Optochemical Sensor Systems for In-Vivo Continuous Monitoring of Blood Gases in Adipose Tissue and in Vital Organs

Merima Čajlaković1, Alessandro Bizzarri1, Gijs H. Goossens2, Igor Knez3, Michael Suppan1, Ismar Ovčina3 and Volker Ribitsch1 1Joanneum Research Forschungsgesellschaft mbH, MATERIALS Institute for Surface Technologies and Photonic, Sensorsystems, Graz, 2Department of Human Biology, NUTRIM School for Nutrition, Toxicology and Metabolism, Maastricht University Medical Centre, Maastricht, 3Department of Cardiac Surgery, University Clinic of Surgery, Medical University Graz, Graz, 2The Netherlands 1,3Austria


Introduction
Sensors capable of determining both oxygen and carbon dioxide partial pressure, as well as pH in blood, are of particular interest with respect to the medical care of critically ill patients.Those parameters are informative indicators of the conditions of a living system and can be valuable in determining the physiologic status of critically ill patients.Tissue oxygenation represents the balance between local oxygen delivery and consumption.Under resting conditions the level varies between and within organs; it is lower in more metabolically active tissues, such as brain (Hou et al., 2003), liver (Vollmar et al., 1992), renal cortex and medulla Whitehouse et al., 2006) and higher in tissues with low metabolic rates, such as the bladder (Rosser et al., 1995(Rosser et al., , 1996)).Changes in blood carbon dioxide levels are associated with myocardial acidosis and alkalosis, ventricular fibrillation and reduced myocardial contractility in addition to providing an indication of cardiac resuscitability (Maldonado et al., 1993;Tang et al., 1996).The respiratory gases themselves reflect cardiopulmonary homeostasis.This is the ability of the cardiopulmonary system to maintain a constant relationship between respiration in the cells of the body, the supply of oxygen from the lungs to the blood and the elimination of carbon dioxide by the lungs.In this way, partial pressures of O 2 and CO 2 are kept at their normal physiological values.Thus, the body constantly consumes oxygen and produces carbon dioxide, and there is a normal balance between the absolute contents and the partial pressures of these gases in the blood (Hahn, 1998).Although this balance can be upset by many clinical or physiological factors, the cardiopulmonary function has a remarkable ability to compensate for these changes, especially to maintain the gases at their normal values.The magnitude of the arterial blood pCO 2 is the key clinical measure of the effectiveness of the lungs to expel CO 2 .Therefore, it is clinically important to measure the inspired and expired partial pressures of both oxygen and carbon dioxide in the critically ill adult and infant patient.It is possible either to take a blood sample and analyse this in the laboratory, what is routinely done in hospitals, or to perform the measurement in vivo.Although a number of methods are used in the clinical setting to measure blood gases, including blood gas analysis and electrodes (Astrup, 1986;Harsten, 1988;Suzuki, 1999), some practical issues restrict their use, e.g. the distance of the analyzer from the bedside, the need for special training of clinicians and staff for proper blood sample collection and handling, potential contamination of blood samples by the biological medias and the sensitivity of the method.Thus, there is huge need and interest to develop more sensitive and practical methodologies such as optical techniques, especially in the field of in vivo intravascular blood-gas analysis (Suzuki, 2001;Tsukada, 2003;Tusa, 2001;Wolfbeis, 1991).In vivo gas sensors should respond rapidly, preferably with a response time of the order of seconds, if the measurements are to have any significant physiological meaning.So far, the only commercially available sensors for the monitoring of adipose tissue oxygen partial pressure pO 2 are needle-type sensors.Those suffer of the typical drawbacks of a point measurement, which is strongly dependent on the position of the sensor tip inside the adipose tissue and show high fragility due to the small dimension of the sensor heads.
Combination of the optical sensor with a microdialysis technique is a powerful tool for online monitoring of metabolism in tissues (Cooney, 1997(Cooney, , 2000;;Ungerstedt, 1991).This is mainly due to the fact that the subcutaneous adipose tissue is easily accessible and the risk of infection is low.The microdialysis probe, which is a hollow fibre which functions as an artificial blood vessel, is inserted into the tissue and is perfused with a dialysis solvent.Exchange of substances between the microdialysis probe and the extracellular fluid (sample medium) occurs over a semi-permeable membrane.Low molecular weight substances in the interstitial fluid surrounding the microdialysis catheter are allowed to diffuse across the membrane and to equilibrate with the perfusion solution.The interstitial fluid leaves the probe via a side arm from which it is collected (Goossens et al., 2011).Optical sensing approach combined with microdialysis have been already used for monitoring of dissolved gases and pH in a flow loop (Cooney, 1997(Cooney, , 2000)).Pasic et al. (Pasic et al., 2006) have described a microdialysis-based glucose-sensing system with an integrated fibre-optic hybrid sensor.An important advantage of this system is that for example adipose tissue pO 2 is measured over a much larger area compared to commercially available needle-type oxygen sensors.
In our first work we reported on optical capillary sensors for continuous measurement of oxygen and carbon dioxide in subcutaneous tissue using microdialysis for extraction of interstitial fluid in subcutaneous tissue (Bizzarri, 2006;Cajlakovic, 2006).The adopted sensing scheme for both sensors was phase fluorometry, which is based on the dynamic quenching due to oxygen of the photo-excited luminophore Platin(II) mesotetra(pentafluorophenyl) porphyrin (Pt-TFPP).In the case of carbon dioxide sensor an ionpairing approach combined with dual luminophore referencing as a detection technique was used for the preparation of carbon dioxide sensitive layers.The analysis of extracted sample by means of the external instrumentation is properly designed for rapid analysis of small volumes (less than microliters) and allowed a less invasive technique than the intravascular approach (Bizzarri, 2006;Cajlakovic, 2006).Recently, we have presented an improved optochemical measurement system for continuous monitoring of adipose tissue oxygen tension in biological fluid using a miniaturized flow-through cell instead of a capillary-based