Chapter 15 Microdialysis based lab-on-a-chip, applying a generic MEMS technology

Chapter 15

Microdialysis based lab-on-achip, applying a generic MEMS technology P. Bergveld, W. Olthuis,*A.J. Sprenkels, D.G. Pijanowska, H.J. van der Linden and S. Bohm

15.1

INTRODUCTION

In reaction to the biocompatibility problems of catheter type chemical sensors, which are meant for in vivo monitoring, the microdialysis concept has been introduced. It is in fact a sampling technique, producing a protein- and cell-free sample containing an imprint of the chemical composition of blood, which can be monitored by sensors outside the body. This chapter treats the integration of a microdialysis probe with a silicon chip, which contains the sensors, as well as calibration facilities to calibrate the sensors and a pump for the liquid handling. Because the chip contains all relevant system components for the chemical analysis, it is in fact a lab-on-a-chip. Glucose as well as potassium concentrations have been measured successfully with the system.

*Corresponding author. Tel.: +31-534892688; fax: +31-534892287. E-mail address: [email protected]. Comprehensive Analytical Chemistry XXXIX, pages 625-663 © 2003 Elsevier Science B.V. All rights reserved ISSN: 0166-526X

625

15.1.1

The need for in vivo monitoring

About 30 years ago, silicon chip technology, developed by the integrated circuit (IC) industry, started also to be exploited in the field of sensor development, especially for biomedical purposes. Because the chips are so rigid and small, they can easily be incorporated into the tip of a catheter to be applied for in vivo monitoring. Especially for chemical blood analysis of critically ill patients in an intensive care room, a strong need for in vivo monitoring exists as an alternative to the usual laboratory tests with blood samples. Only through continuous measurement can trends in a patient's condition, for instance with respect to their blood concentration of potassium ions or lactate concentration, be observed and used as an early warning. Another example is the development of chip-based glucose sensors for diabetic patients, in order to notify hypoglycaemic situations during the night, for glycaemic stabilization during hospitalization or for the control of an implanted insulin delivery system (an ultimate goal of glucose sensor development). Unfortunately, independent of the specific sensor concept or manufacturing technology, all attempts to make (bio)chemical sensors, which are in intimate contact with whole blood for a relatively long period of time have so far failed due to the blood incompatibility of the applied materials. Invariably proteins will adsorb to the sensor interface, ultimately followed by encapsulation with fibroblasts, resulting in interference with the actual sensor function, for instance changing its sensitivity and selectivity. Separating the actual sensor interface from the blood by means of a haemocompatible membrane has not proved to be a successful solution. Usually the response time of the sensor is increased to an unacceptable degree. Moreover, in both situations, with or without an additional membrane, the sensor should be calibrated at regular time intervals, for which up to now no adequate procedure has been developed. Furthermore, the risk of thrombosis, embolism and septicaemia is thought to be too great with the use of invasive sensors, which certainly does not enhance the application when the sensor signal is also not 100% reliable. A solution to the problems as described above is the separation of the sensor (chip) and the sample in such a way that the ions and metabolites to be measured can reach the sensor interface, but the proteins and other large molecules cannot. In addition, the sensor should be placed extracorporeally, where it can be calibrated. Furthermore the sample 626

should be taken from the living organism by means of a minimally invasive procedure. This approach is developed as microdialysis and was first applied in clinical practice about a decade ago. 15.1.2

Microdialysis

The microdialysis concept involves the use of a hollow fibre, made partly of a semi-permeable membrane material. The hollow fibre, having an outer diameter of about 300 [tm, is inserted into the subcutaneous tissue, thus mimicking a blood capillary. The fibre is flushed by a buffered saline solution, the perfusate. At the site of the semi-permeable membrane, ions and water-soluble molecules of size below the molecular cut-off of the membrane will diffuse from the tissue and thus from the surrounding blood capillaries into the perfusate. The dialysate leaving the probe will thus contain an imprint of the chemical constitution of the tissue, which may be an almost 100% faithful representation in case the perfusate flow is stopped for a time that is sufficient for equilibration. For shorter times, in situations of continuous flow of the perfusate, this number will be smaller, depending on the flow rate. This percentage is called the relative recovery (see Section 15.2.2). The actual sensors are placed off-stream and can thus measure the clean dialysate and can also be calibrated by intermediate contact with calibration liquids. Biocompatibility of the sensor surface is no longer a requirement. The technique is regarded as minimally invasive when compared to the insertion of needle-type sensors of millimetre size, because microdialysis probes are much smaller in diameter. Originally microdialysis was introduced by Ungerstedt et al. [1] for monitoring neurotransmitter release in the brain, but has now become a frequently used research tool for the investigation of chemical compositions in many other tissues or body compartments, as introduced by Lonnroth et al. [2]. Up to now no clinical applications have been launched, due to the fact that the commercially available microdialysis probes have to be connected by means of tubes to existing conventional chemical analysis equipment. This equipment is usually too bulky for application near a patient's bed and moreover the tubing will cause too large dead volumes. The same can be said with respect to the introduction of microdialysis in non-medical applications, such as monitoring biotechnological processes, including the automatic control of these processes. 627

Additionally, direct contact between a sensor surface and, for instance, the fermentation solution should be prevented. This can be achieved by applying the microdialysis concept to this technological field also, which is becoming increasingly important. Regardless of the disappointing results of the application of small and rigid chip-based sensors when brought into direct contact with a living system, the sensors as such are well established at present; potentiometric as well as amperometric sensors are commercially available. The same can be said for the microdialysis probes. The sensors are extremely suitable for measuring small volumes, such as the dialysate volumes from a microdialysis probe. Therefore it has been recognized by the first author of this chapter that a combination of the tiny sensors and the tiny microdialysis probes would be a very favorable development. In other words, if the present plastic connector of a microdialysis probe can be replaced by a silicon connector in which the sensors are placed, the best of the two systems are combined. This idea formed the start of a research program 'the comprehensive integration of the microdialysis sampling technique and silicon sensors' in 1996, subsidized by the Dutch Foundation for Fundamental Research on Matter (FOM, Utrecht, the Netherlands). The result at the present moment (2002) is a lab-on-a-chip, containing a sensor array with integrated calibration facilities, to which a microdialysis probe is connected without additional tubing. 15.1.3

