [7] I. Bergman, Rapid response atmospheric oxygen monitor based on fluorescence quenching, Nature, 218 (1968) 396. [S] 0. Burkhard and W.K.R. Barnikol, Ein schneller O,-Detektor fiir die Analyse der Lungenfunktion, Prax. Klin. Pneumol., 37 (1983) 805807. [9] OX Wolfbeis, H. Offenbacher, H. Kroneis and H. Marsoner, A fast responding fluorescence sensor for oxygen,Mikrochim. Actu, I (1984) 153-15s. [ 10 ] J.R. Bacon and J.N. Demas, Determination of oxygen concentrations by luminescence quenching of apolymer-immobilisedtransition-metal complex, Anal. Chem., 59 (1987) 2780-2785. [ll] W.K.R. Barnikol, Th. Gaertner, N. Weiler and 0. Burkhard, Microdetector for rapid changes of oxygen partial pressure (PO,) during the respiratory cycle in small laboratory animals, Rev. Sci. Instrum., 59 (1988) 1204-1208. I:121 H. Karpf, H.W. Kroneis, H.J. Marsoner, H. Metzler and N. Gravenstein, Fast responding oxygen sensor for respiratorial analysis, Proc. Chemical, Biochemical Environmental Fiber Sensors, Boston, MA, USA, 6-7 Sept. 1989, Vol. 1172, pp. 295-304.
[13] A.E. Baron, J.D.S. Danielson, M. Gouterman, J.R. Wan, J.B. Callis and B. McLacblan, Submillisecond response times of oxygenquenched luminescent coatings, Rev. Sci. Instrum., 64 (1993) 33943402.
Christian Kolle wasborn in 1966 in Klagenfurt, Austria. He received the Dip.-Ing. degree in electrical engineering from the Technical University Graz in 1993. He joined the Institute for Chemical and Optical Sensors at Joanneum Research (Graz, Austria) in 1994. He is presently working on the developmentof the instrumentationfor optical oxygen sensors,particularly in the field of breath-gasanalysis. His research interests are in electronics instrumentation and measurement, signal processing and electromagnetic compatibility. Wolfgang Gruber was born in 1961 in Graz, Austria. He received the Dipl-Ing. degree in electrical engineeringfrom the Technical University of Graz, Austria, in 1989. Since 1984 he hasgiven lecturesin electronics at the Wirtschaftsforderungsinstitut in Graz. From 1991 to 1992 he joined the Institute for Optical Sensors,JoanneumResearch,Graz, Austria. Since 1992 he has been with the Institute for Sensor Interfaces and the Institute for Chemical andOptical Sensors, JoanneumResearch,Graz, Austria, where he is working on integrated sensorsystems.He is presently working towards the Ph.D. degreein electrical engineering at the Technical University of Graz, Austria. W&gang Trettnak, born in Graz (Austria) in 1962, received a Ph.D. in chemistry from the Karl-Franzens-University in Graz in 1989. Since 1986 he hasbeen concerned with the development and testing of opto- and electrochemical sensorsand biosensorsfor various parameterssuch as 02, pH, C02, NH3, glucose,lactate and cholesterol. With the exception of a one-year researchproject at the University of Florence in 1992, he hasbeenactive at JoanneumResearch, Graz, Austria, since 1991.
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Karin Biebemik received a Ph.D. degree in chemistry from the Technical University of Graz, Austria. In 1991, she took on a postdoctoral fellowship at the University of Toronto, Canada, where she worked on oxygen reduction electrodes at the Department of Chemistry. In 1995, she joined Joanneum Research at the Institute of Chemical and Optical Sensors, where she is working on the development and characterization of optical sensor membranes. Constanze Dolezal, born in 1960 in Graz, received a degree at the College for Chemotechnical Engineering in Graz, Austria. Since 1994 she has been with the Institute for Chemical and Optical Sensors, Joanneum Research (Graz, Austria), where she is working mainly on the development of optochemical sensors and testing of prototype instruments. Franz Reininger, born in Leibnitz (Austria) in 1961, received a Dipl.-Ing. degree in electrical engineering from the Technical University in Graz in 1989. In 1988 he joined Austria-Mikro-Systeme-International GmbH (Unterprems&ten, Austria) where he did his thesis ‘Evaluation of AD converter and DA converter and dynamic performance’.
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Since 1990 he has been at Joanneum Research in Graz, Austria, where he is involved in designing and building new sensor interfaces for different sensor types, especially for new optical gas sensors. His interests are in dataconversion, signal processing and calibration of special optochemical sensors. In 1993 he became head of the working group of optochemical sensors and in 1995 he became the operative manager of the Institute for Chemical and Optical Sensors at Joanneum Research. Paul O’Leary was born in 1960 in Croom, Co. Limerick, Ireland. He holds a B.A. in mathematics, aB.A.1. in electronic engineering, an M.E.E. in electronic engineering and a Ph.D. in signal processing structures for sensor systems. He founded the Institute for Chemical and Optical Sensors within Joanneum Research in September 1990. Since then he has been responsible for the development and growth of the group. He has 11 years’ professional experience in the development of integrated circuits and complex systems. In 1995 he became a professor and head of the Institute for Automation (University of Leoben, Austria).