sensor coupled to a microdialysis catheter (Bizzarri, 2007;Cajlakovic, 2009).The effectivness and accuracy of the sensors were investigated in vitro using physiological solution (ringer) and in vivo tests in order to examine the magnitude of changes in tissue gas tension in response to evolving shock and resuscitation: 1) on test animals, where hemorrhagic schock model has been adopted to simulate critically condition and 2) on healthy volunteers, where infusion of adrenaline caused the changes in tissue oxygenation.Adipose tissue oxygen tension may be involved in the development of insulin resistance (impaired action of the hormone insulin) in humans.Therefore the continuous monitoring of adipose tissue oxygen tension (pO 2 ) in individuals affected by obesity and/or type 2 diabetes mellitus is highly interesting.Recent cell culture and preclinical experiments suggest that obesity is accompanied by low levels of oxygen in adipose tissue (Ye et al., 2007).However, this concept was recently challenged by Goossens et al. (Goossens et al., 2011), demonstrating increased adipose tissue PO2 in obese insulin resistant compared to lean insulin sensitive men.The optochemical sensor technology combined with microdialysis is particularly suitable for the development of minimally invasive instrumentation for measurement of tissue pO 2 (Goossens et al., 2011).In the present chapter, we present an improved microdialysis-based oxygen-sensing instrumentation with integrated planar sensitive membrane inserted in a miniaturised flow-through-cell (volume app.5μl) for the continuous monitoring of adipose tissue oxygen tension in humans.This instrumentation is suitable for measurement at a flow rate as low as 2 µl/min with a response time, including the flow time from the catheter to the sensitive membrane, of about 10 min.The sensor system has been validated in in vivo experiments, in which adipose tissue oxygen tension in humans has been monitored under resting conditions and after the local administration of a vasoconstrictor and vasodilator (Goossens et al., 2011).As expected, local administration of the vasoconstrictor in adipose tissue decreased both adipose tissue blood flow and adipose tissue pO 2 .Subsequently, local infusion of the vasodilator increased both adipose tissue blood flow and pO 2 values (Goossens et al., 2011).Obtained results showed good agreement between the observed adipose tissue oxygen tension and the expected behavior due to the induced alteration of adipose tissue blood flow.Thus, this system is suitable for in vivo monitoring of pO 2 in adipose tissue.
The second part of this contribution includes the description of the optochemical measurement system for simultaneously monitoring of tissue O 2 and CO 2 partial pressure in brain, heart and liver by means of miniaturized implanted fiber optical sensor.pO 2 and pCO 2 measurements are based on the determination of the luminescence lifetime via phase modulation fluorometry.The instrumentation unit consists of two independent channels for measuring pO 2 and pCO 2 .The O 2 sensor is based on the luminescent dye Ptpentafluorophenyl-porphyrin, while for CO 2 sensor the Resonant Energy Transfer (RET) scheme is used to convert the intensity of the analyte dependent dye indicator into long luminescence lifetime information, which can be measured by phase modulation fluorometry (Lakowicz, 1983).Therewith it was for the first time possible to measure continuously and simultaneously O 2 and CO 2 content in brain, heart and liver during openheart surgery.Fiber optical sensors have been tested in an experimental animal model during cardiac surgery using two different methods as extra corporal circulation: conventional cardiopulmonary bypass (CPB) and minimal extracorporeal circulation systems (mini-EEC).CPB is the most accepted surgical technique to treat cardiac pathologies.It is a safe and established technique with a low mortality rate (Pentila et al., 2001).However, besides neurological events, which may at least be partially related to the use of CPB, and the inflammatory response, the ideal myocardial protection is still a matter of scientific and clinical debate (Chang, 2002;Immer, 2005;Levy, 2003).Mini-ECC system consists of a centrifugal pump, an oxygenator and a modified suction system.The main advantage of mini-ECC is the reduction of tubing length (reduction of the priming volume).It was designed to reduce blood tubing, the contact between blood and air and priming volume during cardiac surgery.The main issue of developing miniaturised extracorporeal circulation systems is the strong need for a CPB concept for cardiac surgery which will cause fewer traumas to blood components and reduce the extensive volume shifts and bubble formation during cardiopulmonary bypass (Kofidis et al., 2008).A closed circuit, where the patient himself/herself serves as a reservoir with a minimized priming volume, could be the ideal alternative to reduce the adverse effects of CPB (Kamiya et al., 2006).In a complex, technically demanding, prospective, randomized and blinded experimental setting, an extracorporeal circulation has been performed using mini-EEC versus standard CPB techniques with main emphasis on differences in organ specific oxygen and carbon dioxide partial pressures and metabolism.In-vivo experiments on animals with implanted optical O 2 and CO 2 sensors were performed in order to be able to establish if there is any difference in O 2 /CO 2 metabolism when using two types of extracorporeal circulation, CPB and mECC, respectively.

The continuous monitoring of adipose tissue oxygen tension (pO 2 ) in humans 2.1 Design of an optical oxygen sensor
The luminophore Platin(II) meso-tetra(pentafluorophenyl) porphyrin (Pt-TFPP) was used because of its long lifetime (∼ 50 ms), strong absorption in green region of the spectrum, which is compatible with green light-emitting diodes (LEDs), its highly emissive metal-toligand charge transfer state and its high photostability.The dye was entrapped in a hydrophobic polymer of polystyrene (4 w. % dye/g solid polystyrene in chloroform).Afterwards the sensor cocktail was spread onto a transparent substrate; either a polyethylene terephthalate foils PET (Mylar foil with thickness 100 µm) or PET foils (50 µm thickness) coated with Al 2 O 3 .Knife coating and screen printing techniques have been used for production.After coating, the oxygen sensitive membranes have been dried at 60 °C for eight hours.In the case of knife coating a successive mechanical reduction of the sensitive spot to a diameter of 2 mm has been applied.This procedure was not necessary in case of the screen printed sensitive membranes.

