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Lab-on-a-Chip Systems for Cellular Assays

Bernhard Wolf, Martin Brischwein*, Helmut Grothe, Christoph Stepper, Johann Ressler, Thomas Weyh

Heinz Nixdorf Lehrstuhl für Medizinische Elektronik Technische Universität München Theresienstrasse 90, Geb. N 3 D-80333 München, FRG. * Contact: Dr. Martin Brischwein Tel.: (089) 289 22344 Fax: (089) 289 22950 E-mail: brischwein@tum.de

München, Mai 2004

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Table of Contents

1 Introduction

2 Chip design and Fabrication

3 Cell Culture on Chips and Accessory Devices

4 Detectable Cellular Output Signals

4.1 Cell Metabolism

4.1.1 Extracellular Acidification

4.1.2 Cellular Oxygen Exchange

4.1.3 Miscellanous Metabolic Parameters

4.2 Cell Morphology

4.3 Electrical Patterns

5 Cell Manipulation on Chips

6 Conclusions and Future Prospects

1 Introduction

Animal and plant cells are dynamic, nanostructured microsystems which have

evolved during billions of years. They contain functional subunits located in distinct

compartments which are interconnected by a complex signalling network. Some of

these compartments can be made visible by electron microscopy (Figure 1).

Figure 1: Transmission electron microscopy photograph of an explanted human tumor tissue. Some of the different cellular nano-sized structures with specific functions are visible.

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For a selective sensing of input signals from the outside the cell uses specific

receptor proteins. These transmit and amplify the signals to elecit an immediate

response or to change gene activities in the cell nucleus. Figure 2 shows

schematically parts of the complex intracellular signalling network. A differentiated

eukaryotic cell can perform 103-104 different chemical reactions simultaneously

(Davidson, 2002) in a volume of a few picoliters. In order to better understand the

dynamic properties of this network and its input-output relationship, approaches of

system analysis have to employed adequate for a description of the massively

parallel structure of cellular signal computing (Kraus, 1995; Hartwell, 1999; Dumont,

2001). From these studies we have learned that multiparametric planar microsensor

arrays might be ideal tools for a on-line and real-time acquisition of complex cellular

reaction patterns and for the realization of biohybrid components (Kraus, 1995).

Figure 2: Schematic representation of the complexity of interacting signal transduction and metabolic pathways. Right panel adapted from (Dumont, 2001).

Not only neurons have computing capabilities, but also cell types from different

organs. In a typical cellular reaction pattern, a weak biochemical stimulation of the

cellular signal perception apparatus is amplified by intracellular signalling cascades to

perform an appropriate functional reaction. The functional analogy in data processing

capabilities between a technical neuron and a eukaryotic cell is illustrated in Fig. 3.

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Figure 3: Analogy of the signal transmission in a technical neuron and in a eukaryotic cell. Both calculate an "output signal" by weighting of different incoming signals. The state of the art in functional analysis of living cells is the use of microtestplates in

combination with various optical readout systems (Taylor, 2001). Components such

as optical microscopes, fiber optics or CCD cameras can be used to detect visible,

fluorescent or luminescent signals from cells or tissues. The advent of cellular

engineering to include reporter elements such as green fluorescent protein has

allowed to detect specific cellular signalling events. Although the ability to array and

to screen cells in high-density formats meets the demands of many users, optical

screening seems less suited for long term monitoring of cultured cells and tissues,

mostly due to toxic or phototoxic properties of many dyes. In contrast to

microelectrode structures, optic setups with light sources, light guides and detectors

are less easily integrated into small, portable lab-on-a-chip systems. Besides, no truly

appropriate optical technique is available for studies on cell adhesion, cell metabolic

activity or electric activity.

The response to input signals from the environment, i.e. the cellular "output signal", is

usually accompanied by a change in the rates of energy metabolism, by changes in

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electric activity and/or by cell morphological alterations. To investigate cellular

signalling properties for diagnostic or screening approaches, these changes can be

monitored in real-time and non-invasively by allowing cells and tissues to make

contact with adequate microsensors. This coupling is usually achieved by growing

cells directly on the surface of silicon- or glass-based sensor chips, which are inert

and non-toxic materials and accepted by many cells as substrate for stable adhesion

and growth.

The chips may integrate different types of sensor structures which have, in the case

of electric microsensors, a potentiometric, amperometric or impedimetric principle of

function. The setup may be regarded as a cell-transducer hybrid which can be used

in different applications such as pharmacology, neurobiology, research in cell biology,

medical diagnostic tests, and in cell-based biosensors.

After having produced a sensor chip (which may integrate microfluidic structures or

not), there are additional challenges to meet the development of practical analytical

tools. For assays and screening applications, the sensor chips have to be arranged in

high density arrays. It requires substantial technological efforts in chip-integrated

electronic sensor control and in data pre-evaluation to reduce the number of

necessary electric connections to the external electronic equipment. Chip packaging,

such as the incorporation of chips into a multiwellplate format with all the electric

connections shielded savely against liquids, must be optimized in order to come up

with an inexpensive and convenient device. Microfluidics must be characterised

properly in terms of flow dynamics and physics of the interactions of aqeous samples

with the different surfaces, regardless of what kind of microfluidic system is chosen.

Finally, the importance of computation of final results from sensor raw data, statistical

design and extraction of information from response libraries will increase along with

increasing array densities.

Most of the cited work on lab-on-a-chip concepts for cell sampling, cell trapping and

sorting, cell treatment and cell analysis derives from the past five years. The majority

of assays on intact cells seems to be suited for a short-term analysis of cell

suspensions without extended cell culturing. An exception may be those microfluidic

carrier devices, which allow the culturing, inspection and characterisation of cells on

a microscope stage (Rädler, 2004; Yang, 2002). On the other hand, existing sensor

chips for a dynamic monitoring of cell parameters such as electric activity or

metabolic rates usually do not contain microfluidic components (Gross, 1994; Wolf,

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1998; Cooper, 1999). It appears that only very few lab-on-a-chip systems for a

simultaneous cultivation and monitoring of cells with integrated sensor functions (i.e.