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Fast optochemical sensor for continuous monitoring of oxygen in breath-gas analysis C. Kolle a9*, W. Gruber a, W. Trettnak a, K. Biebernik a, C. Dolezal a, F. Reininger a, P. O’Leary b a Joanneurn
Research, Institute for Chemical and Optical Sensors, Steyrergasse 17, A-8010 b Institutefor Automation, University ofleoben, A-8700 koben, Austria
Graz, Austria
Abstract An application of an optochemical oxygen sensor to respiratory gas monitoring is presented. A new prototype module has been developed, which is suitable for the determination of oxygen concentration in mainstream breath-gas flow, directly at the patient’s mouth. The sensor is based on the measurement of the quenching of the intensity of photoluminescence of the dye platinum(U) -octaethylporphyrin-ketone by molecular oxygen. The dye has been immobilized in various polymers with and without plasticizer. The characteristics of the resulting oxygensensitive membranes have been determined with emphasis mainly on their oxygen response characteristics, such as calibration graphs, oxygen sensitivities, oxygen resolution and response times. The suitability of the optochemical oxygen sensors for clinical use in respiratory gasexchange measurements has been demonstrated in combination with a prototype instrument. Keywords:
Oxygen sensors; Optochemical sensors; Breath-gas analysis
1. Introduction
Fast real-time measurementsof O2 and CO2 in human breath are fundamental tasks in therapeutic and diagnostic medicine [ 11. For example, the shapeof curves obtained by recording O2 and COz in expired gas is modified by airway obstruction, and the analysis of the deformation provides an indirect measurementof lung disorders [ 21. For severaldiagnostic methods, the determination of carbon dioxide alone is not sufficient. Distribution disordersof gaseousdiffusion to or out of the blood system cannot be analysedby CO2 measurement alone becauseof the high diffusivity of carbon dioxide (about 20 timeshigher than for 0,) [ 31. Thus, the oxygen concentration in breath is an important additional parameter. Oxygen monitoring in pulmonary respiratory measurements involves a variety of clinical applications: adult and neonatal intensive care, anaesthesia,pulmonary function diagnostic (breath-by-breath oximetry) or stress/exercise testing (sports medicine). While in clinical monitoring applications (anaesthesia, intensive care) the measuring range
* Corresponding author. Phone: f43 316 876 226. Fax: f43 316 876 181. E-mail: [email protected]. 0925-4005/97/$17.00 0 1997 Elsevier Science S.A. All rights reserved PIISO925-4005 (97) 00040-3
covers 15 to 100 vol.% of oxygen, the oxygen concentration range in diagnostic medicine (15-21 vol.%) is essentially limited to breathing of ambient air. Generally mainstreammeasurements(the gasmixture is analysedin the primary flow directly at the patient’s mouth) and sidestream measurements(a small stream of gas is extracted from the primary flow) can be distinguished.Oxygen monitoring of gas mixtures in clinical applications is presently performed by instruments based on a variety of methods [ 1,4]. Conventional technical solutions may be classifiedasfollows: l Electrochemical sensors( polarographic or galvanic fuel cells) arecommonly employed in slow-responseapplications andthus restricted to the determination of the averageoxygen concentration. l Paramagnetic sensorsare fast, but this technology is utilized successfully in sidestreamconfigurations only. 0 Mass and Raman spectrometersare applied for sidestream measurements,with less suitability for widespread clinical applications becauseof their high price. l Zirconium oxide sensorsshow fast responsetimes but cannot be usedin the presenceof anaestheticagentsbecause of the high temperature ( > 650°C) inside the measurement
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cell. Thus, this principle is not frequently applied for medical purposes. Optical techniques for measuring oxygen in breath gas have also been suggested. For example, oxygen was measured based on its absorption band at 145 nm [5]. However, various substances such as water vapour or carbon dioxide were described to interfere. Another instrumentation has been presented, which is based on the measurement of oxygen concentration at the very narrow absorption line at 760 nm. A spectrally pure and precisely tuned laser is needed for this purpose, since the absorption of oxygen is very weak at this wavelength [ 61. Other approaches were based on optochemical oxygen measurements. The fast response of the luminescence (no data given) of fluoranthene using polyethylene as a matrix due to the change of oxygen partial pressure is reported [7]. A fast oxygen detector for analysing lung