The microdialysis lab-on-a-chip

The term 'lab-on-a-chip' is reserved especially in the USA for systems that consist of a sandwich of glass plates in which a certain configuration of channels is etched in one or another way. After insertion of a sample into an input port, the chips are placed in a benchtop apparatus, which contains electrodes to contact the various liquid input ports of the channel structure. By means of electro-osmotic forces (EOF pump) the liquids are forced to move through the channels and are separated into different components (electrophoresis), with the result usually monitored by means of an optical system (the actual sensor), which is an integral part of the relatively large equipment, as reviewed by Kopf-Sill of Caliper Technologies Corp. [3]. The chips are presently available commercially for gene analysis, drug screening, etc. Note that not all components of the analysis system have been 628

miniaturized, but only the liquid handling system, whereas the pumping action of liquids is limited to EOF. In Europe the term RTAS has been introduced and is more often used. The term denotes micro total analysis system, a stand-alone system containing all the components necessary for a particular chemical analysis. Such a microsystem comprises for the liquid handling micropumps and valves, and for the actual analysis the necessary micromixers, microreactors, microsensors and the means to provide calibration. Some of these systems are built-up in a modular way, a so-called hybrid approach as described by Wissink et al. [4]. For a higher degree of integration, or in other words, for a complete laboratory on one and the same chip, a monolithic technology is required, as derived from IC technology. For the design of the microdialysis probe with incorporated sensors this monolithic integration is chosen as the starting point. This choice limits the design of the already known components like micropumps and microsensors, because of the different technologies used up to now for these components. For instance, it will be difficult to combine in one and the same process the technology for making piezoelectrically driven pumps for the liquid transport and chemFETs for the sensors. Instead of applying difficult-to-combine technologies for the microdialysis system, an a priori choice should be made for one generic technology. With this requirement in mind the necessary components should be (re)designed or even newly developed. In the case of developing the microdialysis lab-on-a-chip, the chosen generic technology consists of etching the channels and cavities into a silicon substrate, depositing metal electrodes onto a glass wafer and bonding the silicon substrate to the glass cover. With this technology, the following different subsystems have to be realized: · the microdialysis probe is directly connected to the silicon/glass sandwich in order to realize a liquid handling system with minimum dead volumes; · the sensor array, containing potentiometric as well as amperometric sensors and a reference electrode, applies a flow-through concept; · facilities exist to calibrate the different sensors as an integral part of the system; and · a pump for the transport of perfusate/dialysate for continuous as well as stop-flow mode of operation. 629

Figure 15.la shows how the combined subsystems will result in the complete lab-on-a-chip, as patented by FOM [5], whereas Fig. 15.1b shows how the sensors subsequently measure the calibration pulses cal.2 and cal.1 as injected on the chip by the two calibration systems, followed by the actual sensor signal. In the following sections, the development of the first three of the four subsystems mentioned above will only be treated, because the perfusion liquid pump is not yet ready for integration. Nevertheless the fully integrated system will be described in Section 15.6, although for the

a)

calibration chamber

reference electrode

flow-through sensor

,-I 9

sample

m o,

U)

a

b)

_

... n..-F

·I_ time

Fig. 15.1. (a) Drawing taken from the international patent application of a microdialysis probe, integrated with silicon chip; (b) two calibration pulses, followed by the sample signal. 630

time being applied with an external perfusion liquid pump. The integration necessitates partly redesigning the separate components as described in previous sections, as dicted by the actual integration process. 15.2

THE MICROMACHINED DOUBLE LUMEN MICRODIALYSIS PROBE CONNECTOR

Microdialysis probes are at present commercially available. They consist of an ultra thin shaft with, on the indwelling side, the actual semipermeable membrane section and at the other end a plastic connector with small tubes, to which extension tubes have to be connected for introduction of a perfusion liquid and for transport of the dialysate to a remote chemical analyzer. These tube connections are in fact the weak point of the present microdialysis concept, as well as the lack of suitable microchemical analyzers. Replacement of the plastic connector by a silicon chip, which may contain the sensors, will be the first step towards the complete lab-on-a-chip system as described in this chapter. 15.2.1

The conventional microdialysis probe

The construction and functioning of a typical commercially available double lumen microdialysis probe is given in Fig. 15.2. Basically, 0(~ 3

--

cells

dialysate

perfusion liquid

tissue

X-A

connector

I

IM,

E*

-

MO probe

outer tube (400 gm)

innertube semi-penneable (150 in) membrane

Fig. 15.2. Construction of a typical commercially available double lumen microdialysis probe. 631

the probe consists of a small inner tube of about 150 im outer diameter, located within an outer tube of about 400 m outer diameter. The inner tube is connected to a precision pump that drives the perfusion liquid (physiological saline) at a continuous flow rate in the range of 0.2-2 LL/min. The perfusion fluid leaves the inner tube at the probe tip and then flows alongside the semipermeable membrane, via the outer lumen, back to the connector. However, in the application of on-line monitoring, the very low flow rate associated with microdialysis requires swept volumes to be minimized in order to reduce the induced lag time in measurements. The volume between the probe and the analysis system especially requires attention. In a microdialysis probe the total swept volume is made up of two parts. The first part is between the probe tip and the connector, whilst the second part is formed by the connector and its output tube, through which the dialysate is guided to the sensor array. To indicate the effect, a total swept volume of 1 L introduces a lag time in the measurement, i.e. the time to reach the sensor, of the order of a minute. 15.2.2

Theory

Probe recovery As the microdialysis sampling technique yields a sample stream, which reflects to some extent the actual composition at the implantation site, the dialysate could be linked to this actual sample composition. The process of sampling will be described with reference to Fig. 15.2. The probe inlet for the perfusion liquid is connected to a precision pump which drives the perfusion liquid at a continuous pump rate. This perfusion liquid can contain a certain concentration Co (mol/L) of the analyte of interest (for instance an ion like sodium or chloride). On its passage alongside the semi-permeable probe tip, the analyte present at a concentration Cs in the sampling region will diffuse according to the concentration gradient. This gradient, the difference in concentration in the sampling region and the perfusion liquid, Cs - Co, is the driving force for diffusional mass transport from the sample to the perfusion liquid. This mass flux through the semi-permeable membrane enriches the perfusion liquid and results in a dialysate having a concentration Cd. The relative recovery, RR, which describes the relation between the analyte concentration in the dialysate coming from the probe and 632

the actual external concentration at the probing site, is generally given by RR

Cd - C Cs - Co

(15.1)

This relative recovery, often expressed as a percentage, is besides a number of other parameters, inversely dependent on the flow rate. A higher flow rate will for a certain probe geometry result in a dilution of the incoming diffusional mass flux by the perfusion liquid and as a consequence RR will be smaller. Other factors having a large effect on the relative recovery are the type of semi-permeable membrane used and the sample matrix at the probing site (i.e. subcutaneous tissue, blood stream, etc.). Because the relative recovery is needed to calculate the actual sample concentration C from the measured dialysate concentration Cd, identification and understanding of this ratio is of prime importance. Lag time Lag time is defined as the time between a change in composition of the actual sample and the moment that this change is detected by the sensor system (moment of onset of the detector). It is assumed here that the response time of this sensor system is infinitely small, so that the main factor that determines the lag time results from the transport of the dialysate from the probe tip to the detector. As the diameters of the flow channels are small (around 100 Rm) and flow rates are low, it is assumed that dispersion resulting from laminar flow and diffusion will balance. As a consequence the sample will travel as a 'plug' through the fluidic system (flow injection analysis conditions). If the volumetric flow rate of the perfusion liquid is q (m 3/s) and the total swept volume is Vtotal (m3 ), the lag time T (s) induced by the swept volume is given by Vtip + Vconnector

Vtotal

q

q

(15.2)

where Vtip and Vconnector (m3 ), respectively, represent the swept volume in the probe tip and shaft, and the swept volume within the connector. Because the construction of the tip and shaft of the microdialysis probe described in this work is exactly the same as for conventional double lumen probes, the main benefit that can be achieved from 633

the application of microsystem technologies will be in the minimization of the swept volume Vconnector in the connecting part.

