Fabrication of multi-electrode array devices for electrophysiological monitoring of in-vitro cell/tissue cultures

Transcription

1 Fabrication of multi-electrode array devices for electrophysiological monitoring of in-vitro cell/tissue cultures a dissertation by Marc Olivier HEUSCHKEL Diplômé de l Ecole Polytechnique Fédérale de Lausanne Thesis n 2370 Institute of Microsystems EPFL CH Lausanne Switzerland 2001

2 To my family

3 ABSTRACT Considerable progress has been accomplished in the past few decades in the development of multi-electrode arrays for neuronal monitoring under in-vivo and in-vitro conditions. Using microelectronic fabrication technologies, small devices including metallic electrodes (gold, platinum, indium-tin oxide, etc.) on different substrates (silicon, glass, plastics) have been realized and some of them are now commercially available. The precisely controlled dimensions on the micron scale and the small inter electrode spacing of these electrode arrays allow high measurement selectivity and reliability. Results obtained with such multielectrode arrays demonstrate the possibility of long-term stimulation of and recording from cell networks providing a better understanding of cell and network characteristics and functions. This dissertation focuses on the design and fabrication of low-cost planar multi-electrode arrays with two major improvements: the integration of threedimensional electrodes in order to improve recording from acute tissue slice preparations and the integration of surface microchannels for local solution delivery, i.e. local chemical stimulation. Planar platinum and indium-tin oxide multi-electrode arrays were designed and manufactured using photolithographic techniques. Experiments taking advantage of the extracellular electrode array measurement technique were done with organotypic and dissociated rat spinal cord cultures, and rat monolayer cardiomyocyte cultures. The main advantage of protruding electrodes compared to planar electrodes for extracellular measurements consists of tissue preparation penetration allowing improvement of signal recording (larger amplitudes) and stimulation (lower stimulation threshold) even though the outer cell layers of the tissue slice samples were damaged during preparation. On the other hand, a larger electrode area is obtained with 3D electrodes, reducing electrode impedance and the resulting noise level. In order to evaluate improvement due to the use of 3D electrodes instead of planar electrodes, stimulation and recording experiments have been carried out on acute rat hippocampus slices. The introduction of surface microchannels is a key feature for local perfusion and chemical stimulation of tissue without affecting the surrounding tissue or i

4 culture medium. This will increase chemical stimulation selectivity in the cell preparation during experimentation. Another possible application of microchannels is leukocyte (white blood cells) migration characterization. The ability to have active and oriented leukocyte migration represents a decisive protection mechanism for our body in the case of infection. This phenomenon can be studied in small devices composed of two reservoirs connected by a microchannel. Leukocyte migration from one reservoir to the other through the microchannel induced by a chemical signal applied at the other reservoir can be monitored by impedance measurement between two electrodes placed in the channel. ii

5 VERSION ABREGEE Des progrès considérables ont été accomplis durant les dernières décennies dans le développement de réseaux d électrodes pour la stimulation et la mesure de l activité neuronale dans des conditions in-vivo et in-vitro. Des dispositifs composés d électrodes métalliques (or, platine, oxyde d indium et d étain, etc.) sur divers substrats (silicium, verre, plastique) on été réalisés à l aide de technologies de fabrication de la micro-électronique, et certain sont désormais vendu commercialement. Les dimensions contrôlées précisément à l échelle de quelques microns ainsi que le faible espacement entre les électrodes permettent une bonne sélectivité et fiabilité de mesure. Les résultats obtenus à l aide de réseaux d électrodes ont démontrés la possibilité de stimuler et de mesurer des réseaux de cellules pendant une longue durée de temps, permettant une meilleure compréhension des caractéristiques et fonctions de celles-ci. Ce travail se focalise sur la fabrication de réseaux d électrodes planaires et de deux améliorations majeures, c est-à-dire, d électrodes tridimensionnelles pour l étude de tranches de tissus et de l intégration de microcanaux en surface permettant la perfusion et stimulation chimique locale. Des réseaux d électrodes planaires en platine et en oxyde d indium et d étain ont été fabriqués par des techniques de photolithographie. Des expériences tirant parti des avantages de la méthode de mesure extracellulaire à l aide de réseaux d électrodes ont été réalisées sur des cultures organotypiques et dissociées de cellules de la moelle épinière de rat ainsi que sur des monocouches de cardiomyocytes de rat. L avantage principal d électrodes protubérantes comparé à des électrodes planaires est la pénétration de tranches de tissus, améliorant le signal mesuré (amplitudes plus élevées) et la stimulation (seuil de stimulation plus faible), bien que les couches cellulaires extérieures des tranches de tissus ont été endommagées lors de la préparation. De plus, la surface effective des électrodes tridimensionnelles est plus grande ayant pour effet une réduction du bruit des électrodes. Afin d étudier l amélioration amenée par l introduction d électrodes tridimensionnelles, des mesures ont été effectuées sur des tranches aiguës d hippocampe de rat. L introduction de microcanaux en surface est nécessaire pour une perfusion et stimulation chimique locale des tissus sans pour autant toucher le tissu iii

6 environnant ou toute la solution de culture. Ceci permet d améliorer la sélectivité de la stimulation lors d expériences. Une autre possibilité d application de microcanaux est la caractérisation de la migration de leucocytes (globules blancs du sang). La capacité des leucocytes de détecter et de migrer en direction d un site d infection est un mécanisme de protection décisif de notre corps. Ce phénomène peut être étudié à l aide de petits dispositifs composés de deux réservoirs reliés par un microcanal. La migration des leucocytes d un réservoir à l autre, induite par la présence d un signal chimique dans l autre réservoir, peut être quantifiée par la mesure de l impédance entre deux électrodes placées dans le canal. iv

7 TABLE OF CONTENTS Introduction 1 1 Measurement techniques of bio-electrical activity Electrical activity of neurons Neuronal organization The neuron The cell membrane Action potentials Measurement methods of electrical activity In-vivo measurement methods and devices Neural probes Three dimensional shaft electrode arrays Sieve electrodes Cuff electrodes In-vitro measurement methods and devices Cell culturing Cell patterning methods Optical measurement method using potentiometric dyes Glass pipette microelectrodes Multi-electrode arrays (MEA) References 34 2 Multi-electrode array design and fabrication Introduction Multi-electrode array concept and design Multi-electrode array characteristics 48 v

8 2.2.2 The Smart Petri Dish concept Materials Multi-electrode array design External amplification and data acquisition system Fabrication of multi-electrode arrays Transparent multi-electrode arrays (first generation) Second generation of multi-electrode arrays References 63 3 Electrical properties of multi-electrode arrays Electrical MEA model electrode model The electrode/electrolyte interface Parasitic components of the electrode Resulting electrodes characteristics Electrode stimulation properties Experimental electrode characterization Methods and materials Electrode properties function of electrode area Electrode properties versus frequency and material Global noise level of MEAs References 83 4 Investigations using planar multi-electrode arrays Introduction MEA measurement system MEAs for spinal cord cultures Introduction Used multi-electrode array layouts Results: spinal cord slice cultures Results: dissociated spinal cord cultures MEAs for cardiomyiocytes cultures Measurement technique combination 99 vi

9 4.4.2 Analysis of beat rate variability in spontaneously active cardiac cell cultures Structure function relationship Conclusions References Three dimensional multi-electrode arrays The need of three dimensional electrodes Tip-shaped electrodes Bulk wet chemical etching of glass Nature of glass Wet chemical etching in hydro-fluoric acid solutions Substrate preparation before glass etching Characterization of hydro-fluoric acid glass etching Glass microtip results Fabrication of 3D multi-electrode arrays Electrical properties of 3D electrodes Three dimensional versus planar electrodes Planar and 3D MEA electrode recording simulations Experimental verification Biological experiments The rat s hippocampus Preparation of acute hippocampal slices Electrophysiological stimulation and recording method Recordings from rat s hippocampus Stability of recorded signals D versus planar electrode arrays Conclusions References Microchannel fabrication techniques Microfluidic components for MEAs Microchannels 156 vii

10 6.3 Silicon nitride microchannels Process flow of silicon nitride microchannels Fabrication process discussion Compatibility with MEA fabrication technology SU-8 photoepoxy microchannels Process flow of SU-8 photoepoxy microchannels Fabrication process discussion Compatibility with MEA fabrication technology Microchannel connection Leukocyte chemotaxis application Current measurement methods of leukocyte migration Proposed devices Considerations for future devices Summary References 175 Conclusions 179 Appendix 181 Acknowledgments 183 Curriculum Vitae 185 viii

11 INTRODUCTION The understanding of the nervous system and of brain functioning represents one of the most challenging subjects in today s research. Considerable progress of measurement tools has been accomplished in the past 30 years due to the development of microfabrication technologies. More and more efficient measurement techniques of the electrical activity of eletrogenic cells appeared for in-vivo and in-vitro experimentation, allowing to obtain knowledge about basic cell and network characteristics. First, the development of small glass microelectrodes allowed detection of the neuron s intracellular activity with a good signal to noise ratio. However, the glass microelectrode technique is not well suited for long-term recordings and simultaneous multisite monitoring. On the other hand, to get information about complex cell systems like the brain, it is important to have many measurement sites. The introduction of extracellular metallic electrode arrays offered the possibility to monitor large numbers of cells at the same time, connected into a network or not, and for long time periods, thus providing an adequate interface between living cells and electronic circuits. The use of multi-electrode arrays is constantly gaining interest from neurophysiologists because it provides an ideal environment for the study of neural network dynamics and long-term effects of chemical (drugs) and or electrical stimulation. While the potential of multi-electrode arrays was demonstrated, their large scale application requires some improvements. One major drawback of this measurement technique is the overall device cost. As commercially available multi-electrode arrays are actually produced on a single wafer/single device modus, they are very cost effective. Furthermore, the possibility to study acute slice preparations on multielectrode arrays is a key feature in electrophysiology experimentation. In comparison to organotypic cultures (1-2 week cultures), the initial network of the living animal is preserved. However, with acute slice preparations, the multielectrode arrays suffer from low amplitude responses due to damaged cell layers at the border of the slice lying in between the active cells and the measurement electrodes. Moreover, chemical stimulation of tissue preparations can not be made locally when the compound delivery is made from the top, which leads to 1

12 compound diffusion in the culture medium affecting the whole preparation. It results that only global effects induced by chemical stimulation could be observed yet. The objective of the present work was the design and manufacturing of lowcost planar multi-electrode arrays and improvement of the measurement technique by adding two new functions: a) the realization of three dimensional electrodes allowing penetration of slice preparations to improve the signal to noise ratio and single unit activity recording, and b) the integration of microfluidic components, i.e. of surface microchannels allowing local chemical stimulation of tissue samples from the bottom at a measurement site, improving stimulation selectivity in the tissue preparation. However, microchannels can also be used to address other biological research applications. Characterization of leukocyte (white blood cells) migration could be very useful for clinical diagnostic of patients immune system response. It was found that leukocytes migrate following specific chemical gradients to a site of inflammation. This cell migration can be reproduced and studied on a small device composed of two reservoirs connected by a microchannel. Detection of migrating cells can be achieved by impedance measurement between electrode pairs placed at the bottom of the microchannel, allowing to determinate parameters such as how many cells have migrated, the migration speed, etc. The following manuscript is divided into six chapters. The first chapter is focused on the state of the art of neurophysiological measurement techniques and tools of electrical activity in brain tissues under in-vivo and in-vitro conditions. The second chapter is dedicated to the design and fabrication of low-cost planar multi-electrode arrays used for in-vitro stimulation of and recording from cells or tissues in culture. In the third chapter, the electrical characteristics of the multielectrode arrays will be described. Then, chapter four presents results obtained with different multi-electrode array layouts using organotypic and dissociated rat spinal cord cultures, and rat cardiomyocyte monolayer cultures. The fifth chapter introduces the concept of three dimensional electrode arrays, its fabrication and the validation of this concept by experimentation on acute rat hippocampus slice preparations. The sixth chapter is an introduction to new surface microchannel fabrication techniques, based on room temperature PECVD silicon nitride deposition onto a sacrificial photoresist layer and on a multi-layer SU-8 photoepoxy technique, for local perfusion of tissues in culture. A short description of microchannel devices for characterization of leukocyte migration is finally presented. 2

13 1 MEASUREMENT TECHNIQUES OF BIO-ELECTRICAL ACTIVITY The principal eletrogenic cells are neurons from central and peripheral nervous system. Bio-electrical activity recording can be achieved by intracellular measurement and extracellular measurement techniques. However, monitoring of biological activity can be achieved under in-vitro as well as in-vivo conditions. This chapter presents an introduction to bio-electrical events and the state of the art of its main existing in-vivo and in-vitro measurement methods and tools. 1.1 ELECTRICAL ACTIVITY OF NEURONS Multi-cellular animals wouldn t exist without neuronal electrical activity in the brain and in the peripheral nervous system. This intrinsic characteristic of neurons is necessary for the complex information processing of the outside environment sensing and the control of bodily functions. The main questions for a better understanding of how the brain and nervous system work can be answered by measuring different cell and network characteristics. Thus, research at the cellular level (electrical characteristics of neurons, ion gated channels, membrane receptors, toxicology, pharmacology, etc.) and at the network level (information transmission and processing, connectivity in the network, etc.) can be achieved. Other intrinsic characteristics of neurons, as for example the different characteristics of membrane receptors, play also a very important role in neural communication. The best way to discover more about the complex functioning of the neural network demands experimentation on neural response to chemical compounds (drugs, neurotransmitters, etc.) and electrical stimulation. Furthermore, the analysis of neural responses to chemical compounds is a key feature for drug discovery for neuronal diseases. A short introduction to neural activity will be presented below. More information could be found in the related literature [1-6] Neuronal organization The nervous system is composed of two main parts: the central nervous system and the peripheral nervous system. 3

14 1 Measurement techniques of bio-electrical activity The central nervous system is composed of the nervous system components that are located inside bony structures, i.e., the spinal cord and the encephalon including the brain, the cerebellum, and the brain stem. The brain is the regulation center of the body. Information from the peripheral nervous system is analyzed and interpreted in the brain. Finally, it emits signals in response to the obtained information inputs. The peripheral nervous system is divided in the somatic nervous system and the autonomic nervous system. The somatic nervous system innervates the skin, the articulations and the muscles. Its functions are sensing the outside world, and control of the body through muscle activity control. It is composed of afferent sensory neurons which emit signals to the central nervous system about the body status, and of efferent motor neurons which transmit signals from the central nervous system to the periphery, e.g. to muscles. On the other hand, the autonomic nervous system innervates the internal organs and the glands. It controls the inside of the body (heart contractions, arterial pressure, etc.) The neuron Neurons are composed of a cell body also called soma, sensory extensions called dendrites and one axon that propagates information to other neurons. Like all the cells of the organism, a 5nm thick membrane separates the inside of the cell with the outside. The cell body has mostly a spherical shape with a diameter of 20 µm. It is filled with a fluid called the cytosol. Like all animal cells, it encloses several structures such as the nucleus (contains the chromosomes), the endoplasmic reticulum (site of proteins synthesis), the Golgi apparatus, mitochondria (site of cell respiration and the formation of adenosine triphosphate (ATP), which is the energy source of the cell). From the cell body, the dendrites and the axon extend radially. The axon is a structure specialized in information transmission in the nervous system. Its length varies between 1 mm and more than 1 m depending on the type of neurons and is generally much longer than the dendrites. Normally, the axon splits at its extremity in many branches and comes in contact with other neurons on places named synapses. The neuron receives stimulation signals from the terminations on its dendrites. These signals result from stimulation of the sensory part of the dendrites and correspond to information about the body status (pressure, temperature, posture, etc.). 4

15 1.1 Electrical activity of neurons Fig. 1.1 Scheme of a neuron (modified from [6]). Glia cells Around the neurons, glia cells have different protection and regulation functions. Certain glia cells, known as astrocytes, act as support and regulate the extracellular fluid composition. Other glia cells, the oligodendroglia, produce myelin sheaths which wrap around some axons to isolate them. However, the myelin sheath has some discontinuities over short distances where the axon is not isolated, the nodes of Ranvier. The myelin also contributes to accelerate the impulsion propagation along the axon as will be described later on The cell membrane The cell membrane is formed by a double layer of phospholipid molecules which hydrophilic part face the inside and the outside of the cell. The functions of the cell membrane are to isolate the cytosol of the neuron from the extracellular milieu and to contribute to the rest and action potential of the neuron by forming a barrier to water soluble ions and to water itself. However, this membrane is permeable to ions through small membrane ion channels. The cytosol and the extracellular fluid contain ions in a very different concentration as shown in Table 1.1. The most important ions for electrical activity of neurons are potassium (K + ), sodium (Na + ), calcium (Ca 2+ ), and 5

16 1 Measurement techniques of bio-electrical activity Ions Ion concentration in extracellular fluid Ion concentration inside neurons Ratio between the outside and the inside Membrane potential E ion at 37 C K + 5 mm 100 mm 1:20-80 Na mm 15 mm 10:1 62 Ca 2+ 2 mm mm 10000:1 246 Cl mm 13 mm 11.5:1-65 Table 1.1 Intracellular and extracellular ion concentrations. The membrane potential E ion corresponds to the membrane potential that would be obtained if the membrane would be selectively permeable to one specific ion (from [6]). chloride (Cl - ). Small ion channels, which are gated by voltage or chemical components called neurotransmitters, are distributed all over the cell membrane to allow selective permeability to ions. Because of high ion concentration gradients through the cell membrane, a net movement of charges occurs by diffusion (K + ion cross the cell membrane to the outside). On the other hand, no negative ions can cross the cell membrane. The inside of the cell becomes more and more negative and a voltage potential appears between both sides of the membrane due to K + flux to the extracellular fluid. It results an electrical force that tries to keep the K + ions inside the cell. An equilibrium is reached when the diffusion and electrical force are equal and opposite, and when no more K + ions leave the cell. The membrane is essentially permeable to potassium and sodium but the conductivity of potassium channels is 40 times higher then for sodium channels. When computing the membrane potential using the Nerst and the Goldman equations, it yields a membrane potential equilibrium of -65 mv with respect to the extracellular fluid. This equilibrium potential remains as long as the channel conductivity and ion gradients are stable. When perturbations are strong enough to increase the membrane equilibrium to a threshold potential of -40 mv, a membrane depolarization takes places (the potential increases up to 40 mv) and the neuron triggers an action potential, which propagates along the axon to its extremities called synaptic buttons, where it contacts other cells. These potential changes are caused by neurons during either normal activity or are artificially evoked by stimulation. 6

17 1.1 Electrical activity of neurons Fig. 1.2 Sodium/potassium pump which moves ions across the cell membrane against the ion concentration gradient. This pump uses energy obtained by the transformation of ATP in ADP (modified from [6]). Membrane ion channels Membrane ion channels are the assembly of 4 to 6 protein molecules forming a pore crossing the cell membrane. Ionic selectivity is determined by the pore diameter and the nature of the channel s protein composition. Ion channels are selective for one or several specific ions. The gating mechanism is an important property of some ion channels and is the key feature of cellular neurophysiology. Ion channels that are able to be opened and closed in response to the local external environment, more specially in function of membrane voltage or in response to the presence of specific neurotransmitters, are able to control the cell s electrical activity. Ion channels can be in different states: a resting state in which the channel remains closed, an activated state in which the channel is opened, leading to membrane depolarization, and an deactivated state in which a deactivated channel is not only closed but also latched. The deactivated state lasts as long as the membrane is depolarized. After membrane repolarization, it returns to the resting state. The sodium/potassium pump The ion concentration gradient between the inside and the outside of the cell is maintained by the active action of a membrane protein acting as a sodium/ potassium pump. It moves ions against the concentration gradients inside (K + ) and outside (Na + ) the cell. During this process, the amount of charge remains stable inside and outside of the cell. However, the ion pump consumes energy which is obtained by the transformation of adrenosine triphosphate (ATP), 7