µTAS systems with microbioreactors) have been described so far (Hediger, 2001;

Park, 2003). While in many cases cultured cell populations are examined as a whole,

there are various approaches to manipulate and to characterise single cells which

have been reviewed recently (Anderson, 2004).

The field of application for cell-based bioanalytical systems is as large as the variety

of different cell types to investigate. The source and preparation of cells and tissues

is most important for the quality of the analysis. While primary cell cultures often

provide the greatest approximation to the in-vivo phenotype, they also present

potential sources of variability. The discovery of different types of stem cells and the

possibility to induce differentiation in vitro presents a new opportunity to obtain a

source of cells with less functional defects and alterations than found in ordinary cell

lines. On the other hand, a range of applications requires cellular material freshly

derived from individuals. As example, predictive, individualized drug screening in

oncology is performed on tumor tissue biopsies. It appears that miniaturisation and

the possibility to directly follow up the response profiles of short term cultures in the

course of the drug treatments makes chip technologies particularly suitable for this

kind of cell analysis (since the amount of the available tissue material is very limited)

(Ekelund, 1998; Hafner, 2000; Metzger, 2001; Otto, 2003). Again however, the

preparation of the tissue (e.g. consideration of functional interactions between

malignant cell and stroma cells) and the composition of the test medium (e.g.

presence of serum components) have to be considered carefully.

In cell-based biosensors, cellular recognition and signaling capabilities can also be

exploited to detect a physiological response to possibly hazardous substances. In

this case, cells are used for providing the appropriate receptor proteins and for the

generation and amplification of an output signal which is detectable by a physical

transducer. In fact, one of the most important advantages of using living cells as

biological signal discriminators and amplifiers is their ability to regenerate the signal

receptors and amplifying signaling cascades permanently in a native state. Such

sensors which could be specific for toxins might be used, for example, in deterring

biological warfare weapons or they could serve as a biomonitor of waste water

effluents (Schubnell, 1999; DeBusschere, B.D., 2001; Rudolph, 2001; Pancrazio,

2003). For such devices, additional criteria must be met. Apart from the selection of

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the target cell/organism (which may be genetically manipulated) appropriate for the

specific range of compounds to detect, there are numerous issues to be considered.

These include energy management for small, portable devices, sample preparation

(filtration, bioavailability, sterilization), response time, possible effects of physiological

adaptation by sub-toxic concentrations of chemical compounds, life support systems

for robust and invariable sensitivities and well controlled abiotic factors such as

temperature and light exposition. The question of the minimal detectable

concentration is linked not only to the inherent sensitivity of the cellular target, but

also to the determination of any sources of noise in the system. The probability of

false alarm needs to be minimized by correlation analysis of parallel sensors.

The aim of this contribution is to give a general overview on the state of art in cellular

assays on chips. The emphasis will be on technologies and applications related to

the author`s research objectives. These include the current microsensor strategies

for cellular assays but also approaches for manipulation of cells on chips.

2 Design and Fabrication of Chips for Cell Based Assays

In the following section, a brief description of the basics of sensor chip design and

processing is provided. It is confined to the fabrication of ISFETs, LAPS and various

electrode structures on silicon and glass substrates as performed in the available

cleanroom facilities.

For the silicon chip a layout was developed which includes different types of sensors

for multiparametric readout (Figure 4). Four ISFETs, two pO2-FETs, one

amperometric pO2 sensor, an IDES, a temperature diode and a reference transistor

are processed.

Figure 4: (A) Silicon chip layout. The outer dimensions are 7.5 x 7.5 mm. Four pH-ISFETs, one interdigitated electrode structure (electrode width and distance: 50 µm), one amperometric sensor structure and a temperature diode are integrated. The chip is bonded to a printed wire board (24 x 24 mm) and packaged to form a free circular opening (diameter: 6 mm) for sensors and cell culturing. (B) Packaged chip, on PLCC-68 socket, with fluidic setup.

Most of the process steps are standard MOS. Due to the particular application for cell

culture however, special materials and a different order of process steps must be

applied.

Silicon chip fabrication is started with diffusion of the source and drain areas (n-

channel-MOSFETs). Then dielectric material for the gate area is processed by dry

oxidation and additional deposition of a thin silicon nitride layer (10 nm SiO2, 20-30

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nm Si3N4). This non-standard step is included to obtain better cell adhesion on the

nitride layer. Other materials with different physical parameters (dielectric constant,

proton binding characteristics, surface charge) such as Al2O3 and TaO5 are currently

being tested. After defining the active areas by local oxidation and deposition of a

polysilicon protection layer, a titanium layer of a few nm thickness is deposited to

ensure better adhesion of the following sputtered platinum or palladium metallization

(200-300 nm). The metallization layer yields the electrode structures for IDES, pO2-

FET and amperometric pO2-sensor by a lift-off photoresist step. The use of these

metals implements another non-standard step in the fabrication process but it is

necessary for catalyzing electrochemical reactions and for ensuring biocompatibility.

On the ISFETs two passivation layers (silicon nitride and/or silicon oxide) are

deposited and opened on the gate areas. It is important that these passivation layers

provide good adhesion and sealing against liquids to prevent any leakage current.

The completed silicon chips are mounted and bonded on printed wire boards fitting

into standard PLCC 68 sockets. They are encapsulated using a mold that filled with a

biocompatible epoxy resin. Prior to cell culture, the packaged chips are sterilized

either with 70% ethanol or by autoclaving.

Another silicon sensor for pH measurements is the light-adressable potentiometric

sensor (LAPS). As in ISFET fabrication, all process steps are standard MOS, but

fewer steps are required. The fabrication starts with an n- or p-doped silicon

substrate, on which a dielectric layer (SiO2) is deposited by dry oxidation. For better

proton exchange characteristics at the sensor surface, an additional silicon nitride

layer (Si3N4) can be deposited. For an ohmic contact on the backside of the silicon

substrate, a titanium adhesive layer and a gold metalization layer are sputtered. This

metallization layer is structured by a lift-of photoresist step to get an open area, on

which an IR-LED illuminates the silicon chip.

Like the ISFET, the LAPS is an electrolyte-insulator-semiconductor structure. In the

case of the LAPS, a DC bias voltage is applied to this structure. The width of the

depletion layer which is formed at the insulator-semiconductor interface and thus the

capacitance of this interface is affected by the pH-dependent surface potential. The

variation of the capacitance is read out as a photocurrent, induced by the modulated

light of a infrared light emitting diode.

