Biosensors and Bioelectronics 391 (2003) 391 /399 www.elsevier.com/locate/bios
Minimally invasive silicon probe for electrical impedance measurements in small animals A. Ivorra a, R. Go´mez a, N. Noguera a, R. Villa a, A. Sola b, L. Palacios c, G. Hotter b, J. Aguilo´ a,* b
a Centro Nacional de Microelecto´nica (IMB-CSIC), Campus UAB, E-08193 Bellaterra, Barcelona, Spain Department of Medical Bioanalysis, Instituto de Investigaciones Biome´dicas, IIBB-CSIC, IDIBAPS, Barcelona, Spain c Department of Physiology, Faculty of Biology, University of Barcelona, Barcelona, Spain
Received 11 July 2002; received in revised form 4 June 2003; accepted 25 June 2003
Abstract It is commonly accepted that electrical impedance provides relevant information about the physiological condition of living tissues. Currently, impedance measurements are performed with relatively large electrodes not suitable for studies in small animals due to their poor spatial resolution and to the damage that they cause to the tissue. A minimally invasive needle shaped probe for electrical impedance measurements of living tissues is presented in this paper. This micro-probe consists of four square platinum electrodes (300 mm /300 mm) on a silicon substrate (9 mm /0.6 mm /0.5 mm) and has been fabricated by using standard Si microelectronic techniques. The electrodes are not equally spaced in order to optimise the signal strength and the spatial resolution. Characterisation data obtained indicate that these probes provide high spatial resolution (measurement radius B/4 mm) with a useful wide frequency band going from 100 Hz to 100 kHz. A series of in vivo experiments in rat kidneys subjected to ischemia was performed to demonstrate the feasibility of the probes and the measurement system. The impedance modulus and phase were measured at 1 kHz since this frequency is sufficiently low to permit the study of the extracellular medium. The extracellular pH and K were also simultaneously measured by using commercial miniaturised Ion Selective Electrodes. The induced ischemia period (45 min) resulted in significant changes of all measured parameters (DjZj /65%; DpH/0.8; DK /30 mM). # 2003 Elsevier B.V. All rights reserved. Keywords: Bio-impedance; Micro-probe; Silicon needle; Rat kidney; Induced ischemia
1. Introduction It has been demonstrated that electrical impedance is a useful parameter to determine the physiological condition of living tissues (Grimnes and Martinsen, 2000; Rigaud et al., 1996). Some pathologies induce changes in essential tissue parameters, such as the extraintracellular volume ratio or the ionic composition, which are reflected as changes in the passive electrical properties (electrical bio-impedance). One of these pathologies is ischemia (lack of blood supply and consequently lack of oxygen). Animal cells regulate their volume by a Na pump-mediated mechanism. In
* Corresponding author. Tel.: /34-93-594-7700; fax: /34-93-5801496. E-mail address:
[email protected] (J. Aguilo´). 0956-5663/03/$ - see front matter # 2003 Elsevier B.V. All rights reserved. doi:10.1016/S0956-5663(03)00204-5
the case of a decrease in the metabolic energy caused by ischemia, this mechanism fails to regulate the Na equilibrium and an osmotic pressure imbalance is created between the intracellular and the extracellular media. As a result, the extracellular water penetrates into the cell with a subsequent cell swelling (Flores et al., 1972) and an extracellular medium shrinkage which is manifested as a decrease in the electrical conductivity of the tissue at low frequencies. Ischemia is induced during certain surgical practices or during the cold storage of grafts for transplantation. Its influence on the cells determines the further viability of the tissue and, therefore, any parameter able to monitor its evolution, such as bio-impedance, could become in medical practice a useful tool to follow the changes induced by the therapeutic management. Moreover, the use of bio-impedance measurements could be
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useful for biomedical researchers and, in this sense, the possibility to experiment with small animals is very attractive. The impedance probe used in most past in vivo studies consists of a linear array of four metallic needle electrodes placed at a constant inter-electrode separation distance. The outer electrodes are employed to inject an AC current into the tissue while the resulting potential is differentially measured across the inner electrodes (Rush et al., 1963). Although the feasibility of this kind of probe has been widely demonstrated, it implies some important practical drawbacks that restrict its use: (a) the fabrication process results in large tolerances because of the critical