Chloride sensitive electrode In the current micromachined probe, as a case study for fast on-line measurements, a chloride sensitive electrode has been incorporated in the probe connector. This sensor allows the measurement of chloride concentrations in the dialysate. 15.2.3

Experimental

Device construction An exploded view of the proposed micromachined connector/sensor construction is given in Fig. 15.3. Basically it is a sandwich construction of two multi-step KOH etched wafers. In the silicon parts, a number of features can be identified. The actual connecting part is formed by the wide etched U-trench for receiving the outer capillary tube. This trench, which is present in both the upper and lower part, narrows via a funnellike section down to a V-shaped channel in which the inner capillary tube is to be inserted. Also this channel is present in both parts, however, in the lower part it ends in a flow-through cell for sensor integration. Besides channels for the fixing and aligning of the capillary tubes, a number of holes are etched. In the lower part, at the place where the funnel-like structure ends in the V-channel, an opening for the perfusion liquid is located, whilst at the end of the flow-through cell the waste outlet is located. In the upper part, two holes are etched through which, after insertion of the capillary tubes, small amounts of glue can be dosed for fixing and sealing these tubes. The perfusion liquid enters the connector via the lower part and travels in the annular space between the inner and outer tube towards the semi-permeable tube, which forms the tip of the probe. After flowing alongside the membrane, the formed dialysate enters the inner tube and flows back to the connector. After leaving the inner tube, the dialysate flows through the flow-through cell that contains the sensor. Finally after being measured, the dialysate exits the device via the waste output. For the experiments, two probes of different dimensions were assembled, each having identical outer capillary tube lengths 634

holes to app

" sensor position

outlet

-through cell

liquid

Fig. 15.3. Exploded view of the connector/sensor construction. The fluid flow from inlet to outlet is marked with numbers 1-5 (1-3: perfusion liquid, 4: dialysate entering the flowthrough cell, 5: dialysate leaving the system).

(15 mm), but with semi-permeable tube lengths of 5 and 20 mm, respectively. Sensor incorporation For the incorporation of a sensor an easy to integrate Ag/AgC1 electrode was chosen for evaluation of the performance of the proposed system. This electrode for the measurement of chloride ions was placed in the flow cell at the location 'sensor position' as indicated in Fig. 15.3. The mean sensitivity of the resulting electrode in the range between 10 and 100 mM C1- was determined to be 55 mV/pCl. To compare the micromachined probe with a commercial probe, having comparable dimensions (BR-4, 4 mm membrane length, BAS Lafayette, USA), 635

a conventionally produced flow-through sensor for chloride was made by piercing a 0.5 mm internal diameter PVC tube with a 0.2 mm chlorinated silver wire. Measurement set-up Because the target analyte is chloride, as 'sample matrix' a 100 mM KC1 solution was chosen. To provide a stable baseline, a perfusion liquid containing 10 mM KC1 was used. As the main object of the experiment was to compare the lag time of measurements with conventional probes coupled to a sensor and the micromachined probe connector with integrated sensor, probes were switched between two stirred sample solutions. The first sample solution was identical to the perfusion liquid pumped through the probe whilst the actual sample solution was formed by the 100 mM KC1 solution. To evaluate the lag time of a particular probe/sensor combination, first a stable sensor baseline was obtained with the probe placed into the first (stirred) solution and perfused at a fixed rate. Then the probe was transferred swiftly to the stirred sample solution and the sensor output was recorded on a storage oscilloscope. After recording the response, the probe was transferred back to the first solution. This measurement procedure was repeated for different perfusion rates in order to investigate the difference in lag time as well as the relation between perfusion rate and the relative recovery of the probe under study (see Fig. 15.4). 15.2.4

Results and discussion

Conventionalprobe / sensor combination First measurements were performed with the conventional probe connected with minimal swept volume to the wire chloride sensor. The typical microdialysis perfusion flow rate of 1 [LL/min was chosen to result in a reference response with which the responses of the micromachined probe could be compared. The response (not shown) to 100 mM KCl resulted in a final amplitude of 47 mV whilst the lag time was about 60 s. The amplitude is lower than the sensors' sensitivity of 55 mV/pCl, which is the result of the relative recovery of the probe. Using the Nernst equation, a RR of about 0.68 was calculated (Cd = 71 mM). 636

0.01 0 -0.01 -0.02 -0.03 -0.04 -0.05 -0.06 0

5

10

15

20

25

time (s) Fig. 15.4. Measured response to 100 mM KCI for a micromachined probe/sensor (membrane length 5 mm).

Micromachinedprobe/ sensor combination Next the response of the micromachined probe with 5 mm membrane length was measured for increasing perfusion flow rates. From the results plotted in Fig. 15.4, two effects can clearly be seen. The first is the decrease in lag time with increasing flow rate, which is expected from Eq. (15.2). At a flow rate of 12 pL/min, within 1.5 s the output signal starts to change and within 4 s the signal is at its final value of 20 mV, considerably smaller than 47 mV. Due to the high flow rate, the residence time of the perfusion liquid in the semi-permeable probe tip is in the order of a second, and consequently the relative recovery RR is small. These two effects are inversely dependent as can be seen from the same figure. For a flow rate of 1 LL/min, the amplitude is 49 mV, slightly larger than the amplitude obtained with the conventional probe in the previous experiment. This is probably due to a slightly larger permeability of the membrane material. However, the lag time is about 11 s, which is considerably shorter due to the fact that the swept volume in the connecting part was minimized by the application of micromachining. 637

The measured lag time and relative recovery for both the 5 and 20 mm probes are summarized in Fig. 15.5. It can be seen that for both is inversely proportional to the membrane lengths the lag time perfusion flow rate q, as is expected from Eq. (15.2). The relative recovery RR also shows inverse proportionality, which is in agreement with literature. The longer lag time for the 20 mm probe is caused by the larger swept tip volume Vti p being an integral part of the total swept volume Vtota 1. On the other hand, this longer membrane length results in a higher relative recovery RR approaching unity for flow rates q smaller than about 5 L/min. If the probe is operated in this range, the analyte concentration in the dialysate Cd equals the analyte concentration Cs in the sample (see Eq. (15.1)), so that no calibration for the sampling process has to be performed. It can be concluded that a microdialysis probe/sensor combination was constructed in which the swept volume and consequently the lag time has been decreased by a factor of more than five, as compared to a conventional probe connected to a sensor located in the output tube. The minimal volume has been mainly achieved by the application of appropriate micromachining techniques for the manufacturing of a microconnector between a conventionally assembled double lumen probe shaft and tip, and the incorporation of a sensor in silicon. Besides the reduction of lag time, another main advantage of the realized micromachined probes is the fact that they can be I

20

Mn 0:

7

15

E 0

/I/

! 0.40.2

0 0

0.2

0.4

0.6

1/q [min/gl]

0.8

1

1.2

0

i

0.2

0.4

0.6

0.8

1

1.2

1/q [min/gl]

Fig. 15.5. Lag time and relative recovery as a function of flow rate for the 5 and 20 mm micromachined probes. 638

directly integrated with other microfluidic systems in silicon. Another advantage over the use of conventional probes in hybrid combination with microfluidic systems is that all critical fluidic connections are inside the system, from for the user. This feature ensures that no extra lag time is introduced by tubing and additional connectors. As a final advantage, it can be concluded that the resulting microconnector enables the easy (automated) assembly of double lumen microdialysis probes in a highly parallel fashion. This is mainly due to the self-aligning structure for receiving the concentrically positioned capillaries and the fact that the through holes for glue application allow the automatic distribution of the specially formulated epoxy.