function using a fluorescent dye adsorbed on Porapack has been demonstrated to be as fast as 5 ms for 90% signal change [S] . With another fast-responding fluorescent sensor based on covalently immobilized pyrenebutyric acid on a porous glass support, a 90% response time smaller than 20 ms could be achieved [ 91. With an alternative method using [Ru(Ph,phen),]” immobilized in a silicone rubber matrix, a 95% response time smaller than 180 ms is obtained {lo]. In another approach, suitable for respiratory measurements with small laboratory animals, tetraphenyl-porphyrin is adsorbed on Porapack [ 111. In this set-up a 90% response time of 20 ms is reported. A further work with a measuring set-up in the sidestream reports a 90% response time of 31 ms [ 12]. Several luminescent oxygen-sensitive coatings designed for video luminescent barometry were studied with response times in the submillisecond range [ 131. However, it seems that the investigated systems have found no commercial application in breath-gas analysis to date. Up to now no fast-responding oxygen sensor for real-time monitoring of breath-gas in the mainstream is available. In accordance with available CO, modules for mainstream measurements, the required specifications for a mainstream oxygen module comprise a response time of less than 100 ms with a resolution of 0.1 vol.% of oxygen. The primary aim of this study was to investigate the possibilities offered by optical oxygen sensors in designing a prototype instrument for mainstream applications. The optical sensor is based on a new phosphorescent dye, which can be excited with a yellow light-emitting diode as the light source [ 14,151. Oxygen sensor membranes with the dye platinum( II) -octaethylporphyrin-ketone (PtOEPK) embedded in various polymers were investigated with respect to their oxygen response characteristics (calibration graphs, oxygen sensitivities, resolution and response times). Clinical tests (breath-by-breath analysis) were performed with a prototype instrument based on the measurement of luminescence intensity and in combination with a CO, sensor module from Hewlett-Packard (Andover, MA, USA).
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2. Experimental
2.1. Sensor membranes 2.1.1. Poly(viny1 chloride) membranes A solution of 1 g of poly (vinyl chloride) (PVC, high molecular weight; Sigma, Steinheim, Germany) in 10 ml tetrahydrofuran (THF; Merck, Darmstadt, Germany) was mixed with various amounts of solutions of 1 g of plasticizer in 10 ml THF. The following plasticizers were used: bis( 2ethylhexyl) sebacate (DOS), bis (2-ethylhexyl) phthalate (DOP) and tris( 2-ethylhexyl) trimellitate (TOTM) (all from Fluka, Buchs, Switzerland). 1 mg of platinum(II)octaethylporphyrin-ketone (PtO,EPK, Joanneum Research, Graz, Austria) was dissolved in 1 g of PVC or PVC/plasticizer solution. Membranes were cast on a transparent polyester support (175 km Mylar@, DuPont, Bad Homburg, Germany) by using a blade-coating technique and metal spacers with a thickness of 10 or 20 p,m, and left to dry at room temperature for at least 24 h.
2.1.2. Poly(styrene) membranes 1 mg of PtOEPK was dissolved in 1 g of a solution of 15 wt.% of poly( styrene) (average mol. wt. =280 000; Tg= 100°C; Sigma) in toluene. The dye polymer solution was cast onto Mylar@ with a wet-layer thickness of 20 km, and left to dry at room temperature for at least 24 h.
2.1.3. Sensor characterization Basic luminescence intensity measurements were carried out at 25’C and at ambient pressure (970 + 5 hPa) with a SPEX Fluorolog II fluorometer. The excitation and emission wavelengths were 592 and 758 nm, respectively. Dry gas mixtures were provided by a UTAH Medical Products PGM3 gas-mixing device with an accuracy of 0.25% absolute. Nitrogen served as the inert gas component.
2.2. Prototype instrument The prototype device for the determination of oxygen in breath gas consists of an airway adapter mounted on the primary flow tube and an electronic unit for the acquisition and processing of the signals coming from the sensor. To eliminate the condensation of water on the sensor surface, the sensor is heated and the temperature is regulated to 37°C. This airway adapter is connected via a shielded cable (approximately 2 m long) to the main module, which carries out the signal processing required to determine the fluorescence intensity from which the oxygen concentration can be calculated. The main electronics provide both analog (voltage) and digital (serial interface RS232) signals as output.
2.2. I. Airway-adapter construction The geometric dimensions of the measurement cell were chosen to enable easy handling and to facilitatemodifications during test of the optical and electronic components. The principle of the construction is shown in Fig. 1.