18 1 Measurement techniques of bio-electrical activity Fig. 1.3 Schematic plot of an action potential event. The cell membrane is first depolarized by a sodium ion entrance into the cell. Then, the cell membrane repolarizes due to a potassium ion exit from the cell inside to the outside environment. Finally, a hyperpolarization of the cell membrane avoids generation of a second action potential during a refractory period of a few ms (modified from [1]). produced inside the cell by using oxygen and food derivatives in the mitochondria, into adrenosine diphosphate (ATP) Action potentials Neurons are stimulated by the dendrites and synapses formed with other presynaptic neurons. These stimulations are integrated at the soma. When a sufficient membrane potential threshold is reached, an electrical response, in the form of an action potential, is generated and conducted through the axon to its extremities, known as the synaptic buttons. At the synapse, the signal is then transmitted to other neurons by using chemical messengers, the neurotransmitters. Action potential description When a cell is stimulated up to its threshold stimulation potential, the sodium channels open and sodium ions enter into the cell, depolarizing the cell membrane. Due to the instantaneous closure of the sodium channels after membrane depolarization, the potential rise is stopped at a potential between 30 8

19 1.1 Electrical activity of neurons mv and 40 mv. In order to compensate the membrane depolarization, potassium ions exit the cell through potassium channels. On the other hand, potassium channels do not inactivate spontaneously; they remain open as long as the membrane is depolarized and can close only if the membrane potential returns to its resting level. As potassium exits and no more sodium enters the cell, the membrane potential goes back down to the resting level. However, The membrane repolarization do not stops exactly at resting potential. A membrane hyperpolarization occurs until the potassium ion flux outside the cell stops, which induces a refractory period of time in which no new action potential can be activated. Furthermore, it is sometimes difficult for a neuron to initiate a new action potential in the several milliseconds following the refractory period. This time period is called the relative refractory period. The membrane depolarization and repolarization take about 1 ms. The global refractory period lasts several ms. This leads to a maximum activity frequency of about 100 action potentials per second. Action potential propagation To transfer the information to other neurons, action potentials propagate along the axon of a neuron. When the axon is sufficiently depolarized to reach the stimulation threshold, potential dependent sodium channels open and the action potential is initiated. The influx of positive charges depolarizes the next membrane segment in order to reach the stimulation threshold and initiate a new action potential in this axon segment. The action potential propagates this way along the axon to its extremities. The action potential can only travel along the axon to its synaptic terminations in one direction because the membrane locate next behind the action potential became refractory. On the other hand, the action potential propagation is different when the axon is myelinated. The myelin contributes to a faster signal propagation because it facilitates the current conduction between Ranvier nodes. This type of conduction is called saltatory conduction. At the Ranvier nodes, there is a high concentration of voltage dependent sodium channels. When the action potential reaches a Ranvier node, a subsequent action potential is initiated at its location. However, between two Ranvier nodes (a distance between 0.2 and 2 mm), no action potential is initiated. The action potential is thus jumping from one Ranvier node to the next one. When the action potential reaches a synaptic button, the postsynaptic neuron is stimulated by synaptic transmission. 9

20 1 Measurement techniques of bio-electrical activity Fig. 1.4 Schematic of action potential propagation to the right side in a uniform membrane (top) and a myelinated membrane (bottom). When the axon is myelinated, there is a faster action potential conduction due to saltatory conduction. Synaptic transmission At the synapse, the membranes of the pre- and post-synaptic neurons are separated by a distance between 20 nm and 50 nm. This space is filled with a matrix of extracellular proteins which fix both neurons together. In the axon termination, the synaptic button contains several small synaptic vesicles (diameter of 50 nm) filled with neurotransmitters. The neurotransmitters are chemical agents based on amino acids, amines or peptides, allowing communication with the post-synaptic neuron. When an action potential reaches the synaptic button, voltage dependent calcium channels are opened and calcium ions enter into the axon termination. The calcium concentration elevation is a signal for vesicles to release its neurotransmitters by exocytosis. The released neurotransmitters act on the postsynaptic neuron, attaching themselves to specific receptors located on the extracellular part of receptor channels. Although there are more than 100 different receptors, they can be classified into two families: the receptor channels, and the G protein coupled receptors. Receptor channels are channels formed of five subunits defining a pore in the cell membrane. When a neurotransmitter fixes to one of its specific extracellular receptor parts, the pore is opened by torsion of its subunits. The receptor channels have an ion selectivity. Depending on which ion can pass through the channel, 10

21 1.1 Electrical activity of neurons Fig. 1.5 Arrival of an action potential at an axon termination initiating the delivery of neurotransmitters. Neurotransmitters fix onto receptor channels in the postsynaptic neuron and initiate its opening (middle). The consequence of ion entrance is a depolarization (bottom left) or hyperpolarization (bottom right) of the membrane potential called excitatory postsynaptic potential (EPSP) and inhibitory postsynaptic potential (IPSP), respectively (modified from [6]). the membrane of the postsynaptic neuron can be depolarized (Na + ions) or hyperpolarized (Cl - ions). So the effect can be excitatory or inhibitory, respectively. The depolarization potential induced on the postsynaptic neuron by 11

22 1 Measurement techniques of bio-electrical activity a presynaptic neurotransmitter is called the excitatory postsynaptic potential (EPSP). The hyperpolarization potential induced on the postsynaptic neuron by a presynaptic neurotransmitter is called the inhibitory postsynaptic potential (IPSP). G protein-coupled receptors induce slower and more varieties and longlasting post-synaptic effects. The neurotransmitters released by the presynaptic neuron bind on receptors, which activate G proteins (small protein molecules moving freely at the inner side of the membrane), which in turn activate ion channels or enzymes synthesizing special molecules called second messengers. These G protein-coupled receptors induce many metabolic effects as well as changes in the properties of some ion channels. 1.2 MEASUREMENT METHODS OF ELECTRICAL ACTIVITY The measurement methods used to record the potential changes generated by neuronal activity are based on intracellular or extracellular electrode recording configurations. For the intracellular recording method, a fine glass pipette filled with a saline solution connected to a metallic electrode is inserted into a cell and a recording is made versus a reference electrode locate outside the cell. This technique allows measurements of potentials up to 100 mv. However, the main disadvantage is that the pipette damages the cell, limiting the duration of the experiment. Another disadvantage of the intracellular measurement technique is that only a few cells (up to 5) can be monitored simultaneously. In the extracellular measurement configuration, the cell being monitored is placed close to or on a metallic electrode, which records the electrical field around the cell. Lower signal amplitudes are obtained in this recording configuration (10 µv to several mv) compared to an intracellular configuration. Furthermore, the signal amplitude depends on the distance between the cell and the electrode, and the electrical characteristics of the electrodes. However, the main advantage of this measurement technique is that cells are not influenced, allowing long-term experimentation. Moreover, using an array of extracellular electrodes, many recording sites can be monitored at the same time, allowing some insight into complex cell networks. Another way of electrical activity detection is the monitoring using potentialsensitive dyes (chemicals added in the culture medium) through an optical measurement system. The main advantage of this measurement method is the possibility to monitor the entire cell preparation simultaneously. However, its main disadvantage is the dye phototoxicity, which limits experiment duration. 12

23 1.2 Measurement methods of electrical activity Fig. 1.6 Simultaneous intracellular (top trace) and extracellular (bottom trace) potential recording obtained from a MCC Aplysia neuron 24 hours after plating onto a multi-electrode array. An intracellular electrode was used to simultaneously stimulate the cell and to record the resulting action potential. The extracellular electrode was a 12 µm wide platinum black electrode (from [7]). On the other hand, there are two basically different approaches of activity recording: the in-vivo measurements in which recording and stimulation electrodes are implanted in a living animal, and in-vitro measurements in which animals are sacrificed and tissue or cell samples are removed and placed in a culture environment, recordings being carried out by intracellular and or extracellular procedures. In both cases, to get high selectivity measurements of electrical events produced by living cells, the devices should have dimensions close to cell dimensions (10 to 40 micrometers). Most measurement tools used today are small wire electrodes, glass pipette electrodes, and metallic microelectrodes. These electrodes should have the following characteristics: The materials used for the electrode devices (substrates, electrodes and insulation layers) should present no toxic effects. It is very important that the materials remain inert during experiments, otherwise, the measured data would be distorted. 13

24 1 Measurement techniques of bio-electrical activity The electrical characteristics of the electrodes: the electrical impedance of extracellular electrodes should be as low as possible to minimize the noise level, thus improving signal to noise ratio of the obtained data; on the other hand, the electrodes should be suitable for current stimulation of the tissue samples. For glass pipette electrodes, the sealing resistance between the electrode and a cell should be at least about 500 MΩ in order to avoid current leakage. The connection to outside pre-amplification and amplification electronics should be as short as possible to avoid supplementary noise. For in-vivo measurements, the connection to the outside of the body is one of the key features to take care of. The two following sections will introduce the existing measurement devices for both in-vivo and in-vitro experimentation. 1.3 IN-VIVO MEASUREMENT METHODS AND DEVICES Many different types of neural interfaces are used in research and clinical practice for recording and stimulation of neuronal tissues in-vivo. The main applications are measurements from the brain, the central nervous system, nerve regeneration and, cochlear and retina prostheses. The main property of all these electrode types is that they are all implanted by surgery. This requests perfect biocompatibility of used materials for long-term insertion, stimulation and recording. Otherwise, the devices could be rejected from the body and no monitoring should be possible. A non exhaustive list of different types of in-vivo neural interfaces described in literature is as follows: a) wire bundle and needle electrodes [8, 9], b) surface electrodes [10, 11] and shaft electrodes (also called Neural Probes ) [12, 13], c) three dimensional shaft electrode arrays [14, 15], d) sieve electrodes for regeneration of damaged nerves [16, 17], e) cuff [18, 19] and helix [20] electrodes surrounding nerves, and finally f) cochlear [20] and retina [21, 22] implantable electrode arrays Neural probes A goal of many neurophysiological studies is to make long-term connections with the delicate tissues of the central nervous system to increase understanding of these structure functions as they relate to behavior. Such recording capabilities would be especially helpful in the study of motor responses and sensory motor integration. Moreover, the knowledge gained through such long-term 14

25 1.3 In-vivo measurement methods and devices Fig. 1.7 On the left, diagram of a Neural Probe (from [12]). Multiple electrodes are fabricated on a small shaft, which will be implanted into cortex for neural activity monitoring. On the right, picture of a 10 electrodes neural probe on a silicon substrate (from [23]). Fig. 1.8 On the left, diagram of a neural probe connection technique to external signal amplification and data acquisition using a printed circuit board (from [24]). On the right, improvement of neural probe connection to the outside by using a flexible silicon rubber cable (from [25]). connections could be exploited to enhance rehabilitation strategies to overcome the devastating effects of central nervous system injury or disease. For example, long-term connections to the central nervous system could provide signals suitable for controlling prostheses or other external devices. Measurement principle Multiple electrodes fabricated on a small sharp ended shaft are inserted into cerebral cortex parts like visual cortex [26], auditory cortex [27] or somatosensory cortex [28] in order to stimulate and record neural activity. The 15

26 1 Measurement techniques of bio-electrical activity experimental procedure is composed of probe implantation, long-term monitoring, and finally histological verification of measured tissue [29]. Unfortunately, this technique suffers from some weaknesses like the connection between these shafts and the outside world [25, 30], the delicate procedure of insertion into neural tissue without causing major tissue damage [31, 32], and material biocompatibility. However, neural activity could be measured with signal amplitudes varying from a few microvolts to several millivolts depending on the type of electrodes, implantation procedure, and type of tissue. Devices and characteristics Historically, most microelectrode arrays have been fabricated from bundles of fine wires. However, the major limitations are the difficulty to produce such electrode arrays with reproducible mechanical and electrical properties, and to maintain it stable during implantation. Thus, the use of a more rigid device was desirable. The first so-called Neural Probe penetration micro-electrodes, which contained several measurement sites arrayed on a thin shaft, were designed about 30 years ago to be inserted perpendicularly into the cortex for neural activity monitoring [33, 34]. Since then, neural probes were manufactured on glass [26], silicon [12, 35], and flexible material [13, 36]. CMOS multiplexing and amplification electronics were integrated onto neural probes in order to simplify signal output by reducing the number of cables [37, 38]. Furthermore, first attempts to integrate micro fluidic channels on the neural probes were successful [39] Three dimensional shaft electrode arrays One limitation of shaft electrode arrays is that the measurement electrodes are only displaced in a line along the shaft. Thus, implantation of a single probe into cerebral cortex will only capture activity patterns in function of electrode depth. Another weakness of this measurement configuration is that brain motion occurs during experimentation, changing the cell locations with regard to the electrodes. However, to better understand cortical signal treatment, three dimensional probes, which allow examination of spacial distribution of cortical sensory activity, should be implanted instead of single shaft neural probes to get as much informations as possible on neural connectivity. Devices and characteristics The Utah Intracortical Electrode Array (UIEA) is one approach to obtain information in parallel from multiple sites in the cortex (see Fig. 1.9). This device is composed of a 10x10 matrix of tip-shaped electrodes (one tip is one electrode) 16

27 1.3 In-vivo measurement methods and devices Fig. 1.9 Picture of the 100 electrodes Utah Intracortical Electrode Array (from [40]). The different electrodes are separated by glass as represented on the right picture (from [15]). Even this device is called a 3D electrode array, the measurement area is defined in a plane because the height of all electrodes is the same. Fig On the left, diagram of two dimensional shaft electrode array assembly technique: shaft electrodes are fixed onto a substrate and electrical connections are achieved for signal processing and output. On the right, SEM picture of a 3D shaft electrode array composed of an assembly of 2D shaft electrode arrays (from [14]). with a height up to 1.5 mm realized from a silicon substrate. The electrode tips are fabricated by dicing kerfs into a thick silicon wafer, which is then etched to obtain sharp tip extremities. The tips are then coated with gold and polyimide as electrode and insulation layer material. The electrodes were first separated with a pn junction [41] that was later replaced by a glass layer [15, 42] because of its 17

28 1 Measurement techniques of bio-electrical activity better insulation properties. Unfortunately, a limitation of the UIEA is that all electrodes are located at the same height, which means that measurements can only be achieved in a plane in the cortex. Based on a similar fabrication technology, a real 3D electrode array was obtained by dicing at different heights into a silicon substrate [43, 44]. The main limitations of all these devices are the implantation into the cortex and the generated tissue damage [45]. However, long-term measurements could be achieved over more than 6 months with signal amplitudes in the range of 100 µv [46, 47]. Another approach is to use several 2D shaft probes disposed one next to the other. 3D electrode arrays can be realized by assembling 2D neural probes on a rigid support [14] as shown in Fig With this method, more complex electrode arrays with up to 128 shanks were fabricated [48]. To allow easy signal output, CMOS signal processing and multiplexing circuitry has been integrated onto the shanks. Recent experiments driven on guinea pig cortex showed response amplitudes up to 500 µv, which is an improvement compared to other 3D electrode arrays [49] Sieve electrodes Sieve electrodes are used for stimulation of and recording from regenerating peripheral nerves. They are confined to applications of nerve stumps and injured peripheral neural axons that have lost the interconnection to the effector organ, e.g. a muscle. The major advantages of electrode implantation at peripheral nerve fracture are the possibility to test the functionality of small groups of regenerated axons and to learn how to control limb (in particular hand) prostheses. After interruption of an axon due to nerve injury, the nerve cells respond in a predictable fashion [50]. First changes are noted in the cell body as early as several hours after the injury. It can be noted that the cell nucleus migrates from the center of the cell to the periphery, and substance, that will be transported to the injury level are manufactured in the cell body. Furthermore, metabolic changes occur that help recovery of the injured axon. Nerve regeneration electrode principle After peripheral nerve severing, a rigid or flexible electrode array is implanted in between the two nerve stumps as shown in Fig for stimulation and recording of neural activity. These electrode arrays include many holes through which the cut axons are allowed to regenerate, which can be stimulated by electrical stimulation. 18

29 1.3 In-vivo measurement methods and devices Fig Diagram of nerve regeneration electrode principle (from [22]). The devices are placed between two cut nerve stumps. Regenerating axons are allowed to growth through holes manufactured into the sieve electrode device to reconnect to the other nerve stump. The device substrate can be rigid (left) or flexible (right). Fig Pictures of a silicon based (left, from [16]) and a flexible polyimide based (right) sieve electrode array. Devices and characteristics Most sieve electrode arrays are manufactured onto silicon substrate to generate a rigid interface in between two nerve stumps [16, 17, 51-53]. However, some sieve electrodes were also made of polyimide substrate [54, 55], thus generating more flexible devices. Holes with diameters varying from a few micrometers up to 100 µm allow reconnection of severed axons through the electrode array chip. Some metallic electrodes are patterned around a few holes 19

30 1 Measurement techniques of bio-electrical activity for electrical stimulation and recording of selective regenerated axons in order to inspect the axon regeneration and to learn how to control future prostheses. This electrode array is placed in between two silicone tubes used as nerve regeneration guide. Experiments made 2-3 months after sieve electrode implantation showed a partial recovery of muscle force after nerve injury, suggesting that axons regenerated through the sieve electrode devices and made contact with the other nerve stump [16, 52, 55]. Mechanically evoked action potentials with amplitudes up to 200 µv could be measured with sieve electrodes [51, 52, 56, 57]. It was shown in a recent study about nerve regeneration through sieve electrodes [58] that small hole diameters of 30 µm allow a higher regeneration yield than larger holes of 90 µm for the same device transparency due to less axon deflection at the holes. On the other hand, biological problems with glia proliferation and neuron survival are still remaining. However, It can be expected that future sieve electrodes will allow better nerve regeneration yield than those achieved yet Cuff electrodes For the same application than sieve electrodes, cuff electrodes are used to establish non-invasive interfaces to peripheral nerve bundles that lack electrical activation as, for example, a consequence of spinal cord injury. Cuff electrode measurement principle Cuff electrodes can record an electric signal generated by the superposition of single fiber action potentials events propagating in the nerve. Extracellular current flow generated by traveling action potentials is restricted and confined by the insulating cuff, which is placed around the nerve. One or more electrode contacts along the inner side of the cuff are used to record potential differences due to longitudinal resistive potential drops. The geometry and configuration of cuff electrodes have a filtering effect on the waveform of each recorded action potential. Larger signal amplitudes can be achieved with the use of subthreshold anodic currents between the longitudinal cuff electrodes slowing down the action potential propagation, which is analogous to an increase of cuff electrode length [59]. It results a longer intercontact delay, hence, larger signal outputs improving measurement signal to noise ratio. Finally, the measured signals have an approximate Gaussian statistic distribution [60]. Using an autocorrelation function determined by the shape of single action potentials, it is possible to detect and classify the active fibers in the 20

31 1.3 In-vivo measurement methods and devices Fig On the left, split cylinder cuff design principle: two pieces of biocompatible isolation are wrapped around the nerve and sutured shut. On the right, spiral cuff electrode design principle (from [19]). Fig On the left, single electrode configuration of spiral cuff electrode (from [19]). On the right, multiple electrode configuration cuff electrode (including connector and cable) for fascicle selective stimulation. nerve composing the measured signal. This allows control of selective muscle stimulation with neural cuff electrodes prostheses. Cuff electrodes are used for electrical stimulation of nerve trunk regions and therefore multiple muscles. For optimal stimulation, individual electrode contacts should have a sufficient area made of suitable metal to avoid reversible electrochemical processes during stimulation. The electrode configuration is also of importance. Multiple electrodes with longitudinal separation similar to space constant of target nerve fibers are needed for nerve stimulation [18]. Multiple electrodes placed uniformly around the nerve will form a circumferential contact allowing muscle activation by selective nerve fascicles (bundles of myalinated 21

32 1 Measurement techniques of bio-electrical activity axons) stimulation [61]. Selective muscle activation depends on stimulation electrode configurations. Experiments on cat sciatic nerves demonstrated the possibility of selective muscle stimulation using different electrode stimulation configurations with a 12 site cuff electrode [62]. Devices and characteristics First available cuff electrodes designs were based on a split cylinder shape characterized by electrodes set within flexible pieces of biocompatible insulation that are wrapped around the nerve and sutured shut to secure them. The major drawback of these cuff electrode designs is nerve damage due to mechanical stress applied on the nerve by the sutured cuff. At the end of the 1980 s, practical spiral cuff electrode designs appeared [19]. Spiral cuff electrodes are not sutured shut and therefore are self sizing around the nerve. This self-conformation to nerve diameter has the advantage of avoiding mechanical stress and thus nerve damage. Experiments carried out on adult cats showed no evidence of nerve damage after a seven month implantation, i.e. axonal degeneration and or abnormal immunological reactions, due to the spiral cuff electrodes. Spiral cuff electrodes are manufactured using silicone rubber sheets in combination with platinum foils, welded to stainless steel wires [62]. The number of channels embedded in the silicone is strongly limited in these devices. A new class of cuff electrodes made of flexible, micro-machined thin-film electrodes embedded into polyimide substrates of various shapes were recently developed [21]. The advantages of these new cuff electrodes are the low cost and reproducibility of micromachining manufacturing and the fabrication of a large number of integrated stimulation sites (up to 18). Recent developments of cuff electrodes allow the integration of multiplexer circuitry reducing the number of necessary interconnection leads to external stimulator with a factor of 3 [63]. Resulting cable reduction should reduce the risk of cable breakage which is one of the main cause of cuff electrode implant failure. 1.4 IN-VITRO MEASUREMENT METHODS AND DEVICES Today, in-vitro experimentation is widely used because of its broad application possibilities on various tissue or cell preparations, like dissociated neurons, brain slices, retina, spinal cord slices, and heart cells. Unlike in-vivo measurement methods, neuronal networks in-vitro show more basic features of natural neuronal networks and cell properties. Since the degree of organization and complexity of network cultures is lower compared to animal 22