Potentiometric methods aimed at the measurement of absolute concentrations (e.g.

of protons) are dependent on the availability of a high-quality reference potential.

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Obviously, the integration of Ag/AgCl half cells on chips with electrolyte solutions

saturated with Ag+ ions is problematic and indeed, the advantages of planar,

potentiometric sensors are often attenuated when planar sensor structures have to

be combined with conventional "tube and wire" reference electrodes. Fortunately,

some applications tolerate a certain degree of signal drift, which may be caused by

simplified, so called "pseudo-reference" electrodes (e.g. micro-planar Ag/AgCl

structures, directly immersed into the cell culture medium). An example for the use of

such chips is the monitoring of changes in pH due to cell metabolism within short

time intervals (e.g. some minutes). However, it should be noted that the

determination of absolute values of pH is not the objective of such configurations.

Its excellent optic properties makes glass an attractive substrate material for cell

chips. It provides the option that cells are studied with optic microscopes or other

photonic techniques in addition to measurements with electric sensors. An economic

advantage in comparison to silicon-based sensor technologies are the lower

fabrication costs, especially at small and medium production rates. Using

photosensitive glass materials (e.g. Foturan, Schott), several process steps can be

facilitated. The chip made in the author`s laboratory (Figure 5) includes two

amperometric three-electrode-structures and two IDES. The metal structures are

fabricated using the same platinum lift-off process as described above. Afterwards, a

single passivation layer is deposited (Si3N4, about 500 nm) to prevent electric

shortcuts in liquids. Sensors for pH are either integrated by a glass-silicon hybrid

technology (insertion of ISFET-silicon chips into the glass chip) or by deposition of

proton-sensitive layers of metal oxides such as Al2O3, TaO5 or RuO2. A special

passivation (hard mask) is deposited on the glass prior to etching in order to protect

the surface.

Figure 5: Glass chip layout. The outer dimensions are 24.0 x 33.8 mm, the cell culture area has 12 mm in diameter. Two interdigitated electrode structures and two amperometric electrode structures are integrated. For pH-measurement, a 3.5 x 7.0 mm silicon chip can be inserted to form a glass-silicon hybrid. Alternatively, pH-sensitive metal oxide structures such as ruthenium oxide or iridium oxide can be deposited using thin film technology. Electric connections are made with needle contacts.

Based on such glass chips, the fabrication of chip arrays with densities approaching

those which are necessary for screening purposes can be started. In a first step, 24

IDES were integrated on a common glass plate forming the bottom of a 24 well plate.

The structures were processed on a glass plate covered with an ITO layer (indium tin

oxide, transparent and conductive) by photoresist-masked etching. As the number of

electric pathways and contacts for sensors rapidly increases with the density of such

plates, an electronic readout system was developed which allowed a successive

impedance measurement in each well (Ressler, 2003).

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Figure 6: 24-well plate. In this layout, sensors for electric impedance are combined with optic sensors for pH and oxygen (PreSens GmbH, Regensburg, Germany, blue and red spots). This plate allows functional assays on cell metabolic activity and cell morphologic alterations.

3 Cell Culture on Chips and Microfluidic Systems

Cells require a defined environment in order to survive and to respond reproducibly

to signals from the outside. As a general rule it is desirable to mimic the cell`s

physiological conditions. The major conditions that have to be maintained in culturing

cells are (1) physicochemical properties of the culture medium, (2) temperature, and

(3) sterility.

Physicochemical properties of the cell culture medium include pH, oxygen partial

pressure, osmolarity, and defined concentrations of nutrients. In the course of cell

culturing, metabolic waste products and secreted compounds (e.g. proteins) become

additional constituents. For medium- and long-term assays on chips, a liquid handling

system is necessary to adjust and to maintain the required physico-chemical

parameters and, eventually, drug concentrations of the medium. To ensure sensitive

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measurements of metabolic rates, small, closed microvolumes should surround the

cell cultures: Only when a sufficiently high ratio of cell number to medium volume is

achieved, those rates can be recorded precisely (Owicki, 1994). This brings about an

additional requirement, namely the repeated exchange of cell culture media (with

cycle times of some minutes) while retaining the cells in the microchambers.

Significant advances have been recently made in the development of micro-scale

devices for biomedical purposes. One category out of these "lab-on-a-chip" systems

aims at the fabrication of micro-containments for a miniaturized cell culture. The

surfaces of the resulting cell microarrays can be patterned with molecular structures

with selective cell adhesion properties (Curtis, 2001; Koh, 2003; Weigl, 2003).

The ability to decrease the consumption of valuable reagent compounds and

(primary) cells is obviously the most attractive feature of cellular lab-on-a-chip

devices. Some basic design considerations for microfluidic systems to meet

microenvironmental requirements of cells have been summarized recently (Walker,

2004). In microfluidic systems however, there is a limit to the miniaturisation of

microchannels since beyond a critical size aggregation of cells and debris may clog

up the system. With flows driven by electroosmotic forces, possible side effects on

cellular physiology exerted at high electric field strenghts must be considered, since

electric fields in the range of several hundreds of volts per centimeter are required

(Bousse, 2000; Wheeler, 2003). Another difficulty which often moderates the

advantages of microfluidic systems is the necessary interfacing between microchip

and "macroworld". Samples and media usually have to be introduced manually or

with laboratory pumps. A more practicable way for interfacing cell chips with

"macrofluidic" components, suitable for high throughput purposes, may be the use of

pipetting robots: This allows the exchange of culture media, addition of drug

solutions, the adjustment of microvolumes and the aspiration of cell suspensions in

combination with cell chip arrays (Farinas, 2001).

There are also efforts to stabilize and to conserve mammalian cells on solid

substrates for the preparation of cell-based biosensors with extended shelf lives.

Such efforts may become meaningful for applications such as environmental

monitoring, where robust and portable devices are most important (Bloom, 2001).

Figure 7 (A): Front view of a screening device equipped with six channels (modules), each with cell culture vials on silicon sensor chips connected to a fluid perfusion system. An internet port allows for on line transfer of data. Figure 7 (B): Single modules can be operated independently as portable cell signal analyzers. For remote control applications, sensor data can be transmitted by wireless communication to a mobile phone.