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2.2.2. -
cable entry
pie-amplifier and heating electronics
slots for heating transistors
02 sensitive spot
airway adapter 60 mm
luminescence excitation LED
I’I
optical filters
L--
.A-
Fig. 1. Mechanical construction of the airway ments: (a) side view; (b) front view.
adapter for clinical
measure-
The airway adapter is made of black polyoxymethylene. The tube diameter fulfils the requirementsto match the commercially available COP-detection modules (Novametrix, Wallingford, CT, USA; Hewlett-Packard, Andover, MA, USA). Tb.e O,-sensitive membrane (10 mm diameters)is located at a window in the wall of the tube. The optical cell is made of aluminium (black, anodically oxidized) and definesthe geometry of the optical measurementhead while serving asa holder for the optical filters (coloured glass) and optoelectronic components (light-emitting diode, silicon photodiodes). The cell geometry hasbeen designedto maximize the signal yield while minimizing stray light. A robust aluminium metal housingguaranteesmechanicalstability and shielding of the primary electronics (preamplifier) during clinical test.
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System electronics
A modulated system has been chosen, where a fixed-frequency excitation signal is used. The applied synchronous detection techniquemeasuresonly the information (luminescence) which is at the samefrequency asthe excitation signal. Therefore this measurementmethod eliminates errors introduced by ambient light interference and electrical drifts. The block diagram of the implementedhardware arrangementis shown in Fig. 2. The front-end configuration comprisesa yellow LED with a suitable optical bandpassfilter (BG 39, Schott, Mainz, Germany) serving as excitation light source; the emitted fluorescenceis detected by a silicon photodiode. The radiation of the LED, which is scattered and reflected by the membrane, is simultaneously monitored via a second photodiode. This enablesa monitoring and correction of LED intensity fluctuations over time. The transimpedanceamplifiers perform the pre-conditioning of the detector signals.A heating-type thermostat,whichutilizes twopowertransistors, is implementedto eliminate temperaturevariations. Temperature sensingis accomplishedusing a small thermistor bead mounted inside the aluminium cell. The cell temperature is regulated to 37 + O.l’C. In order to reduce noise and interference, a shieldedcable is usedto connect the amplifier to the main module. The input stageof the main module consistsof a passive lowpassfilter and an active bandpassfilter for each channel. The essential part of the filter stage is a Sallen-Key bandpassfilter with positive feedback. The primary objective is to improve the signal-to-noise ratio by limiting the noise bandwidth and eliminate the d.c. component in order to match the signal amplitude to the input range of the following stage.A synchronousdemodulation systemis usedto obtain the intensity signal of the luminescenceand the excitation LED. This synchronous detection is performed via a suitable integrated circuit [ 161. It comprisesa programmablegain amplifier, the synchrondemodulator and a 20-bit C/A analog-to-digital converter (ADC) with an additional 1280-tap FIR filter integrated on one chip. A single-chip microprocessoris the central unit of the instrument and provides a digital serialoutput (RS232 and CAN) and an analog voltage output with an update rate of 30.52 Hz. The excitation square-wave signal (244 Hz) is generated by modulating a precision current source. A specialset-up of the prototype wasdesignedto compare different sensorsdue to their temporal responsecharacteristics and the noisebehaviour. This modified arrangementcomprisesa smallsamplechamberwith thin gas-tubefittings and replacesthe mainstreamairway adapter.Additional miniature solenoid valves are employed to generatesuddenchangesof oxygen concentrations.The response-timetestswere carried out at a temperatureof 37 & O.l”C with a flow rate of approximately 500 ml min- ‘. During the teststhe averageexcitation irradiance was set to 50 p,W cm-‘. For clinical tests (single-breath analysis in mainstream), a CO2 module (HP M1016A with transducer HP 1436012,
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,.-----.----......REFERENCECWANNEL ;
opicilr
filters \
Photodetector ,“I!
Adtier r~
7 r
; * i I
hmp%er F&i
> I
SIGNAL CiANNEL t; f
Fig. 2. Block diagram of-the electronic signal processing (for explanation, see text).
Hewlett-Packard, Andover, MA, USA) wasconnectedto the O2airway adapterof the prototype instrument. Both gasconcentrationswere recorded simultaneouslyin the samebreathgasstreamby using a personalcomputer. An additional small temperature sensor (BetaTHERM microCHIP thermistor probe diameter 0.018”, BetaTHERM Corporation, Shrewsbury, MA, USA) was located close to the sensingmembranedirectly at the primary flow. With this temperaturemeasurementthe temperaturebehaviour of the sensorduring breathing wasapproximated anda corresponding correction of the luminescence intensity signal was applied.
tests have been performed with the well-known poly(styrene) [ 14,151matrix.