33 1.4 In-vitro measurement methods and devices models, in-vitro studies allow an improved analysis and understanding of network properties. On the other hand, the possibility of measuring a piece of cell membrane or single cells gives a better insight into intrinsic cellular characteristics. Tissue or cell samples are removed from sacrificed animals and placed under culture conditions, i.e. in a culture chamber (plastic dish, an electrode array culture chamber, etc.) filled with a nutritive solution and placed under controlled conditions (temperature, ph, etc.) over short or long time periods. When studying characteristics of simple neuronal networks, control of the cell organization is of importance. Geometrically defined neuronal networks can be obtained by cell patterning methods promoting or inhibiting cell adhesion. Unfortunately, such generated networks will show characteristics that are inevitably not the same than those in animals or man. At the single cell level, in-vitro experimentation allows research of more basic characteristics of cell function by studying cell response to different compounds having effects on membrane ion channels or cell receptors. Different in-vitro measurement methods of neuronal activity like optical monitoring with voltage sensitive dyes, intracellular recording with glass micro electrodes and extra cellular recording with multi-electrode arrays can be used for in-vitro experimentation. Each method has its advantages and drawbacks that limits its application. The next sections will introduce cell culturing methods, cell patterning possibilities and the principal in-vitro measurement techniques Cell culturing Cell culturing is important for later cell or tissue experimentation. Indeed, the cell environment will play a major role in cell survival, thus cell activity and properties. The culture conditions are specific to the cell species, age, dissection and dissociation technique, the number of cells and the culture media. Thus many different cell culture techniques have appeared over the last few decades. The two most common ways of cell culturing are the culturing of dissociated cells and the culturing of tissue slices called explant or organotypic cultures [64, 65]. Dissociated cell cultures These cultures are prepared from suspensions of individual cells obtained by dissociation of embryonic neuronal tissue. When plated onto an appropriate substrate, the neurons begin to extend and to form a dense network. Under favorable conditions, such cultures can be maintained for months, during which time the cells acquire most of the properties of mature neurons. 23

34 1 Measurement techniques of bio-electrical activity Fig On the left, dissociated chick motor neurons culture made on a polymer surface following a technique described elsewhere [66, 67] (courtesy of Prof. Bertrand s Group, University of Geneva). On the right, hippocampal slice culture after 30 days in-vitro. Bar, 250 µm (from [68]). Organotypic cultures Small tissue pieces of interest are cut using a vibratome and attached to an appropriate substrate usually coated with adhesion molecules that stick to the tissue slice. The need for diffusion of nutrient and oxygen to the center of the tissue slice limits its thickness to about one millimeter. Organotypic cultures can also be maintained for months. Culture chambers A variety of plastic culture dishes including traditional Petri dishes and multi well plates is commercially available. To avoid microbial growth, the dishes are cleaned and sterilized. An elegant approach of cell culturing and monitoring is to use substrateintegrated measurement electrodes. The so called multi-electrode arrays are based of insulated metallic electrodes on a silicon, glass or polymer substrate which builds the bottom part of a culture chamber. It allows extra cellular measurement on cells grown directly onto the measurement electrodes. Other culture chambers specially designed for measurements under specific conditions were used by different scientists for dissociated cell cultures [69, 70] and organotypic cell cultures [71, 72] Cell patterning methods Cell positioning and cell patterning are important topics, especially for dissociated cell cultures, in order to build small neuronal networks following 24

35 1.4 In-vitro measurement methods and devices defined geometries and to allow good multiple extra cellular measurements through embedded electrode arrays. Without cell positioning or patterning, cells are deposited randomly in the cell culture chamber. In case of extracellular measurements with a predefined electrode array in the culture chamber, the probability that one or more cells lie on an electrode site becomes very small (depending on cell number). However, the possibility to create areas with favoured or inhibited cellular adhesion allows patterning of cells in order to build a geometrically defined cell network. Using small cellular networks, it may be possible to study basic network characteristics allowing better understanding of cell communication and information propagation. Thus, cell patterning on measurement electrodes allows improvement of overall measurement conditions. Unfortunately, patterns generated by adhesion molecules and/or based on hydrophobic characteristics of the substrate surface will disappear with time because of cell fragments accumulating on the substrate, which allows cell attachment. The main existing cell patterning approaches are described below. Cell outgrowth and migration following substrate topography Cell alignment and guided cell outgrowth can be achieved using very fine structures at substrate surface [73-75]. Arrays of lines with micrometer and submicrometer scale dimensions are fabricated by photolithography processes. The smallest structures were realized by X-ray printing and etched by reactive ion etching in an oxygen plasma. The response of different cell types to topographically modified surfaces is both alignment to the direction and elongation following the formed lines. Unfortunately, guidance of this type is strongly dependent on cell type and cell-cell interactions. Photolithography patterning methods Photolithography patterning of adhesion promoting molecules seems to be one of the simplest ways of cell patterning to form small neural networks. It has been observed that amines promote, whereas alkanes prevent, cell adhesion. Generation of cell patterns by amines and alkanes by a photolithography process is easy to achieve [65, 76, 77], allowing precise neural network geometries. Silanes are hydrophobic compounds that can also be used to form cell adhesion preventing areas [78]. Another possibility is patterning of silanes by photolithography on substrates followed by exposure to laminin, which adsorbs preferentially to hydrophobic areas, to generate laminin patterns promoting cell adhesion and growth [79]. Many other adhesion promoting proteins can also be patterned by photolithography processes. 25

36 1 Measurement techniques of bio-electrical activity Laminin patterns made by laser ablation One characteristic of laminin is that when irradiated by UV light, it looses its adhesion promoting properties. Thus, one method of cell patterning by laminin can be realized by placing laminin on the whole substrate and then exposing the substrate to UV light through a mask defining the final pattern [80, 81]. Soft-lithography The soft-lithography is a recent fabrication technique allowing replication of patterns using polydimethylsiloxane (PDMS) stamps made by pouring on negative master structures [82, 83]. This replication technique allows easy patterning of substrate surfaces with adhesion proteins like polylysine [84]. Substrate material properties Materials with surface properties preventing any kind of sticking like thin films of amorphous Teflon AF can also be used for cell patterning applications. Indeed, a thin Teflon AF layer on a silicon-silicon dioxide substrate can be patterned by plasma etching through a thin aluminium layer to re-expose the silicon dioxide regions that promote cell adhesion [85, 86]. Other emerging patterning methods Many other different techniques allowing cell patterning appeared in the last decade. The use of dielectrophoretic forces to repulse the cells can generate simple geometries [87]. Simple microfluidic structures can also be used to restrict cell positioning during culture establishment [88]. Finally, the use of perforated substrates with integrated electrodes around the holes allow cell positioning by fluid aspiration [89, 90] and cell attachment by antibodies coated onto the electrodes [91] Optical measurement method using potentiometric dyes Over the past 20 years, it has become possible to detect and to measure rapid changes in membrane potential of neurons by using voltage-sensitive dyes as optical probes from multiple sites [92-95]. Two kinds of voltage-sensitive dye signals have been obtained with bath applied dye. First, there are recordings of the action potentials in individual neurons in invertebrate ganglia, and second, recordings of population signals that have been obtained in vertebrate preparations representing the average of the change in membrane potential in a number of neurons whose light falls onto a single detector. It has been demonstrated that specific types of neurons in the vertebrate spinal cord can be labeled by calcium-sensitive dyes [96, 97]. However, voltagesensitive dyes have potential advantages over calcium-sensitive dyes: they have a better temporal resolution, they allow monitoring of neurons with small calcium 26

37 1.4 In-vitro measurement methods and devices currents, and are superior at detecting small subthreshold events or inhibitory potentials. Measurement principle The measurement principle is based on application of fluorescent voltagesensitive dyes to cells in culture. The cells are illuminated only when recording by a light flash lasting about 100 to 250 ms. A photodiode array is used as light detector and transforms the light signal into an electrical signal [98]. Two kinds of noise are significant in optical measurements: the shot noise, i.e. noise due to statistical fluctuations in the arrival of photons at the photo detector, and the dark noise, i.e. noise due to mechanical vibrations and other optical disturbances in the light path. Optical measurements with signal to noise ratio less than 50 and signal amplitudes up to 4 mv were obtained with large cell bodies of µm in diameter [92]. One major advantage of optical techniques is the possibility of its combination with intracellular or extracellular measurement techniques, i.e. with glass electrodes or transparent device integrated electrodes respectively. However, current optical recording techniques require the use of high numerical aperture lens to obtain the best signal to noise ratio. This tends to limit the field of view over which recordings can be made. Moreover, optical recording perturb the neurons under investigation due to possible pharmacological or photo dynamic effects related to the use of dyes. High-intensity illumination is required which may cause rapid photo bleaching of potentiometric dyes and may lead to problems associated with phototoxicity. This optical measurement technique is unsuitable for long-term recording of neurons because of fast signal decrease over time [99] Glass pipette microelectrodes Introduced in 1940, the use of glass pipette electrodes is the widest applied neural activity measurement method in the neuroscience community. The patch clamp measurement technique developed in the end of the 1970 s by Erwin Neher and Bert Sakmann (Nobel Price 1991) is based on the use of a saline-filled glass pipette with micrometer scale diameter. These electrodes allow puncture of cell membrane with a high sealing resistance and low impedance, enabling direct monitoring of ionic cell membrane channels. A detailed description of patch clamp techniques can be found in literature [ ]. Measurement principle The patch clamp technique referred to whole-cell or patch clamp of a small membrane patch allowing recording of macroscopic whole-cell or microscopic single-channel currents flowing across cell membrane through ion channels. The 27

38 1 Measurement techniques of bio-electrical activity Fig Schematic of a conventional micropipette recording (left) and a tightseal whole cell recording (right) configuration (from [105]). Fig On the left, picture of a single-barreled hard glass pipette tip. The diameter of the tip opening is 0.98 µm (modified from [105]). On the right, picture of an in-vitro brain slice recording chamber with mechanical micromanipulators holding a recording microelectrode (from [100]). main targets are membrane ion channels that can be voltage dependent (Na +, K +, Ca 2+, Cl - channels), receptor activated (neurotransmitters, hormones, exogenous chemical mediators), and second-messenger activated (G proteins, kinases). Other electrical parameters like cell membrane capacitance may be monitored as well. 28

39 1.4 In-vitro measurement methods and devices The measurement of small ionic currents in the picoampere range requires a low noise recording technique. This is achieved by tightly sealing (seal resistance higher than 10 9 Ω) a glass pipette electrode onto the membrane of an intact cell, The current flowing through ion channels enclosed by the glass pipette tip are measured by means of a patch-clamp amplifier. The attached cell may be used to record single channel activity or can be isolated from its environment by withdrawing the pipette from the cell retaining the integrity of the seal. Different measurement configuration can be used like cell-attached recording, whole-cell recording, inside-out recording and outside-out recording. Glass pipette electrodes are fabricated by pulling small glass tubes [106]. The obtained glass pipettes are then filled with saline and mounted onto a patchclamp setup as shown in Fig Today, several patch clamp measurement systems are commercially available. Applications of this technique are classification of ion channel types and characterization of their biophysical properties in terms of conductance, voltage dependence, selectivity, and pharmacological profile. However, this technique is not well adapted for studies of neural networks because it is technically impractical to study several cells simultaneously. In best case, up to 5 cells can be measured at the same time. It is also a rather destructive process, so that the probability of making a number of successive penetrations of a cell to determine the variation of its activity pattern over time is very small. Moreover, penetration of the cell with a glass pipette alters the cell s properties and slowly kills it, limiting duration of investigations within a time scale of 1 hour Multi-electrode arrays (MEA) In contrast to glass pipette electrodes, multi-electrode arrays are used for extracellular measurement of neural activity. Potential changes of cell membrane can be measured by external electrodes by sensing the outside electrical field around the cells. However, the signal amplitude that can be recorded is a factor 100 to 1000 smaller than the intracellular potential changes. In comparison with glass pipette electrodes, the main advantage of this measurement method is the possibility of simultaneous monitoring of several cells in culture, and thus to be able to obtain informations about cell connectivity in the tissue or cell network. The possible applications follow scientific interest in spontaneous and evoked electrophysiological activity over long time periods and tests of biological effects of toxicological and pharmaceutical agents. Another research 29

40 1 Measurement techniques of bio-electrical activity Fig Diagrams of extracellular measurement principle. Potential variations of a cell lying over an electrode (left) can be measured through an embedded electrode. The electrical circuit is closed by a reference electrode (at ground), which is put into the saline solution of the culture chamber (right). direction is the use of cells as biosensor for environmental condition change sensing [107, 108]. Measurement principle Multi-electrode array devices consist of a biocompatible culture chamber with embedded planar electrodes lying at the bottom of the chamber. Cells or tissue preparations can be deposited in this chamber and kept alive on the measurement electrodes. Electrical monitoring and stimulation of the biological activity is made through the measurement electrodes. Obtained signals are amplified and sent to a data acquisition system for later analysis. One major advantage is that this measurement method can be combined with intracellular measurements with glass pipette electrodes and or with optical measurement techniques using potentiometric dyes (only in case of using transparent electrodes). Many types of cell preparations, i.e. dissociated heart muscle cells [ ], dissociated [ ] and organotypic [116] spinal cord cultures, retina [ ], acute [ ] and organotypic [72, 123, 124] hippocampus slice cultures, different dissociated ganglia cells [ ], and dissociated cortical cells [129, 130] were investigated with this measurement technique. Obtained signals vary with the type of cell preparations. Even though signal amplitudes measured from heart muscle cells are the range of several millivolts, measured neuronal response amplitudes are in a range between 100 µv and 1 mv. Devices and characteristics The first reported micromachined multi-electrode array adapted for extracellular monitoring of electrical activity was fabricated around 30 years ago [112] and was composed of 30 platinized gold electrodes arranged in two lines. Chick embryo cardiac myocytes could be monitored and signal amplitudes of 30

41 1.4 In-vitro measurement methods and devices Fig Left, picture of the first multi-electrode array layout composed of 30 electrodes disposed on two lines 50 µm apart. On the right, enlarged picture of these electrodes (from [112]). Fig Pictures of a commercially available multi-electrode array (Multi Channel Systems) composed of 60 titanium nitride electrodes on a quartz substrate insulated with a silicon nitride layer (the distance between two electrodes is 200 µm). more than 1 mv were obtained. Following this first achievement, several designs of multi-electrode arrays composed of different materials were manufactured and used for heart cell and neuronal monitoring. 31

42 1 Measurement techniques of bio-electrical activity Fig Scanning electron microscopy of surface guidance structures with gold electrodes. The structure walls are made of a polyimide layer covered by a silicon nitride layer used as polyimide etching mask (from [65]). Fig On the left, SEM picture of a neurochip well with integrated platinized gold electrode (from [131]). On the right, picture of an array of 16 wells of a neurochip. Some neurites of a dissociated hippocampal neuron culture are slightly visible (from [132]). Most realized designs are based on a regular matrix of 4x4 [131, ], 5x5 [110], 6x6 [113, 126], 8x8 [109, 121, 136], 3x6 [137] and 4x8 [138] electrodes, or on a regular hexagonal grid with 61 electrodes [118, 139]. However, other design geometries can be found for more special applications: 30 electrodes on an ellipse in order to follow hippocampal geometry [72], and 64 electrodes on a circle [130]. Different materials were used as substrate, electrode and insulation material. Even though most devices are based on glass or quartz substrates, some multielectrode arrays are embedded on silicon [116, 124, 134] and polymer [72, 120] 32

43 1.4 In-vitro measurement methods and devices substrates. Electrodes could be found with different shapes, dimensions and materials. All described electrodes have a square or round shape with dimensions in between 10 µm and 50 µm. These sizes are a compromise regarding both biological and electrical considerations, i.e. the electrodes should be close to cell size to get a good selectivity when a cell lies on the electrode and it should have a sufficient large surface in order to detect electrical signals with an acceptable signal to noise ratio. Thus the choice of electrode material is important to obtain the best measurement conditions. Most electrodes realized so far are based on indium-tin oxide, gold, iridium and platinum electrodes. However, to improve the electrical properties without changing the electrode dimensions, the electrodes surface morphology is modified by deposition of a porous material, e.g. electroplating of platinum black or deposition of a titanium nitride layer. Commonly used insulation layer materials are silicon nitride and silicon dioxide, as well as and polymers like polyimide and polysiloxane. Only two systems based on multi-electrode arrays are commercially available. The first one from Multi Channel Systems [140] is based on a 60 electrodes array composed of titanium nitride electrodes (10 or 30 µm wide and spaced by 100 or 200 µm) on a quartz substrate insulated by a silicon nitride layer [123, 136]. The second one from Panasonic [141] is based on a 64 electrode array composed of platinized indium-tin oxide electrodes (50 µm wide and spaced by 150 µm) on a glass substrate insulated with a polyimide layer [121]. The drawbacks of the multi-electrode array measurement technique with dissociated cell cultures are difficulties in stimulating single neurons due to bad confinement to the electrodes, tracking the same cells over long periods not being possible due to cell migration and, amplitudes of extracellular measured signals drop off rapidly with distance. Therefore, there is a need for cell patterning on multi-electrode arrays. Some electrode arrays integrate a geometrical structure at surface of the device for specific cell guidance [137, 142]. Another patterning approach is the use of wells with integrated measurement electrodes [131, ]. An alternative to passive electrodes is the use of active electrodes, i.e. the use of field effect transistors [143, 144]. In this configuration, the cells form the gate of a transistor controlling the current flux through the transistor. The used material is doped silicon as substrate forming the source and the drain of the transistor and silicon dioxide or silicon nitride as gate instead of metal electrodes. 33

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56 1 Measurement techniques of bio-electrical activity 46

57 2 MULTI-ELECTRODE ARRAY DESIGN AND FABRICATION Multi-electrode array design and manufacturing technologies are highly dependent on biological samples and the underlying experimentation goals. Adequate culture chamber design is also of importance to insure good cell survival conditions (workspace, material biocompatibility, cell adhesion, temperature control and adequate medium perfusion system). Materials, specially for the electrodes, vary following the requested electrical and optical characteristics. The electrode array device fabrication costs depend on device dimensions and the used microfabrication technologies. In this chapter, low-cost planar multi-electrode arrays combining adequate cell culture and measurement conditions for electrophysiology experimentation are presented. 2.1 INTRODUCTION Until now, no standards of multi-electrode arrays have been defined. As a result, most research groups fabricate custom-made multi-electrode arrays and data acquisition systems which show many similarities. Furthermore, existing commercial data acquisition systems are not well established yet, due to overall system and device costs. One of the goals of the present work was to democratize the multi-electrode array monitoring technique by reducing the device costs. This goal can be realized by adapting microelectronic fabrication techniques, whose main advantage is to fabricate many chips on one single substrate. A multi-electrode array layout should be flexible in order to satisfy scientists expectations and to be able to allow several different types of experimentations. To achieve devices answering such a layout flexibility, the interface to an external data acquisition system should remain the same for each device, but the electrode array layout should be modified, implying that only the chip manufacturing differs from one design to another. One important aspect of electrode array design are the electrode characteristics, which are defined by characteristics of the materials used in this study and the electrode area, as will be shown in the Chapter 3. The design and manufacturing techniques of such multi-electrode array devices will be presented in the next sections. 47