Measurement principle
The luminescence lifetime of a luminophore is an intrinsic property and is independent of for example detector sensitivity and fluctuations in light intensity therefore phase fluorometry was used as sensing scheme.In this case the luminophore is excited by sinusoidally modulated light at a given frequency ω and the emitted sinusoidal luminescence is also modulated at the same frequency.The lifetime of the luminophore causes a time lag between the absorbance and emission, expressed by the phase angle θ and a decreased intensity relative to the incident light, called demodulation.The relationship

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between the lifetime τ and the phase shift φ for a single exponential decay is given by equation ( 1) and is illustrated in Fig. 1: where f is the modulation frequency.An optimal frequency need to be selected because signal-to-noise (SNR) ratio decreases and phase sensitivity increases with increasing modulation frequency.The used luminophore was Pt-TFPP, whose luminescence is dynamic quenched (reduced) in the presence of oxygen and subsequently the lifetime as well.The optimal frequency for this dye was chosen 4.5 kHz.The instrumentation measures the phase shift of the luminescence with respect to the excitation signal.This is related to the luminescence lifetime of the dye, which in turn depends on the pO 2 of the tissue in contact with the surface of the sensor probe.The pO 2 can be calculated from the measured phase shift of the luminescence using the so called false-light model (Bizzarri et al., 2006) as described by equation ( 2): where φ is the phase shift measured at a given oxygen partial pressure pO 2 , φ0 is the phase shift in absence of O 2 , K is the quenching parameter and f 0 is the false-light parameter (the fraction of the luminescence emitted by the sensor, which is quenched by oxygen).

Miniaturised flow-through cell and optoelectronic measurement unit
The fibre optic measurement system consists of an optoelectronic read-out module for measurements with optical fibres (core diameter 1000 µm) and of a miniaturised flow-through cell (volume app. 5 µl), where a planar oxygen sensitive membrane (diameter of the sensitive spot 2 mm) is allocated.The measurement set-up is shown in Fig. 2. The flow through cell is equipped with inlet tubing of inner diameter of 0.15 mm and outer diameter of 0.30 mm (Rometch GmbH).The outlet tubing and the flow chamber have an inner diameter of 2.0 mm.The outer diameter of the outlet tubing is 3.0 mm as presented in Fig. 3.The inner diameter of the inlet was chosen to be 0.15 mm in order to match the inner diameter of the inlet of the microfluidic catheter.This geometry has been chosen to reduce the risk of overpressure during measurement, which would cause an overestimation of the dissolved pO 2 .The length of the inlet is 20 mm whereas the outlet is 32 mm long.The tubings are made of stainless steel in order to prevent or to reduce the permeation of ambient oxygen to non-significant values.For the extraction of interstitial fluid microdialysis catheter CMA60 was used.The outlet tubing of catheter was first cut as close as possible to the microdialysis membrane (at a distance of about 2 mm) and then connected with the flow through cell by means of steel tubing (diameters 0.15/0.30mm).The length of the steel tubing was standardised at 90 mm in each experiment.In this way the ambient oxygen diffusion through the permeable tubing materials of the catheter can be neglected.This has been proven in oxygen recovery rate measurements, which have been carried out in laboratory.These measurements have been already exhaustively described in previously reported work (Cajlakovic et al., 2009).The standardization of the length of all the connection tubing allows that the flow time from the catheter to the sensitive membrane is in all experiments the same.

Sensor performance and laboratory calibration protocol
The membranes have been produced by knife coating technique onto PET substrate and by means of screen printing onto PET substrate coated with Al 2 O 3 as oxygen barrier.
All sensor membranes have been firstly laboratory characterized in Ringer solution (physiological solution used as perfusate in many clinical procedures) at temperature held constant at 37 °C.The temperature regulation was realised by using a water bath thermostat (Lauda RC 6 CP).The measurement medium has been equilibrated with a gas mixture of nitrogen and oxygen at defined concentrations (typically 0, 5, 10 and 21 % O 2 ) using a mass flow controller (Bronkhorst, The Netherlands).A paramagnetic gas purity analyser (Servomex, U.K.) was served as reference instrumentation.The solution was then pumped by means of a peristaltic pump into a miniaturised flow-through cell at the flow rate of 2.0 µl/min (Minipuls 3, Gilson, Germany).One example of a typical sensor response curve, phase angle as a function of different O 2 concentrations, is shown in Fig. 5. Two cycles have been performed, whereas the first complete cycle of dissolved O 2 was used for sensor calibration purpose, and the following cycle is used to check the specifications of the oxygen measurement (Table 1).Both types of membranes perform satisfactorily with the same measurement specifications.

Delay time
The time needed for the extracted interstitial fluid to reach the sensitive oxygen membrane in the flow-through cell is indicated as delay time.This delay time is defined as the time required for the sensor output to change from its previous state to a final settled value within a tolerance band of the new value.It is a common practice to measure the delay time as the time needed for the sensor to reach the 90 % of the final output.This is usually indicated as the t 90 of the sensor.Fig. 6 shows the delay time t90 for an oxygen change from air saturated solution (circa 21 % O 2 ) to oxygen free solution (about 0 %).The time needed to achieve the 90% of the final value as been recorded.Six different membranes of both types have been investigated and the averaged t 90 could be estimated as 11.7 ± 4.6 min.Even if, the two different types of sensitive membranes (based on uncoated or Al 2 O 3 coated PET substrate) perform satisfactorily with regard to the measurement specifications, it could be observed that, in single membranes based on the uncoated PET substrate, the response time increased above the specified time of 10 min, due to ambient oxygen diffusion through microscopic damages of the coated surface, caused by the mechanical reduction of the sensitive spot.