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Figure 8: Experimental setup for a 24-well plate. The glass bottom of the cell culture plate is equipped with an array of 24 sensor units. Each unit consists of a indium-tin-oxide (ITO-) structure for electric impedance measurement and optic sensor spots for pH- and oxygen measurement (read out with optic interfaces at the bottom of the plate). A pipetting robot (not shown) coordinates microvolume adjustment and medium exchange.

4 Detectable Cellular Output Signals

Currently, chip-based sensors are capable of detecting four major cellular output

signals (Figure 9): (1) Changes in the rate of proton extrusion are detected with pH-

microsensors, namely ion sensitive field effect transistors (ISFETs) or light activated

potentiometric sensors (LAPS). (2) Changes in the rate of cellular oxygen exchange

are detected with planar oxygen sensors such as amperometric oxygen sensors or

so called pO2-FETs (an ISFET-electrode combination, see below). (3) Changes in

cell morphology and membrane associated electrical effects are monitored with

impedance sensors such as interdigitated electrode structures (IDES). (4)

Electrophysiologic activity of cells, i.e. rapid changes in the membran potential of

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electrically active cells or currents of ion channels are detected with microelectrode

arrays or with patch-clamp chips.

In most systems, only one parameter is recorded on a given chip. However, there are

efforts to develop multiparametric chips combining several complementary sensors

for a comprehensive understanding of dynamic cellular behavior (Brischwein, 2003;

Lorenzelli, 2003).

Figure 9: Parameters of living cells, which are currently monitored by chip-based sensors

4.1 Cell Metabolism 4.1.1 Extracellular Acidification For the analysis of proton fluxes across the cell membrane pH-ISFETs with gate

regions of about 10 x 100 µm2 (LxW) are used. In such transistors, a low current

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passing through a semiconductor channel between two electrodes ("source and

drain") is controlled by the voltage of a third, non metallic "gate-area” on top of this

channel. In pH-ISFETs, the gate is covered with an isolating material which binds

selectively protons, resulting in a pH-dependent potential (potential between gate and

source connection, UGS). With silicon nitride as gate isolator material, the pH-

sensitivity is between 40 and 55 mV/pH. Due to the pH-sensitivity of the structures

the rate of extracellular acidification of cells growing directly on the sensors can be

detected (Baumann, 1999; Martinoia, 2001). Since the typical noise level of pH

recordings by ISFETs is about 0,5 mV, pH variations of about 0,01 pH can be

detected. Most important for a low noise level is a fluid connection to the external

reference electrode. In a cell culture setting a simple and practicable method is to use

the culture medium itself as the electrolyte solution for a Ag/AgCl- reference

electrode, if precautions are made to avoid an intoxication with silver ions.

Figure 10: A measurement sequence reflecting a cyclic pattern of extracellular acidification of LS 174 T cells (a cell line derived from human colon adenocarcinoma) growing on an ISFET on a silicon sensor chip. Addition of Triton X-100 leaves a residual signal which is subtracted upon raw data evaluation. For evaluation, the slope of the graph (dU/dt) during the flow-off intervals is calculated.

The method of monitoring extracellular acidification rates employs a stop-and-flow

mode of a fluid perfusion: A drug substance is introduced during a first flow intervall,

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followed by a stop interval for acid accumulation. The next flow interval re-

equilibrates the medium through renewed perfusion. A series of such stop-and-flow

cycles affecting the measured UGS of an ISFET is shown in Figure 11. Values were

measured every 6 seconds. For the transformation of ISFET raw data into

quantitative information about extracellular acidification, the slope of the graph

(voltage vs. time) during the stop phase is calculated by linear regression

(Brischwein, 2003).

An alternative chip-based technology to analyse the pH is the LAPS, which has been

employed successfully for microphysiometry. A description of this transducer

technology and of the physical chemistry and cell biology underlying the

measurement of extracellular acidification is given in (Owicki, 1992; Owicki, 1994). In

a comparative study, LAPS and ISFETs showed similar features with respect to pH-

sensitivity and drift (Fanigliulo, 1996). Studies have also been done to evaluate the

feasibility of an imaging of surface potentials (which could be influenced by the cell

metabolic activity) with highly integrated sensors (George, 2000). Since cell cultures

or tissue specimen (e.g. slices of explanted tumor biopsies) are often heterogeneous,

the possibility of resolving measured pH values spatially would improve the assay. A

reasonable setup for obtaining multisite records can be an ISFET array which

monitors pH values directly beneath the cells or a CMOS camera (Stepper, 2003).

An optical technique for monitoring extracellular acidification rates called "Bead

Injection Spectrophotometry" has also been developed (Lähdesmäki, 2000) and

realized as a semi-automated "lab-on-a-valve" system (Erxleben, 2004). This

approach uses microbeads loaded with pH-indicators. The absorbance of the

indicator changes upon pH variations caused by the extracellular acidification of cells

is detected with a spectrophotometer.

4.1.2 Cellular Oxygen Exchange The measurement of dissolved oxygen in a cell biological context is a convenient way

to analyse respiratory activity. Usually amperometric Clark-type sensor electrodes

are employed. The technical conditions necessary for these recordings can be

estimated from the typical rate of oxygen consumption in mammalian cell cultures,

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which is roughly between 10-16 and 10-17 mol/cell*sec (Schlage, 1983; Peterson,

1985; Casciari, 1992). Despite the development of technological solutions for planar

miniaturized amperometric oxygen sensors (Lambrechts, 1992) and despite the

potential advantages of planar, chip-integrated sensors for monitoring in cell cultures,

only a few published reports exist with an application in cell biology (Amano, 1999;

Brischwein, 2003). One of the reasons is probably the poor long term stability of most

of these sensors. Moreover, the fabrication technology, which includes the

deposition of reference electrode and membrane structures, is not compatible with a

standard CMOS process.