58 2 Multi-electrode array design and fabrication 2.2 MULTI-ELECTRODE ARRAY CONCEPT AND DESIGN Multi-electrode array characteristics As described in the first chapter, several electrode array designs with different electrode materials and shapes were realized on various substrates for experimentation on different biological samples. The main characteristics that all in-vitro electrode array devices fulfill are the use of a culture chamber to preserve the cells alive under best possible conditions, biocompatible materials, a high number of electrodes (mostly more than 30) and low impedance electrodes allowing low noise measurements. It is apparent that several characteristics are required for electrophysiology experimentation: The biocompatibility of all used materials is very important to ensure good survival conditions of biological samples. Specially, all interface materials should show no toxic effects on the cells to allow long-term experimentation. The culture chamber should have adapted dimensions to provide adequate cell culture conditions. Moreover, the device workspace dimensions should be large enough for easy handling and simple positioning of cells or tissue slices on the defined measurement area. The surface characteristics of the material should allow good adhesion of biological samples. Cells should like the material. The electrical characteristics of electrodes should allow measurement of signals with a good signal to noise ratio, i.e. the noise level should not be larger than 20 µv. To achieve these characteristics, the electrode materials should be precious metals or porous metals showing larger electrode areas. The electrode insulation layer should have a low dielectric constant or have a large thickness to reduce parasitic capacitances between the culture solution and the electrodes. The device should be transparent. This will allow easy observation of the biological samples by inverted microscopy mostly used by neurophysiologists. It is easier to observe the samples from the bottom than from the top because of the culture chamber solution. Moreover, transparency of the measurement electrodes will allow the combination of optical (potentiometric dyes) and electrical measurements. The device should become a consumable, which means that the fabrication costs should be low. It should become single use to warrant measurement quality and reproducibility. 48

59 2.2 Multi-electrode array concept and design To combine all these characteristics in multi-electrode array design with standard microelectronic fabrication technologies is the challenge of the present work The Smart Petri Dish concept The Smart Petri Dish concept developed in this work consists of the use of 35 mm Petri dishes as culture chamber with integrated measurement electrodes. The advantages of this concept are easy use of the electrode arrays for established culturing protocols and easy handling for experimentation. As in existing multi-electrode arrays, microfabricated planar electrodes on glass or silicon are integrated into a cell culture chamber. There are three ways of culture chamber electrode array device integration: a) to pattern the electrodes on a substrate which will become the bottom part of the culture chamber and to use a glass ring to define the culture chamber walls, b) to place an electrode array into a Petri dish which requires good isolation of the output leads to external data acquisition system, and c) to fix an electrode array under a Petri dish allowing access to the measurement workspace through a hole in the bottom part of the Petri dish, leaving the electrode output leads outside the Petri dish. Commercially available multi-electrode arrays [1, 2] are based on option a). This solution allows easy fabrication of multi-electrode array devices, but is costly due to the use of large substrates (about 5 cm x 5 cm). The option b) and c) allow for smaller chip dimensions which reduces the overall fabrication costs. Indeed, it is more favorable to manufacture many small chips on one single substrate than only one single chip. The introduction of an electrode array in a dish is less elegant than it would be to fix it under the dish. However, when an electrode array is placed in a Petri dish, it requires isolation of most chip parts from the culture medium, which will considerably reduce the culture chamber dimensions. When an electrode array is fixed under a Petri dish, the main problem is the connectivity. An elegant way for signal output is the use of printed circuit boards (PCB). The electrode array is first mounted with a conductive glue to a PCB allowing easy output of the measurement signals to external data acquisition system. The bottom part of a Petri dish is then glued on top of the PCB. Only a silicone sealant is necessary to insure that the culture chamber is watertight and to isolate the electrical output leads from the culture solution. An advantage of this method is that these packaging techniques are low-cost and well known. Moreover, today PCB technology allows versatile design possibilities, i.e. it is very easy to adapt the devices to different existing interfaces. 49

60 2 Multi-electrode array design and fabrication Fig. 2.1 Diagram of two main possibilities of electrode array integration into a Petri dish: top, an electrode array becomes the bottom part of the culture chamber, and bottom, an electrode array is fixed under a Petri dish and is accessible through a hole made in the dish. The solution that was chosen was to fix the multi-electrode array chip on a PCB under a Petri dish Materials Many materials used in microfabrication technology are biocompatible. The most current used substrate, electrode and insulation materials are presented below. Substrate materials The most current substrate materials are silicon, glass and plastics. Silicon is the standard material used for integrated circuits and its fabrication technology is well known. However, its main drawback is that it is not transparent. The use of glass as substrate material instead of silicon allows easy observation during experimentation using inverted microscopy because of glass transparency. Moreover, glass has good electrical insulation properties allowing direct electrode deposition on substrate. Plastics are also used as substrate materials specially when flexible substrates are required. Most of them are also transparent and non conductive but are less resistive to chemicals and temperature than silicon and glass. Conductor materials Electrodes are fabricated from noble metals such as smooth gold and platinum or porous metal layers like titanium nitride, iridium oxide and black platinum for its good electrical characteristics. Best electrical characteristics can be achieved with porous materials due to its larger electrode area surfaces for the 50

61 2.2 Multi-electrode array concept and design same electrode dimensions. Indium-tin oxide, which is a transparent conducting oxide, can also be used. Insulation layers The most common insulation layer materials are silicon nitride [3] and polymers like polyimide [4] and polysiloxane [5]. However, silicon dioxide can also be used but its major drawback is its permeability to ions. The same phenomenon has been observed for polyimide, allowing only short term measurements to be performed [6, 7]. Silicon nitride and silicon dioxide layers are also more cost effective than polymer insulation layers Multi-electrode array design The multi-electrode array design is based on insulated thin conductive electrodes on a substrate. One important aspect to take care of are the electrode characteristics (electrode impedance at electrode/electrolyte interface and the noise level), which depend on material and electrode dimensions. Design of multi-electrode arrays Glass was chosen as substrate material because of its transparency, biocompatibility and its insulation characteristics. As described in the first chapter, existing multi-electrode arrays have different number of electrodes and electrode dispositions (matrix of electrodes with 8x4 [8], 6x6 [9] or 8x8 electrodes [3, 10, 11], a hexagonal matrix [12, 13], or a disposition depending on the biological tissue [14, 15]). A simple matrix of 6x6 or 8x8 electrodes was chosen as multi-electrode array design. This design has the advantage that is allows many different applications because of simple electrode displacement. The electrode dimensions should be in the dimension range between 10 µm x 10 µm and 100 µm x 100 µm. Small electrodes with dimensions close to cell dimensions allow best cell activity measurement but only if a cell is well positioned on top of the electrode and if the electrode impedance is small. Larger electrode areas allow the use of other materials and show better measurement characteristics when recording tissue slice activity. The chosen dimensions are 40 µm x 40 µm. The electrode material could be indium-tin oxide, platinum or black platinum, respectively from the highest to the lowest electrode impedance values. The spacing of the electrodes was chosen at 200 µm (center to center) which defines a square workspace about 1 mm 2 or 2 mm 2. The chosen material for the insulation layer used is an epoxy polymer called SU-8 [16-18]. This material seems to be biocompatible, i.e. there is no evidence of material toxicity to the cells as shown by long-term rat spinal cord 51

62 2 Multi-electrode array design and fabrication cultures lasting more than one month (results obtained at University of Bern) and chicken embryonic motoneurons (results obtained at University of Geneva). Moreover, its main advantages are: a) easy to pattern because it is a photosensitive epoxy resist, b) the SU-8 is fully transparent in visible light, and c) it is a low-cost material. Backend Process Following the smart Petri dish concept, it was chosen to fix the multielectrode array on a PCB under a Petri dish. After substrate dicing, the resulting chips have to be mounted on an external PCB for signal output. Different connection techniques exist to fix the electrode leads of the multi-electrode array to a PCB: a) wire bonding with thin gold or aluminium wires, b) the flip chip method consisting of a reflow of metallic connection bumps to generate the electrical contact, c) gluing with anisotropically electroconductive adhesive films, and d) gluing with a conductive glue deposited by screen printing on one of the two parts. The wire bonding method is a very current technique in industry for realization of electrical connections. Unfortunately, the chosen packaging configuration of our device, in which two parts are directly in contact, do not allow wire bonding. The flip chip method is an easy way to connect two electrical leads in contact. This method is also currently used in industry and well known. Unfortunately, the use of metal bumps and high temperature reflow is inadequate for our application [19, 20]. The gluing with anisotropically conductive adhesive films is a delicate process. The used films are based on epoxy polymer charged with metallic particles. The principle of these glues is based on film compression. Under pressure, specially at the electrode pads, the metallic particles are pressed together and an electrical contact in the vertical axe is created, leaving the other parts without electrical contact in the horizontal plan. This gluing process requires a temperature of about 150 C and a constant pressure over all the connection area. Tests made with anisotropic glues showed that if the pressure is not homogeneous, electrical conduction in the horizontal plan may occur generating a shunt between several electrodes. This method was taken into account only for the first electrode array generation, but was then replaced by a more reliable screen printing process. Screen printing of conductive epoxy based glue charged with silver particles is a very simple method to contact the electrode array devices with the PCBs. The pattern of the connection pads is reported on a metal mesh through which one the glue is applied onto the PCB [21, 22]. This well known method is accurate, 52

63 2.2 Multi-electrode array concept and design Fig. 2.2 Top: diagram of anisotropically conductive adhesive film principle (data sheets of 5303R film from 3M Ltd.): a charged glue is pressed and heated between two parts. Electrical contact is achieved in the z axe at points that had a strong compression. Bottom: diagram of screen printing principle (from [21]): conductive glue (ink) is printed through a stencil mask on one part. The gluing process is completed by placing the two parts together and glue hardening by a bake. reproducible and low cost. It has therefore been used for multi-electrode array manufacturing. The final backend steps are the gluing of the bottom part of a Petri dish, in which a circular hole was stamped previously, onto the printed circuit board and the sealing of the culture chamber External amplification and data acquisition system A multi-electrode array device is not sufficient for electrophysiology experimentation. Signal amplification and data acquisition on a computer are two other features needed to build a complete measurement setup. 53

64 2 Multi-electrode array design and fabrication Fig. 2.3 Multi-electrode array interface for external signal amplification of the first multi-electrode array generation. The interface is composed of a printed circuit board. The device is held in position by a metallic ring. A ring cap generates a pressure on the device to insure good electrical contact between the printed circuit board and the multi-electrode array device. Fig. 2.4 On the left, a data acquisition system composed of a multi-electrode array interface and real-time data acquisition hard- and software. On the right, a multi-electrode array interface including 60 channel amplification stage close to the measurement electrodes (from [2]). The amplitude of biological signals measured with extracellular electrodes varies between a few microvolts to a few millivolts. These small signals must be amplified with a gain of at least 1000 in order to make possible computer data acquisition. Moreover, the distance between the amplifier electronics and the measurement electrodes should be as short as possible to avoid noise capture and 54

65 2.3 Fabrication of multi-electrode arrays following noise amplification by the amplifiers. Another point of care is the electrode polarization. A capacitive coupling between the electrodes and the amplifiers will eliminate the DC signal due to electrode polarization and let through only the AC part of the signal. Furthermore, there is a need of having an interface in which the multielectrode array devices are connected to for the experimentation. This interface should allow easy signal output and integrate the amplification electronics on multiple channels (one channel per electrode). This system should also allow the possibility of selecting a channel (an electrode) for electrical stimulation of the cells/tissue slice in culture. Computer data acquisition requires adequate hardand software allowing the sampling, display, analysis and to save all obtained data. Two commercially available systems [1, 2] fulfill all the needs for signal amplification and computer data acquisition for storage and later data analysis. As shown below, an interface for amplification of output signals was designed for the first generation of multi-electrode array devices. A second generation of multi-electrode array devices was adapted to the commercially available data acquisition system from Multi Channel Systems GmbH, Reutlingen, Germany. 2.3 FABRICATION OF MULTI-ELECTRODE ARRAYS Fabrication process and the resulting devices for two different designs of planar multi-electrode arrays for electrophysiology measurements will be described in the following sections. The first one, following the Smart Petri Dish concept, was designed with its own custom made interface for signal amplification and data acquisition. The second one, adapted to the Multi Channel Systems data acquisition interface, was designed following commercially existing multi-electrode arrays. The main advantage of adapting the multielectrode arrays to this system was the presence of adequate data acquisition and analysis software required for electrophysiology experimentation. Due to the small electrode dimensions, microelectronic fabrication technologies were used for the manufacturing of the multi-electrode arrays in a clean room environment. The basic used microfabrication technologies are well described elsewhere [23-25] Transparent multi-electrode arrays (first generation) Glass substrates and transparent conductive indium-tin oxide measurement electrodes are used to realize transparent multi-electrode arrays. However, 55

66 2 Multi-electrode array design and fabrication Patterning of negative photoresist for lift-off process. Titanium (50 nm)/platinum (150 nm) deposition by sputtering process. Stripping of photoresist. Patterning of positive photoresist as etching mask. Etching of indium-tin oxide layer in hydrochloric acid. Patterning of an SU-8 epoxy insulation layer. Fig. 2.5 Process flow of planar indium-tin oxide/platinum electrode MEAs. because of its better electrical characteristics, the fabrication process of platinum electrodes at the measurement sites on indium-tin oxide leads will be described instead. 56

67 2.3 Fabrication of multi-electrode arrays Fabrication process The fabrication of multi-electrode arrays is divided into two main parts: a) the microfabrication of the multi-electrode array chips and b) the back end assembly steps of the electrode array device. Float glass plates (thickness of 700 µm, diameter of 10 cm) covered with a thin layer of indium-tin oxide (thickness of 100 nm, 20 Ω/square) were used as substrate material. A platinum layer was deposited at the measurement areas by a lift-off process. First, AZ5214 photoresist (Image Reversal Photoresist) from Clariant was deposited onto the substrate on a spin coater and solvents were evaporated by a 1 minute bake at 110 C. The photoresist was then exposed to UV light at 365 nm defining the electrode pattern. A 1 minute bake at 117 and a supplementary exposure to UV light at 365 nm inverted the photoresist pattern. It was then developed in S351 developer from Shipley. The resulting pattern was the negative of the mask pattern. The walls of the photoresist were negative which allowed easy release after metal deposition. A metal layer, 50 nm of titanium as adhesion layer and 150 nm of platinum was deposited by sputtering in a vacuum chamber at room temperature onto the substrates. The underlying photoresist was then dissolved in acetone removing the metal deposited on its surface. The next step consisted of the patterning of the indium-tin oxide layer by positive photolithography using S1813 photoresist from Shipley. The S1813 resist was deposited onto the indium-tin oxide/platinum surface on a spin coater and the solvents were evaporated by a 1 minute bake at 115 C on a hot plate. The photoresist was then exposed to UV light at 365 nm defining the electrodes and their leads. It was then developed in S351 developer from shipley. The indium-tin oxide was then etched by wet chemical etching in a hydrochloric acid solution (HCl 37%) for 4 minutes, rinsed in deionized water and let dry. The last fabrication operation was the deposition and patterning of an SU-8 epoxy insulation layer with a thickness of 5 µm. SU-8 60/40 (epoxy/solvent proportions) photosensitive resist was deposited on a spin coater (5000 rpm, 40 s) and baked (95 C, 15 minutes) for solvent evaporation. The SU-8 was then exposed to UV light at 365 nm. Cross linking of the illuminated SU-8 parts was achieved by a polymerization bake (95 C, 15 minutes). Finally, the nonpolymerized epoxy areas were removed in a PGMEA (poly-glycol-methyl-etheracetate) development bath for 1 minute opening the electrode measurement areas. After electrical control of electrode impedances, the electrode arrays were released by dicing. The obtained chips were then mounted onto a printed circuit board by gluing with an anisotropic conductive adhesive film Threebond 3370 from 3M Ltd. The 57

68 2 Multi-electrode array design and fabrication Chip dimensions 1.5 cm x 1.5 cm Electrode dimensions squares of 40 µm x 40 µm Spacing of electrodes 200 µm center to center Number of electrodes Matrix of 6x6 electrodes Table 2.1 Dimensions and electrode configuration of first design multi-electrode arrays film was deposited manually on the electrode output pads of the chips. The printed circuit board was then placed in contact with the chip and the gluing was done for 45 s at 150 C with a pressure of 20 kg/cm 2. Then, the bottom part of a Petri dish, in which a hole of 1 cm in diameter was stamped, was glued with double side tape on the other side of the printed circuit board, defining the culture chamber of the multi-electrode array. Finally, the obtained culture chamber was sealed with a biocompatible two-component silicone (DC 184 from Dow Corning). Realized structure The dimensions of the fabricated multi-electrode arrays are indicated in Table 2.1. The multi-electrode array layout and a picture of the workspace of the multi-electrode array is shown in Fig A bottom and a top view of a finished device are shown in Fig The chip mounting onto the printed circuit board with anisotropically electroconductive adhesive film was difficult to achieve with good reproducibility. The deposition of the adhesive film on the chips was also very time consuming because of the geometrical electrodes pads disposition (square configuration). Moreover, the applied pressure during the gluing had to be controlled precisely over all the chip surface to insure conduction between the electrodes pads and the printed circuit board pads, and avoid horizontal shunt between the electrodes. These points led to a change of the chip gluing technique. This device was designed to fit into the output interface shown in Fig The signal amplification and data acquisition system for this interface had to be designed and realized. To overcome this very time consuming development of electronics and software, the multi-electrode arrays were modified and adapted to a commercially available signal amplification and data acquisition system. The next section will describe these adapted multi-electrode arrays. 58

69 2.3 Fabrication of multi-electrode arrays Fig. 2.6 Picture of the design and the workspace of the multi-electrode array. 36 electrodes are disposed in a 6x6 matrix in the center of the chip. Electrode dimensions are 40 µm x 40 µm and interspace between two electrodes is 200 µm. Indium-tin oxide/platinum electrodes on glass substrate are insulated by an 5 µm thick SU-8 epoxy layer. Fig. 2.7 Resulting multi-electrode array composed of 36 planar platinum electrodes. The microstructures are mounted on a printed circuit board. A Petri dish (Ø 35 mm) glued with double side tape to the printed circuit board is used as culture chamber. A two component silicone sealing insures that the culture chamber remains watertight. 59

70 2 Multi-electrode array design and fabrication Patterning of negative photoresist for lift-off process. Titanium (50 nm)/platinum (150 nm) deposition by sputtering process. Stripping of photoresist. Patterning of an SU-8 epoxy insulation layer. Fig. 2.8 Process flow of planar platinum MEAs Second generation of multi-electrode arrays Two different metals were used as electrode material, which means that two different fabrication processes were used. Indium-tin oxide (ITO) and smooth platinum electrodes were used with the same multi-electrode array design. Thus the fabrication process is at least similar to that of the transparent multi-electrode arrays. The following process flow includes only platinum electrodes. Fabrication process The fabrication of multi-electrode arrays is divided in two main parts: a) the fabrication of the multi-electrode array on a glass substrate and b) the back end assembly of the electrode array device. Float glass plates (thickness of 700 µm, diameter of 10 cm) were used as basic material. Platinum electrodes were deposited by a lift-off process as described in Chapter First, AZ5214 photoresist from Clariant was patterned onto the glass substrate. Then, a 50 nm thick layer of titanium for adhesion and a 150 nm thick layer of platinum were deposited by sputtering in a vacuum chamber at room temperature on the substrates. The underlying photoresist was then dissolved in acetone removing the metal deposited on its surface. 60

71 2.3 Fabrication of multi-electrode arrays Chip dimensions 1.5 cm x 1.5 cm Electrode dimensions squares of 40 µm x 40 µm Spacing of electrodes 200 µm center to center Number of electrodes Table 2.2 Dimensions of planar multi-electrode arrays. matrix of 8x8 electrodes without the 4 corners Deposition and patterning of an SU-8 epoxy insulation layer with a thickness of 5 µm defined the measurement electrodes and insulated the electrode leads. After electrical control of the electrode impedances, the electrode arrays were released by dicing. The obtained chips were then mounted onto a printed circuit board by a gluing process. A 20 µm thick layer of E212 (from Epotecny) two component epoxy glue charged with silver particles was screen printed onto the printed circuit board electrode contact pads. Then the two parts were put together, baked 3 hours at 80 C and allowed to slowly cool down to room temperature. The bottom part of a polystyrene box (diameter of 27 mm), in which a hole of 1 cm in diameter was stamped, was then glued on the other side of the printed circuit board, defining the culture chamber of the multi-electrode array. Finally, the culture chamber was sealed with a biocompatible two component silicone (DC 184 from Dow Corning). Realized structure The dimensions of the fabricated multi-electrode arrays are indicated in Table 2.2. The multi-electrode array layout and a picture of the workspace of the multi-electrode array is shown in Fig A bottom and a top view of a finished device are shown in Fig The electrode array protrudes under the PCB. When the chip is inserted into the commercial data acquisition interface from Multi Channel Systems, the glass substrate comes in contact with the bottom of the interface, which is formed of the heat control of the culture chamber. The electrical connection between the interface and the multi-electrode array is achieved by spring contacts. When using this multi-electrode array on the data acquisition interface, a PCB spacer has to be placed prior to the multi-electrode array into the system. The function of this spacer is to avoid a short circuit between the electrode leads and the heating resistor located directly under the chip, and mechanical damage of the chip during the use of the device. If the MEA is placed into the interface without this spacer, the interface spring contacts will exert a pressure on the chip 61