positioning and alignment of the electrodes, (b) the damage caused to the tissue is considerable since each probe causes four punctures, (c) the presence of a conductive layer (e.g. blood) on top of the tissue under study shunts the electrodes and seriously disturbs the measurements (Steedijk et al., 1993) and (d) the strong dependence of apparent resistance on insertion depth (Tsai et al., 2000) makes crucial the probe fixation to the tissue. Because of these facts, bio-impedance in vivo studies have been mostly limited to moderate size animals and the few studies that have been carried out with small animals made use of different impedance probes with important practical limitations (Jossinet et al., 2001; Raicu et al., 1998). In this paper, a needle shaped probe with four aligned platinum electrodes on its surface is presented (Fig. 1). This configuration not only lessens the limitations of the classical plunge probe but provides another interesting advantage: the probe can be easily fabricated by using standard microelectronic technologies which implies high reliability, low costs and the possibility to integrate other sensors and electronics on the same probe. That kind of geometry has been employed previously for neural activity recording or stimulation (Blum and
Charles, 1988; Drake et al., 1988; Yoon et al., 2000), but, as far as the authors know, it has not been applied yet for impedance measurements. Thus, new design strategies had to be considered in order to optimise the probe performance for bio-impedance measurements. As it has been pointed out, ischemia induces tissue changes that can be sensed by bio-impedance measurements. Here, in order to demonstrate the feasibility of these novel impedance probes to carry out in vivo studies with small animals, a series of in vivo experiments in rat kidneys subjected to ischemia was performed. Other two parameters known to be influenced by the ischemia: the extracellular pH (Dmochowski and Couch, 1966; Mo¨ller et al., 2000) and the extracellular potassium concentration (Mayevsky et al., 2001; Reeves and Shah, 1994) were also monitored and their results and practical constraints were compared with those obtained with the impedance measurements.
2. Materials and methods 2.1. Bio-impedance probes 2.1.1. Probe design The starting point for the probe design was the fourelectrode structure with constant inter-electrode separation distance (IESD). For this structure, the voltage drop at the inner electrodes is inversely related with the IESD (Steedijk et al., 1993) and, therefore, it is desirable to make the probe as long as possible to enhance signalto-noise ratio. However, since the needle shaft length was restricted in order to minimise the tissue damage and to obtain high spatial resolution, a modification was investigated: to bring the inner electrodes near to the outer electrodes (Ivorra et al., 2001). It can be deduced intuitively that this strategy increases the voltage drop
Fig. 1. (a) Impedance probe after wafer sawing. The platinum electrodes are placed at non-constant inter-electrode separation distance to optimise the probe performance. (b) The packaging covers the bonding pads and reduces the needle shaft to 9 mm.
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for a given current and resistivity. Fortunately, it is also beneficial in terms of spatial resolution. By applying the same methods used by Suesserman and Spelman (1993) the signal strength (voltage drop, V ) is easily related with the injected current (I), the sample resistivity (r ), the separation distance between the outer and the inner electrodes (a ) and the separation distance between the inner electrodes (b ) for an isotropic and uniform infinite medium: V
1 4p
r
b a(a b)
I
(1)
This expression shows what was expected intuitively: it is possible to improve the voltage signal without increasing the length occupied by the electrodes (2a/b). Concerning the spatial resolution, it is possible to analytically derive, by using the image method (Robillard et al., 1979), an expression that relates the apparent resistivity (measured resistivity, r ?) with the infinite extent medium resistivity (r ) when there exists a medium transition at a distance x from the nonconstant IESD electrode array (see Fig. 2): G? r? r 1K (2) G 1 1 G (3) a ab 1 1 ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi q G? pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi (4) 4x2 a2 4x2 (a b)2 K
r2 r1 r2 r1
(5)
where K (‘‘reflection coefficient’’) depends on the sample resistivity (r1) and the resistivity of the boundary medium (r2).
Fig. 2. Estimated r ?/r (apparent resistivity/infinite extent medium resistivity) of a conductive medium bounded by a non-conductive medium located at x .