15.3

THE PASSIVE AND THE ACTIVE CALIBRATION SYSTEM

For micrototal analysis systems to work in a reliable fashion it is important that all steps in the system, like fluid handling, separation and detection are performed in a reproducible and accurate manner. Particularly the detection requires attention as every microsensor exhibits time-varying sensitivity and offset. Therefore an integrated calibration system is a useful microcomponent. In flow-through sensors, such as those used in FIA-based systems, calibration can be performed by formation of reference plugs in the fluid stream leading to the microsensor. When two plugs of different concentration are used, reference responses are generated in the microsensor, which can be used to correct both offset and sensitivity variations. A microcalibration unit can be made using passive or active control. 15.3.1

Passive control of a calibration plug

A passive microcalibration unit can be made by using the passive diffusion of reference substances from a hydrogel into an adjacent microchannel. The hydrogel is filled with a known concentration of the reference compound and placed in direct contact with the carrier stream that leads to the sensors to be calibrated. A reference plug is formed by stopping the flow for a time long enough to allow the reference compounds to diffuse from the hydrogel into the microchannel and form a reference plug. Important parameters for 639

the calibration unit are the time needed for reference plug formation and the number of calibrations that can be performed before the unit is depleted. A prototype of a passive microcalibration unit has been made that comprises a microfluidic channel, a hydrogel loaded with pH buffer and an ISFET downstream of the hydrogel for detection of the pH changes generated by the plug formation, as schematically shown in Fig. 15.6. Results from measurements obtained with the prototype have shown that reference plugs can be formed in a time period of a few seconds. Also, it has been shown that sequential reference plugs show a decrease in the reference compound concentration due to depletion of the hydrogel, which is a drawback of the system. Another drawback is the continuous out-diffusion of reference compound from the hydrogel resulting in an offset in the measurement signal and an interference with the sample to be analyzed. It can be concluded that instead of passive plug formation, a method of active fluid dispensing on demand is needed for accurate calibration of the sensors. By using an active pumping mechanism that provides undiluted calibrant only when this is needed, accurate reference plugs can be generated without generating an offset in the measurement signal and without interference to the sample.

hvdlronl

carrier out

Fig. 15.6. Perspective view and cross-section of the realized flow-through system for calibration of an integrated pH sensitive ISFET. 640

15.3.2

Active control of a calibration plug

As an alternative to the passive formation of a calibration plug, an active microdosing system has been developed that can be used as a microcalibration unit. The system is based on the displacement of a fluid from a reservoir by the electrochemical production of gas bubbles. This principle was used earlier in a portable miniature pump for drug infusion by Koch et al. [6]. Because the electrochemical gas production can be accurately controlled, the pump can dose small volumes with high accuracy. The microdosing system consists of a channel structure micromachined in silicon, closed by a Pyrex® cover with noble metal electrodes. In the silicon structure, the liquid to be dispensed is stored in a meander-shaped channel that starts in an electrolyte-filled reservoir above which two platinum electrodes are positioned. By sending an electric current through the liquid via the electrodes, gas bubbles can be produced. As these gas bubbles expand in the reservoir, the liquid stored in the meander is driven out into the channel structure leading to the microsensor to be calibrated. Figure 15.7 shows a realized double calibration unit, where the meander structures and the gas pump reservoirs can be easily distinguished. Theory If in the system shown in Fig. 15.7 the appropriate conditions are chosen (electrolyte composition, electrode material, cell configuration) then the total volume of gas bubbles produced and consequently the amount of fluid that is dispensed, can be controlled by means of the on-time of the current source and the current amplitude. If two noble metal electrodes are chosen, the following reactions take place in water at neutral pH, Anode:

2H 2 0 -* 4H+ + 4e + 02 (g)

(15.3)

Cathode:

2H2 0 + 2e -* 20H- + H 2(g)

(15.4)

From both half-reactions, the rate of the gas production dV/dt (m3/s) and consequently the dosing rate q (m 3/s) at atmospheric pressure can be derived if a constant current i (A) is assumed dV 3 i q = dt = V. dt 4 k

(15.5) 641

Fig. 15.7. Photograph of a realized dual-pump. Note that the left pump is almost completely filled with generated gas whilst the right pump and meander is filled with electrolyte.

where F is Faraday's constant (96.49 x 103 C/mol) and Vm is the molar gas volume at 25°C and atmospheric pressure (24.7 x 10 - 3 m3 /mol). If a current pulse of amplitude i and duration At (s) is applied, the total dosed volume AV is given by AV = qAt

(15.6)

Results The dosed volume as a function of the amplitude i of a current pulse of 250 ms duration, applied to the electrodes, is shown in Fig. 15.8, together with the theoretical expected dosed volume. It can be seen that the dosed volume corresponds well with the theoretical expectations. A problem with this system is the catalytic back reaction of the generated oxygen/hydrogen mixture at the noble metal electrodes (platinum). This back reaction results in a decrease in the gas volume in the reservoir and in a back-flow of dosed volumes. Also the produced gas sometimes remained dissolved in the electrolyte and did not give rise to evolving gas bubbles, which resulted in errors in the dosed volume. 642

300 250 -' 200

>

150 100

I

·

5(

--

Experimental Theory

0 0

2

4

6

8

Current pulse amplitude (mA) Fig. 15.8. Measured and calculated dosed volume as a function of the actuation current (250 ms actuation, r2 = 0.996).