3.1. Oxygen calibration graphs The Stern-Volmer plots obtained with four different oxygen-sensitivemembranes(pure PVC membranesand PVC membranescontaining 25 vol.% of three different plasticizers) and oxygen concentrationsbetween 0 and 100vol.% are shown in Fig. 3. Only the calibration graph of the pure PVC membraneis linear and can be describedby the ideal SternVolmer equation:
IO 7= 1 +&.dw
(1)
3. Measurement results Three different plasticizers (DOS, DOP, TOTM) were identified for application in breath-gasanalysis. These not only fulfil the toxicity requirements,which are of essential importance for medical applications, but also show good physical and chemical properties. The characteristicsof the oxygen-sensitive membranes(calibration graph,oxygen sensitivity, oxygen resolution and responsetime) are affected by the type of polymer matrix, the type of plasticizer, the fraction of plasticizers in the polymer and the thicknessof the sensitive layer. The influence of matrix effects on the oxygen calibration curves hasbeensufficiently describedfor membranescontaining polystyrene, pure PVC and PVC with varying fractions of DOS [ 151. Since the various plasticizers may influence the oxygen responsecharacteristics,PVC sensor membranescontaining DOS, DOP and TOTM and pure PVC membraneswere comparedin this work. The time responseand the noise behaviour have not yet been investigated. Detailed measurementswith different polymer compositions were carried out. Preliminary clinical
where [O,] is the oxygen concentration in vol.%, KS, the Stern-Volmer quenchingconstant,I0 the luminescenceintensity in the absenceof oxygen andI the luminescenceintensity measuredat [O,] .
0
20
40
oxygen cmoentration
60
80
Go
[vol. %]
Fig. 3. Stem-Volmer plots obtained with oxygen-sensitive membranes consisting of pure PVC and of PVC with 25% plasticizer (TOTM, DOP and DOS) in dry oxygen/nitrogen mixtures aQWC.
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In all other cases(PVC with plasticizer), at least a twocomponent model has to be applied to describethe curve of the calibration graphssufficiently [ 171: IO -=
I
[
fl
l-f1
1+~svl[o,I
+ 1+~svz[o21
-I
1
(2)
where fi is the fractional contribution of KS”,, and (1 -fi) the fractional contribution of KSv2. To compare the sensitivities of the investigated sensor membranesover the full measurementrange, the tist derivative of the function Z/I, =f( 0,) is estimated by taking the tangents in known function values (measuredoxygen concentrations). This step is mathematically described as follows:
(3) The values obtained by these approximations (slope is negative) are shown on a logarithmic scalein Fig. 4. These graphs demonstratefor each sensormembranethe normalized luminescenceintensity change (I, = III,) for an increment of 1% in oxygen concentration over the full measurement range. As can be seen from Fig. 4, the PVC membrane with 25% DOS shows the highest sensitivity in the range O-10% oxygen, whereasthe pure PVC membrane showsthe highest sensitivity in the range 25-100% oxygen. Theseresults are not obvious from the graphsin Fig. 3, but are important for the choice of the most suitable sensor membranefor a given oxygen concentration range.
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level is strongly dependenton the oxygen level becauseof the non-linear Stem-Volmer relationship. Table 1 shows a comparison of the response-timecharacteristicsof different membraneswith the r.m.s. noise at one given oxygen concentration. Fig. 5 servesto illustrate the time responseof a selection of three different membranecompositions.
3.3. Clinical tests Before the breath sequenceof a patient was investigated, two calibration gases(one gasconsisting of 12% O2and 5% COZ in N2 and a secondof ambient air) were introduced into the measuringset-up. With these two calibration points, the oxygen concentration was calculated under the assumption of a linear Stern-Volmer equation (Eq. ( 1) ) in the range 1220.9 vol.% of oxygen. A comparison of the oxygen concentration measured with our prototype instrument and the carbon dioxide concentration measuredwith a commercially available insbument is shown in Fig. 6. The characteristic patterns are marked by alternating inspiratory and expiratory phasesduring breathing of a patient. While the capnogram (graphic presentation of COPconcentration) reflects the carbon dioxide exhalation, the oxygram (graphic presentationof O2concentration) corresponds to the oxygen uptake through the lungs. The capnogram B
180
e e g
170 HI
P 150
3.2. Response time and noise measurements The responsetimesfor 90% of the total signalchange( tgo) of various sensormembranes,which were recorded with the described special set-up, are summarized in Table 1. The measuredstandard deviation is in accordancewith the root mean square (r.m.s.) noise of the oxygen signal. This noise
140
?! $
130
6 0
120 ,9 time [s] (b)
- 180 dY t 170 2 160 E 150 ;
140 a 130 I 6 120 e time [s] Cc) - 180 fi f 170 I 160 B 150 \;;
oxygenconcentration [vol. %] Fig. 4. Oxygen sensitivities of oxygen-sensitive membranes consisting of pure PVC and of PVC with 25% plasticizer (TOTM, DOP and DOS) covering the full measurement range of O-100 vol.% oxygen (see also Fig. 3).
g 120 0 52
53
54
55
56
-57 58 time [sl
59
60
61
62
Fig. 5. Time response characteristic of a PVC membrane (20 pm wet layer thickness) with (a) 50% DOS fraction, (b) 25% DOS fraction and (c) without plasticizer (37°C).