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The r ?/r ratio versus the transition distance with a non-conductive medium (K /1) is represented in Fig. 2 for different electrode distances and a constant array length. It can be observed that effective measurement volume is reduced when a is reduced. Thus, bringing the inner electrodes closer to the outer electrodes also enhances the spatial resolution of the probe. From what has been said above, it could be deduced that the best situation would be to place the inner and the outer electrodes at the minimum possible separation distance. However, it must be taken into account that the most sensitive volume of the probe will be the volume surrounding the separation between an inner and an outer electrode. In the case that this distance is reduced too much, the impedance measurements will depend on the relative position of these electrodes in the cellular structure of the tissue and not on the homogeneous properties of the tissue surrounding the probe. Hence, a separation distance quite larger than the cell size is recommended in order to avoid the heterogeneity of the living tissue at cellular level. In the designed probe, the minimum separation distance between electrodes is 300 mm. The electrode/tissue, or electrode/electrolyte, interface impedance determines the performance of any tissue impedance measuring system, even if the fourelectrode method is applied (Palla´s-Areny and Webster, 1993). This undesired impedance causes measurement errors, especially at low frequencies, that could be understood as noise or signal distortion. In order to minimise these errors, the electrode/tissue impedances should be reduced and matched as much as possible. Since the conductance of these interfaces is directly related to the area, the most evident way to reduce their impedance is to enlarge the electrode surface. In this sense, the square shape is a good choice because it makes good use of the shaft length given for each electrode. In the vicinity of the square corners the current density will not be uniform and non-linear tissue electrical properties could become manifested, however, since the tissue impedance measurement covers a greater volume and the injected current is very low ( B/10 mA), this effect will unlikely influence the measurements. About the probe materials it must be said that the technological process was a decisive factor in regard to the substrate material (Si) and the isolating materials (SiO2 and Si3N4). The chosen material for the electrodes was platinum because of its biocompatibility and low electrode /electrolyte interface impedance compared with other materials such as stainless steel or silver. Furthermore, the Pt electrodes can be coated with a porous layer of ‘‘black platinum’’ which largely reduces the impedance (Geddes, 1972).
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2.1.2. Probe fabrication The fabrication of the bio-impedance probes was carried out at the Centro Nacional de Microelectro´nica (CNM) clean room facilities. The technological process consists of two photolithographic steps starting from a thermal oxidation to grow a thick field layer (800 nm) on 4-in. ( /10 cm) P-type 100 Si wafers with a nominal thickness of 525 mm. The first photoresist layer is applied and patterned on the wafer surface in order to pattern a double Ti/Pt layer (30/150 nm) by using the so-called lift-off technique. Then, two Low Pressure Chemical Vapour Deposited (LPCVD) layers of SiO2 and Si3N4 (300/700 nm) acting as passivation layers are deposited and patterned using the second photolithographic level to open the electrodes and the bonding pads. After the clean room processes, the wafer is sawed by successive parallel and oblique cuts which result in a significant amount (/500) of needle shaped probes (Fig. 1a). Then, each probe is fixed on a tiny Printed Circuit Board (PCB) with gold contacts and wires connected to the electrodes through the PCB by wedge bonding (Fig. 1b). The packaging process ends with complete covering of the PCB with an epoxy resin (H77 from Epoxy Technology, Billerica, MA, USA). 2.1.3. Electrode platinisation A well known method to reduce the impedance of the platinum electrodes is to electrochemically coat them with a porous layer of black platinum as described by Geddes (1972). At low frequencies, this method can reduce the interface impedance ten or more times. However, the resulting black platinum surface is fragile and becomes easily detached when the electrode is inserted into the tissue. Here, the original technique was modified in order to improve the mechanical stability: after a first electroplating using a solution containing platinum chloride (0.025 N hydrochloric acid/3% platinum chloride/0.025% lead acetate) the electrodes are mechanically cleaned with tissue paper and a second electroplating with the same solution is performed. The mechanical cleaning seems to remove the slightly attached platinum deposits and only the strong attached deposits remain on the electrodes acting as anchors for the second electroplating. Up to a point, this method is equivalent to the one described by Marrese (1987), in which ultrasound agitation is used while the electroplating is performed. 2.1.4. Impedance probe characterisation The characterisation of the probes was performed by using a commercial impedance analysis system (SI 1260, Solartron Analytical from The Roxboro Group plc, Cambridge, UK) after the fabrication process has been completed. Dry inter-electrode and electrode-pad im-