Validation protocol
To validate the sensor instrumentation for in vivo use in humans, a healthy, normal-weight male subject participated in an experiment carried out at Maastricht University Medical Centre in The Netherlands.The subject was asked to refrain from drinking alcohol and to perform no strenuous exercise for a period of 48 hours before the in vivo experiment.On arrival, a microdialysis catheter (CMA 60, CMA microdialysis AB, Stockholm, Sweden) was inserted in the abdominal subcutaneous adipose tissue 6 cm right from the umbilicus after the skin was anesthetized by a lidocaine/prilocaine cream (EMLA, AstraZeneca BV, Zoetermeer, The Netherlands).After insertion, the probe was perfused with Ringer solution (Baxter BV, Utrecht, The Netherlands), supplemented with 50 mM ethanol, at a flow rate of 2.0 µl/min (CMA400 micro infusion pump, CMA microdialysis AB, Stockholm, Sweden).Afterwards, the system was allowed to equilibrate.After pO 2 had reached stable values, the probe was consecutively perfused with a vasoconstrictor and vasodilator, each for 60 min.
There was a wash-out period of 120 min (Ringer infusion) before the infusion of the vasodilator was started.

Results and discussion of in vivo measurement of adipose tissue oxygen tension in humans
A stable baseline pO 2 of 52.5 mmHg was reached after 2-3 hours (Fig. 7, time -30 to 0 min).
As hypothesized, local administration of a vasoconstrictor in adipose tissue (t0-60 min) decreased both adipose tissue blood flow and adipose tissue pO 2 (37.3 mmHg, Figure 7).
During the wash-out period (t60-180min) the probe was perfused with Ringer solution, and pO 2 values returned to near-baseline values again (53.5 mmHg).Subsequently, local infusion of a vasodilator increased both adipose tissue blood flow and pO 2 values to 74.8 mmHg (t180-240min).This experiment clearly showed that pO 2 values can be reliably measured in vivo in humans, and can be modified by changes in tissue blood flow.These results have been extensively described elsewhere (Goossens et al., 2011).
The rapid increase of the measured pO 2 at the beginning of each drug administration (spikes in Fig. 7) is caused by ambient oxygen, which has been introduced in the fluidic during the change of the infusion solution.
After the in vivo experiments several oxygen sensitive membranes were recalibrated in the laboratory, and the deviations of the pO 2 values obtained by the second calibration were calculated and plotted against the actual pO 2 values (Fig. 8).Importantly, the change in the calibration of the sensitive membrane after the in vivo experiment causes very small discrepancy (<0.5 mmHg) from the actual pO 2 values.This result is excellent considering that the oxygen sensitive membranes have been recalibrated about one year after the first laboratory calibration and after in vivo experiments of a typical duration of 5 -8 hours.

Conclusion
We have developed and validated a microdialysis-based oxygen-sensing instrumentation with integrated miniaturized sensitive membrane for the continuous monitoring of adipose tissue oxygen tension in humans.In comparison to previously reported instrumentation (Cajlakovic et al., 2009), this system can be used at a flow rate as low as 2 µl/min with a response time of about 10 min (including the flow time from the catheter to the sensitive membrane).The production technology of the oxygen sensitive membranes has been improved, allowing the pO 2 measurement in the range between 0 and 300 mmHg with accuracy better than ±1 mmHg and resolution better than 1 mmHg.The system has been validated in vivo, where adipose tissue oxygen tension has been monitored in humans under resting conditions and during local administration of vasoactive compounds.Local administration of the vasoconstrictor in adipose tissue decreased adipose tissue pO 2 and subsequently, local infusion of the vasodilator increased pO 2 in adipose tissue (Goossens et al., 2011).The agreement between the observed adipose tissue oxygen tension and the expected behaviour after modulation of adipose tissue blood flow allows concluding that the system is suitable for in vivo monitoring of pO 2 in humans.Several sensitive membranes were recalibrated after in vivo experiments in the laboratory and deviations smaller than ± 0.5 mmHg were observed.This indicates that the membranes have an excellent storage and operating stability, and that contact with biological fluid does not affect the read-out of the optochemical measurement instrumentation.

Acknowledgement
The in vivo experiments in humans were carried out at Maastricht University Medical Centre, Maastricht, The Netherlands, under supervision of Dr. G. H. Goossens.

Optochemical measurement system for simultaneously monitoring of O 2 and CO 2 partial pressure in vital organs by means of miniaturized implanted fibre optical sensor
The measurement system is composed of two channels allowing simultaneously measuring of O 2 and CO 2 partial pressures (pO 2 and pCO 2 ).The optoelectronic measurement unit for O 2 is same as described in chapter 2 (Paragraph 2.3).For this application optical fibres were coated with sensitive layers (O 2 and CO 2 ).In the case of using an optical fibre as support for the sensitive layer, pulses of light emitted by a green (in the case of O 2 sensor) or blue LED (in the case of CO 2 sensor) are carried via an optical fiber to excite the dye at the sensor tip.
On interaction with measured parameter the luminescence is quenched and is transmitted back within the same fibre to the detection unit.The instrumentation measures the phase shift of the luminescence with respect to the excitation signal.