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Table 1 Response time and noise (r.m.s.)
measurements
of different
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(37°C)
Sensor polymer
Wet layer thickness WC
Response time tsw (ms) step 123-177 hPa O2
Noise r.m.s. ( =std.dev., at 177 hPa O2
PVC:DOS 50:50 PVC:DOS 75:25 PVC:DOS 75:25 PVC:TOTM 75:25 PVC:DOP 75:25 Polystyrene PVCDOS 90: 10 PVC without plasticizer
20 10 20 20 20 20 20 20
66 a7 100 140 200 800 1070 2400
0.83 0.54 0.28 0.26 0.27 0.21 0.28 0.31
35
Fig. 6. Characteristic
40
45
pattern of simultaneously
50
obtained
60 time (2.)
65
55
capnography
and oxygraphy
70
80
75
during
hPa)
human breathing
of ambient
air
37,o 365
360 355 35a-l..,.I~~~~1.~‘,i.~‘~I’~“i’~“i~~”1~”’I”’~I 0 5 IO 15
Fig. 7. Detailed view of one breath cycle (in normal conditions record is a mirror image of the carbon dioxide curve).
the oxygen
should be a mirror image of the oxygram, Fig. 7 shows a single breath cycle (detail view of Fig. 6). From this curve it can be seen that the O2 recording seems to be slower in response. With the airway adapter we have observed temperature changes associated with breathing due to cooling of the sensor surface during inspiration, although the temperature is regulated. The ambient air flow across the sensor decreases the temperature of the oxygen-sensitive layer and thus causes changes in the luminescence intensity. The temperature controller is not able to compensate this temperature deviation immediately. Particularly in the case of hyperventilation, the
20 25 time (s)
34
35
40
45
Fig. 8. Comparison of (a) an uncompensated oxygen graph with (b) a realtime temperature-compensated oxygen curve. The lower graph is the monitored temperature at the sensor surface, which is taken for compensation.
luminescence intensity variations due to temperature effects can lead to significant errors in the measurement. The lower graph in Fig. 8 reflects the temperature change at the sensor surface during respiration. These temperature data were used to implement a compensation algorithm to obtain a stable inspiration baseline. In this approach a linear approximation is supposed to establish a model for the temperature behaviour of the sensor in the expected temperature range. Both the raw data recordings of breath-by-breath analysis and the mathematically treated data can be seen in Fig. 8.
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4. Discussion A number of oxygen-sensitive membranes have been investigated with respect to their utility in respiration-gas measurements. The results reflect that there are several influencing parameters which must be considered for the selection of a breath-gas sensor. The calibration graph obtained with an oxygen-sensitive membrane is influenced by the type of polymer used, the type of plasticizer and the concentration of the plasticizer. Addition of plasticizer to PVC leads to an increased oxygen sensitivity of the membrane. But whereas the calibration graph was linear with pure PVC, it became clearly non-linear in the presence of the plasticizer, thus complicating the calibration of the sensor. The sensor needs a goodresolution over the whole concentration range of interest to meet the requirements of a reliable measurement. Whereas PVC membranes with 25% DOS showed very high sensitivity to oxygen below 10 vol.%, it was found that pure PVC gave the best results (and consequently highest resolution), in comparison to plasticized PVC membranes, at O2 concentrations above approximately 25%. Preliminary results have shown that the storing of the membranes at 25°C over 1 month did not change the sensor characteristics significantly. Current investigations regarding the behaviour at 37’C are in progress. Another feature which is of utmost importance for breathgas analysis is the response time of the sensor. In this case the added plasticizer improves the response characteristics. A problem associated with the used set-up is the fact that it is very difficult to quantify the true response time of the oxygen-sensitive membranes. The measured data represent a sum of the response characteristics of the pneumatic set-up, the speed of gas-concentration change and the limited bandwidth of the electronic detection unit. Several parameters influencing the response time can be deduced from Table 1: the type of polymer and plasticizer, the fraction of plasticizers in the polymer and the thickness of the sensitive layer. These factors are also inseparably linked with the required resolution (limited by the noise) of the measuring system (Table 1). The obtained noise depends to a high extent on the excitation light intensity. A further increase of this intensity may cause increasing photobleaching of the dye. This photobleaching is a limiting factor of the stability and calibration requirements of the optochemical oxygen sensor at the moment. With the frequently used system the maximum error due to photobleaching of the dye (PtOEPK PVC:DOS 75:25) within 1 h of continuous measurement is smaller than 0.1 vol.% oxygen. Thus, to meet the requirements of ameasuring stability of &- 0.1 vol.% 02, a recalibration is necessary every hour. An obvious reduction of excitation light intensity again leads to high demands on the quality of signal processing to meet the resolution requirements of an efficient fast oxygen analyser. As a result of all these facts, it can be concluded that it is a difficult task to select an adequate combination of sensor composition and instrumentation set-up.