pedance measurements were performed to obtain parasite capacitances and parasite resistances. To characterise the electrode /electrolyte interface impedance, the probe was immersed in physiological saline solution (0.9% NaCl, resistivity at 298 K/71.3 V cm) (Oehme, 1991) and impedance spectroscopy was obtained. For each electrode couple, a frequency scan from 10 Hz to 1 MHz was performed at a constant voltage amplitude of 10 mV. These tests were done before and after the series of in vivo experiments to verify the stability of the black platinum deposit on the electrodes. With the aid of a front-end to enhance the input properties of the SI 1260 (Gersing, 1991), four-electrode measurements in physiological saline solution from 10 Hz to 1 MHz were also performed in order to assess the useful frequency band. Equation (2) was deduced under the assumption that the electrodes are point-sized and this is not the case here since the electrode sizes are comparable to the separation distances (Fig. 1). For this reason, an experimental set up was created to study the accuracy of the model: the probe was introduced in a large container filled with 0.9% NaCl and r was measured at 1 kHz, then, a PVC block (resistivity much larger than the saline solution, K /1) attached to a micro-positioner was displaced along the perpendicular axis to the silicon probe and the r ? values were obtained for various distances (x ) and compared with those computed with the model. 2.2. Potassium and pH probes To measure the pH and K extra-cellular levels, commercial miniaturised Ion Selective Electrodes (ISE) and glass reference electrodes were acquired (references 800, 601 and 401 from Diamond General Development Corp., Ann Arbor, MI, USA). The response of these electrodes was checked in lab using known solutions before the in vivo experiments. Especial effort was made to test sensibility, linearity, cross-interferences and drifts for the expected measurement ranges (pH from 6 to 8 and K from 1 to 40 mM). 2.3. Instrumentation for in vivo tests As it is described later on, an in vivo test to validate the measurement frequency for impedance measurements was performed using Solartron SI1260 and a front-end (Gersing, 1991). However, since this commercial system is not portable and it was desired to perform multiple measurements of multiple parameters during the in vivo studies, it was necessary to develop a custommade instrumentation system (Go´mez et al., 2001) whose main features are summarised in Table 2. During all the in vivo studies the system was programmed to
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acquire the samples from all the active channels at each 10 s.
2.4. Experimental procedures
2.4.1. Validation of the measurement frequency At frequencies below 1 Hz and at low current densities, the electrical bio-impedance is mainly related with the extracellular medium since the cell membranes behave as dielectric layers with very low conductance (Grimnes and Martinsen, 2000). In the case that the conductivity of the extracellular medium do not change significantly, this fact can be exploited to study the extracellular volume changes. Unfortunately, due to the electrode/electrolyte interface impedance, it is practically impossible to work at these frequencies without important measurement errors. In some cases, however, it is possible to use higher frequencies and consider that the results are very close to those obtained at a very low frequency. Generally, 1 kHz is considered to be a sufficiently low frequency to study the extracellular medium (Osypka and Gersing, 1995) because, at this frequency, most of the current does not penetrate into the cells. At the same time, this frequency is not too low to induce measurement errors because of the electrode/ electrolyte interface impedance. Thus, a priori this frequency seemed a proper choice but, taking into account the singularity of the tissue under study in the related experiments, impedance spectroscopy was used to test the suitability of this frequency. It was tested to check if at this frequency it would produce results similar to those obtained from studies at lower frequency (/1 Hz). The test is based on the fact that most living tissues exhibit relaxation processes that can be modelled by the Cole /Cole equations (Grimnes and Martinsen, 2000), and from that model it is possible to obtain an impedance value at 0 Hz (R0): Z R
DR ; 1 (jvt)a
DR R0 R
(6)
Following the same procedure that is described in the next section, the right kidney of a rat was subjected to acute ischemia (60 min) followed by a reperfusion period (25 min) while the impedance was monitored using the needle probe and the SI1260/font-end system. The impedance values were obtained from 10 Hz to 1 MHz at every 50 s. With the resulting data, the Cole parameters for the low frequency relaxation were automatically fitted with ad hoc developed MatlabTM routines. Then, after verifying a good Cole fitting, it was possible to compare the R0 value with impedance modulus at higher frequencies.