Generally, both three electrode systems controlled by a potentiostatic electronic

circuit and two electrode systems are feasible. Using a three electrode system, a net

consumption of oxygen in the vicinity of the cathode and a consumption of metallic

silver in reference electrodes can be avoided. However, this would not be serious

points if the cathodic current is limited to a few nanoamperes and the life time of the

reference electrode is at least equivalent to the life time of the whole chip. Among the

different technological solutions for planar amperometric oxygen sensors tested in

our laboratory, a simplified palladium (Pd) electrode structure without gas permeable

membrane (but with a poly-HEMA hydrogel cover to prevent cells from directly

growing on the electrodes) yielded promising results. As a (pseudo-) reference

electrode, a bare Pd electrode is connected. The constant concentration of chloride

ions in cell culture media is able to sufficiently stabilize the reference electrode

potential. However, such a sensor configuration cannot be used for the determination

of absolute pO2 values but only for the analysis of relative changes within short time

periods. A measurement with cells growing on the bare electrode structures is shown

in Figure 11. The advantage of avoiding additional membrane structures would

clearly be the feasibility of a low cost, CMOS-compatible fabrication of those sensors.

Figure 11: A measurement sequence reflecting a cyclic pattern of oxygen depletion due to cell respiration in a cell culture of LS 174 T cells, which has been grown to confluence directly on an amperometric sensor structure. During the stop intervals of the fluid systems, the cellular oxygen consumption results in decreasing current values. Addition of Triton X-100 disrupts cell membranes and cell respiration ceases. For evaluation of sensor raw data an algorithm simular to that used for ISFET evaluation (see Figure 10) was used.

Another method to analyse cellular oxygen consumption is based on the formation of

OH- ions during oxygen reduction and the corresponding pH shift (Sohn, 1996;

Lehmann, 2001). If the pH-sensitive gate area of an ISFET is surrounded by an

electrode, the ISFET will detect an increase of pH when the palladium- (or platinum-)

electrode catalyzes oxygen reduction at a cathodic potential (-750 mV vs. Ag/AgCl).

In order to prevent deterioration of electrocatalytic activity, intermittent anodic

potentials are applied. In a recent work (Ekelund, 2003) a cytosensor

microphysiometer was modified in order to measure cellular oxygen consumption

rates. However, the voltammetric measurement was not performed with planar

electrodes on the (LAPS-) silicon chip but with nafion-coated platinum wires inserted

into the plunger head of the microphysiometer. Oxygen determination based on

quenched luminescence was used by another group for a respirometric assay

(Alderman, 2004). Since this assay is performed in sealed microchambers with 3 µl

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volume, the sensitivity is reported to exceed considerably the sensitivity of former

optical respiration assays.

4.1.3 Miscellaneous Metabolic Parameters A further detection concept for assaying cellular metabolism is directed to the

measurement of metabolic heat with micro-machined, planar calorimetric sensors.

The transducer principle is based on the Seebeck effect. The employed material

combinations (e.g. p+ polysilicon/aluminium) have a high Seebeck coefficient. The

sensitivity is further increased by the arrangement of a high number of thermocouples

in series to form a thermopile structure and by limiting the heat conduction between

thermopile junctions which is achieved by etching thin membraneous conductors. If a

cell culture is grown on one side, metabolic heat will cause a temperature difference

and thus a measurable voltage difference (Verhaegen, 1999; Johannessen, 2002).

Obviously there are other approaches to employ (amperometric) sensors for central

metabolic species such as glucose for cell based assays (von Woedtke, 2002). Such

efforts resulted in the fabrication of a microfluidic system with an integrated

microbioreactor, which is reported to yield glucose consumption and lactate release

rates of hepatocyte cell lines. For detection, a fiber optic equipment was connected to

the chip and a NAD+-coupled assay was performed (Schulz, 2002). Multianalyte

sensor systems with planar electrochemical glucose and lactate detection have also

been developed, which could be employed in vitro for cell based assays (Petrou,

2003).

4.2. Cell Morphology Further critical parameters in cell culture that can be assessed with sensors on chips

are the number and the growth of viable cells as well as changes in cell adhesion or

cell morphology. In the past, impedance measurements proved to be well applicable

for the analysis of a variety of such parameters, including cell growth, cell attachment

and spreading, cell motility and cell migration, barrier function of confluent layers,

cell-substratum spacing and cell death (Giaever, 1993; Ehret, 1997; Ehret, 1998,

Wegener, 2000 a; Jager, 2002; Luong, 2001; Arndt, 2004; Yang, 2004). In a recent

report it was even shown, that a glucose sensing device with a rather linear response

up to 14 mM glucose could be realized based on an impedance sensor structure

grown with fibroblasts (Tlili, 2003). In all these configurations, adherent cells were

cultured directly on electrode structures and a low-amplitude alternating current in a

frequency range of about 10-20 kHz was applied. The measured electric impedance

is the ratio of a sinusoidal voltage applied to a pair of electrodes to the sinusoidal

component of the current. Unless the system is purely resistive, the impedance is a

complex value, because the current and voltage have different phase angles. The

cytoplasmic membranes of living cells are effective insulators for alternating current

at the frequency used.

The measured impedance values reflect the process of cell spreading/cell adhesion

and subtle rearrangements of the cytoskeleton, which is linked to cell-cell and cell-

matrix junctions. For example, changes in intracellular calcium ion concentration are

known to alter the structure of the cell cytoskeleton, resulting in morphological

changes that can be detected by impedance sensors. Figure 13 illustrates the

experimental setup while Figure 14 shows an exemplary experiment with the

recordings of two sensors (interdigitated electrode structures, IDES) on the same

glass chip, detecting the morphological response of a confluent monolayer of Hela

cells to stimulation of the histamine (H1-) receptor.

Figure 12: Geometry of IDES (left side) and IDES-based impedance measurement on adherent cell cultures (right side). The insulating properties of cell membranes increase the measured electric impedance in the frequency

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range used (10 kHz). If the cell morphology changes upon drug stimulation, this change is usually reflected by a change in impedance.

Figure 13: Effect of histamine on the capacitive component of the IDES-impedance. The glass chip has two identical IDES, both were grown with a confluent monolayer of Hela tumor cells. Although no distinct morphological alteration can be detected with the microscope, a reversible effect occurs, possibly on cell adhesion. The insert diagram shows the biphasic response of impedance in detail. Histamine does not cause measurable effects neither on extracellular acidification nor on cell respiration (data not shown). At the end of the experiment the cells were killed with 0,1% Triton X-100.