72 2 Multi-electrode array design and fabrication Fig. 2.9 Picture of the layout and the workspace of the multi-electrode array. Electrode dimensions are 40 µm x 40 µm and interspace between two electrodes is 200 µm (center to center). Platinum electrodes on glass substrate are insulated by a 5 µm thick SU-8 epoxy layer. Fig Resulting multi-electrode array composed of 60 planar platinum electrodes. The microstructures are mounted on a printed circuit board. A polystyrene box (Ø 27 mm) is used as culture chamber. A silicone sealant insures that the culture chamber remains watertight. generating torsion of the PCB. This torsion should be avoided because of irreversible chip damage due to break of electrical contacts between the chip and the PCB (ungluing of the chip from the PCB). 62

73 2.4 References 2.4 REFERENCES [1] [2] [3] W. Nisch, J. Böck, et al., A thin film microelectrode array for monitoring extracellular neuronal activity in vitro, Biosensors & Bioelectronics, Vol. 9, pp , 1994 [4] G. W. Gross and F. U. Schwalm, A closed flow chamber for long-term multichannel recording and optical monitoring, Journal of Neuroscience Methods, Vol. 52, pp , 1994 [5] M. Meister, J. Pine and D. A. Baylor, Multi-neuronal signals from the retina: acquisition and analysis, Journal of Neuroscience Methods, Vol. 51, pp , 1994 [6] S. A. Boppart, B. C. Wheeler and C. S. Wallace, A flexible perforated microelectrode array for extended neural recordings, IEEE Transactions on Biomedical Engineering, Vol. 39, N 1, pp , 1992 [7] J. F. Hetke, J. L. Lund, et al., Silicon ribbon cables for chronically implantable microelectrode arrays, IEEE Transactions on Biomedical Engineering, Vol. 41, N 4, pp , 1994 [8] J. L. Novak and B. C. Wheeler, Multisite hippocampal slice recording and stimulation using a 32 element microelectrode array, Journal of Neuroscience Methods, Vol. 23, pp , 1988 [9] G. W. Gross, E. Rieske, et al., A new fixed-array multi-microelectrode system designed for long-term monitoring of extracellular single unit neuronal activity in vitro, Neuroscience Letters, Vol. 6, pp , 1977 [10] L. J. Breckenridge, R. J. A. Wilson, et al., Advantages of using microfabricated extracellular electrodes for in vitro neuronal recording, Journal of Neuroscience Research, Vol. 42, pp , 1995 [11] H. Oka, K. Shimono, et al., A new planar multielectrode array for extracellular recording: application to hippocampal acute slice, Journal of Neuroscience Methods, Vol. 93, pp , 1999 [12] W. G. Regehr, J. Pine, et al., Sealing cultured invertebrate neurons to embedded dish electrodes facilitates long-term stimulation and recording, Journal of Neuroscience Methods, Vol. 30, pp , 1989 [13] A. Litke and M. Meister, The retinal readout array, Nuclear Instruments and Methods in Physics Research, Vol. A 310, pp , 1991 [14] L. Stoppini, S. Duport and P. Corrèges, A new extracellular multirecording system for electrophysiological studies: application to hippocampal organotypic cultures, Journal of Neuroscience Methods, Vol. 72, pp ,

74 2 Multi-electrode array design and fabrication [15] P. Thiébaud, N. F. d. Rooij, et al., Microelectrode arrays for electrophysiological monitoring of hippocampal organotypic slice cultures, IEEE Transactions on Biomedical Engineering, Vol. 44, N 11, pp , 1997 [16] N. LaBianca and J. D. Gelorme, High aspect ratio resist for thick film applications, SPIE: Advances in Resist Technology and Processing XII, Vol. 2438, Santa Clara, USA, February , pp , 1995 [17] H. Lorenz, M. Despont, et al., SU-8: a low-cost negative resist for MEMS, Journal of Micromechanics and Microengineering, Vol. 7, pp , 1997 [18] H. Lorenz, M. Laudon and P. Renaud, Mechanical characterization of a new high-aspect-ratio near UV-photoresist, Microelectronic Engineering, Vol. 41/42, pp , 1998 [19] J. Eldring, E. Zakel and H. Reichl, Flip-chip attachment of fine-pitch GaAs devices using ball-bumping technology, The International Journal of Microcircuits and Electronic Packaging, Vol. 17, N 2, pp , 1994 [20] T. A. Frieswijk, J. A. Bielen and W. L. C. Rutten, Development of a solder bump technique for contacting a three-dimensional multi electrode array, The 17th Annual International Conference of the IEEE Engineering in Medicine and Biology Society, Vol. 2, Montreal, Canada, pp , 1995 [21] R. M. Swerdlow, The step-by-step guide to screen-process printing, Prentice-Hall Inc., Englewood Cliffs, New Jersey, USA, 1985 [22] F. Lambert, La sérigraphie, EMI, Vol. 182, pp , 1974 [23] S. M. Sze, Semiconductor Devices: Physics and Technology, John Wiley & Sons, New York, USA, 1985 [24] M. Madou, Fundamentals of Microfabrication, CRC Press, New York, USA, 1997 [25] P. Rai-Choudhury, Handbook of microlithography, micromachining, and microfabrication, SPIE Optical Engineering Press, Bellingham, Washington, USA,

75 3 ELECTRICAL PROPERTIES OF MULTI-ELECTRODE ARRAYS The electrical properties of measurement electrodes are of importance to insure good measurement and stimulation conditions, i.e. electrodes should have a high signal to noise ratio and should have high charge injection capabilities. In this chapter, an electrode model, the resulting electrode properties and the electrode noise level are presented. Experimental impedance and phase shift measurements allowing determination of global electrical electrode properties (resistance, capacitance) are presented and discussed. 3.1 ELECTRICAL MEA MODEL To insure best signal measurement conditions, it is of importance that the electrode design takes into account the electrode properties and the resulting noise level and signal to noise ratio. To evaluate the electrical characteristics of MEAs, it is necessary to have an electrical model corresponding to the realized devices. The main parts of such an electrical model are of course, the electrode leads, the electrode/electrolyte interface, parasitic electrode components, biological cells lying over the electrode (signal sources), and the culture medium. Fig. 3.1 and Fig. 3.2 show an electrical model of one MEA electrode. The electrode properties are mainly determined at the electrode/electrolyte interface, where an electrical double layer is formed by the electrons in the metal and by the ions close to its surface, resulting in a capacitive component. As documented in literature [1, 2], the impedance of metal electrodes is modelled by a charge transfer resistance in parallel with a capacitance. Some parasitic components like the electrode lead resistance, the insulation capacitance between the electrode lead and the culture solution, and the coupling capacitance between the electrode leads play only a minor role on the global electrode properties. The electrode/electrolyte resistance polarization encountered by a current spreading out from the electrode into the conductive culture solution is also of minor importance. The solution is used as potential reference and is thus connected to the ground. 65

76 3 Electrical properties of multi-electrode arrays Fig. 3.1 Diagram of a MEA device including the culture chamber filled with a physiologic solution and a cell lying over an electrode. A simplified electrical scheme of the build circuit is also shown. Fig. 3.2 A simplified electrical diagram of one MEA electrode. The electrode is mainly composed of a charge transfer resistance R transfert and an interfacial capacitance C interface. Parasitic electrode components are the lead resistance R lead and, insulation C insulation and electrode coupling C coupling capacitances. When a cell is disposed above an electrode, the electrical model of the cell, the cell membrane, and a sealing resistance have to be taken into account in the global model. This cell will be the source of signals (action potentials) that are 66

77 3.2 electrode model measured through this electrical circuit. For the following electrode model description, we focus on the physical properties of the electrodes and omit the cell. This simple electrical model can be considered to be a good approximation of reality. On the other hand, the model of an electrode at electrode/electrolyte interface is more complex than the simple RC circuit introduced so far and will be described in the next section. Since the global electrode model can be considered as a low pass filter, a high electrode impedance value will increase a filtering effect on the measured signal. This implies that low impedance values are best for cell activity monitoring, in particular to obtain a good signal to noise ratio allowing post analyzing of the measured data. For electrical cell stimulation through the electrodes, a low impedance allows also a better charge delivery for a given excitation voltage. 3.2 ELECTRODE MODEL The following small signal electrode model, based on physical mechanisms occurring at the metal/electrolyte interface, is described in more detail in the literature [1-4] The electrode/electrolyte interface The interfacial capacitance When a metal object is placed into an ionic solution, a space charge layer rapidly builds up at the interface due to chemical reactions, which represents in terms of electrical circuit element a capacitance that can be modeled using several increasingly complex theoretical approaches. The equilibrium for a given electrode potential is reached when a sufficient charge separation is achieved at the interface so that reduction and oxidation reactions proceed at equal and opposite rates. One approach, the Helmholtz- Perrin model, considers a layer of oriented water molecules in contact with the metal surface, the hydration sheath of inner Helmholtz plane shown in Fig The next layer consisting of hydrated cations is referred to as the outer Helmholtz plane. It is in this region where electrochemical reactions take place. The charges located at the electrode interface can be modeled as a parallelplate capacitance, also called the Helmholtz capacitance C H defined as: C H = ε 0 ε r A d (3.1) 67

78 3 Electrical properties of multi-electrode arrays Fig. 3.3 Illustration of the nanostructure of the space charge layer at an aqueous electrode showing the hydration sheath, a monolayer of oriented water molecules at the metal surface and the next layer of hydrated cations. Adsorbed species are not taken into consideration here (from [4]). where ε 0 is the dielectric permittivity of free space ( x10-12 F/m 2 ), ε r is the relative dielectric permittivity of the medium (78.54 for water at 25 C), A is the electrode surface, and d is the distance between the inner and the outer Helmholtz plane, being about 5 Å. It results for a 40x40 µm 2 electrode a capacitance of C H = 2.2 nf. This model assumes a fixed capacitance. In practice, however, the capacitance is voltage dependent. As the potential increases, the ions are pushed in closer vinicity of the electrode surface. The Gouy-Chapman model of diffuse space charge introduces an increase in capacitance with increasing applied potential. The derived space charge diffusion capacitance C D is defined by: C D ε 0 ε r A = zv 0 L cosh D V th (3.2) 68

79 3.2 electrode model where z is the charge of the ion, V 0 is the potential at electrode, V th is the thermal voltage (26 mv at 25 C), and L D is the Debye length (distance over which a small pertubation of potential decays e -fold) defined by: ε L 0 ε r V D = th (3.3) 2n z 2 q where n is the bulk number concentration of the ion (0.154 mol/l times Avogadro s number 6.02x10 23 ) and q is the charge of an electron (1.6x10-19 C). If we assume that V 0 =50 mv and z=1, then the resulting value of this capacitance is C D =30 pf. Finally, the interface capacitance C i is best defined by the Stern model, which combines the Helmholtz-Perrin model with the diffuse space charge layer model of Gouy-Chapman in putting these two capacities in parallel: C i = C H C D (3.4) In our case, the value of the interface capacitance is about C i =29.6 pf. This approximation of the interfacial capacitance of a metal electrode in solution is not perfect but provides a better understanding of the physical mechanisms of the electrical interface. The resistive mechanism of charge transfer The movement of charge into or out of an electrode interface depends on several mechanisms. The charge transfer through the Helmholtz double layer, the interface diffusion and the chemical reactions play an important role for the definition of a model close to reality. However, the model most used follows two main principles: the charge transfer and diffusion, which becomes very important when high currents are applied. At equilibrium, the current due to oxidation reaction is equal and opposite to the reduction current. The absolute value of this current for a given surface area is referred as the exchange current density J 0. The applied potential has an influence over the oxido-reduction reaction which defines the current through the interface. For small applied potentials, the total charge transfer resistance R t can be expressed as: R t = V th J 0 za (3.5) Assuming an exchange current density of J 0 =8 A/m 2 for platinum [5], the value of the charge transfer resistance is R t =2.03 MΩ. Unfortunately, this model 69

80 3 Electrical properties of multi-electrode arrays Fig. 3.4 Schematic representation of the electrical circuit model of the electrode interface showing the double-layer interface capacitance C i, the charge transfer resistance R t, and the parallel Warburg circuit elements R w and C w (from [4]). is not sufficient to explain the electrode behavior due to diffusion at high frequencies. It was shown that when the frequency increases, the ion concentration gradient increases at the electrode interface, increasing the current that can be passed through the electrode. Warburg proposed in 1899 that the impedance due to diffusion Z d varies with frequency as: k Z d = (3.6) F where k is a constant determined by the electrochemistry and mobilities of the reactant species involved and F is the frequency in hertz. Warburg solved the diffusion equation under the assumption that the depth of penetration of the concentration wave into the bulk solution was small compared to the thickness of the diffusion layer. His solution was a frequency dependent impedance modeled as a series RC pair, in turn in series with the charge transfer resistance R t. This RC combination can be transformed into a parallel RC network that appears to provide for a DC current path past the double-layer capacitance, which appears to be more intuitively satisfying (see Fig. 3.4). The Warburg parallel RC components are defined as: R w V th 1 = z 2 qn A πd F (3.7) 70

81 3.2 electrode model and 1 C w = πFR w (3.8) where D is its diffusion coefficient. At a frequency of F=1000 Hz and a diffusion coefficient D=2.07x10-9 m 2 /s, the values for R w and C w are R w =429 kω, C w =371 pf respectively. It is important to notice that the frequency dependence of Warburg parallel RC elements are only valid if the penetration depth of the concentration wave is small compared to the diffusion depth. For very low frequencies, this assumption is no longer valid, the Warburg impedance approaches infinity when the frequency tends to zero. (R w = ). Spreading resistance Resistance polarization arising from the bulk resistance of the electrolyte encountered by a current spreading out from the electrode into the electrolyte can be included in the circuit model through the simple addition of a series resistor R s. It is determined by the geometric surface of the electrode and is defined as ρln( 4) R s = (3.9) πa for a planar square electrode where ρ is the resistivity of the solution (80 Ωcm) and a is the side length of the electrode (40 µm). The resulting value of R s is R s =8825 Ω. This resistance is negligible compared to the charge transfer resistance of the electrode. Model extension for porous surfaces For porous materials, the model should be extended by a pore-geometry circuit model. A supplementary impedance Z p must be added in parallel with the previously developed electrode interface model in order to incorporate the effects of a porous surface into the microelectrode circuit. The basic premise of this pore model is that the series impedance and interfacial admittance per unit length are constant within a pore and that under this assumption, a pore is analogous to a uniform transmission line. It is admitted that this supplementary impedance has a magnitude proportional to 1/ F and a constant angle of -45. However, in practice, detailed pore models are not very useful because one cannot accurately measure or predict pore sizes, depth or numbers. In general, a variable scaling factor k is used to fit an expression representative of the pore impedance to experimental data. The pore impedance can be defined as: 71

82 3 Electrical properties of multi-electrode arrays Fig. 3.5 Schematic representation of the electrode/solution interface including the interface capacitance C interface, the transfer resistance R transfer, the Warburg impedance Z Warburg, the pore impedance Z pores, and the spreading resistance R spreading (from [4]). k Z p = ( 1 j) (3.10) F The global electrode/electrolyte interface model is shown in Fig Due to frequency dependence, Z Warburg tends to zero at very high frequencies. This implies that the electrode circuit model can be approximated by the spreading resistor R spreading and the double layer interface capacitance C interface. The high frequency current is entirely the result of the displacement of charge rather than the transport of charge across the double layer. When the impedance is measured at high frequencies, it remains a constant impedance and phase shift determined by the interface capacitance and the spreading resistor added to the electrolyte resistance between the electrode and the reference electrode in the bath. On the other hand, at very low frequencies, the Warburg impedance becomes very large and dominates the total electrode impedance. In this case, the circuit model is reduced to the Warburg elements. Here, the current due to the transport of charge across the double layer is large compared to the charge displacement current Parasitic components of the electrode Some parasitic components are present at the electrode as shown in Fig These main parasitic components are the electrode lead resistance, the coupling 72

83 3.2 electrode model capacitance with neighboring electrodes, and the passivation layer capacitance between the electrode lead and the culture medium. The electrode lead resistance R L is defined by: ρ R m l L = (3.11) wt where ρ m is the resistivity of the electrode (10.6x10-6 Ωcm for platinum), l is the lead length (7 mm), w and t are the width (50 µm) and the thickness (100 nm) of the electrode lead respectively. The resulting value of R L for the electrodes of the fabricated MEAs is R L 150 Ω. This value is negligible compared to the electrode/electrolyte interface. An accurate approximation for the electrode coupling C c and passivation layer C p capacitance was described by Sakurai and Tamaru [6]. The proposed electrode coupling capacitance assumes that a given electrode lead interacts only with its two nearest neighbors and is given by: C c = 2ε 0 ε r l 0.03 w t d d t d s -- d 1.34 (3.12) where ε r is the permittivity of the insulation layer (3.6 for SU-8 epoxy), d the passivation layer thickness (5 µm of SU-8) and s is the intertrace spacing (200 µm). For the designed MEA electrodes, the value of C c is C c =0.9x10-3 ff. The proposed passivation layer capacitance takes into account sidewall contributions and field effects between two electrodes and is given by: C p = 1.15ε 0 ε r lw t ε d 0 ε r l -- d (3.13) The value of C p is C p =1.47 pf. Therefore, the parasitic capacitances C c and C p are negligible compared to the interface capacitance C i ( 30 pf) of the electrode/ electrolyte interface Resulting electrodes characteristics The electrode noise level is essentially composed of the thermal noise, which is defined as: U th = 4kTBR tot (3.14) where k is the Boltzmann constant (1.38x10-23 J/K), T is the temperature in Kelvin, B is the frequency bandwidth, and R tot is the global electrode resistance. 73

84 3 Electrical properties of multi-electrode arrays However, the electrode circuit configuration corresponds to a low pass filter and the frequency bandwidth can be approximated by the cutting frequency F c defined as: 1 F c = R tot C tot (3.15) where R tot and C tot are the global electrode resistance and capacitance respectively. To avoid DC signal component on the signal amplifier stage, a high pass filter is incorporated in the data acquisition interface. The overall electrode circuit configuration corresponds thus to a band pass filter. Finally, the resulting signal to noise ratio of the electrodes is defined as the measured signal magnitude U m divided by the electrode noise level U th : U m U th SNR = (3.16) 3.3 ELECTRODE STIMULATION PROPERTIES A detailed description of stimulation through metal electrodes, the underlying mechanisms and possible effects on the electrodes and the tissue can be found in literature [2, 7, 8] and will not be detailed here. When cell stimulation occurs through the microfabricated MEA electrodes, the global electrode capacitance is of importance. Electrical stimulation of biological tissue with metal electrodes requires the flow of ionic charge in the biological tissue, which can be induced by a capacitive and a faradic mechanism. The capacitive mechanism involves alternate attraction and repulsion of ions in the tissue fluid in response to changes in the electrostatic charge on the metal surface (electrode double layer). This mechanism is an ideal mechanism because no chemical changes occur in the tissue. However, the amount of charge that can be injected by this mechanism is only small. Faradic mechanisms of charge injection involve electron transfer across the electrode-tissue interface and therefore necessitate that some chemical species be either oxidized or reduced. This mechanism is divided into reversible and irreversible electrode reactions. Reversible faradic reactions involve only species that remain bound to the electrode, i.e. oxide formation and reduction, and H- atom plating. Unfortunately, the reversible faradic processes are limited. On the other hand, irreversible faradic reactions involve chemical species that do not remain bound to the electrode surface, i.e. electrode corrosion, water electrolysis to form H 2 or O 2 provoking a ph shift, and the oxidation of chloride ion. All 74