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The results obtained from the clinical tests are very promising with respect to the agreement between oxygen and carbon dioxide measurements. The rapid rise of CO2 concentration at the beginning of the expiration (Fig. 7)) corresponding to expiration air, seems to be time shifted to the decrease in O2 concentration. This is due to the specific geometric set-up of the detectors. The steeper slope of the carbon dioxide response characteristic is true. However, these initial experiments have been made with relatively slow-responding sensors (poly( styrene) membranes). Further tests with the faster oxygen-sensitive membranes (Table 1) have not yet been performed. Using these membranes, we expect an exact conformity of the time response of COZ and O2 measurements. The experiments only cover the requirements for lungfunction tests in the measurement range 15-21% OZ. For example, for intensive care or anaesthesia, measurements in the full range up to 100% oxygen are necessary and possible cross-sensitivities to anaesthetic agents have to be investigated. Corresponding work relating the choice of a suitable sensor membrane in combination with an optimized electronic set-up for a required measurement range is in progress. With the additional temperature-compensation method suppression of an essential temperature influence could be obtained. However, this method is not yet mature. One of the keys to this problem certainly is to tune the temperature response time of the oxygen sensor membrane with the response time of the utilized temperature sensor in the gas stream. Another approach to compensate temperature effects could be the combination of the optochemical oxygen sensor with a luminescence-based temperature sensor [ 181.
5. Conclusions
and possible improvements
A prototype instrument for respiration gas exchange based on an optochemical oxygen sensor has been developed and verified experimentally. It can be emphasized that this detection principle is a viable alternative to other oxygen-measurement techniques in medical gas-exchange measurements, with the optical sensors offering some essential features: 0 The measurement can be carried out directly in the primary flow. l The sensor is small and cheap. Sterilization problems, which are evident with any medical application, can be prevented by changing the sensor spot after each measurement. 0 The oxygen sensor works in the concentration range O100% oxygen, opening applications in many fields of clinical gas analysis. l In comparison with other methods (e.g., paramagnetic), the sensor is not affected by humidity, nor is it poisoned by CO,. l The sensors require no moving parts. Consequently, they are more stable and cheaper to construct than paramagnetic instruments. It is possible to apply the sensors in vibrating and harsh environments.
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In the future we intend to intensify’our investigations in the selectionof a proper polymer or polymer/plasticizer combination, especially with respect to sensorstability due the photobleaching of the dye. To overcome temperatureproblems,the temperaturecharacteristicsof the sensormembranes have to be investigated in order to find a reliable temperature compensation.Moreover, and looking further ahead,we may expect that there areconsiderableadvantagesin replacingthe intensity-measuring principle with a luminescencelifetimemeasuringtechnique [ 19-2 11.
Acknowledgements The authors wish to thank ProfessorK. Harnoncourt and his colleagues (II. Medizinische Abteilung, LKH Graz) for their assistanceduring the clinical tests.
References
instrum.
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[14] D.B. Papkovsky, G.V. Ponomarev, W. Trettnak and P. O’Leary, Phosphorescent complexes of porphyrin ketones: optical properties and application to oxygen sensing, Anal. Chem., 67 (1995) 41124117. [15] P. Hartmann and W. Trettnak, Effects of polymer matrices on calibration functions of luminescent oxygen sensors based on porphyrin-ketone complexes, Anal. Chem., 68 (1996) in press. [ 161 D.A. Kerth and D.S. Piasecki, An oversampling converter for strain gauge transducers, IEEE J. Solid-State Circuits, 27 (1992) 1689/ 1696. [ 171 M.R. Eftink and C.A. Ghiron, Exposure of tryptophanyl residues in proteins. Quantitative determination by fluorescence quenching studies, Biochem., 15 (1976) 672-680. [ 181 J.N. Demas and B.A. DeGraff, On the design of luminescence based temperature sensors, SPIE Proc., Vol. 1796, 1992, pp.71-75. [ 191 J.R. Lakowicz, Principles of Fluorescence Spectroscopy, Plenum, New York, 1983. [20] M.E. Lippitsch, J. Pusterhofer, M.J.P. Leiner and OS. Wolfbeis, Fibre optic oxygen sensor with the fluorescence decay time as information carrier, Anal. Chim. Actu, 20.5 (1988) 1-6. [21] W.R. Gruber, P. O’Leary and OS. Wolfbeis, Detection offluorescence lifetime based on solid state technology and its application to optical oxygen sensing, SPIE Proc., Vol. 2388,1995, pp. 148-158.