As the dosed liquid volume equals the volume of generated gas bubbles, precise dosing can be obtained by closed-loop control, with the amount of generated gas as the control parameter. 15.3.3

Closed-loop controlled electrochemically actuated microdosing system

The dosing accuracy of the above-mentioned active dosing system could be improved by incorporating a feedback system in which the gas bubble volume is measured and used to control the dosing action. In this set-up, a pair of interdigitated electrodes is used to electrochemically generate gases and to simultaneously measure the generated gas volume, via an impedance measurement of the gas/liquid mixture in the reservoir. As this measured gas bubble volume equals the dosed volume, active control of dosed volumes can be achieved. Using this set-up, accurate dosing of volumes in the nanoliter range with an accuracy of 5 nL was reached. The realized system is depicted in Fig. 15.9. The system has separate chambers for the generation of hydrogen and oxygen gas, which are in ohmic contact by a chevron that acts as a diffusion barrier for 643

rough hole -meander B

filling port 0 - hydrogen

P~

bubble

5

interdigitated S/A electrodes - bond pads

Fig. 15.9. Proposed dosing system with feedback control, geometry and electrical connections.

the generated gases. These two measures have been taken to compensate the catalytic back-reaction of the gases at the platinum electrodes. To link the cell impedance with the total gas bubble volume a void volume fraction has to be defined: VH, (or VO2)

(15.7)

Vreservoir

where Vreservoir (m3) is the internal volume of the corresponding bubble

is zero before any gas bubbles are reservoir. By this definition, generated and can in theory increase to unity, which represents the case of a bubble reservoir completely filled with gas bubbles. If the impedance is determined at a high enough frequency (200 kHz) to eliminate the influence of the double layer capacitancies, the generated gas volume can be calculated from the following equation, where Relectrolyte is the electrolyte resistance, Rseries is the resistance of 644

the connecting wires and the planar electrode structure and R(e) is the real part of the cell impedance: VH 2 (or V 0 2 )

Vreservoir(1

=

-

Relectrole R(E) - Rseries

(15.8)

)

The feedback loop was controlled with a computer on which a proportional control routine was implemented for continuously controlling the amplitude of the current through the electrodes based on the measured cell impedance (i.e. the applied electric current was proportional to the difference between the preset value and the actual value of the gas volume determined via the measured resistance Relectrolyte )

Figure 15.10 shows the graph for a dosing step of 100 nL. In the graph the controlled actuation current as well as the graph representing the injected volume are shown. It can be seen that as the dosed volume goes up the current is ramped down when the dosed volume approaches the set-point of 100 nL. By using this approach it is possible to accurately dose volumes in the nanoliter range with 5 nL accuracy and to overcome the unpredictability in bubble formation associated with the open loop active dosing system. The resulting device can, because of the remarkably simple