Optoelectronic measurement unit for CO 2
As in the case of the O 2 measurement channel a measurement system for CO 2 is based on a collimating optics to convert the light of excitation source (Nichia NSPB500) into a parallel beam, which can be focused by a GRIN lens into the optic fibre.The luminescence collected by the optic fibre is also transformed by the same GRIN lens into a parallel beam, which is deflected by a dichroic filter at 45 ° (Unaxis B51) with respect to the optical axis to the silicon photodiode as a photo-detector (Hamamatsu S5106).Dichroic filters transmit selected ranges of the visible spectrum and reflect the complementary ranges, thus they may be used to separate the radiation emitted by the excitation source from the luminescence.Additional glass optical filters (SCHOTT BG12, thickness 2 mm for the excitation source and SCHOTT KV550, thickness 3 mm as emission filter) are used to separate the luminescence from stray light and optical background due to the excitation source.For the compensation of the reading of the measurement unit for changes of the electronic parameters of the components with fluctuations of the operating conditions (e.g.forward current through the LED and temperature) a reference light source is used (HLMA-KL00-I0000 (Avago Technologies US Inc.).The excitation light of the LED was sine wave modulated at a frequency f of 20 kHz.

Optical fiber tip micro fabrication by chemical etching
The most efficient configuration of optical fibre for collecting the luminescence emission and to reduce the optical background was achieved by coating the sensitive layers onto the sharpened tip of optic cables with core diameter of 400 and 200 µm.For this purpose the fiber tips have been micro-machined by chemically etching process (tube etching) of multimode optical fibers (Si core and hard polymer cladding, Thorlabs).The schematic illustration of the forming of a sharpened fiber tip is presented in Fig. 9.After the etching process was completed, the tip of the optical fiber was rinsed with distilled water and cleaned with acetone.The height of the obtained cone tip is about 1.5 mm.All the etching experiments were carried out at room temperature.The obtained form of an optical fibre efficiently resulted in improved signal-to-noise ratio of about one order of magnitude for both types of sensors thus resulted in an increased luminescence signal.

Preparation of O 2 fiber optical sensor
The sensitive cocktail is composed of Pt-TFPP dissolved in polystyrene (4 w. % dye/g solid polystyrene in chloroform).The optical fibres were coated by dipping only the sharpened tip of the fibre into a sensitive cocktail.After coating, the oxygen sensors have been dried at 60 °C for eight hours.

Preparation of CO 2 sensing cocktail
The carbon dioxide sensitive cocktails were prepared by mixing donor and acceptor dye, along with a phase transfer agent, tetraocytlammonium hydroxide (TOAOH) in ethylcellulose solution and coating a sensitive thin film onto the tip of etched optical fiber.
The polymer solution of ethyl cellulose (EC) was prepared by dissolving 4.4 g of EC in the 40 g of a toluene/ethanol mixture (80/20, v/v), giving a viscous 11 % (w/w) solution.Ru(dpp) nanobeads incorporated in oxygen impermeable polymer, polyacrilonitrile (PAN) have been suspended in polymer (EC) solution.Acceptor was prepared by dissolving 2.03 mg of phenol red (PR) in 1.5 ml of toluene/ethanol mixture to which 20 µl of TOA-OH (20 % solution in methanol) solution was added.The rate of energy transfer and therefore the sensitivity of the sensor are almost independent of the concentration of the donor in the matrix.However, an increase in donor concentration enhances the luminescence intensity and therefore the signal-to-noise ratio.In contrary, the concentration of the deprotonated form of acceptor is a decisive factor upon which the rate of energy transfer strongly depends since the overlap integral is proportional to the concentration of the acceptor in Förster equation.Owing to this, the signal change depends not only on the degree of protonation by carbon dioxide but also on the overall concentration of the acceptor in the sensor membrane.Therefore, the final cocktail composition consisted of 13.8 mmol Ru(dpp) and 3.08 mmol PR per kg of solid polymer.The cocktails were saturated with carbon dioxide before coating of the fiber tip.The resulting fiber optical sensors were dried for 4 hours at 70 °C before an additional ion impermeable layer of Teflon with added black graphite as an optical and protective isolation was applied.The dried sensors were stored in desiccator over sodium carbonate before use.Before the calibration was proceeded the optical fibers have been exposed to solution equilibrated with CO 2 (20 %) and N 2 in order to obtain faster stable response of the sensor.