85 3.4 Experimental electrode characterization irreversible faradic mechanisms present a toxic effect onto tissue lying close to the electrode and should thus be avoided. The choice of the material is thus of importance when high current stimulation should be realized through electrodes. It was shown that noble materials like iridium, rhodium, platinum and palladium are extremely resistant to electrolysis [9]. However, porous materials, specially platinum black and iridium oxide show the best stimulation characteristics due to larger electrode capacitance [7, 10-12]. The stimulation signal is also of importance in order to avoid irreversible reactions at the electrode surface. Experiments showed that the best stimulation signal is composed of a biphasic current waveform consisting of two consecutive pulses of equal charge but opposite polarity [7]. Furthermore, when stimulation occurs with a low frequency, the chemical species released during irreversible faradic reactions have more time to diffuse through the tissue and the culture solution, which reduces the negative effect onto the tissue. 3.4 EXPERIMENTAL ELECTRODE CHARACTERIZATION Methods and materials Experimental measurements of electrode properties were obtained from different MEAs integrating electrodes of different dimensions and materials. Impedance (Z) and phase shift (θ) measurements were done with an impedance/ gain-phase analyzer from HP (HP4194A). The MEAs were first filled with a 0.9% NaCl solution. An OPEN compensation was made for cable impedance compensation. A large counter electrode (piece of platinum) at ground potential was immersed in the solution and the measurements were achieved by closing the electrical circuit in making contact with the signal output pads of the MEA devices, and by applying a 50 mv, 100 khz tension to the electrodes. In this configuration, the electrode is small in comparison to the counter electrode. Its impedance will dominate and the measured impedance is essentially that of the MEA electrode. Five electrodes were measured for each different experience. Mean values and the relative error of the obtained data were then calculated and plotted. The global electrode resistance and capacitance can be extrapolated from the impedance and phase shift measurement. As shown in Fig. 3.6, the impedance and the phase shift can be expressed as the modulus of a vector (complex domain) and the angle between this vector and the X axis respectively. The 75

86 3 Electrical properties of multi-electrode arrays Fig. 3.6 The impedance Z and the phase shift θ plotted as a vector in the complex domain. The projections of this vector onto the axes give the global resistance R (Z ) and capacitance C (Z ) of the electrode. projections of this vector onto the X and Y axes define the global electrode resistance and capacitance respectively by the two following equations: R = Z cos(θ) (3.17) 1 / C = (2 π F) Z sin(θ) (3.18) Electrode properties function of electrode area This experiment consisted of an impedance and phase shift measurement at a fixed frequency (1 khz) of platinum and indium-tin oxide electrodes versus geometrical electrode dimensions. In order to make this experiment, another MEA layout was designed and realized (see Chapter for the fabrication process). It is composed of 30 circular and 30 square electrodes with electrode dimensions going from 5 µm x 5 µm up to 150 µm x 150 µm as shown in Fig As can be seen on the obtained results shown in Fig. 3.8., the electrode impedance varies with electrode area as Z = k/a. This is in accordance with Eqn (3.1), Eqn (3.2), Eqn (3.5), and Eqn (3.7). However, an electrode impedance higher than 1 MΩ should be avoided because of increasing noise level and signal to noise ratio decrease. To insure a sufficient measurement quality, electrode dimensions should be at least 25 µm x 25 µm for smooth platinum and 40 µm x 40 µm for indium-tin oxide electrodes. 76

87 3.4 Experimental electrode characterization Fig. 3.7 Picture of the layout (on the left) and the realized test MEA (on the right) with circular and square platinum and indium-tin oxide electrodes. The electrode side dimensions vary between 5 µm and 150 µm Electrode properties versus frequency and material An impedance and phase shift measurement versus frequency of different types of electrodes was achieved with the following MEA devices and corresponding electrode materials: MEA devices (see Chapter 2.3.2) with smooth platinum electrodes deposited by a sputtering process. Electrode dimensions were 40 µm x 40 µm. MEA devices (see Chapter 2.3.2) with smooth platinum electrodes covered with an electroplated platinum black layer. This black platinum layer was deposited using a solution containing 1% of platinum chloride H 2 PtCl 6 / 6H 2 O, 0.01% of lead acetate Pb(COOCH 3 ) 2 / 3H 2 O and % of HCl [13]. For good electroplating results, the deposition was done at room temperature with a current density of 30 ma/cm 2 during 2 min. Electrode dimensions were 40 µm x 40 µm. MEA devices (see Chapter 2.3.1) with indium-tin oxide electrodes manufactured from commercially available glass substrates covered with a 100 nm thick indium-tin oxide layer. Its layer resistivity is below 20 Ω/square. These glass substrates were provided by Merck Balzers AG, Liechtenstein. However, no platinum was deposited at the electrode sites. Electrode dimensions were 40 µm x 40 µm. 77

88 3 Electrical properties of multi-electrode arrays Fig. 3.8 Experimental results of impedance and phase shift measurements (at 1 khz) for smooth platinum and indium-tin oxide electrodes versus electrode dimensions. 78

89 3.4 Experimental electrode characterization Fig. 3.9 Electrode resistance and capacitance calculated from the data plotted in Fig. 3.8 versus electrode dimensions. 79

90 3 Electrical properties of multi-electrode arrays Fig Experimental results of impedance and phase shift measurements for four different material electrodes versus frequency. Note that the TiN electrodes have smaller geometrical dimensions (710 µm 2 ). 80

91 3.4 Experimental electrode characterization Fig Electrode resistance and capacitance calculated from the data plotted in Fig versus frequency. The values were normalized to a 1 mm 2 electrode area. 81

92 3 Electrical properties of multi-electrode arrays Commercially available MEA60 devices from Multi Channel Systems, Germany [14]. The electrodes of these MEAs are made of titanium nitride (TiN) deposited by a sputtering process. The electrodes are circular with a diameter of 30 µm. Measurements were done by sweeping the frequency from 100 Hz to 20 khz on the previously described MEA electrodes. The electrode resistance and the electrode capacitance were calculated and normalized to 1 mm 2 area, considering that the resistance is inversely proportional to the area while the interface capacitance is directly proportional to it. The obtained results shown in Fig and Fig are in good agreement to results described in literature [1, 2, 10, 15, 16]. The electrode impedance decreases in the measured frequency range as Z = k/ F. This is due to the Warburg resistance (see Eqn (3.7)), which is frequency dependent. In the frequency range of 1 MHz, the Warburg impedance becomes small compared to the interface capacitance, which becomes determinant for the electrode impedance. At higher frequencies, the impedance remains at a few kω, representing the resistance polarization effect and the solution resistance. On the other hand, the phase shift also tends to a value representing the interface capacitance and remains at this value. However the shapes of the traces show similarities for the different electrode materials. The obtained results show that porous materials (black platinum and titanium nitride) have better electrode properties due to a larger effective area due to porosity. It results a smaller electrode resistance and a larger capacitance, which reduces the noise level and increases the current injection capability of the electrodes. Furthermore, due to a larger interface capacitance for porous electrodes, the Warburg impedance reaches faster the point at which it becomes comparable to the interface capacitance. Therefore, the electrode resistance value is shifted to lower frequencies compared to non porous material electrodes. Note that the values are shifted one decade to the left in case of smooth platinum that is covered with black platinum (porous) in comparison to smooth platinum electrodes as can be shown in Fig (top) Global noise level of MEAs The global noise level that can be measured through the Multi Channel Systems MEA interface and data acquisition system is directly displayed by the data acquisition software. The measurement of the noise level is thus very easy to achieve and is the simplest way to control MEA integrity and electrode quality. The noise level obtained for the different materials is presented in Table 3.1. The best results were obtained for the porous materials (platinum black and TiN). 82

93 3.5 References Electrode material and dimensions Noise level in [µv] Indium-tin oxide, 1600 µm Smooth platinum, 1600 µm Platinum black, 1600 µm Titanium nitride, 710 µm Table 3.1 Noise level measured by the Multi Channel Systems MEA interface and data acquisition system for different electrodes. These values can also be obtained by mathematical calculation of the low pass cut-off frequency and the thermal noise described in Eqn (3.15) and Eqn (3.14) using the data from Fig These noise level values are satisfactory for monitoring of biological cells as long as the signal to noise ratio value remains large, i.e. as long as the magnitude of biological signals measured through extracellular electrodes are in the range of several hundreds of microvolts. 3.5 REFERENCES [1] R. S. C. Cobbold, Transducers for biomedical measurements: principles and applications, John Wiley & Sons, New York, 1974 [2] L. A. Geddes, Electrodes and the measurement of bioelectric events, John Wiley & Sons Inc., New York, 1972 [3] E. T. McAdams, A. Lackermeier, et al., The linear and non-linear electrical properties of the electrode-electrolyte interface, Biosensors & Bioelectronics, Vol. 10, pp , 1995 [4] G. T. A. Kovacs, Introduction to the theory, design, and modeling of thinfilm microelectrodes for neural interfaces, in D. A. Stenger and T. M. McKenna, Enabling technologies for cultured neural networks, Academic Press, Inc, San Diego, 1994 [5] R. C. Weast and ed., CRC Handbook of chemistry and physics, CRC Press, Boca Raton, Florida, USA, 1988 [6] T. Sakurai and K. Tamaru, Simple formulas for two- and three-dimensional capacitances, IEEE Transactions on Electron Devices, Vol. ED-30, N 2, pp. 183, 185, 1983 [7] L. S. Robblee and T. L. Rose, Electrochemical guidelines for selection of protocols and electrode materials for neural stimulation, in W. F. Agnew and D. B. McCreery, Neural protheses: fundamental studies, Prentice Hall Biophysics and Bioengineering Series, New Jersey,

94 3 Electrical properties of multi-electrode arrays [8] S. B. Brummer and T. J. Turner, Electrical stimulation of the nervous system: the principle of safe charge injection with noble metal electrodes, Bioelectrochemistry and Bioenergetics, Vol. 2, pp , 1975 [9] R. L. White and T. J. Gross, An evaluation of the resistance to electrolysis of metals for use in biostimulation microprobes, IEEE Transactions on Biomedical Engineering, Vol. 21, N 11, pp , 1974 [10] A. Blau, C. Ziegler, et al., Characterization and optimization of microelectrode arrays for in vivo nerve signal recording and stimulation, Biosensors & Bioelectronics, Vol. 12, N 9-10, pp , 1997 [11] R. Fröhlich, A. Rzany, et al., Electroactive coating of stimulating electrodes, Journal of Materials Science: Materials in Medicine, Vol. 7, pp , 1996 [12] T. M. Silva, J. Rito, et al., Electrochemical response of irridium oxide for implanted neural stimulation electrodes, Journal of Materials Science: Materials in Medicine, Vol. 7, pp , 1996 [13] H. Oka, K. Shimono, et al., A new planar multielectrode array for extracellular recording: application to hippocampal acute slice, Journal of Neuroscience Methods, Vol. 93, pp , 1999 [14] [15] A. R. Varlan and W. Sansen, Characterisation of planar electrodes realised in planar microelectronic technology, Medical & Biological Engineering & Computing, Vol. 34, N 7, pp , 1996 [16] S. J. Carter, C. J. Linker, et al., Comparision of impedance at the microelectrode-saline and microelectrode-culture medium interface, IEEE Transactions on Biomedical Engineering, Vol. 39, N 11, pp ,

95 4 INVESTIGATIONS USING PLANAR MULTI-ELECTRODE ARRAYS Multi-electrode arrays were realized for a collaboration with two research groups at the University of Bern, Switzerland. Investigations on spinal cord cultures and cardiomyiocytes monolayer cultures using a custom made signal amplification and data acquisition system, and provided multi-electrode array devices are presented below. 4.1 INTRODUCTION This chapter reports the work done in collaboration with Dr. J. Streit and Dr. S. Rohr from the Department of Physiology at the University of Bern, Switzerland. Due to specific culture conditions and cell patterning in culture, different multi-electrode arrays were needed for adequate experimental conditions. A custom-made MEA measurement system described below was developed at the University of Bern including a culture chamber, a multi-electrode array chip being only the bottom part of this culture chamber. Therefore, the main difference of the MEAs fabricated for this collaboration in comparison to the previously described MEAs is that no back end packaging was necessary. The electrode array chips are directly used for cell culturing and patterning. Two different biological research purposes, i.e. the investigation of rhythmic activity in organotypic and dissociated spinal cord cultures and the investigation of heart beat variability and conduction characteristics in cardiomyiocytes monolayer cultures will be presented in this chapter. 4.2 MEA MEASUREMENT SYSTEM Planar multi-electrode arrays with different layouts were designed and manufactured for a MEA measurement system developed at the University of Bern. The main difference between these MEAs and those described in Chapter 2.3 is the chip dimension, and thus the number of available electrodes. Chip dimensions are 21 mm x 21 mm allowing integration of 68 recording electrodes and 4 ground electrodes. Due to special culturing conditions of organotypic 85

96 4 Investigations using planar multi-electrode arrays Fig. 4.1 Picture of a 72 channel data acquisition MEA interface. A MEA is inserted in the middle of the interface and builds the bottom part of a special plexiglas culture chamber. Small steel shanks insure electrical contact between the interface and the MEA chip for signal output. spinal cord cultures by a roller tube method [1-3] and to a photolithographic patterning of cardiomyiocytes monolayer cultures, these MEAs are not mounted under a printed circuit board. The layouts of each realized MEA devices are customized for specific experimentation purposes, and will thus be described in the corresponding following sections. The whole measurement system is composed of an interface and a data acquisition setup. The interface integrates a MEA culture chamber assembly and signal output circuitry. The external data acquisition system is composed of a signal amplification circuitry and computer data acquisition hard- and software. MEA interface In Fig. 4.1, the MEA interface is shown. Due to square MEA chip geometry, 18 electrode outputs are placed on each chip side. Small steel shanks insure electrical contact between the MEA chip and the external circuitry. For the experiments, a MEA is placed into a special plexiglas chamber which is mounted in the center of the interface, and superfused. External data acquisition system Each electrode is AC-coupled to an individual custom-made amplifier. The amplified signals are digitized at a rate of 6 khz with 12 bit resolution (NI-DAQcard, AT-MIO-64E-3, National Instruments) and stored on a computer hard disk 86

97 4.3 MEAs for spinal cord cultures for off-line analysis. The Analog/Digital-card is controlled by a custom made Labview program allowing also the display of the measured channels. 4.3 MEAS FOR SPINAL CORD CULTURES Introduction Local spinal cord networks are able to generate a rhythmic output which is directed to muscles used for locomotion. Such patterns consist of alternating activity produced by central pattern generators (CPGs) for flexor and extensor muscles. Based on extensive studies on different spinal cord preparations, the functional architecture of CPGs is proposed to be that of coupled unit oscillators. In the intact spinal cord preparation of neonatal rat, fictive locomotion can be induced by the application of NMDA (N-methyl-D-aspartate) and or serotonin (5-hydroxyztyptamine) or by increasing the extracellular K + concentration. Besides fictive locomotion patterns, spinal networks produce a second kind of rhythmic activity when disinhibited by the application of γ-aminobutyric acid (GABA A ) and glycine receptor antagonists [4]. Such patterns are characterized by slow synchronous bursting. Recent findings have shown that there are strong interactions between these two types of rhythmic activity. It is an open question whether the same spinal network and network mechanisms underlie all of the expressed pattern. Although some attempts have been made to localize CPGs in the transverse plane of the spinal cord using activity dependent staining and calcium imaging [5], however, not much is known about the source and the propagation of activity. Therefore the possible role and location of pacemaker cells and the identity of networks involved in various forms of rhythmic activity remain unclear. One approach to address such questions is the use of organotypic slice cultures offering direct access to networks of spinal interneurons. However, recordings from one or two sites in the spinal cord do not allow determination of the source and propagation of activity. Therefore, multisite recording by growing spinal cord slice cultures directly on a planar MEA was used. With this method the activity patterns of differently induced rhythm and the activity propagation in the slice cultures could be investigated. The spaciotemporal similarities of the activity patterns suggest a common rhythm generating network. Part of the rhythmic phenomena in spinal networks have been proposed to be based on recurrent excitation in a random excitatory network with frequencydependent synaptic depression, which implies that a specific network architecture is of minor importance for the rhythmic phenomena. Dissociated cell cultures of 87

98 4 Investigations using planar multi-electrode arrays rat embryonic spinal cord grown on multi-electrode arrays were used as randomized networks, and their patterns of rhythmic activity were compared to those of the slice. In addition, neuronal discharge in the absence of synaptic transmission and its contribution to rhythmic activity patterns was investigated. All these organotypic and dissociated spinal cord culture findings will be described in detail elsewhere [6, 7] Used multi-electrode array layouts Multi-electrode arrays were composed of a glass substrate (thickness of 700 µm, dimension 21 mm x 21 mm), indium-tin oxide electrodes (thickness of 100 nm, dimension of 40 µm x 40 µm), and an SU-8 polymer insulation layer (thickness of 5 µm). The recording site is composed of 68 electrodes disposed on a hexagonal grid (see Fig. 4.2) or disposed on a 6x12 matrix without the 4 corner electrodes (see Fig. 4.3) with an electrode center interspace of 200 µm. The recording electrodes have an impedance of 1 MΩ at 1 khz in extracellular solution. Four large counter electrodes (~5 mm 2 ) are integrated on the MEA around the recording site. The device is fully transparent allowing good observation of the tissue culture on an inverted microscope. Some of the MEAs were realized with smooth platinum electrodes, i.e. the effective electrode measurement area was covered with a titanium adhesion layer (50 nm) and a platinum (150 nm) layer. The platinum electrodes have lower impedances (between 250 kω and 350 kω at 1 khz) resulting in lower noise. However, platinum covered electrodes are not transparent Results: spinal cord slice cultures Organotypic cultures of spinal cord including dorsal root ganglia and skeletal muscle were obtained from 14-day old rat embryos (E 14) and cultured for up to 3 weeks using a roller tube technique described in literature [1-3]. The experiments were performed at room temperature after 10 to 15 days invitro. For the experiments, a MEA with an organotypic spinal cord culture was placed into a special plexiglas chamber, mounted on the measurement system interface described in Chapter 4.2 and superfused with a bath solution of the following composition in mm: NaCl 145, KCl 4, MgCl 2 1, CaCl 2 2, HEPES 5, Na pyruvate 2, glucose 5 at ph 7.4. High K + solution was obtained by replacing up to 3 mm NaCl by KCl. The bath solution was exchanged every 10 to 15 min during the recording session, which usually lasted for 5 to 9 hours. All recordings were made in the absence of solution flow. All chemical agents were applied by exchanging the bath solution for a solution containing the drug. The following 88

99 4.3 MEAs for spinal cord cultures Fig. 4.2 On the left, picture of the 72-electrode MEA layout. Four large ground electrodes are disposed around the measurement area composed of 68 electrodes disposed on an hexagonal grid with an interspace of 200 µm. On the right, picture of the obtained electrode array composed of indium-tin oxide leads and 40 µm x 40 µm large smooth platinum electrodes. This MEA was realized for organotypic spinal cord cultures. Fig. 4.3 On the left, picture of the 72-electrode MEA layout. The measurement area is composed of 68 electrodes disposed on a 6x12 matrix with an interspace of 200 µm. On the right, picture of the obtained electrode array composed of 40 µm x 40 µm large indium-tin oxide electrodes. This MEA was realized for dissociated spinal cord cultures. 89

100 4 Investigations using planar multi-electrode arrays Fig. 4.4 Shape of recorded signals and event detection. A) Upper trace: voltage recording of spontaneous activity in control solution with one electrode at two time resolutions. One electrode records spikes simultaneously from several neurons due to its size. The spikes appear mainly in clusters, which probably represent quasisynchronous action potentials in a group of neurons. Lower trace: Time markers found by an event detector, which recognizes single isolated spikes as one event and translates overlapping spikes into trains of events. B + C) Two simultaneous voltage recordings in the presence of bicuculline (20 mm) and strychnine (1 mm) at three time resolutions. During episodes of spike clusters, large voltage changes of several seconds appear at some electrodes chiefly located dorsally (lower traces in B). In addition, oscillating voltage changes can be observed mainly at electrodes located ventrally (upper traces in B). The highest time resolution reveals the presence of one spike cluster per oscillation period (from [7]). 90