Controlled dosing 120

60

100

E

-

50

60

3

~~~~~0

afi~~~30 100

20-

10

20

30

40

2 0 .............................. .....-................................................. ............................---................

50

60 ...................

70

80 ............. .............

90

1

...............

0

time (s) Fig. 15.10. Dosed volume of 100 nL and controlled current as a function of time. 645

design and implementation, be easily integrated within silicon or polymer based microfluidic systems for the accurate dispensing of liquids such as reagents or calibrants.

15.4

THE FLOW-THROUGH POTENTIOMETRIC AND AMPEROMETRIC SENSOR ARRAY

A number of microsensors has been developed for the purpose of being built into microchannels. However, the integration of these sensors can impose some problems if they have to be closely integrated within microfluidic systems. A specific disadvantage of these sensors is that they have mostly a planar construction. Therefore, especially for the microdialysis lab-on-a-chip system, a new sensor concept has been developed. In this concept the microchannel itself is an integral part of the sensor geometry and is formed by a length of tubular semi-permeable membrane. This semi-permeable tubing, similar to the tubing used in the microdialysis probe, forms a tubular frit for the separation of the dialysate to be analyzed and a cavity filled with an appropriate solution containing a specific receptor (ionophore, ion exchanger or enzyme). It results in a stable interface between two phases, allowing only the passage of the small molecules of interest. The major advantage of this geometry is the fact that off the shelf ion-selective cocktails or enzyme solutions can be applied, yielding sensors for a wide variety of analytes. As a consequence no specific immobilization of the receptors is required. In this case the tubing prevents leakage of enzyme out of the cavity into the dialysate, as the pores in the tube only allow passage of small molecules. Based on potentiometric and amperometric detection, sensors for ions such as Na + , K+ , Li+ , H + and for metabolites such as lactate, glucose and glutamate were implemented, respectively. 15.4.1

The flow-through potentiometric sensor array

The flow-through potentiometric sensor is, regarding its geometry, comparable to the liquid membrane ion selective electrode. The layout of the potentiometric flow-through ion selective electrode is shown in Fig. 15.11. The sensor is formed by a half-cell consisting of the sample 646

ion selective filling semi-permeable micro tube

7:~

sample

internal electrolyte N_

-Ag/AgCI electrode Fig. 15.11. Layout of the flow-through potentiometric sensor.

liquid in the tube and the ion selective filling, the contacting electrolyte and the Ag/AgCl electrode in a cavity around the tube. Note that the flow-through sensor is in fact identical to a conventional liquidmembrane ISE, but inside out. Therefore, the operational mechanism is also identical. In the half-cell, two liquid/liquid interfaces are present. The first interface is formed between the sample solution and the lipophilic ion selective (liquid) membrane, whilst the second exists between this membrane and the internal reference solution. It is assumed that the semi-permeable dialysis membrane is inert and only functions to form a stable interface and to limit leakage of the (liquid) membrane to the sample liquid, comparable with the porous plug of a conventional ISE. The outer boundary potential (Eb) is described by the well-known Nernst equation: Eb= E'+ RTln (s ai

ziF

E 0o

=

i-

zF

1im

airm

(15.9)

(15.10)

where R is the universal gas constant (8.31 J mol 1 K- 1 ), F is Faraday's number (9.65 x 10 4 C mol-1 ), zi the valence of ion i, T (K) is the absolute temperature, ai,s and ai,m (mol/L) are the activities of ion i in, respectively, the sample and the membrane phase, whilst /i° s and 1/tm 647

(Jmol-1) represent the standard chemical potential of ion i in, respectively, the sample and membrane phases. The main function of the ion selective species in the membrane (ionophores or ionexchangers) is to provide a constant activity aim of the specific ion in this phase. Since the diffusion potential resulting from a concentration gradient of ions within the membrane (Ed) and the inner reference boundary potential (Er) are considered to be constant and the membrane potential (Em) can be expressed as the sum of the contributing potentials: Em = Er +Ed + Eb

(15.11)

it follows that Em = E ° +

RT ln(ai,s) zjF

(15.12)

where E ° is the reference potential (i.e. Em for ai,s = 1 mol/L). From this equation it can be concluded that for a monovalent ion at the standard temperature of 298 K, the half-cell potential theoretically changes about 59 mV per decade of change in ion activity (for low concentrations, the ion activity can be approximated by the ion concentration). Based on this potential Em, which is measured against a reference electrode also placed in the sample solution (Section 15.4.2), the actual ion activity can be determined. As a proof of principle the sensor body was machined from a piece of Perspex® by conventional precision engineering. To yield a stable Ag/AgCl electrode, the silver surface of the rod was carefully cleaned prior to chemical chlorination by a 1% FeCl 3 solution for about 30 min. Next a length of semi-permeable tubing (300 [Lm outer diameter) adapted from an artificial kidney (regenerated cellulose, Filtral® 6, AN69 HF, Hospal, France) was inserted and glued with epoxy in a drilled horizontal hole. To form liquid type ion selective membranes, off the shelf ion selective cocktails for sodium, potassium, lithium and hydrogen ions, obtained from Fluka, were applied. The performance of various potentiometric sensors, operated in a continuous flow mode, are summarized in Table 15.1. The sodium and lithium sensors showed excellent sensitivities, close to the theoretically expected value of about 59 mV/decade. The sensitivities of the potassium and pH sensors are somewhat lower. The response time of all sensors, except that of the pH sensor, is in the order of 10-15 s, which is sufficiently small for most applications. 648

TABLE 15.1 Potentiometric sensors' performance Sensor

Sensitivity (mV/dec)

Response time (s)

Detection limit log a

Range (mM)

Na K+ Li + H+

59.2 50.5 58.0 52.7

15 10 10 60

<-3.5 <-4 <- 2.5 <-4

0.1-500 0.1-1000 5-100 pH 4-10

15.4.2

The flow-through reference electrode

+

Based on the same generic electrode design, a reference electrode was successfully implemented, enabling the construction of a sensor array with integrated reference electrode. In this case, the sensor cavity as shown in Fig. 15.11 was filled with a 3 M KCl solution only and the tube membrane acts as the frit of a conventional reference electrode. The stability was measured with respect to a conventional reference electrode and appeared to be within 1 mV. Due to the relatively large contact area, provided by the tubular membrane section, the sensors and reference electrode exhibit a very low impedance. All measurements could be performed without applying shielded wires. However, to extend the lifetime of the reference electrode, it is preferable to use a gelled KCl solution and to limit the surface of the (frit)membrane. 15.4.3

The flow-through amperometric sensor

One of the limiting factors in biosensor fabrication is immobilization of the enzyme. This is due to the inhibition of enzyme during the immobilization procedure as well as the limited enzyme loading within a membrane or onto a sensor surface. Moreover, the immobilization methods are in some cases rather elaborate and it is also suggested by Danilich et al., Seo et al., and Almeida et al. [7-9] that enzyme immobilization affects the enzyme activity and stability. The flow-through amperometric sensor is designed as a threeelectrode electrochemical cell in the form of a microreactor, shaped around a semi-permeable membrane tubing, similar to the potentiometric sensors as described in Section 15.4.1. A platinum spiral wire 649

working electrode is wound around the dialysis tube, whereas an Ag/AgCl reference electrode is placed at the bottom of the cell, as is the case in the potentiometric sensors. An additional Pt counter electrode is placed downstream. The layout of the amperometric sensor is shown in Fig. 15.12. Based on this construction, sensors for glucose, lactate and glutamate determination were developed. The microreaction cell was filled with an appropriate aqueous enzyme solution to achieve an oxidation reaction of the analyte, catalyzed by the enzymes: glucose oxidase (GOD), lactate oxidase (LOD) and glutamate oxidase (GltOD), respectively, according to: GOD

Glucose + 02 + H 2 0 -

Gluconolactone + H 2 0 2

(15.13)

LOD

L-Lactate + 02 + H 20 -* Pyruvate + H 2 0 GitOD

L-Glutamate + 02 t

(15.14)

2

Q-Ketoglutarate + NH 3 + H 2 0 2

(15.15)

The electrochemical oxidation of hydrogen peroxide, being the product of all enzymatic reactions, was utilized for the actual measurement, according to: H 2 02 - 02 + 2H + + 2e

(15.16)

The current resulting from this reaction was measured between working and counter electrodes. The platinum working electrode was biased at about 0.65 V versus the Ag/AgCl reference electrode. Pt counter electrode electrode

Pt work electrode

T

HO. A 2A

analyte (substrate)

(

-

I. ~-, If __ reactor with free enzyme

Li W -

-

-

--------I )I 1LI -----

Al

I