Measuring principle of CO 2 optical sensor
For the realisation of an optical lifetime-based CO 2 sensor, whose luminescence and lifetime are modulated by pH dependent efficiency of resonance energy transfer, some criteria need to be considered at selection suitable donor-acceptor pairs: (a) a pH-sensitive indicator is combined with long-lived transition metal complex; (b) an absorbance spectrum of deprotonated form of analyte dependent pH-indicator has to overlap with the luminescence spectrum of the inert luminophore, added as a donor so that efficient energy transfer takes a place; (c) the distance between donor and acceptor does not exceed Förster distance (in the range 1-10 nm), (d) lifetime of the donor in microsecond range (e) high quantum yield of the donor and (f) excitation of the acceptor and donor at a single band of wavelength and detection of donor luminescence at a common wavelength using a single photodetector.In order to convert a change in absorption spectrum of acceptor in presence of CO 2 into lifetime information, a long-lifetime luminescent donor need to be co-immobilised in the same matrix.Depending on the pH, the acceptor changes its colour and deactivates the excited state of the donor.This effect changes both the luminescence intensity and the lifetime of the donor.Ru(II) diphenyl-phenanthroline (Ru(dpp)) was selected as a donor due to the high quantum yield (0.6) and long lifetime in range of microseconds (∼ 5.5 µs).Its absorption spectrum perfectly overlaps the emission of bright blue LEDs and its red emission spectrum overlaps the absorption of the deprotonated form of pH-indicators in red region of spectrum.Among the investigated pH indicators the main problem was their poor photostability (for example for the case of cresol purple and bromothymol blue) and due to the pronounced photobleaching these indicators turned out to be not suitable for the fiber optic measurement system.Even the strategy of optimising the ratio between the concentrations of pH indicator in comparison to the concentration of the donor did not allow realising a sensor with satisfactorily stability properties.Therefore on the basis of enough overlapping of the absorption spectrum of deprotonated form of phenol red (i.e. in absence of CO 2 ) with luminescence of Ru(dpp) phenol red was chosen as an acceptor (Fig. 10) .This dye provides enough intensity when the energy transfer with Ru(dpp) takes place, has pK a value in the physiological range (6.8) and expand the measurement range up to 20 % of CO 2 (corresponding to 147 mmHg).(Mills et al., 1992).To stabilize the deprotonated form of the indicator dye in the matrix and to form a lipophilic bicarbonate buffer system in the polymer, tetraoctyl ammonium hydroxide was used.The excess amount of TOA-OH is neutralised by atmospheric carbon dioxide to form a lipophilic hydrogen carbonate buffer in the solid matrix, and this can tune the sensitivity and improve the storage stability of the sensor.The sensing mechanism of an acceptor-Q + OH ion-pair is described by (3): ( ) where [TOA + A -. xH 2 O] is the complex of the deprotonated pH indicator dye ion paired with the quaternery ammonium cation TOA + and a number x means the hydrophilic sites present in the polymer matrix.
In the case of Resonance Energy Transfer (RET) as sensing scheme the average lifetime is converted into a parameter (phase angle) that reflects the concentration of deprotonated pHindicator (acceptor).The CO 2 dependent lifetime information is detected using the technique of phase fluorometry.In this technique, the lifetime data (τ) are converted to phase shift data (φ) as it was described by equation ( 1).
Efficiency of energy transfer E is expressed as: where k RET is rate of resonance energy transfer and k other is the sum of the rates of all other deactivation processes including luminescence.τ max reflects the lifetime of donor (complete protonation of acceptor at very high pCO 2 , when no energy transfer occurs) and τ is a lifetime in the absence of carbon dioxide.By rearrangement of the equation 4 the ratio τ max /τ can be expressed as: From Förster theory it follows, that the rate of energy transfer k T is proportional to the spectral overlap integral J between donor emission and acceptor absorption (k T ∝ J).When the pH-indicator is completely completely protonated, there is no spectral overlap with the donor and overlap integral J is proportional to concentration of deprotonated dye D -: This model corresponds to the situation of dynamic quenching of luminescence described by the Stern-Volmer equation.Therefore pCO 2 was calculated from the measured phase shift of the luminescence using as described by equation ( 7):

Laboratory calibration protocol
Before of the in vivo measurements, the fiber optical pO 2 and pCO 2 sensors have been calibrated in laboratory.In the case of pO 2 sensors, a glass vessel filled with Ringer solution and kept in a water bath (Lauda RC 6 CP thermostat) at the constant temperature of 37 °C has been equilibrated with a gas mixture of nitrogen/oxygen at defined concentrations (typically 0, 5, 10 and 21 % O 2 ) using a mass flow controller (Bronkhorst, The Netherlands).Similarly, in the case of the pCO 2 sensors the Ringer solution has been equilibrated with a gas mixture of nitrogen/carbon dioxide at defined concentrations (typically 0, 5, 10, 15 and 20 % CO 2 ).A gas purity analyzer (Servomex) with both a paramagnetic O 2 sensor and a CO 2 infrared sensor has been used as reference to analyse the composition of the gas mixture dissolved in the Ringer solution.The fiber optic sensors have been put in contact with the Ringer solution and the phase shift at each defined gas concentration has been recorded.An example of the phase shift response at exposed gas concentration is shown in Fig. 11 for pO 2 sensors and in Fig. 12 for pCO 2 sensors, respectively.It can be seen that for both sensors usually two cycles of dissolved O 2 and CO 2 concentrations have been measured.The first cycle has been used for sensor calibration purpose, whereas the following cycle is used to check the specifications of the oxygen and carbon dioxide measurements calculated with calibration parameters obtained from equations 2 and 7, respectively.

In vivo pO 2 and pCO 2 measurements on animal by use of fiber optical sensors
In vivo experiments on test animals have been performed at the Medical Univeristy Graz, Clinical Department of Heart Surgery.Fourteen pigs (30.7 + 2.5 kg) were investigated by monitoring the pO 2 and pCO 2 content in tissue of the three vital organs: brain, heart and liver in order to examine the oxygen delivery to the tissue of these organs during simulated heart surgery operations.Heart surgery intervention is necessary to heal and treat cardiac defect or valvular heart diseases.During this kind of surgeries an extracorporeal blood circulation can be carried out using the so-called "heart lung machine" (HLM) which temporarily takes over the functions of the heart and lungs during surgeries, maintaining the circulation of blood in the body.A heart lung machine maintains perfusion to other body organs and tissues while the surgeon works in a bloodless surgical field.
The aim of test on animals was the simultaneously evaluation of pO 2 /pCO 2 exchanges in three vital organs: brain, heart and liver using two different methods of extra corporal circulation: conventionally cardiopulmonary bypass (CPB) and Minimal Extra Corporal Circulation (MECC) (see Fig. 13).The questions that appeared were: are the optical fiber sensors suitable for the evaluation of metabolic variables and organ specific O 2 /CO 2 exchanges during heart surgery and furthermore, is it on the basis of the obtained results capable to estimate difference between CPB and MECC?  15 min after end of using of HLM T 10 30 min after end of HLM (end of experiment) Table 2. Protocol of in vivo experiments