As the most simple equivalent model for the system composed of electrode, cell layer

and cell culture medium a circuit with a resistor and a capacitor in parallel can be

selected. The results can be represented either as a pair of Capacitance Cpar and

Resistance Rpar or as a pair of the absolute value of the impedance ΙZΙ and the phase

angle. Although refined equivalent circuits of the cell-electrode interface can be

designed, the value of such models for the interpretation of measured impedance

data is limited and depends on additional experimental information (Wegener, 1999).

A completely different transducer type which has been used for the detection of

morphological changes of adherent cells is a piezoelectric quartz crystal

microbalance: Perturbations of the cytoskeleton of endothelial cell layers are reported

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to cause a shift in resonance frequency and resonance admittance (Wegener, 2000

b; Marx, 2001).

4.3 Electrical Patterns Inspirated by the pioneering work of G. Gross (Gross, 1979), several groups started

to use multi-electrode arrays on planar substrates for the multi-site analysis of activity

patterns of explanted networks of neuron cells or muscle cell cultures (Pine, 1980;

Hämmerle, 1994; Hofer, 1994; Hämmerle, 1994; Blau 1997; Bove, 1998; Stenger,

2001; Stett, 2003 a). For the first time it was feasible to address experimentally

questions such as how the nature of a neuron`s connections with its neighbours

affects its ability to generate action potentials. Arrays with different size, spacing and

geometry of electrodes are meanwhile offered commercially (even arrays with three-

dimensional, protruding electrodes are available). The arrays are produced on glass

using varying photolithographic techniques. Cultures of explanted primary cells as

well as whole brain tissue slices (Egert, 2002) can be placed on the arrays to monitor

spontaneous - or microelectrically stimulated - electrophysiological activity with

remarkable sensitivity. Sampling rates of at least 10 kHz are necessary for a

sufficient time resolution. However, the network of connections between

neighbouring neurons is typically so complex, that it is difficult to assign the inputs of

any given cell. In order to overcome this problem, attempts are made to guide

neurons on the chip surface by patterning organic compounds attracting or repelling

the cells (Stenger, 1998). Microelectrode arrays were also utilized to set up a

differential cell-based biosensor with both genetically engineered and wild-type

mammalian cells of the same type (Aravantis, 2001). Starting with the work of

Fromherz, who succeeded to record electrical activity of a leech neuron by placing it

on top of a silicon field effect transistor (Fromherz, 1991), the use of ISFETs for

electropyhsiological measurements is now emerging (Offenhäuser, 1997;

Offenhäuser, 2001; Besl, 2002). The noise level of measurements, however, seems

to be consistently higher as compared to microelectrodes. More recently, the

fabrication of an array of 16384 sensors on a single silicon chip was published,

allowing records of cellular electric activity with high spatio-temporal resolution

(www.infineon.com/bioscience).

Figure 14: "Neurochip" layout with FETs and microelectrodes. This layout was designed to test both transducers simultaneously with a single neuron network for their ability to detect cellular electric activity.

Figure 15: Living neuron (phase contrast microscopy) near terminal of a cruciform, thin film recording electrode. Indium-tin oxide conductors are 8 µm wide (picture with friendly permission from G.W. Gross).

The method preferred for many electrophysiological experiments is patch clamp. For

many years this technique working with highly sophisticated mechanical setups was

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left up to trained personal. Recently, the invention of planar glass chips (patch clamp

chips) with µm-apertures paved the way for a simplified instrumentation, for

automation and thus for considerable increases of experimental throughput in

electrophysiological screening (Fertig, 2002; Stett, 2003 b; Brueggemann, 2004).

Cells are drawn to the small apertures in glass chips (formed e.g. by laser ablation)

by applying a controlled negative pressure from beneath the glass. This is followed

by the development of a very high shunt resistance ("gigaseal"), the prerequisite for

successfull measurements. Using a more advanced and costly technology,

micronozzles can be structured in silicon substrates for the immobilisation and

electrical characterisation of cells (Lehnert, 2002). The focus of patch-clamp

technology is the characterisation of ion channels. Malfunctions of ion channels are

involved in the molecular pathophysiology of many diseases. These channels are,

therefore, most important targets of pharmaceutical drug screening.

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Parameter Technology Examples, selected References

ISFETs (1) Baumann W H et al (1999): Microelectronic sensor system for microphysiological application on living cells. Sens Act B 55, 77-89 (2) Martinoia S, Rosso N, Grattarola M, Lorenzelli L, Margesin B, Zen M (2001) Development of ISFET array-based microsystems for bioelectrochemical measurements of cell populations. Biosens Bioelectron 16:1043-50

LAPS (1) McConnell, H. M.; Owicki, J. C.; Parce, J. W.; Miller, D. L.; Baxter, G. T.; Wada, H. G.; Pitchford, S. (1992): The cytosensor microphysiometer: biological applications of silicon technology. Science 257, 1906-12 (2) Metzger R et al (2001): Towards in-vitro prediction of an in-vivo cytostatic response of human tumor cells with a fast chemosensitivity assay. Toxicology 166, 97-108

extracellular acidification

optic sensors (1) Erxleben, H.A., Manion, M.K., Hockenbery, D.M., Scampavia, L., Ruzicka, J. (2004): A novel approach for monitoring extracellular acidification rates: based on bead injection spectrophotometry and the lab-on-valve system. The Analyst 129, 205-12

amperometric sensors

(1) Amano Y et al (1998): Measuring Respiration of Cultured Cell with Oxygen Electrode as a Metabolic Indicator for Drug Screening. Hum Cell 12, 3-10 (2) Brischwein, M., Motrescu, E.R., Otto, A.M., Cabala, E., Grothe, H., Wolf, B. (2003) Functional Cellular Assays with Multiparametric Silicon Sensor Chips. Lab on a Chip 3, 234-240

oxygen consumption

optic sensors

(1) Alderman, J., Hynes, J., Floyd, S.M., Krüger, J., O`Connor, R., Papkovsky, D.B. (2004): A low-volume platform for cell-respirometric screening based on quenched-luminescence oxygen sensing. Biosensors and Bioelectronics 19, 1529-35