HO' , - I 1 | reference 0 electrode

I semi-permeable membrane

Fig. 15.12. Layout of the flow-through amperometric sensor. 650

The sensors were evaluated in both continuous-flow mode as well as in an FIA system (Table 15.2). Due to the fact that the enzyme is present in its native form, the total enzyme activity can easily be changed by addition of an extra amount of the enzyme into the reaction cell. This results in changes to the sensor sensitivity, e.g. the sensitivity for the glucose sensor increased by only 35% when the glucose oxidase activity in the reaction cell was increased by a factor of 3.2. All amperometric sensors show a high sensitivity and a large range covering all physiological ranges [10]. The stability of the sensors was remarkably good: one lactate sensor was used intermittently for about 30 days without any significant reduction in sensitivity. The performance of the amperometric sensors operating in an FIA system is shown in Table 15.3. The experimental values of travel time (tA) and peak width (tB) are somewhat higher than those obtained from the semi-empirical model of Vanderslice et al. [11], in which the detector response has been assumed to be very fast. This is a result of the increased response time of the sensor caused by the fact that the analyte has to diffuse through the semi-permeable membrane into the reaction cell towards the working electrode. Taking into account that the total distance for diffusion is about 100 pm, which gives a diffusion time of about 10 s (diffusion coefficient for lactate = 1.04 x 10 - 9 m 2/s), the experimental values for tA and tB are in agreement with theory. Dynamic response of the lactate sensor operating in an FIA system is shown in Fig. 15.13. The reproducibility of the sensor response is expressed by a relative standard deviation (RSD) for 50 consecutive sample injections.

TABLE 15.2 Summarized amperometric sensors' performance for continuous mode Sensor Glucose Lactate Glutamate

Enzyme activity (U/ml)

Slope (nA/mM)

Corr. coef.

Linear range (mM)

Detection limit (AM)

100 320 80 5

188 297 240-250 76

0.9990 0.9988 0.9995 0.9984

0-26 0-26 0-3 0-5

50 25 15 651

TABLE 15.3 Amperometric sensors' performance, operating in an FIA system Sensor

Peak height (nA/mM)

RSD (%)

Range (mM)

tA. tB experimental (s)

tA, tB theoretical (s)

Glucose Lactate Glutamate

39.2 42.9 8.35

<5 <2 1.5

0-10 0-9.8 0-5

52, 130 50, 120 62, 135

22-44, 55 116 47-94, 51-103 28-56, 70-140

Measurements were performed in 20 mM phosphate buffer at pH 7.3.

THE INTEGRATED MICRODIALYSIS-BASED LAB-ON-A-CHIP

15.5

In previous sections, the results of all separate parts of the microdialysis lab-on-a-chip have been discussed. In particular a microdialysis probe (Section 15.2), a calibration facility (Section 15.3) and a potentiometric and amperometric ion and enzyme sensor (Section 15.4) have been 0 -50

I IIm i I II

-100 2 mM

I I

-150 E -200 5mM

-250

:m~'~~~

-300

I

7.15 m||

-350 9.18 mm

-400 0

1000

2000

3000

4000

5000

6000

time s] Fig. 15.13. Calibration peaks for lactate for the lactate sensor operating in an FIA system. 652

developed separately using different precision micromachining techniques. From the beginning it was clear that an integrated lab-on-a-chip cannot be realized by applying the different technologies that were used for making the various parts because of conflicting process conditions. Therefore it was decided that all system components should be made with one and the same generic technology, namely a stack of a silicon wafer and a glass wafer that are anodically bonded. In this concept, the silicon wafer contains the double lumen tube connections, meander-formed cavities that contain the calibration solutions, a dosing pump chamber, a peristaltic perfusion-liquid pump and various microcavities for potentiometric and amperometric ion and enzyme sensors as well as for the reference electrode. The glass wafer with a thin patterned Pt layer contains all the necessary electrodes and electrical connections for the sensors and the calibration facility. Both wafers are anodically bonded, yielding a hermetically sealed liquid handling system. 15.5.1

The complete integrated microdialysis lab-on-a-chip

The final integrated system design is shown in Fig. 15.14, where all components as described in the previous sections can clearly be distinguished, except for the peristaltic pump. At the place of this micropump a glass tube is temporarily inserted into the input channel (perfusate in) for connecting an external perfusion pump. Perfusion liquid is pumped via this input channel to the microdialysis probe (left) at a highly continuous flow rate in the range of 0.22 uL/min. This fluid leaves the inner glass tube at the probe tip and flows along the semi-permeable membrane through the liquid channels to the sensor area. During the passage in the probe tip small molecular substances can diffuse through the membrane and enrich the perfusion liquid to form a continuous stream of dialysate. In the sensor area the same process takes place through the second semi-permeable membrane in the potentiometric and amperometric ion and enzyme sensor as well as in the reference electrode. Finally the dialysate leaves the chip to a waste container. On the left side of Fig. 15.14 two calibration facilities are also connected to the main liquid channel between the probe and the sensor area, and are used for the calibration of the sensors. In addition, various 653

dialysis tube ID 200 pm

filling hcsles glass tube D 100

vpe

Ln

filfa hoes fpm sensor

perfusate microdialysis probe ==

waste

0 Sili-con

- filling

hol for pump

3-

P electrodes\

0O Ag/AgCl

hoes for sensors

- filling

10 mm Fig. 15.14. Layout of the final chip.

filling holes for the pumps and the sensors can be distinguished. On the Pyrex glass part bonding pads are visible for the connection of the various components with appropriate measuring and controlling electronics. The microdialysis probe The actual microdialysis probe is a double lumen type and consists of an inner silica tube (150 Lm OD) surrounded by a semi-permeable tube adapted from an artificial kidney (regenerated cellulose, MWCO 20 kD, Filtral® 6, AN69 HF, Hospal France, 200 Fm ID). The layout of the integrated probe is slightly modified to suit the adapted generic technology of a glass-silicon wafer stack compared with the selfaligning silicon-silicon structure as described in Section 15.2. The effective volume of the microdialysis probe tip amounts to approximately 13 nL/mm, which equals the sensor volume (100 nL) if the length of the probe is 8 mm. The calibrationfacility The complete design comprises a double calibration facility for two (different) calibrants. Each consists of a meander-shaped etched 654

channel that contains the appropriate calibrant solution. The actual pump, which can dose these calibrants into the central channel between the microdialysis probe and the sensor area, consists again of an etched cavity filled with electrolyte and two gas-generating electrodes. These electrodes are shaped as a pair of interdigitated electrodes providing the possibility of closed-loop dosing by measuring the gas/electrolyte fraction in the pump chamber [12]. The volume of each meandershaped calibrant reservoir is 1500 nL, which is sufficient for about 15 calibration cycles for a single sensor. The sensor area The sensor area (Fig. 15.15) comprises an amperometric sensor, a potentiometric sensor and a reference electrode. For all three sensors a second microdialysis tube (ID = 200 m) is used to separate the dialysate from the internal sensor liquids. Each sensor is 2 mm wide, thus the volume of the sample inside the microdialysis tube in each sensor is approximately 100 nL. The middle sensor in Fig. 15.15 is the potentiometric sensor filled with an ion selective cocktail that is in contact with a KCl solution and an internal Ag/AgC1 electrode. Leaving out the ion selective cocktail results in a reference electrode. By replacing the ion selective cocktail by an enzyme solution and adding a Pt working electrode on the glass nearby the microdialysis tube, the same design has been adopted for use as an amperometric sensor. Pyrex wafer silicon wafer

I idows for Reference elect

ig sensor's cell

Potentiometric sen

filling holes for glue

Amperometric sensor

(a)

over · \·

nlcroiayss ruoe

Fig. 15.15. The sensor area comprising an amperometric sensor, a potentiometric sensor and a reference electrode: (a) top side; (b) bottom side. 655

Advantages of this approach are the generic technology to make sensors with incorporated reference electrodes, capable of measuring a wide variety of analytes for which cocktails are commercially available, and the fact that sub-microliter samples can be measured. In addition, the electrical impedance of the proposed flow-through sensors is small due to the relatively large sensing surface formed by the tube wall. The perfusion liquidpump At present an external syringe pump is used to provide the system with perfusion liquid in a continuous flow regime. In the meantime a perfusion liquid pump has been developed that can be integrated into the system using the same generic technology: etched channels and cavities through which a tube can be led. The cavities contain a temperature- or pH-sensitive hydrogel that can swell and shrink as a function of the applied stimulus. The necessary temperature or pH actuators of Pt electrodes exist, deposited onto the cover plate. Actuation by current pulses results in a controlled swelling or shrinking of the gel in the cavities. With three cavities in a row a peristaltic pump is developed, based on pH-sensitive hydrogels. Figure 15.16 shows the design of such a pump. The tube consists of a silicone rubber fiber of which the wall thickness is etched back to about 20 pm, in order to make it more elastic and easier to close by the swelling hydrogel. First results are promising, but the pump is not yet ready for integration.