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[6] Fast Oxygen Gas Sensor, data sheet, Oxygraf, Palo Alto, CA, USA, 1994. [7] I. Bergman, Rapid response atmospheric oxygen monitor based on fluorescence quenching, Nature, 218 (1968) 396. [S] 0. Burkhard and W.K.R. Barnikol, Ein schneller O,-Detektor fiir die Analyse der Lungenfunktion, Prax. Klin. Pneumol., 37 (1983) 805807. [9] OX Wolfbeis, H. Offenbacher, H. Kroneis and H. Marsoner, A fast responding fluorescence sensor for oxygen,Mikrochim. Actu, I (1984) 153-15s. [ 10 ] J.R. Bacon and J.N. Demas, Determination of oxygen concentrations by luminescence quenching of apolymer-immobilisedtransition-metal complex, Anal. Chem., 59 (1987) 2780-2785. [ll] W.K.R. Barnikol, Th. Gaertner, N. Weiler and 0. Burkhard, Microdetector for rapid changes of oxygen partial pressure (PO,) during the respiratory cycle in small laboratory animals, Rev. Sci. Instrum., 59 (1988) 1204-1208. I:121 H. Karpf, H.W. Kroneis, H.J. Marsoner, H. Metzler and N. Gravenstein, Fast responding oxygen sensor for respiratorial analysis, Proc. Chemical, Biochemical Environmental Fiber Sensors, Boston, MA, USA, 6-7 Sept. 1989, Vol. 1172, pp. 295-304.
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Christian Kolle wasborn in 1966 in Klagenfurt, Austria. He received the Dip.-Ing. degree in electrical engineering from the Technical University Graz in 1993. He joined the Institute for Chemical and Optical Sensors at Joanneum Research (Graz, Austria) in 1994. He is presently working on the developmentof the instrumentationfor optical oxygen sensors,particularly in the field of breath-gasanalysis. His research interests are in electronics instrumentation and measurement, signal processing and electromagnetic compatibility. Wolfgang Gruber was born in 1961 in Graz, Austria. He received the Dipl-Ing. degree in electrical engineeringfrom the Technical University of Graz, Austria, in 1989. Since 1984 he hasgiven lecturesin electronics at the Wirtschaftsforderungsinstitut in Graz. From 1991 to 1992 he joined the Institute for Optical Sensors,JoanneumResearch,Graz, Austria. Since 1992 he has been with the Institute for Sensor Interfaces and the Institute for Chemical andOptical Sensors, JoanneumResearch,Graz, Austria, where he is working on integrated sensorsystems.He is presently working towards the Ph.D. degreein electrical engineering at the Technical University of Graz, Austria. W&gang Trettnak, born in Graz (Austria) in 1962, received a Ph.D. in chemistry from the Karl-Franzens-University in Graz in 1989. Since 1986 he hasbeen concerned with the development and testing of opto- and electrochemical sensorsand biosensorsfor various parameterssuch as 02, pH, C02, NH3, glucose,lactate and cholesterol. With the exception of a one-year researchproject at the University of Florence in 1992, he hasbeenactive at JoanneumResearch, Graz, Austria, since 1991.
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Karin Biebemik received a Ph.D. degree in chemistry from the Technical University of Graz, Austria. In 1991, she took on a postdoctoral fellowship at the University of Toronto, Canada, where she worked on oxygen reduction electrodes at the Department of Chemistry. In 1995, she joined Joanneum Research at the Institute of Chemical and Optical Sensors, where she is working on the development and characterization of optical sensor membranes. Constanze Dolezal, born in 1960 in Graz, received a degree at the College for Chemotechnical Engineering in Graz, Austria. Since 1994 she has been with the Institute for Chemical and Optical Sensors, Joanneum Research (Graz, Austria), where she is working mainly on the development of optochemical sensors and testing of prototype instruments. Franz Reininger, born in Leibnitz (Austria) in 1961, received a Dipl.-Ing. degree in electrical engineering from the Technical University in Graz in 1989. In 1988 he joined Austria-Mikro-Systeme-International GmbH (Unterprems&ten, Austria) where he did his thesis ‘Evaluation of AD converter and DA converter and dynamic performance’.
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Since 1990 he has been at Joanneum Research in Graz, Austria, where he is involved in designing and building new sensor interfaces for different sensor types, especially for new optical gas sensors. His interests are in dataconversion, signal processing and calibration of special optochemical sensors. In 1993 he became head of the working group of optochemical sensors and in 1995 he became the operative manager of the Institute for Chemical and Optical Sensors at Joanneum Research. Paul O’Leary was born in 1960 in Croom, Co. Limerick, Ireland. He holds a B.A. in mathematics, aB.A.1. in electronic engineering, an M.E.E. in electronic engineering and a Ph.D. in signal processing structures for sensor systems. He founded the Institute for Chemical and Optical Sensors within Joanneum Research in September 1990. Since then he has been responsible for the development and growth of the group. He has 11 years’ professional experience in the development of integrated circuits and complex systems. In 1995 he became a professor and head of the Institute for Automation (University of Leoben, Austria).