glucose/ lactate exchange

optic sensors (1) Schulz, C.M., Scampavia, L., Ruzicka, J. (2002): Real-time monitoring of lactate extrusion and glucose consumption of cultured cells using a lab-on-valve system. The Analyst 127, 1583-88

changes in cellular morphology

electric impedance sensors

(1) Ehret R et al (1997): Monitoring of cellular behaviour by impedance measurements on interdigitated electrode structures. Biosensors & Bioelectronics 12, 29-41 (2) Wegener, J et al (2000): Electric Cell-Substrate Impedance Sensing (ECIS) as a Noninvasive Means to Monitor the Kinetics of Cell Spreading to Artificial Surfaces. Exp Cell Res 259, 158-166

patch-clamp chips (1) Fertig, N. et al (2002): Whole Cell Patch Clamp Recording Performed on a Planar Glass Chip. Biophysical Journal 82, 3056-62 (2) Stett, A.; Bucher, V; Burkhardt, C.; Weber, U; Nisch, W. (2003): Patch-clamping of primary cardiac cells with micro-openings in polyimide films. Medical & Biological Engineering & Computing, 41, 233-240

microelectrode arrays

(1) Gross, GW et al (1995): The use of neuronal netwoks on multielectrode arrays as biosensors. Biosensors & Bioelectronics 10, 553-567 (2) Egert, U et al. (1998): A novel organotypic long-term culture of the rat hippocampus on substrate-integrated multielectrode arrays. Brain Res Prot 2, 229-242

electric activity

FET arrays (1) Ingebrandt, S et al. (2001): Cardiomyocyte-transistor-hybrids for sensor application. Biosensors & Bioelectronics 16, 565-570 (2) www.infineon.com/bioscience (3) Stepper, C., Wolf, B., Wiest, J., Loeser, M., Brischwein, M., Grothe, H., Hansch, W., Schmitt-Landsiedel, D.: Technological Pre-Investigations on the Realization of a Local Resolution Microphysiologic Cell Chip for Medical Diagnostics and Pharmaceutical Screening. Proceedings of Sensor 2003 11th International Conference, 13.-15. May 2003, Nürnberg, Germany. Part II, p. 335-338

Table 1: Current methods for cell-based assays on chips and a selection of relevant publications

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5 Cell Manipulation on Chips Engineering advances in microfluidics have allowed the construction of versatile

devices transporting substances through microchannels of a glass or plastic chip by

electroosmosis, capillary forces or pressure. Both bacteria and eukaryotic cells have

been successfully transported in microfluidic systems (Weigl, 2003). The number of

publications describing different types of cell manipulations, including complex

multistep procedures, such as cell separation (Fiedler, 1998; Fu, 1999; Huang,

2002), cell fusion (Stromberg, 2001; Chiu, 2001), electroporation (Huang, 2003), cell

lysis (Gao, 2004, Chiu, 2001), and incubation in different media performed (Yang,

2002) in such devices is rapidly increasing. A recent review discusses the use of

microtechnologies as a very useful tool for cell manipulation and analysis (Anderson,

2003). Microfluidic structures on chips for cell manipulations are frequently combined

with analytical functions in order to realize cellular assays. These assays include both

a chemical analysis of cell constituents after cell lysis (Schilling, 2002; McClain, 2003;

Gao, 2004) and the analysis of intact cells, the latter most often based on labeling

with specific fluorescent dyes and read out with a fluorescence detector (Matsubara,

2004; Yang, 2002).

For cell counting, the principle of cytometry, focussing cells into a single line for

individual analysis has been adapted to microfluidic devices by different groups (see

e.g. Gawad, 2001; McClain, 2001; Fu, 2002; Wolff, 2003). Commercially available

products are already utilized for cell staining and cell analysis on chip systems

(Buhlmann, 2003). Versatile microfabricated flow-cells as a basis for applications in

cell analysis have been described by the group of Vellekoop (Vellekoop, 2003).

Unlike the Coulter Counter principle, which relies on measuring the volume displaced

by passing cells, the polarisation response of cells passing through the channel of an

integrated microfluidic device in an electric field region was reported to give

information about the DNA content of the cell (Sohn, 2000).

A typical application of cell chips is the detection and quantification of viable bacteria

in aqueous solutions with electrode structures. In the present chip setup (Figure 16),

an interdigitated electrode structure is used both to accumulate the bacteria in their

vicinity using dielectrophoretic forces between the cells and the electrodes (prior to

the measurement) and to detect the bacteria by impedance measurement. The grade

of accumulation can be changed by the voltage applied to the sensor electrodes.

Figure 17 shows an accumulation of E. coli at an electrode structure at 2 V and 10

kHz. Varying the voltage causes a change in impedance, which depends also on the

concentration of the bacteria.

Figure 16: E.coli bacteria, stained with a fluorescent dye, are attracted to the electrodes by dielectrophoretic forces. Typical values for voltage and frequency are 1-2 Volts and 10-100 kHz.

There are numerous examples for the use of dielectrophoretic forces in lab-on-a-chip

devices for cell separations and cell trapping (Gray, 2004; Gambari, 2003). Using

various types of electric fields, it is possible to move, separate, fuse, perforate, or

deform cells. Such methods are broadly applied in biotechnology. In contrast to the

effects of weak electric and electromagnetic fields, in these techniques rather strong

fields are used, with an energy input significantly larger than the energy of thermic

noise. For example, dielectrophoretic field cages can be created with electrode

structures for handling of single cells on chips. In addition, such field cages are useful

for a dielectrical characterisation of cells by electrorotation (Gimsa, 2001). To avoid

various drawbacks associated with dielectrophoresis (such as unwanted

electrochemical effects), "electrodeless" devices have been described (Chou, C,

2003; Cummings 2003), creating a constriction and thus a high gradient of an electric

field in a conductive solution. Another non-contact method for cell manipulation is

based on superimposed magnetic AC-fields, acting on cells labelled with magnetic

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31

nanoparticles and independent of any material constants of the extracellular liquid

(Koch, 2004).