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2.4.2. In vivo studies The study was conducted under the supervision of the IIBB ethics commission and conformed to the EU guidelines for handling and care of laboratory animals. Male Wistar rats (Ifa Credo, Spain) weighting 250/300 g were anaesthetised with sodium pentobarbital (50 mg/ kg). Indwelling polyethylene cannulas (PE-50, Clay Adams, Sparks, MD) were inserted through the left carotid artery into the aorta (Popovic and Popovic,) for blood sampling and saline infusion (1 ml/100 g per h). The trachea was intubated by using polyethylene tubing (PE-240), and ventilation was maintained using a Harvard animal respirator. PaCO2 values were kept between 4.7 and 5.4 by ventilatory control while PaO2 was controlled between 14 and 20 kPa by adequate oxygen/air mixture control. The abdominal area was covered with saline soaked gauze at 310 K (37 8C) and a plastic cover to minimise dehydration of the exposed tissues. Animals were constantly exposed to radiant heat, maintaining abdominal and kidney surface temperatures at 3099/1 K. To induce kidney ischemia and reperfusion, laparotomy was performed and the left renal pedicle was dissected and occluded with a non-traumatic microvascular clamp. Rats were randomised into two groups: sham-operated (Control group, C, n/5) and those subjected to 45 min ischemia followed by 60 min reperfusion (I/R group, n/5). The impedance probes were orthogonally inserted into the kidney by direct puncture. The penetration depth of the centre of the electrode array was 5 mm. The probe was kept into the tissue by its own shape and no bleeding was appreciated. The shape and volume differences between the kidneys from the series of rats induce sample volume differences that are manifested as differences in the impedance modulus. For this reason, the impedance modulus measurements were normalised to their initial value. ISE were inserted with the help of a cutting edge butt needle allowing controlled deep puncture (bleeding was negligible). The sensitive tip lain at the boundary area between outer medulla and cortex. The reference electrode was placed into the abdominal cavity. All the electrodes were held by loose supports in order to keep constant the angle of the tip. The sensitivity of the pH and K ISEs was calibrated before each experiment using two isotonic solutions of known pH and K concentrations (sol. 1: 1 mM K, pH 8.78; sol. 2: 4 mM K, pH 6.74) prepared from a NaCl buffer as described by Cosofret et al. (1995). The ISEs were kept in these calibration solutions for more than an hour in order to reduce the stabilisation time after the insertion into the living tissue. Baseline arterial plasma pH and K measurements were used to set the timezero values of tissue pH and K .
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3. Results and discussion
Table 2 Instrumentation system main features
3.1. Bio-impedance probes characterisation The results from the impedance probe characterisation after the manufacturing are summarised in Table 1. The electrode/electrolyte interface impedance (Fig. 3) becomes very high at frequencies below 100 Hz and that can involve important tissue impedance measurement errors, especially in a heterogeneous tissue where each electrode can have completely different interface impedances. On the other hand, at frequencies beyond 100 kHz, the capacitive coupling of the wires (including the coaxial wires from the probes to the instrumentation) is strongly manifested (Fig. 4). Thus, it can be considered that the useful frequency band goes from 100 Hz to 100 kHz. The interface impedance test after the series of in vivo experiments resulted in minor changes (impedance modulus increase at 100 Hz B/20%) confirming the stability of the black platinum deposit. The results from the spatial resolution characterisation are shown in Fig. 5. It can be observed that the experimental measurements follow the predicted values by using the model. Therefore, the model works properly although the electrodes cannot be considered as point electrodes. From these results it is possible to provide a value for the spatial resolution: 4 mm. That is, any medium disruption beyond 4 mm from the centre of the probe will cause measurement errors below 1%. The use of microelectronic materials and fabrication processes reduces production costs and offers the possibility to integrate signal conditioning electronics or additional sensors. This could be used, for instance, to expand the useful frequency band by integrating voltage buffers on the silicon. Furthermore, a temperature sensor could be integrated on the same substrate to
Bio-impedance meter Number of channels Ground isolation impedance Oscillator frequency Injected current amplitude CMRR at 1 kHz Input common impedance Input differential impedance
10 50 pF 100 Hz to 125 kHz B/5 mA 88 dB /50 MV//7pF /50 MV//2pF
Ion meter (voltmeter ) Number of channels Ground isolation impedance Input impedance
16 50 pF /10 TV
isolate changes caused by physiological reasons from those caused by temperature changes. 3.2. Validation of the measurement frequency After inducing ischemia by arterial clamping, the R0 increased from 1.7 to 12.5 kV in 60 min. The impedance modulus at 1 kHz followed this evolution with a reduced sensitivity resulting in a 25% of maximum relative difference. This fact could suggest the use of lower frequencies to reduce this difference, however, it must be taken into account that measurement errors are observed for frequencies below 300 Hz and that the range of possible interface impedances is amazingly wide. Therefore, in order to ensure quality for all the measurements it seems that 1 kHz is a proper choice. Fig. 6 shows the results obtained from this experimental test before and after 10 min of induced ischemia and the superimposed Cole fittings. The Cole models were adjusted for low frequencies (B/10 kHz) since the relaxation arc presented distortion for higher frequencies due to other tissue relaxation constants and because of the needle parasitic capacitances.