101 4.3 MEAs for spinal cord cultures agents were used: bicuculline and strychnine (all from Sigma). After a period of about 45 min, during which an increase of the mean rate of activity and of the signal to noise ratio was observed, the activity remained stable for the rest of the experiment. Spontaneous activity In control solution (1 mm Mg 2+ ) activity consisted of single spikes and spike clusters as shown in Fig The single spikes were not synchronized between different electrodes, whereas spike clusters were simultaneously recorded at most electrodes from the left and right sides of the spinal cord cultures. The start of the spike clusters was delayed at the different electrodes, showing that a wave of activity spread throughout the network. Such synchronized spike clusters are therefore called activity waves. The majority of the activity recorded in spinal cord slice cultures was made up by activity waves. They typically were short lasting (88±4 ms) and could be grouped into burst waves (see Fig. 4.5 and Fig. 4.6). Burst waves appeared in an irregular way and lasted one to 5 seconds, and occasionally for up to several tens of seconds. Activity was usually more pronounced in the ventral parts of the spinal cord slices, especially around the central fissure. Rhythmic activity Rhythmic activity is most reliably induced in slice cultures through disinhibition by strychnine and bicuculline, the GABA A and glycine receptor antagonists. Under such conditions, more or less regular, slow bursting of individual motoneurons or interneurons as well as burst-like contractions of cocultured muscle fibers appeared [4, 8]. In the present study, bursting network activity with all these characteristics has now been found with MEA recordings of spinal networks in the slice cultures. The multisite recordings showed that both the bursts and the oscillations are synchronized over the whole network. In addition, simultaneous recordings of activity from the spinal network and co-cultured muscle fibers reveal that muscle fibers contract during network activity. Therefore, the network activity is directed to motor neurons innervating the muscle fibers, in the early phase of the activity waves. Under bicuculline and strychnine, activity consisted of bursts at low rates (1 to 10 per minute) and in many cases of intraburst oscillations at high rates (2 to 8 per second). The sources of burst were mainly situated ventrally on both sides of the central fissure. The pathways of network recruitment from one source were variable between bursts, however, in average the bursts propagated first form one to the other ventral side of the spinal cord and finally to the medial part within less than 80 ms (see Fig. 4.7). 91

102 4 Investigations using planar multi-electrode arrays Fig. 4.5 Spontaneous activity in an organotypic spinal cord culture (after 13 days in -vitro) under control conditions grown on a MEA. Activity was recorded from electrodes with a number or a star (from [7]). Rhythmic activity cannot only be induced by disinhibition but also by decreasing extracellular Mg 2+ to 0 mm or by increasing extracellular K + to 7-8 mm, a common effect of these two last procedures, which among others is an increase in neuronal excitability. Under these conditions, the duration and period of the activity waves were similar to those under disinhibition, where the duration and period of bursts was shorter. This suggests at least in part that the same neurons were involved in the generation of these patterns (although not proof since single unit resolution was lacking). In addition, the bursts spread over the same areas of the slices with also a high density of burst sources around the central fissure, excluding the involvement of spatially separated networks in the slice Results: dissociated spinal cord cultures Dissociated cultures were made from the whole spinal cord of rats at embryonic age 14 (E 14). The fetuses were delivered and prepared in the same way as described previously for organotypic slice cultures [1]. The same plexiglas chamber, measurement system and medium was used as for the slice cultures. The following agents were used: Bicuculline, strychnine, 92

103 4.3 MEAs for spinal cord cultures Fig. 4.6 Top: Event raster plot of the electrodes 1-9 from Fig. 4.5 at two different time resolutions. The activity consists of single spikes (s), of single activity waves (w) and of bursts (b) with different duration (c = spike cluster). A burst consists of several activity waves appearing at frequencies of 1-8 Hz. Not all waves propagate over all recording sites. The activity waves are highly synchronized between both sides of the culture. Bottom: Network activity plot. It shows the number of events detected from all selected channels in Fig. 4.5 within a sliding window of 10 ms (events/10 ms) of the recording shown at top (from [7]). 93

104 4 Investigations using planar multi-electrode arrays Fig. 4.7 Location of the burst sources and propagation of the activity in the presence of bicuculline (20 mm) and strychnine (1 mm). The lines (isochrones) show how long it takes that the wave front of the burst reaches the different electrodes (in ms). A) Burst sources: the size of the black circles shows at which percentage (numbers on circles), bursts started at the corresponding places. Bursts started in less than 5 % of cases at the electrodes marked with a star. B) Variation in the propagation of three individual bursts starting at the same source (*). This variation in the propagation of individual bursts beginning at the same source is observed in all protocols. Sometimes the wave fronts of the bursts do not propagate in a sliding way but jump from one area to an other ( ). C) Mean propagation pattern (5-7 bursts) of the bursts starting at the three sites indicated with stars. The bursts propagate first from one side to the other side of the slice, then to the medial and dorsal parts. D) Mean propagation of the oscillating activity waves of the last part of the bursts. This propagation shows a similar pattern as the wave front propagation of the bursts but is slower and does not reach the whole network. Isochrones were extrapolated from the latencies at the individual electrodes using a conventional algorithm (from [7]). 94

105 4.3 MEAs for spinal cord cultures AP-5 (APV), and CNQX (all from Sigma). One recording session usually lasted for 5 to 7 hours. Native spontaneous activity Experimental data were obtained from 29 cultures and a total of 329 electrodes. In control solution containing 1 or 2 mm Mg 2+ all cultures showed spontaneous activity. Following an initial period of an increase in the mean rate of activity during the first 20 to 30 min of the recording sessions, the activity remained rather stable for the rest of the experiments. In the majority of the cultures such spontaneous activity showed no evident rhythmic structure and was therefore considered as random and consisted of local activity, waves and bursts. Local activity was defined as spatially restricted activity, one or a few spikes seen at one or few electrodes. At low rates of activity, reliable synchrony in the activity of two or more channels on a time-scale < 1 ms only rarely occurred, suggesting that the activity of one cell was usually not seen by more than one electrode. Waves consisted of episodes of consecutive spikes in most of the channels lasting for 50 to 100 milliseconds (see Fig. 4.8 C and D). Most of the bursts contained fluctuating continuous activity at most of the channels lasting for hundreds of milliseconds to seconds. Furthermore, in a small subset of cultures (3 of 29), rhythmic patterns as described below appeared as native patterns of activity in normal control solution. Intrinsic neuronal activity and its modulation during bursting To investigate the role of intrinsic spontaneous activity of neurons for burst initiation, network activity in the presence of blockers of synaptic transmission was measured. Synaptic transmission was blocked using strychnine, bicuculline and CNQX (10 µm) for blocking AMPA receptors, in most cases in combination with APV (50 µm) for blocking NMDA receptors. However, the application of bicuculline and strychnine induces rhythmic activity, blockade of synaptic transmission caused the disappearance of bursts and waves together with a drastic decrease in network activity rate and in the network recruitment (see Fig. 4.9 A and B). The asynchronous activity remaining under such conditions was attributed to intrinsic neuronal activity (see Fig. 4.9 B). There was usually a good correspondence between the level of intrinsic activity and the probability of burst initiation at individual sites (see Fig. 4.9 D and E). 24 of 34 burst sources (71%) showed intrinsic activity compared to 22 of 66 non sources (33%), showing that the probability of burst initiation was significantly higher at sites of intrinsic activity. Of a total of 46 sites with intrinsic activity, 24 (52%) were burst sources, suggesting that roughly half of the intrinsically active sites were powerful enough to activate the whole network. 95

106 4 Investigations using planar multi-electrode arrays Fig. 4.8 Native spontaneous activity in dissociated spinal cord cultures. A) Spinal culture on a MEA viewed at two scales. The indium tin oxide electrodes are covered with platinum and have a size of 40 µm x 40 µm. The relative sizes of neurons and electrodes are shown on the right. B) Extracellular voltage recordings from the electrodes indicated in A at two time resolutions. C) Event raster plot showing the time markers of activity of all channels shown in B. Individual spikes are represented by one marker, trains of overlapping spikes are represented by series of markers. D) Network activity plots showing the activity of all channels shown in C summed within a sliding bin of 10 ms (from [6]). 96

107 4.3 MEAs for spinal cord cultures Fig. 4.9 Intrinsic neuronal activity and its modulation during bursting in dissociated spinal cord cultures. A) Event raster plots of bursting activity induced by disinhibition in the presence of low (0 mm, left) and high Mg 2+ (2 mm, right). Note the activity in the interburst intervals, especially in high Mg 2+, which disappears following the bursts and recovers after various delays between 5 and 15 s. B) Event raster plots of activity recorded following the addition of CNQX and AP-5 in the same experiments as in A. Note that asynchronous intrinsic neuronal activity persists in many channels in the absence of synaptic transmission. C) Superposition of network activity plots of A (black) and B (gray), averaged within a sliding window of 2 s. Note that in low Mg 2+ the network activity in the burst intervals is much smaller than the intrinsic neuronal activity, whereas in high Mg 2+ it reaches the level of intrinsic activity after an initial suppression after the bursts. D) Burst initiation sites found for bursting in low Mg 2+ as shown in A. E) Sites of intrinsic activity recorded in low Mg 2+ as shown in B. The size of the diamonds represents the level of intrinsic activity at the given channel. The smallest dots represent channels only showing activity in the presence of synaptic transmission (from [6]). 97

108 4 Investigations using planar multi-electrode arrays Intrinsically active neurons have been found in cultured neurons of the spinal cord. However their location and their role in spinal pattern generation remained unclear. Intrinsically active cells were found at a high percentage of the locations where bursts started, suggesting that they initiated the bursts. The fact that not all of the burst sources showed intrinsic activity is easily explained by the low spatial resolution of multisite recording. A burst source at one specific electrode does not necessarily mean that this electrode actually sees the intrinsically active cell. The site could also be in the close neighborhood not seen by one of the electrodes. Conclusions To test the hypothesis that rhythmic activity in the spinal cord may be based on immanent properties of unstructured neuronal networks, MEAs have been used to compare the patterns of rhythmic spontaneous activity in organotypic slice cultures and in dissociated cultures of the same origin. It was found that disinhibition as well as certain protocols of increased excitation induced spontaneous bursting activity in dissociated cultures of embryonic rat spinal cord. The bursts were initiated at various sites from which the fronts of activity propagated to the whole network with a high variability between individual bursts. Such burst sources corresponded well to the sites of intrinsic spontaneous activity in the absence of synaptic transmission, suggesting that such intrinsic activity is the source of bursting. In summary, it is proposed that the patterns of regular activity in dissociated as well as in slice cultures of the spinal cord are determined by mutually linked network parameters, and not by a specified architecture of the network. These parameters are intrinsic neuronal activity, modulation of excitability, recurrent excitation through excitatory synaptic coupling and synaptic depression. When these factors are optimally adjusted, the network functions as a rhythm generator. The use of the MEA measurement technique in this study was of importance because no other measurement technique would provide data from the overall spinal cord culture that might be sufficient to establish bursting source location and burst propagation. Furthermore, the use of MEAs provides data from several cells at the same time, which reduces the number of needed experiments to obtain the same amount of data in comparison to the patch clamp technique. 98

109 4.4 MEAs for cardiomyiocytes cultures 4.4 MEAS FOR CARDIOMYIOCYTES CULTURES Measurement technique combination The three most widely used techniques for recording electrical activity in cardiac preparations are the conventional use of classical glass micropipettes, the use of optical methods based on potentiometric dyes and the use of extracellular electrodes. With glass micropipettes, it is possible to record the transmembrane potential during prolonged periods of time. However, it is technically difficult to record from many sites simultaneously. On the other hand, the use of voltagesensitive dyes allows to record transmembrane potential changes from many sites with photodetector arrays. Unfortunately, such recordings are limited in time due to the onset of the phototoxic effects of the dyes. For multisite long-term recordings, methods using arrays of extracellular electrodes were developed and are successfully used to record from the intact heart or from pieces of cardiac tissue, at the expense of recording extracellular rather than transmembrane potentials. Multi-electrode arrays designed for long-term multisite recordings from cardiac cell cultures and for electrical stimulation of the preparations are composed of a glass substrate (thickness of 700 µm, dimension 21 mm x 21 mm), indium-tin oxide electrodes (thickness of 100 nm, dimension of 40 µm x 40 µm), and an SU-8 polymer insulation layer (thickness of 5 µm). Different recording site geometries composed of stimulation and recording electrodes were realized for cell stimulation and signal conduction experiments. The main used MEA design is composed of two lines of 32 electrodes. The distance between the electrodes is 300 µm. A pair of large platinum stimulation electrodes with dimensions of 250 µm x 50 µm is placed at the end of both sides of an electrodes line (see Fig left). The recording electrodes have an impedance of 1 MΩ at 1 khz in saline solution (KCl 0.9 %). The stimulation electrodes are covered with a titanium adhesion layer (50 nm) and a platinum (150 nm) layer. The device is transparent allowing good observation of the heart cell cultures by inverse microscopy. Due to the transparency of the MEAs, electrical and optical monitoring of heart cell activity can be combined. Neonatal rat heart cells were grown onto the realized MEAs and extracellular potentials could be measured (see Fig. 4.10) under incubating conditions for up to two days. Moreover, optical recording of transmembrane voltage was performed simultaneously using a voltage sensitive dye. The possibility to combine these two measurement techniques permits specific measurement of impulse propagation using voltage sensitive dyes 99

110 4 Investigations using planar multi-electrode arrays Fig The image in the center of the figure depicts two extracellular electrodes of an array on which a monolayer of neonatal rat heart cells was grown. The preparation was mounted in the measurement system which was run for 46 hours under incubating conditions. Extracellular action potentials (V e ), recorded from these electrodes at the end of this period are shown on the left. The preparation was then mounted in a setup for optical recording of transmembrane voltage. Optical signals of transmembrane action potential upstrokes (V t units normalized on action potential amplitude, APA) were recorded using a voltage sensitive dye from the sites marked on the image as circles and are plotted on the right. The traces to the right (dv t /dt) correspond to the time derivatives of the optically recorded upstrokes. It illustrates the possibility of long-term experiments and the combination of extracellular recordings with transmembrane optical recordings through the transparent indium-tin oxide leads and electrodes (courtesy of Dr. Rohr, University of Bern). (measurement occurs only at light flashes) and the continuous monitoring of the activity at several locations giving a global insight into the culture activity. 100

111 4.4 MEAs for cardiomyiocytes cultures Analysis of beat rate variability in spontaneously active cardiac cell cultures The analysis of beat rate variability in cell cultures is motivated by the fact that the analysis of heart rate variability in-vivo (HRV) can be used to assess the balance of the autonomic regulation of the cardiovascular system and to provide, non-invasively, markers suited for the assessment of prognosis and for riskstratification of cardiac patients. It is widely accepted that extracardiac influences (nervous, humoral, hemodynamic) participate in HRV. However, it is not known to what extent intrinsic properties of cardiac tissue might also play a role. Therefore, beat rate variability was assessed in spontaneously beating monolayer cardiac cell cultures which are largely devoid of the above mentioned extracardiac influences. Fig A illustrates a typical unipolar extracellular recording with three spontaneous action potentials. The first action potential is also shown on an expanded time scale in Fig B. The extracellular signals have a typical biphasic shape. The activation time for each action potential was defined as the time of occurrence of the minimum peak of the first derivative (vertical line). From the knowledge of successive activation times it is possible to reconstruct beat rate time series as illustrated in Fig C for a 1 hour recording period. The oscillations in the trace correspond to variations of beat rate. To analyze these variations in the frequency domain, the power spectrum of the beat rate time series was computed with a fast Fourier transform. The power spectral density (PSD) corresponding to the trace in Fig C is shown in Fig (note the double logarithmic axis). There was no peak in the power spectrum, which means that the variations in the culture beat rate did not occur with characteristic frequencies. This is in contrast to HRV which is characterized by periodic components related to the respiratory cycle and to baroreflex regulation. The spectrum exhibited a descending pattern which could be fitted with a line. Such a linear relationship between log(psd) and log(frequency) indicates that the power spectrum follows a power law: PSD ~ f -b. The exponent b (corresponding to the slope of the line in Fig. 4.12), which characterizes the power law, amounted to 1.3. A similar power law behavior could be found in 12 out of 13 experiments with an average exponent of 1.3±0.2 (mean±sd, range 1.0 to 1.7). Such power law behavior was previously described in HRV in vivo (with an exponent close to 1 in healthy individuals). Therefore, the present results suggest that the power law behavior of HRV might not only be based on extracardiac influences but also on intrinsic properties of cardiac tissue. 101

112 4 Investigations using planar multi-electrode arrays Fig Long-term recording of the spontaneous activity of a monolayer culture of neonatal rat heart cells. A) recording of the extracellular potential with an indium-tin oxide electrode showing three spontaneous action potentials. B) The first action potential from A) is shown on an expanded time scale with its time derivative. The activation time was defined (vertical line) as the time of occurrence of the negative peak of dv e /dt. C) plot of spontaneous beat rate over one hour under incubating conditions, as computed from successive activation times. The fluctuations of the beat rate are intrinsic to the preparation (from [9]). 102

113 4.4 MEAs for cardiomyiocytes cultures Fig Plot of the power spectral density of the beat rate time series plotted in Fig C (from [9]). Fig On the left, picture of the MEA layout composed of two lines of 32 ITO electrodes disposed every 300 µm, and 4 pairs of platinum stimulation electrodes. On the right, picture of two stimulation electrodes (50 µm x 250 µm) covered with a monolayer cardiomyiocytes pattern (courtesy of Dr. Rohr, University of Bern). More information about power-law behavior of beat-rate variability findings has been published elsewhere [9]. 103

114 4 Investigations using planar multi-electrode arrays Fig Top: pictures of a line- and comb-shaped pattern of dissociated cells from neonatal rat ventricular tissue on a line of ITO electrodes (courtesy of Dr. Rohr, University of Bern.) Bottom: simultaneous measurement of 16 electrodes along the electrode line after a stimulation pulse. The stimulation artefact is shown at measurement start. An evoked response propagates along the cell pattern geometry from one electrode to the next. It is shown that the propagation velocity is reduced by a factor 2 by a comb-shaped pattern compared to a line pattern (from [10]) Structure function relationship It is known that geometrical discontinuities in excitable tissue represent socalled current-to-load mismatches. For example, in the situation where a parent strand releases branches, the current generated in the parent strand from which an action potential is approaching will disperse in the daughter branches, which will result in a local slowing of conduction. If the parent strand releases branches at regular intervals over its entire length, the resulting conduction pattern will, at a 104

115 4.5 References macroscopic scale, appear slower compared to conduction in an unbranched strand. Fig illustrates such a pattern of slow conduction induced by branching tissue geometry. In this experiment, the technique of the patterned growth of cardiac cell cultures [11, 12] was combined with the use of MEAs. Patterns of activation were assessed with a row of 16 extracellular electrodes (at 300 µm intervals) along a control unbranched strand (100 µm wide, 10 mm long, Fig. 4.14, left) and along a multiple branching preparation (a 10 mm long and 100 µm wide main strand releasing 500 µm long and 100 µm wide branches to both sides at equidistant intervals of 300 µm, Fig. 4.14, right). The preparations were stimulated at one extremity with bipolar electrodes as shown in Fig (right). The traces recorded by the rows of 16 electrodes are shown on the right side of the schematic drawings of the preparations, for both experiments shown in Fig The initial synchronous deflection corresponds to the stimulation artifact and the second propagated deflection to the conducted action potential. At this macroscopic scale, both activation patterns appeared linear. However, conduction was about 45% slower in the multiple branching preparation. This experiment is in accordance with results obtained in a previous study [13] and shows that this type of slow conduction along branching substrates can be supported over macroscopic distances. Therefore, this mechanism of slow conduction might be involved in the establishment of slow conduction in cardiac tissue with a discontinuous architecture such as the physiologically slowconducting atrioventricular node (where so-called dead-end pathways were functionally described) or in infarct scars ( mottled myocardium ), where slow conduction increases the risk of reentrant arrhythmias Conclusions These two examples illustrate the development of a technique using microelectrode arrays to investigate electrical activity in cardiac cell cultures. With the possibility to perform long-term recordings, this method represents a promising tool for the study of the long-term behavior of cardiac tissue in vitro, like, e.g., developmental aspects, electrical remodeling or rate-dependent phenomena as a function of tissue structure. 4.5 REFERENCES [1] U. F. Braschler, A. Iannone, et al., A modified roller tube technique for organotypic cocultures of embryonic rat spinal cord, sensory ganglia and 105