For many applications, the use of mathematical methods is necessary in order to

obtain optimal electrode geometries. For example, a glass chip was constructed with

a novel sensor structure for measuring the concentration of bacteria in an aqueous

solution using dielectrophoretic forces. The four sensor electrodes have a sandwich

structure, with two similarly designed electrode layers separated by an isolating layer.

In each layer two similarly modelled electrodes are arranged in two insulated layers.

The electrodes of one single layer are active for the collection mode of the sensor. In

the measuring mode, two of the back-to-back lying electrodes are used to measure

the concentration-dependent impedance. The collecting property can be achieved by

a combination of high electric field strength and a large field gradient. A minimal

capacity is necessary for a high sensitivity of the sensor. The electrode structure was

therefore optimized by three-dimensional numerical field calculation. With this

program the dielectric properties of the glass substrate, the electrodes, and the

aqueous solution were simulated. The result of optimization was an electrode

arrangement consisting of one even and two zigzag electrodes (Figure 17).

Figure 17: Electric field distribution in a distance of 20µm above the electrode surface. These data were used to optimize the electrode geometry.

6 Conclusions and Future Prospects

Micro- and nanotechnologies are rapidly expanding into biomedical applications.

Among these applications, live cell studies are becoming increasingly important not

only in pharmaceutical drug discovery, but also in clinical diagnostics and monitoring

of environmental toxicants and pathogens. There is a considerable demand for

practical and low-cost tools for cell studies in all these areas. Cells are the minimum

functional and communicating unit of any living system and the ultimate target of any

drug. Although cellular signal amplification mechanisms often result in impressive

cellular responses, such "output signals" are not easily detected with conventional

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analytical methods, at least in small cell numbers. A substantial aim of micro- and

nanotechnology is to integrate the "microsystem" of the living cell with technical

microsystems providing the necessary physiological environment, the microfluidic

systems for controlled supply of media and drug solutions, and transducers for a

sensitive recording of subtle changes in cellular behavior. If not only a short-term cell

culture is involved, it will be important for the lab-on-a-chip system to create an in-

vitro cellular microenvironment as close as possible to in-vivo conditions. Groups

working with cell-based assays are frequently facing problems with a gradual cell

culture deterioration in the time course of the studies. Therefore, conditions including

the protection of cell culture media from evaporation, adjustment of the gas

composition and measures against contamination must be guaranteed. The structural

versatility of microfluidic systems should help to approach this demand. Chip devices

integrating microbioreactors, microfluidics and transducer functions are expected to

emerge within the next few years.

Research and development in cell based assays strongly focusses on detection

principles based on fluorescence optics. This is due to a steady improvement of live

cell stains directed to a great variety of cellular targets and goes along with a similar

progress in instrumental technology adapted to a very versatile readout of

fluorescence parameters. On the other hand, the equipment for optical detection

(optic fibers, microscope lenses) sometimes appears to attenuate the advantages of

small and cheap cell chips. Compromises are obviously necessary. Moreover,

photobleaching and (photo-)toxic properties inherent to virtually any dye counteract a

monitoring of cells on a time scale longer than a few minutes.

With microelectric transducers, special attention has to be paid to a practical solution

for electric contacting and to a condensation and pre-processing of the generated

data, particularly with respect to multiparametric sensor chips and high-densitity

plates. To achieve this aim, substantial technological efforts in chip processing are

necessary: One of the most important benefits of CMOS technology is the possibility

of on-chip circuitry for signal amplification, data analysis and sensor self-testing. On-

chip sensor multiplexing is a precondition for the construction of 96- or 384- multiwell-

arrays since the number of necessary electric connections becomes unmanagable.

Although single-parameter assays are reasonable for many purposes, it would

increase the efficacy of cell studies if it were not necessary to combine data from

different assays which might have been performed under slightly different

34

experimental conditions. Multiparametric cell chips would help to reveal various

aspects of cellular events within a single assay, and they would do so in a dynamic

and real-time mode. With glass chips, the combination of electric sensors with optic

sensors, e.g. for pH and oxygen is obvious. If required, even high-resolution light

microscopy providing imaging information can be combined with on-line sensor

monitoring of the cells.

Although miniaturization of cell chips has its advantages by saving valuable (primary)

cells and drug compounds and although the size of single sensor elements on chips

can be as small as about 100 µm2 (e.g. the sensitive gate area of a pH-ISFET or a

microelectrode for the extracellular detection of action potentials), it is mostly not

intended to analyse single cells with electric transducers. It should be emphasized,

that single-cell measurements which do not take into account the typical social

context of cells must sometimes be assessed critically. This fact is highlightened by

neuronal networks on microelectrode arrays, which begin to develop after

explantation and which can only be reasonably assessed as whole functional units.

Nevertheless, a direct analysis of single cells on chips which does not involve long-

term cell culturing is possible and even a direct chemical analysis of single cells (e.g.

cell lysis followed by PCR) is expected to emerge in the next years (Anderson, 2004).

In metabolic assays, efforts dedicated to metabolic imaging using highly integrated

two dimensional chips arrays are only in their beginnings. Generally, the strength of

metabolic (and morphologic) assays is their unspecificity, allowing to detect a great

variety of cellular responses. This is because cell metabolism and cell morphology

itself are closely coupled to the signaling apparatus and can thus be regarded as

first-step signal transducers. For the interpretation of results, a multiparametric and

kinetic analysis may provide first indications on a drug`s mode of action. For

interpretations on a molecular level, methods such as the comparison of different,

transformed cell lines or the use of specifically acting biochemical inhibitors or

antibodies have to be applied. Thus, it is likely that cell-chip systems will find their

place in the screening of pharmaceudatical agents for activity or toxicity, as an

additional source of information about cellular behavior. An example for a

multiparametric record of tumor cells treated with a cytotoxic drug is given in Figure

18.

Figure 18: Effect of chloroacetaldehyde (50 µM) on LS 174 T cell cultures (human colon adenocarcinoma cell line, two cultures have been run in parallel), monitored with pH-, oxygen and impedance sensors. The drug was added twice for one hour to test for reversibility of drug effects. A strong, but reversible effect on cell respiration is observed. Impedance measurements however, reveal that cell death is not the predominant effect (cell death would be accompanied by early morphologic changes and cell attachment, detectable with IDES). At the end of the experiment, the cells were killed with 0,1% Triton X-100.

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