Table 1 Summarised results from the probes characterisation Parameter
Conditions
Electrode-pad resistance: (connection resistance) I/ V/ V/ I/ Inter-electrode capacitance Inter-electrode impedance modulus in saline solution
TA /298 K
Cell constant (k/r /R ). R , measured resistance; r , resistivity Spatial resolution
TA /298 K TA /298 K, 0.9% NaCl, VOSC /100 mVp 10 Hz 100 Hz 1 kHz 10 kHz 100 kHz TA /298 K, 0.9% NaCl, VOSC /100 mVp ErrorB/1%
Minimum Typical
Maximum
1050 V 900 V 850 V 600 V
1200 V 1000 V 1000 V 700 V 5 pF
1300 V 1100 V 1050 V 800 V 6 pF
5 kV 3.5 kV 3.6 kV 3.4 kV 3.2 kV
8 kV 5 kV 3.8 kV 3.5 kV 3.3 kV 0.32 cm 4 mm
25 kV 7 kV 4 kV 3.6 kV 3.4 kV
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Fig. 3. Inter-electrode impedance modulus (a) and phase (b) measured in NaCl 0.9%.
Fig. 4. Measured impedance modulus (a) and phase (b) of a NaCl 0.9% solution using the impedance probe (four-electrode measurements).
3.3. Induced ischemia in rat kidneys Fig. 7 shows the plots of mean values for normalised impedance modulus, phase angle, tissue pH and interstitial potassium ion concentration. It can be observed that renal vascular occlusion induced immediate and rapid changes in monitored variables. Impedance phase and pH decreased (Fig. 7B and C), and impedance modulus and K increased (Fig. 7A and D). Rapid initial changes in the four parameters were followed with a declining rate change as ischemia progressed. Reperfusion produced an increase in kidney pH as well as in phase (Fig. 7C and B) and a drop in kidney potassium and impedance modulus (Fig. 7D and A). Post-ischaemic impedance modulus experienced a sharp decay, reaching a minimum within the first 10 min of reperfusion, but with second increase episode, with peak values preceding a slow decrease. It is important to note that the impedance do not return to its original values after the reperfusion. This could indicate that some kind of permanent damage to the tissue is not detected by the measurement of pH or K . For low-frequencies ( B/10 kHz), the impedance changes that are exhibited by organ tissues during the course of ischemia are mainly attributed to cell swelling, which narrows the extracellular space, and the gap junctions closure (Gersing, 1998). The presence of gap junctions in the kidney is much more insignificant than
those in other organs (e.g. heart or liver). Thus, it is quite reasonable to consider that the impedance at low frequencies is mainly modified by the extracellular / intracellular volume changes. This assumption agrees with previous histologic observations of rat kidney during ischemia (Flores et al., 1972). In those observations, cell swelling is noted when the tissue is made hypoxic. This cell swelling results in a decreasing extracellular volume which in turn results in an increasing impedance modulus (the current path is narrowed)
Fig. 5. Expected and experimental r ?/r values for the implemented silicon probe.