Results of in vivo experiments
Since the brain tissue perfusion is one of the most important substrate for energy metabolism delivered by the blood, an example of pO 2 and pCO 2 profiles measured by fiber optical sensor in this organ are shown in Fig. 14.Despite the very low pO2 values at the beginning of the measurement, a slightly decrease in pO 2 value during HLM can be observed, followed by an increase up to 55 mmHg during recirculation.On the contrary, pCO 2 increases up to 40 mmHg during HLM and after recirculation reaches the same value as at the beginning (10 mmHg).These measurements confirmed a heterogeneous distribution of local pO 2 and pCO 2 values in brain.The tissue concentration is determined by the balance between supply and utilisation and the anatomical distribution of concentration and within the brain was reported to be wide, depending on the depth of penetration of the sensor into the tissue (Luebbers & Baumgaertl, 1997).The pO 2 and pCO 2 responses as shown in Fig. 14 have an opposite trend which is in agreement with medical point of view that anaerobic condition of tissue correlate with increase of CO 2 content.
The supply of brain with oxygen depends on the cerebral perfusion and on difference between arterial and venous O 2 partial pressures.CO 2 partial pressure has a significant influence on the blood circulation in the brain.When the arterial CO 2 partial pressure p a CO 2 increases, pCO 2 in brain increases proportional with this increment.If on the contrary Hyperventilation causes a decrease of pCO 2 down to 20 mmHg, than a cerebral Hypoxia and increased cerebral Lactacidose is provoked due to the vasoconstriction.
Despite the fact that brain are very active organ with higher oxygen and energy demand, low pO 2 values below 10 mmHg has been often measured.Such sensor behavior could be obtained due to the influence of the narcosis on the brain activity and due to the position of the sensor tip in brain.Latter can result in holes by implantation of the optical sensor probes into brain region since the brain consists of the tissue and the hollow spaces.As a consequence of the positioning of the optical fiber in hollow spaces, lower pO 2 values can be obtained during measurements in brain.
Fig. 15 and 16 show oxygen and carbon dioxide profile in liver and heart, respectively.In vivo measurements in liver reflected lower pO 2 values before HLM was put into an operation (pO 2 values app. 5 mmHg) and afterwards a decrease of pO 2 vales in the range of 1.5 mmHg.After corporal recovery, the pO 2 values reach the value at the beginning of the measurement.Similarly, pCO 2 values in liver are at beginning of 8 -10 mmHg, and during HLM operation an increase in pCO 2 has been noticed (app.13 mmHg).After recirculation pCO 2 reaches the values similar at the beginning of the experiment (8 mm Hg).
Fig. 14.In vivo measurement in brain obtained with inserted pO 2 and pCO 2 fiber optical micro sensors: red line indicates the start of functionality of HLM, while green line indicates the return to the normal intra-corporeal circulation.The x scale shows the real time of the measurement, so it can be seen that the experiment has started at about 9:30 am and ended few minutes after 13:00 for an overall duration of less than 4 hours.The above Figures indicate on the similar trends of pO 2 and pCO 2 curves for both extracorporeal circulation systems.It is shown that under standardised conditions pO 2 curves in brain were absolutely higher in the case of mini-ECC.In the case of liver it was found that under standardised conditions intrahepatic parenchymal pCO 2 curves were absolutely higher and they are significantly correlated to arterial and central venous lactate in CPB.

Conclusions
In the presented work a novel combined optochemical measurement system which allows a simultaneously measurement of dissolved oxygen and carbon dioxide in tissue of vital organs is described.The tissue oxygenation is frequently monitored with electrochemical micro sensors (e.g.polarographic needle electrodes) which one limitation is that due to the oxygen consumption during pO 2 measurements, the probe needs to be moved through the tissue (Lowry & Fillenz, 2001;Luebbers, 1996, Nair, 1987).On the contrary fiber optic probe can be left in the same position in the tissue for a prolonged period of time because is noninvasive.This makes it possible to directly temporal changes in pO 2 as well as the oxygenation-modifying interventions over time in a tissue.Furthermore, optochemical sensors based on luminescence intensity show higher sensitivity and the use of phase modulation fluorometry as measuring techniques for both oxygen and carbon dioxide sensors enable the realisation of optical sensors with excellent characteristics of reliability and measurement stability possible superior to those of electrodes.Further advantage of optical sensors over the electrodes is ease of miniaturisation.In presented work tapered optical fiber tips obtained through chemically etching improved the signal-to-noise ratio of about one order of magnitude.Based on this fact further miniaturisation can be achieved, using a fiber optical cables with smaller diameter (e.g.50 µm).This type of the optical sensors also offers a possibility to use them for endoscopy.Furthermore, the described optical sensors are almost calibration free using only one point recalibration.
Despite the fact, that the last years are witnessed to an increased number of optochemical oxygen sensors based on phase modulation fluorometry in field of medical applications, especially for analysing the spatial heterogeneity in oxygen tension in tumor tissue using optical oxygen probes OxyLite (Oxford Optronix, Oxford, UK), at present and up to our knowledge there has been no reports about fiber optical CO 2 sensors for measuring pCO 2 content in tissue of organs.
The described optochemical instrumentation has been used in in vivo tests for simultaneously detection of oxygen and carbon dioxide exchange in brain, heart and liver using two different methods as extra corporal circulation: Minimal Extra corporal Circulation (MECC) and Cardiopulmonary bypass (CPB).It has been turned out that a significantly higher demand of blood transfusion is required and higher lactate values (correlated to higher pCO 2 values) are observed when using classical method CPB.The continuous acquisition of metabolic parameters (pO 2 and pCO 2 ) has showed that in the case of using MECC as extra corporal circulation, significantly higher pO 2 values are obtained, while the use CPB resulted in significantly higher pCO 2 in liver and heart tissue.
It is demonstrated that with a help of miniaturised non-invasive fiber optical sensors for continuously measurement of tissue pO 2 and pCO 2 a significant better O 2 /CO 2 metabolism is achieved in the case of MECC as extra corporal circulation.Furthermore, MECC produced less lactate and showed significantly favourable results concerning O 2 /CO 2 metabolism and transfusion demand.