Flexible microfiber

tched avity Silico botto wafer

I Pyrex top wafer

I Temperature sensorheater combination

I Hydrogel

Fig. 15.16. Design of the peristaltic micropump. 656

15.5.2

Experimental

Device preparation The complete lab-on-a-chip systems are produced in the cleanroom of the MESA+ Research Institute of the University of Twente, Enschede, The Netherlands. Figure 15.17 shows a realized glass wafer on top of a matching silicon wafer, prior to the anodic bonding process. The original layout as given in Fig. 15.14 can clearly be recognized. All channels and cavities are etched using reactive ion etching (RIE) in a double-sided polished 100 mm silicon wafer. The filling holes for the internal liquids are etched through the wafer in a second step. In the same processing step, several holes are etched for the fixation glue of the microdialysis tubes connected to the anti-creep channels as will be discussed below. All Pt electrodes are deposited on a 3-inch Pyrex glass wafer using standard sputtering, photolithography and wet chemical etching. Then both wafers are anodically bonded and subsequently diced in separate systems. After fixing the system on a small PCB (Fig. 15.18) the silica tube and both microdialysis tubes are introduced into the appropriate channels and fixed with epoxy through the various glue holes. After application of

Fig. 15.17. Photograph of a silicon and a glass wafer just before anodic bonding. 657

Fig. 15.18. The complete lab-on-a chip, mounted on a small PCB, with a double lumen microdialysis probe on the right side and perfusion liquid input connection as well as the dialysate waste output connection on the left side.

a small drop of Ag/AgCl paste on the three Pt electrodes inside the sensor cavities through the windows in the silicon substrate, these windows are closed with a small microscopic cover glass using UV curing glue. Finally the different liquids are injected into the appropriate cavities through the filling holes, which are subsequently closed with epoxy. The anti-creeping system For fixation of the various tube sections epoxy is applied through holes in the silicon. However, this technique worked well during assembly of separate components but appeared to be troublesome with respect to the assembly of a complete system. At some places the epoxy should close the gap between a tube and the surrounding silicon-glass channel, but should not creep into the cavities, nor cover any of the electrodes. Therefore an anti-creep design is developed, consisting of overflow capillaries in the silicon at these critical points. Figure 15.19 shows a photograph of one of the anti-creep subsystems, where the epoxy (glue) is injected through a hole into the silicon just beneath the center of the picture. It fixes the microdialysis tube near the working electrode of an amperometric sensor, whereas the surplus of glue is taken up by the capillary structure. 658

Fig. 15.19. Details of the glue anti-creep system.

Measurement set-up The measurement set-up to characterize the microdialysis system is shown in Fig. 15.20. The perfusion liquid is delivered by a microdialysis pump (CMA 102, CMA Microdialysis AB, Sweden) via the inlet to the actual probe that is immersed in the sample solution. The calibration facilities P1 and P2, containing the specific calibrants, are controlled by the connected PC, while the amperometric sensor is measured with a potentiostat (EG&G Model 263A) in a three-electrode configuration. The potentiometric sensor is connected to a highimpedance instrumentation amplifier. Measurements are recorded on a standard PC. 659

sample -\

microdialysis probe

Fig. 15.20. Measurement set-up for the characterization of the microdialysis lab-on-achip.

15.5.3

Measurements

As an example, the results of experiments with an amperometric sensor and a potentiometric sensor will be discussed. Testing the amperometric sensor was carried out by making it sensitive for glucose by filling the sensor cavity with the enzyme glucose oxidase (GOD) with an activity of 15,500 U/g solid (Sigma). The perfusion solution was a 45 mM phosphate buffer, driven through the system by the external pump with a constant flow rate of 2 [LL/min. All measurements were carried out by immersing the microdialysis probe for 25 s in small sample reservoirs (Fig. 15.20) containing solutions of different concentrations of glucose ranging from 1.25 to 25 mM (Fig. 15.21). This was repeated twice for each sample. Between consecutive samplings, the probe was placed in a 45 mM phosphate buffer solution. 660

_Pn

-50 -70 _

-90

It -110 -130 -150 -170 1200

2200

3200

4200 5200 time /ls Fig. 15.21. Typical FIA mode response of the described glucose sensor.

A similar experiment has been carried out for testing the potentiometric sensor (Fig. 15.22). In this case the sensor was sensitive for potassium, after application of an ionophore cocktail based on valinomycin (60031) from Fluka. The microdialysis probe was immersed in a sample solution of 0.1 mM KC1, while again a perfusion

A

C

0.

!

0

400

800

1200

1600

2000 2400 2800 3200 3600 4000 timels Fig. 15.22. Typical results of the potassium sensor in continuous-flow mode. 661

solution of a 45 mM phosphate buffer at a constant flow rate of 2 pL/min was used. The concentration of the potassium ions in the sample solution was changed by a standard addition method from 0.1 up to 160 mM KCl. The final step from 160 to 31 mM KCl corresponds to dilution of the sample.

15.6

GENERAL CONCLUSIONS

Based on the proof of principle of the different subsystems, as described in Sections 15.2-15.4, the ultimate integrated system, as described in Section 15.5, has been produced. It is clear that this is nearly the same as initially intended and sketched in Fig. 15.1. Only the perfusion liquid pump is missing in the final design, but the chip contains enough space to integrate a small peristaltic pump. Also more sensors can easily be integrated as well as the relevant additional calibration units, although it might also be possible to calibrate with a mixture of all components to be measured, with only the existing two calibration reservoirs. All measurements have been carried out with standard laboratory equipment, which is not especially designed or configured for this typical lab-on-a-chip. In a final product all the necessary functions should be integrated into a single dedicated IC or ASIC. Independent of the original aim of the project to develop a small bedside monitoring system, the different subsystems can also be applied individually, such as the microdosing system and the flow-through sensor array. In addition the microdialysis concept is not limited to biomedical applications, but can also be applied with advantage in other application fields, such as biotechnology and environmental monitoring. More detailed information is given in open literature [12-21].

REFERENCES 1 U. Ungerstedt and C. Pycock, Bull. Schweiz. Akad. Med. Wiss., 30 (1974) 44. 2 P. L6nnroth, P.A. Jansson and U. Smith, Am. J. Physiol., 253 (1987) E228. 3 A.R. Kopf-Sill, Micro Total Analysis Systems, Kluwer, Dordrecht, 2000, ISBN 0-7923-6387-6, p. 233. 4 J. Wissink, A. Prak, M. Wehrmeijer, R. Mateman, Proc. Vol. 2 MICRO.tec 2000-VDE World Microtechnologies Congress, EXPO Hannover, 2000, p. 51. 662

5 Int. Patent 'Microdialysis probe integrated with Si-chip' PCT/NL99/00057, 1999. 6 M. Koch, A.G.R. Evans and A. Brunnschweiler, J. Micromech. Microengng, 8 (1998) 119. 7 M.J. Danilich, D. Gervasio, R.E. Marchant, Ann. Biomed. Engng, 21, 655. 8 H. Seo, K. Itoyama, K. Morimoto, T. Takagishi, M. Oka and T. Hayashi, Eur. Polym. J., 34(7) (1998) 917. 9 N.F. Almeida, E.J. Beckman and M.M. Ataai, Biotechnol. Bioengng, 42 (1993) 1037. 10 H.A. Harper, Review of Physiological Chemistry, Lange Medical Publications, Los Altos, CA, 1975. 11 J.T. Vanderslice, K.K. Stewart, A.G. Rosenfeld, et al., Talanta, 28 (1981) 11. 12 S. B6hm, B. Timmer, W. Olthuis and P. Bergveld, J. Micromech. Microengng., 10 (2000) 498. 13 S. Bohm, W. Olthuis, P. Bergveld, Proc. of the SPIE Conf. on Industrial and Environmental Sensing, Boston, USA, 1999, p. 24. 14 S. B6hm, D. Pijanowska, W. Olthuis and P. Bergveld, Biosen. Bioelectron., 16 (2001) 391. 15 A.J. Sprenkels, W. Olthuis, P. Bergveld, First Annual International IEEEEMBS Special Topic Conference on Microtechnologies in Medicine & Biology, Lyon, France, October 2000, p. 326. 16 S. B6hm, W. Olthuis and P. Bergveld, Sens. Actuat. B, 63 (2000) 201. 17 P. Bergveld, Biomed. Microdev., 2 (2000) 185. 18 S. B6hm, W. Olthuis and P. Bergveld, Mikrochimi. Acta, 134 (2000) 237. 19 S. B6hm, W. Olthuis, P. Bergveld, Proc. of Eurosensors XIII, The Hague, the Netherlands, September 1999, p. 165. 20 S. Bbhm, W. Olthuis, P. Bergveld, Proc. of the Int. Mech. Engng Congress ME'99, MEMS, Nashville, USA, November 1999, p. 391. 21 S. Bohm, W. Olthuis, P. Bergveld, Proc. of Transducers 99, Sandai, Japan, June 1999, p. 880.

663