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(not shown in Fig. 7). These facts reinforce the usefulness of the impedance measurements for experimental or clinical applications. The real time monitoring of living tissues with these novel impedance probes can become a useful tool for biomedical researchers and physicians, not only for rapid detection of ischemia, but also for following the changes induced by the therapeutic management. The electrical bio-impedance monitoring offers some important advantages when compared with other parameters: on-line monitoring, robustness, no calibration is needed, no inherent drift and fully biocompatible materials.
Fig. 6. Nyquist plot evolution from a kidney rat subjected to ischemia (frequency sweeping plots before and after 10 min of induced ischemia). The modelled values (k) have been adjusted to the experimental values (I) for the arc observed at low frequencies.
and in a decreasing impedance phase (current is forced to go through dielectric membranes and, therefore, the reactive part is increased). From these results (Fig. 7) it is important to note that the scatter of the impedance modulus is significantly lower than the scatter of the pH and K values. Furthermore, the impedance values are stabilised much faster (5 min) than the pH and K readings (20 /30 min) after the probes are inserted into the tissue
4. Conclusion A miniaturised four-electrode probe for electrical bioimpedance measurements has been developed by using microelectronic materials and fabrication processes. It consists of four square platinum electrodes (300 mm/ 300 mm) on a silicon substrate (9 mm /0.6 mm /0.5 mm). The electrodes are placed at non-constant interelectrode distance in order to optimise the signal-tonoise ratio and the spatial resolution. The results from the characterisation indicate that these probes provide high spatial resolution (measurement radius B/4 mm) with a frequency band from 100 Hz to 100 kHz.
Fig. 7. Evolution of the measured tissue parameters in kidneys subjected to 45 min of renal vascular occlusion and 60 min of reperfusion. (A) Normalised modulus impedance. (B) Phase angle. (C) Tissue pH. (D) Potassium ion concentration. Results are expressed as mean9/standard deviation (n/5).
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The capability of the probe to measure bio-impedance in small animals has been demonstrated. Furthermore, the obtained in vivo results reinforce the usefulness of the bio-impedance as a marker of the tissue condition.
Acknowledgements This work has been supported by the European Commission through projects ESPPRIT-LTR-23485, IST-1999-13047 and QLK6-CT-2000-00064 and by the Spanish government through projects TIC98-1634-CE, TIC2000-2486-CE, FISS 01/1691 and SAF 2000-3090CE.
References Blum, N.A., Charles Jr, K.K., 1988. Multi-microlectrode probe for neurophysiologial experiments. In: Proceedings of the Fourteenth Annual Northeast Bioengineering Conference, Durham, NH, USA, pp. 15 /18. Cosofret, V.V., Erdo¨sy, M., Timothy, A.J., Buck, R.P., 1995. Microfabricated sensor arrays sensitive to pH and K for ionic distribution measurements in the beating heart. Analytical Chemistry 67 (10), 1647 /1653. Dmochowski, J.R., Couch, N.P., 1966. Electrometric surface pH of the ischemic kidney and the effect of hypothermia. Journal of Surgical Research 6 (1), 45 /48. Drake, K.L., Wise, K.D., Farraye, J., Anderson, D.J., BeMent, S.L., 1988. Performance of planar multisite microprobes in recording extracellular single-unit intracortical activity. IEEE Transactions on Biomedical Engineering 35 (9), 719 /731. Flores, J., DiBona, D.R., Beck, C.H., Leaf, A., 1972. The role of cell swelling in ischemic renal damage and the protective effect of hypertonic solute. The Journal of Clinical Investigation 51, 118 / 126. Geddes, L.A., 1972. Electrodes and the Measurement of Bioelectric Events. Wiley-Interscience/Wiley, New York, pp. 32 /35. Gersing, E., 1991. Measurement of electrical impedance in organsmeasuring equipment for research and clinical applications. Biomedizinische Technik 36 (1-2), 6 /11. Gersing, E., 1998. Impedance spectroscopy on living tissue for determination of the state of organs. Bioelectrochemistry and Bioenergetics 45, 145 /149. Go´mez, R., Noguera, N., Ivorra, A., Villa, R., Aguilo´, J., Milla´n, J., Lo´pez, J., Palacios, L., Sola, A., Hotter, G., 2001. Instrumentation system for in vivo organ studies. In: Proceedings of 24th International Semiconductor Conference, Sinaia (Romany), 9 /13 October, vol. 1, pp. 261 /264. Grimnes, S., Martinsen, O.G., 2000. Bioimpedance and Bioelectricity Basics. Academic Press. Ivorra, A., Aguilo´, J., Milla´n, J., 2001. Design considerations for optimum impedance probes with planar electrodes for bioimpe-
399
dance measurements. In: Proceedings of 24th International Semiconductor Conference, Sinaia (Romany), 9 /13 October, vol. 1, pp. 269 /272. Jossinet, J., Moulin, V., Matias, A., McAdams, E.T., 2001. In vivo spectrometry of dunning prostate tumours. In: Proceedings of the XII Conference on Electrical Bio-Impedance, Oslo (Norway), 17 / 21 June, pp. 285 /288. Marrese, C.A., 1987. Preparation of strongly adherent platinum black coatings. Analytical Chemistry 59 (1), 217 /218. Mayevsky, A., Nakache, R., Luger-Hamer, M., Amran, D., Sonn, J., 2001. Assessment of transplanted kidney vitality by a multiparametric monitoring system. Transplantation Proceedings 33, 2933 / 2934. Mo¨ller, H.E., Gaupp, A., Dietl, K.-H., Buchholz, B., Vestring, T., 2000. Tissue pH in human kidney transplants during hypothermic ischaemia. Magnetic Resonance Imaging 18, 743 /751. Oehme, F., 1991. In: Go¨pel, W., Jones, T.A., Kleitz, M., Lundstro¨m, J., Seiyama, T. (Eds.), Sensors, A Comprehensive Survey, Chemical and Biochemical Sensors, Part I, vol. 2. VCH Verlagsgesellschaft mbH, Weinheim, pp. 314 /332. Osypka, M., Gersing, E., 1995. Tissue impedance spectra and the appropriate frequencies for EIT. Physiological Measurement 16, A49 /A55. Palla´s-Areny, R., Webster, J.G., 1993. AC instrumentation amplifier for bioimpedance measurements. IEEE Transactions on Biomedical Engineering 40 (8), 830 /833. Raicu, V., Saibara, T., Irimajiri, A., 1998. Dielectric properties of rat liver in vivo: a noninvasive approach using an open-ended coaxial probe at audio/radio frequencies. Bioelectrochemistry and Bioenergetics 47, 325 /332. Reeves, W.B., Shah, S.V., 1994. Activation of potassium channels contributes to hypoxic injury in proximal tubules. Journal of Clinical Investment 94 (6), 2289 /2294. Rigaud, B., Morucci, J.-P., Chauveau, N., 1996. Examples of applications. Critical Reviews in Biomedical Engineering 24 (4-6), 323 /351. Robillard, P.N., Poussart, D., 1979. Spatial resolution of four electrode array. IEEE Transactions on Biomedical Engineering BME-26 (8), 465 /470. Rush, S., Abildskov, J.A., McFee, R., 1963. Resistivity of body tissues at low frequencies. Circulation Research XII, 40 /50. Steedijk, P., Mir, G., Van Del Velde, E.T., Baan, J., 1993. The fourelectrode resistivity technique in anisotropic media: theoretical analysis and application on myocardial tissue in vivo. IEEE Transactions on Biomedical Engineering 40 (11), 1138 /1148. Suesserman, M.F., Spelman, F.A., 1993. Quantitative in vivo measurements of inner ear tissue resistivities: I. In vitro characterization. IEEE Transactions on Biomedical Engineering 40 (10), 1032 / 1046. Tsai, J.-Z., Cao, H., Tungjitkusolmun, S., Woo, E.J., Vorperian, V.R., Webster, J.G., 2000. Dependence of apparent resistance of fourelectrode probes on insertion depth. IEEE Transactions on Biomedical Engineering 47 (1), 41 /48. Yoon, T.H., Hwang, E.J., Shin, D.Y., Park, S.I., Oh, S.J., Jung, S.C., Shinn, H.C., Kim, S.J., 2000. A micromachined silicon depth probe for multichannel neural recording. IEEE Transactions on Biomedical Engineering 47 (8), 1082 /1087.