Acknowledgement
The research work was part of the Zukunftsfonds funded by the Styria Country Government A3 Science and Research, Reference A3-22M1/206-4.The aim of this project was to develop a combine optochemical measurement system for determination of oxygen and carbon dioxide for medical applications.
The tests on animals have been made in cooperation with Univ.Prof. Dr. Igor Knez and Dr. med.univ.Elisabeth Beran at the Medical Univeristy of Graz, Clinical Department for Heart Surgery (Austria).

Fig. 1 .
Fig. 1.Principle of phase fluorimetric technique.Luminescence emission is delayed in phase angle Φ and demodulated relative to the excitation (m).

Fig. 2 .
Fig. 2. Measurement system for planar sensing membranes.The flow-through cell has a size of 20 x 13 x 4.8 mm (length x width x heigth).The measuring chamber of the flow-throughcell has a diameter of circa 1.8 mm and a height of about 2.8 mm, thus resulting in a volume of about 5 µl.The sensing membranes have a round sensitive spot of 1.8 mm in diameter.

Fig. 3 .
Fig. 3. Schematic draw of the flow through cell

Fig. 5 .
Fig. 5. Example of the experimental protocol used for the characterisation of the miniaturised planar O 2 sensitive membrane.The measurement is carried out in Ringer solution equilibrated at defined O 2 concentrations (0, 5, 10 and 21 % O 2 ) at the constant temperature of 37 °C.

Fig. 6 .
Fig. 6.Measured delay time between 21 and 0 % of O 2 in case of one oxygen sensitive membrane To perform this measurement, two different glass vessels filled with Ringer solution has been equilibrated with air and nitrogen.The catheter CMA 60 connected by steel tubing to the flow through cell has been first immersed in the vessel containing air saturated Ringer solution and the peristaltic pump was started with flow rate of 2 µl/min.After a stable value of the oxygen sensor instrumentation has been achieved, the pump has been shortly stopped to change the catheter in the oxygen free Ringer solution and the flow started again.The time needed to achieve the 90% of the final value as been recorded.Six different

Fig. 7 .
Fig. 7. Measured abdominal adipose tissue pO 2 with optochemical sensitive membrane under stable baseline conditions (local Ringer administration up to time 0), during local administration (2 µl/min) of a vasoconstrictor between 0 and 60 min, followed by wash-out period with Ringer infusion (from 60 to 180 min).At 180 min an infusion of a vasodilator was started up to 240 min.The experiment was carried out in the laboratory of Dr. G.H. Goossens (Maastricht University Medical Centre, The Netherlands)(Goossens et al., 2011).

Fig. 8 .
Fig. 8. Deviations of the pO 2 values measured for single oxygen sensitive membrane after in vivo experiments compared to the calibration values.

Fig. 9 .
Fig. 9. Top: (a) immersion of a fiber tip into the etchant-organic solution (40 % hydrofluoric acid-toluene); (b) pull out of the fibre until a meniscus is formed at the interface; (c) the meniscus height reduces progressively until the portion of the fiber below the organic solvent is completely etched (d).Below: formed a cone-shaped tip with an angle between 15 -20 °.

Fig. 10 .
Fig. 10.Luminescence emission spectrum of Ru(dpp) 3 2+ (solid line) and absorption spectra of phenol red in the absence (dashed line) and presence (dotted line) of CO 2 (20 %).The hatched area indicates the spectral overlap of the donor emission and acceptor absorption.The CO 2 sensor is solid type sensor, what means that no buffer was used.Instead this the pH indicator (A) forms a hydrated ion-pair (A − Q + •xH 2 O) by an addition of quaternary ammonium base (Q + OH − )(Mills et al., 1992).To stabilize the deprotonated form of the indicator dye in the matrix and to form a lipophilic bicarbonate buffer system in the polymer, tetraoctyl ammonium hydroxide was used.The excess amount of TOA-OH is neutralised by atmospheric carbon dioxide to form a lipophilic hydrogen carbonate buffer in the solid matrix, and this can tune the sensitivity and improve the storage stability of the sensor.The sensing mechanism of an acceptor-Q + OH ion-pair is described by (3):

Fig. 11 .Fig. 12 .
Fig. 11.Oxygen sensor response curve, phase angle as a function of different O 2 concentrations; below: calibration curve obtained by the first cycle of dissolved oxygen concentrations.

Fig. 15 .
Fig. 15.In vivo measurement in liver obtained with inserted pO 2 and pCO 2 fiber optical micro sensors: red line indicates the start of functionality of HLM, while green line indicates the return to the normal intra-corporeal circulation.For the comparison of all performed in vivo experiments on animals using either CPB or MECC as extracorporeal circulation system the measured pO 2 and pCO 2 values were statistically arranged with a two way analysis of variance (ANOVA) model.The measured parameters were normalized to the initial values at a level of 100 % due to the high variances of initial values of those variables measured with optical sensors (pO 2 , pCO 2 ).Those results are shown in Fig.17to 19 for the brain, liver and heart, respectively.

Fig. 16 .
Fig. 16.In vivo measurement in heart obtained with inserted pO 2 and pCO 2 fiber optical micro sensors: red line indicates the start of functionality of HLM, while green line indicates the return to the normal intra-corporeal circulation.

Fig. 18 .
Fig. 18.Comparison of the values measured with optical fibers and normalized to the initial values at a level of 100 % for in vivo measurements in liver.

Fig. 19 .
Fig. 19.Comparison of the values measured with optical fibers and normalized to the initial values at a level of 100 % for in vivo measurements in heart.

Table 1 .
Achieved specifications of the O 2 measurement with planar sensitive membranes compared to the optimum requirements defined in the project