116 4 Investigations using planar multi-electrode arrays skeletal muscle, Journal of Neuroscience Methods, Vol. 29, pp , 1989 [2] B. H. Gähwiler, Organotypic monolayer cultures of nervous tissue, Journal of Neuroscience Methods, Vol. 4, pp , 1981 [3] J. Streit, C. Spenger and H.-R. Lüscher, An organotypic spinal cord - dorsal root ganglion - skeletal muscle coculture of embryonic rat. II. functional evidence for the formation of spinal reflex arcs in vitro, European Journal of Neuroscience, Vol. 3, pp , 1991 [4] J. Streit, Regular oscillations of synaptic activity in spinal networks in vitro, Journal of Neurophysiology, Vol. 70, N 3, pp , 1993 [5] M. O'Donovan, S. Ho and W. Yee, Calcium imaging of rythmic network activity in the developing spinal cord of the chick embryo, Journal of Neuroscience, Vol. 14, N 11, pp , 1994 [6] J. Streit, A. Tscherter, et al., The generation of rhythmic activity in dissociated cultures of rat spinal cord, submitted at European Journal of Neuroscience, 2001 [7] A. Tscherter, M. O. Heuschkel, et al., Spatiotemporal characterisation of rhythmic activity in spinal cord slice cultures, submitted at European Journal of Neuroscience, 2001 [8] J. Streit, Mechanisms of pattern generation in co-cultures of embryonic spinal cord and skeletal muscle, International Journal of Develomental Neuroscience, Vol. 14, N 2, pp , 1996 [9] J. P. Kucera, M. O. Heuschkel, et al., Power-law behavior of beat-rate variability in monolayer cultures of neonatal rat ventricular myocytes, Circulation Research, Vol. 86, pp , 2000 [10] J. P. Kucera, M. O. Heuschkel, et al., Monitoring of electrical activity in cardiac cell cultures with extracellular electrode arrays, International Workshop on Computer Simulation and Experimental Assessment of Electrical Cardiac Function, Lausanne, Switzerland, December 7-8, pp , 1998 [11] S. Rohr, D. M. Schölly and A. G. Kléber, Patterned growth of neonatal rat heart cells in culture, Circulation Research, Vol. 68, pp , 1991 [12] S. Rohr and J. P. Kucera, Involvement of the calcium inward current in cardiac impulse propagation: induction of unidirectional conduction block by nifedipine and reversal by Bay K 8644, Biophysical Journal, Vol. 72, N 2, pp , 1997 [13] J. P. Kucera, A. G. Kléber and S. Rohr, Slow conduction in cardiac tissue: II. Effects of branching tissue geometry, Circulation Research, Vol. 83, pp ,

117 5 THREE DIMENSIONAL MULTI- ELECTRODE ARRAYS Cutting thin tissue slices from brain or spinal cord generates cell damage and cell death at the first cell layers. To overcome problems raised by these dead cell layers, it was assumed that three dimensional electrodes should penetrate in the tissue and become closer to the active sites. In this chapter, tip-shaped electrode fabrication methods are first presented. Measurements done with 3D multi-electrodes arrays on acute rat hippocampal slices showed that larger signal amplitudes can be obtained, which will open new opportunities in pharmacology and toxicology research on brain slice preparations. 5.1 THE NEED OF THREE DIMENSIONAL ELECTRODES The study of electrical measurements on acute or fresh tissue slice preparations allows to improve research methodology by maintaining synaptic organization. Acute preparations can be measured in a time window of 1 to 8 hours after animal death in comparison to organotypic slice cultures which need a culture time of more than one week before electrical cell activity could be measured. Another advantage of acute slices versus organotypic slices is that the animal model is preserved. When cells are cultured over a long time period, the connections between the cells change, and after one week, the whole network is different than the initial network that was present in the animal. Unfortunately, it is more difficult to measure activity in acute slice preparations than in organotypic cultures. This is the reason why acute slice preparations are not much used nowadays for pharmacological and/or toxicological experimentation. Why is it so difficult to record the activity in acute slice preparations? The main reason is that the active cells of the slice are far from the measurement electrodes. The tissue samples are prepared by cutting thin slices with thicknesses between 300 and 500 µm. However, cutting the tissue damages the first cell layers at the slice border and forms an electrically passive layer. This means that active cells can only be found about fifty microns inside the slice. 107

118 5 Three dimensional multi-electrode arrays Fig. 5.1 Diagrams of slice preparations on planar and three dimensional electrode arrays. Left: dead cell layers (dashed lines) lies between the active cells (on top) ant the electrodes (black). Right: because of the three dimensional shape of the electrodes, the distance between the active cells and the electrodes becomes very short due to tissue penetration allowing better measurement conditions. When electrical measurement is achieved with planar electrodes, these dead cell layers play the role of a parasitic shunt between the active cells and the electrodes. Moreover, in extracellular measurements, the signal that is measured is the electrical field around the cell. Knowing that the electrical field decreases with the square of the distance, the electrical signal that can be measured becomes very small. The main idea to overcome this problem is to make electrodes that are able to penetrate the tissue to get as close as possible to the active cells. This can be achieved by using adequate three dimensional protruding electrodes. Knowing that the active cells can be found about 50 µm inside the slice preparation, tipshaped electrodes with a height between 50 µm and 100 µm should improve MEA recording in acute slices. 5.2 TIP-SHAPED ELECTRODES Today, microstructured tips are used in scanning force and probe microscopy. These microtips are mostly made by bulk silicon etching in KOH solutions and the obtained tip-height is seldom higher than 10 µm [1-4]. DNA injection into plant and animals using 150 µm high silicon tips was also reported [5]. Another silicon micromachining method developed at university of Twente, the Black Silicon Method (BSM) [6], allows easy fabrication of tips up to a height of 200 µm by SF 6 /O 2 plasma etching. This method is also well adapted for scanning force and probe microscopy applications but can also be applied for transdermal drug delivery [7]. Selective epitaxial vapor-liquid-solid growth (VLS) of silicon wires onto thin gold layers allows also fabrication of 10 µm high and 2 µm wide tips [8]. 108

119 5.2 Tip-shaped electrodes Fig. 5.2 Comparison of tip-shapes obtained using wet chemical etching of glass in a hydro-fluoric acid solution (on the left) and wet chemical etching of silicon in a KOH solution (on the right). The height of the tips is about 25 µm. However, other fabrication techniques using other materials than silicon allow easy fabrication of microtips. Hollow microneedles made of NiFe fabricated using an SU-8 epoxy mold of a silicon mold insert for transdermal drug delivery [9] and glass tips made by wet chemical hydro-fluoric acid etching for near-field optical [10] and atomic force [11] microscopy were reported. As mentioned above, the desired electrode height required for slice penetration and active cell measurement is between 50 µm and 100 µm. For these dimensions, the best microtip fabrication method is the bulk micromachining of silicon or glass. Wet chemical etching of silicon or glass with KOH or hydrofluoric acid solutions respectively through a metallic mask allows easy fabrication of tips at substrate surface. Although, the cristallographic structure of the silicon and the amorphous structure of glass are different, the etch process characteristics and the obtained shapes of tips are very similar. Electroplated platinum hillocks with a height of 20 µm and silicon microtips fabricated by deep wet chemical etching of silicon by KOH with a height of 50 µm were reported [12-14] to improve multi-electrode arrays for hippocampus organotypic culture measurements. The main weakness of these electroplated hillocks is the mechanical fragility of the hillock base, which limits the number of its utilizations due to easy breakage. On the other hand, silicon etched tips show good mechanical properties and allowed measurement of the cellular activity. 109

120 5 Three dimensional multi-electrode arrays However, it was decided to use glass as the MEA substrate to preserve chip transparency. The next section will introduce glass wet chemical etching characteristics and results. 5.3 BULK WET CHEMICAL ETCHING OF GLASS Nature of glass Glass is the product of the fusing or melting of crystalline materials at elevated temperatures to produce liquids which have subsequently been cooled to rigid condition without crystallization. The chemical composition is largely inorganic, with silica (SiO 2 ) being the most important material. Silicate glasses are composed of three dimensional molecule networks, the basic structural unit being a silicon-oxygen tetrahedron in which a silicon atom is bonded to four oxygen atoms. Some other atoms such as sodium are present in glass. These atoms are ionically bonded to oxygen and disrupt the continuity of the network since some oxygen atoms are no longer bonded to two silicon atoms. Several inorganic oxides can be incorporated into silicate glasses. Some elements which can replace silicon atoms are called network formers. Mono- or divalent cations do not enter the network but form ionic bonds with nonbridging oxygen atoms and are called network modifiers. A full description of glass properties can be found in literature [15, 16]. Properties Composition float glass SiO %, Na 2 O 13.9%, K 2 O 0.4 %, CaO 8.4%, MgO 4.4%, Al 2 O %, FeO %, SO 3 0.3% Pyrex (Corning code 7740) SiO 2 81%, Na 2 O 4%, Al 2 O 3 2%, B 2 O 3 13% Density of glass 2.49 g/cm g/cm 3 Refractive index Transmission at 560 nm 91.1 % 91.8 % Thermal expansion coefficient C C -1 Table 5.1 Main characteristics of float glass and Pyrex (Corning code 7740). 110

121 5.3 Bulk wet chemical etching of glass Two different glasses were used for the characterization of microtip manufacturing process by wet chemical etching in HF solutions. The first glass is the borosilicate glass known commercially as Pyrex (Corning, glass code 7740) and lime glass commonly known as float glass. Vitreous boric oxide with a boric-oxygen triangle network replaces parts of the three-dimensional network of vitreous silica in borosilicate glasses, which have a low alkali oxide composition. The thermal expansion of borosilicate glasses matches the expansion of various metals and silicon. This property provides good bonding properties to silicon and is the cause of its broad use in microfabrication applications. The lime glasses are the oldest and the most widely used glasses. Lime glasses usually contain silica and alkali oxides (principally sodium oxide) and lime (mostly calcium and magnesium oxide). Usually, a small amount of alumina is included in the lime glass formulation to improve working characteristics and chemical durability Wet chemical etching in hydro-fluoric acid solutions Glass is distinguished by its great resistance to almost all chemicals at usual temperatures. Only hydro-fluoric acid (HF) makes an immediately noticeable attack on glass by bringing the glass main component, the silica, into solution according to the following chemical reaction formula: SiO 2 (s) + 6 HF (l) H 2 SiF 6 (aq) + 2 H 2 O (l) (5.1) However, the HF attack is somewhat more complicated because it is affected by the fact that reaction products due to the other glass components settle at the surface and can thus influence the etching process as is continues. Glass etching principle The basic reactions that take place are the silica structure dissolution and the dissolution of the other glass components called glass leaching. Network modifiers such as Na 2 O, K 2 O and CaO are incorporated into glass by breaking a siloxane bond, forming non-bridging oxygens by the following reaction: Na 2 O + Si-O-Si 2 SiO - Na + (5.2) In this way, the three-dimensional structure is partially broken down. Cations from the network-modifying oxides are bonded ionically to the silicate network and are therefore relatively mobile. In aqueous solutions, these mobile monovalent and bivalent ions like Na + and Ca 2+, respectively, are leached out of the glass and replaced by ion-exchange of H 3 O + which forms SiOH groups in 111

122 5 Three dimensional multi-electrode arrays Fig. 5.3 Diagram of glass etching principle. The glass surface is leached (a ion exchange takes place) and then the silicon dioxide latter is dissolved (modified from [16]). the silica structure. The formation of these hydrated silica film precedes the dissolution of the silica structure. Dissolution of silica is characterized by a two-part etching mechanism [17, 18]: First, H + will open the SiO 2 network at the surface, and then the fluorine species F - and HF 2 - react with the silicon to form SiF 4 which is soluble in water and forms H 2 SiF 6. The adsorption processes of HF molecules, HF 2 - and H + ions determine the reaction rate. Adding more H +, i.e. by adding HCl, to the etchant enhances the etch rate dramatically. Indeed, it seems, that the creation of silanol bonds at the surface of the SiO 2 is the limiting reaction at high HF concentrations. Experiments showed that bath agitation induces only little variation on the silica etching rate [18, 19]. Moreover, trace elements such as Na, Ca, Ni, Co and Mg are reported to accelerate the attack, and Cu, Fe, Pb, Zn, Sn and Al are reported to decrease the glass attack. Solubility of fluorides and hexafluorosilicates The low solubility of alkaline earth and lead hexafluorosilicates generated during glass etching can cause the precipitation of these compounds, particularly when multi component glasses incorporating theses bivalent cations are etched in HF solutions. These compounds are Na 2 SiF 6, K 2 SiF 6, MgSiF 6, Ca 2 SiF 6, etc. These hexafluorosilicates are more or less soluble in water, but not soluble in hydro-fluoric acid. 112

123 5.3 Bulk wet chemical etching of glass Fig. 5.4 Bulk glass under-etching of a chromium mask during tip manufacturing. Due to isotropic glass etching properties, the mask is underetched symmetrically until detachment. The glass etching process should be stopped at mask detachment to get very sharp tips. When a precipitation occurs, the glass etching becomes less effective. Indeed, when a crystal is generated, the underlying glass is protected from the etch solution. The etch rate is thus globally reduced and heterogeneities can appear at the glass surface. A more detailed description of this phenomenon can be found in literature [15, 17, 20] Substrate preparation before glass etching A mask with a regular pattern of different dimensions should be deposited onto the glass substrates in order to characterize glass etching properties in HF solutions. Since most photoresists are etched or removed from substrate (bad adhesion) in HF solutions, the best way to mask glass for bulk wet chemical etching in HF solutions is the use of a metallic mask. Moreover, the use of a mask will allow easy measurement of etch depth by measuring the height difference between the substrate surface and the etch front. The isotropical mask underetching occurring in HF solutions will form sharp tips at substrate surface as shown in Fig Only copper, silver, gold and platinum, which are commonly used for metallization, are not attacked by hydro-fluoric acid. However, most metals deposited as adhesion layer like titanium, tantalum and chromium are etched in HF solutions. Tests showed that chromium deposited by sputtering at high temperature (455 C) has an improved adhesion to the glass substrates and seems to resist hydro-fluoric acid when covered by another metal like copper or gold. First, masks were made by sputtering a 150 nm chromium layer at 455 C, followed by a chromium/gold or chromium/copper evaporation process. After patterning these metal layers, etching tests showed the following results: 113

124 5 Three dimensional multi-electrode arrays Chromium/copper mask: this mask showed a good resistance to hydro-fluoric acid. Copper is an effective protection layer when it has a thickness larger than one micron (no pin holes in the layer). No important under-etching due to chromium dissolution was observed. Chromium/gold mask: This mask was less efficient than the chromium/copper mask due to pin holes in the gold layer. Through these pin holes, hydro-fluoric acid was able to etch the underlying chromium layer and afterwards the underlying glass, which results in the appearance of cavities at the glass surface under the mask. Another way to protect the chromium layer is the use of an adequate photoresist that resists hydro-fluoric acid solutions. This solution is more lowcost because a second metal deposition can be avoided. The SC100 negative tone photoresist from Olin Corporation presents all requested properties for hydrofluoric acid etching of glass. Tests showed that this photoresist resists hydrofluoric acid exposure quite well. After all of these considerations, the masks for the hydro-fluoric acid etching of glass were made of a 150 nm sputtered chromium layer deposited at 455 C covered by the SC100 photoresist. The photoresist was deposited by spinning on a Karl Suss RC8 spin coater at 3000 rmp for 30 s in order to obtain a 1 µm thick layer. A 85 s prebake at 90 C hardened the resist. Exposure (100 mj/cm 2 ) to UV light at 365 nm followed by development in Waycoat Negative Resist Developer (WNRD) from Olin corporation and rinse in n-butyl acetate allowed patterning of the resist. Finally, a 45 s hardbake at 135 C insured good chemical stability of the resist. After patterning the photoresist, the last mask manufacturing step was the etching of the non-protected chromium areas in a (NH 4 ) 2 {Ce(NO 3 ) 6 } + HClO 4 solution (Lodyne CR-7S from Laporte Electronics) Characterization of hydro-fluoric acid glass etching Pyrex and float glass samples were etched in HF solutions with HF concentrations of 5%, 10%, 25%, and 49% at three different temperatures (20 C, 30 C and 40 C) in order to characterize the glass etch rate. Wafer samples were etched in a 22 liter solution bath with some bath agitation. The etch depth was measured every five minutes or less in order to determinate the glass etch rate. Results obtained are similar to reported silicate glass etching characteristics [19]. The dissolution rate is linear versus time but depends strongly on HF concentration and temperature. The etch rate increased with higher HF concentrations and higher temperatures. 114

125 5.3 Bulk wet chemical etching of glass Sputtering of 150 nm of chromium at 455 C onto glass substrate. Deposition and patterning of SC100 photoresist. Etching of chromium layer through SC100 photoresist mask. Fig. 5.5 Description of mask realization for glass hydro-fluoric acid etching. A chromium and a photoresist layer are patterned onto the glass substrate. Etch rate The Fig. 5.6 shows the glass etch rate in function of HF concentration and temperature for Pyrex and float glass. The curves show that the etch rate increases more rapidly at higher HF concentrations, an effect which can be explained by assuming that higher polymeric H n F n+1 - ions are present and that they are more reactive towards the siloxane bonds [20]. The borosilicate glasses like Pyrex have a composition including higher concentrations of network-forming oxides (B 2 O 3 and P 2 O 5 ). These networkforming oxides form a separate network which is intimately mixed with the silicate network. The breakage chemistry of these bonds is not necessarily the same as for siloxane bonds, which affects the etch rate of these glasses [20]. B 2 O 3 present in Pyrex induces an increase of the etch rate in HF solutions in contrast to buffered oxide etches. This is due to higher dissolution of =B-O-Si and =B-O- B= bonds that are more rapidly attached by H + ions. However, the high concentration of SiO 2 limits the etch rate close to the silicate bond dissolution rate. The float glass etch rate is higher than the Pyrex etch rate. This is due to higher network-modifying oxide (Na 2 O, K 2 O and CaO) concentrations, which opens the silica structure and increase the number of broken siloxane bonds. This effect becomes larger with an increase of network modifiers. The etching of float glass was more delicate than Pyrex etching because of precipitation of hexafluorosilicates on the sample surface. Fig. 5.7 shows a 115

126 5 Three dimensional multi-electrode arrays Fig. 5.6 Pyrex (top) and float glass (bottom) etch rate versus HF concentration for three different temperatures (20 C, 30 C, and 40 C). The Pyrex etch rate is slower than the float glass etch rate due to glass composition and resulting etching reactions. 116

127 5.3 Bulk wet chemical etching of glass Fig. 5.7 Scanning electron microscopy picture of hexafluorosilicates precipitation appearing at substrate surface during float glass etching in high concentrated HF solutions. Fig. 5.8 Ratio between horizontal and vertical glass etching speed versus HF solution concentration for Pyrex and float glass. 117

128 5 Three dimensional multi-electrode arrays Fig. 5.9 Scanning electron microscopy pictures of Pyrex microtips etched in a 5% HF (left) and 49% HF (right) solution. A set of square chromium masks was used to generate a pyramidal tip shape. For high HF concentrations, the high mask underetching ratio causes very flat tips. picture of this hexafluorosilicates precipitation at substrate surface. As can be seen in this picture, the silica dissolution front seems to keep a uniform and smooth shape. These precipitations occurred mainly at mask locations, i.e. at places were the masks were underetched. This should be due to less solution agitation, reducing the hexafluorosilicates dissolution capability. This phenomenon increases with HF concentration and samples were covered with so much hexafluorosilicates (at HF concentration of 49%) that no clear surface remained. To reduce hexafluorosilicate precipitation at the mask locations, the samples were rinsed in water for 2 minutes after every 5 minutes lasting glass etching step. Furthermore, the wafer samples were rotated of 90 in the wafer basket in order to obtain an homogenous glass attack. Mask underetching Wet chemical etching of glass in HF solutions should be isotropic. When a glass substrate is etched through a chromium mask window, a semi-circular wall should be formed at the edge of the mask window when the masking film has a good adhesion. In this ideal case, the ratio between the horizontal and the vertical etch deepness values 1. However, the experimental results obtained didn t show this isotropic underetching properties of glass etching. The results shown in Fig. 5.8 demonstrate that other effects play a role at mask window border. For the Pyrex, there is an increase of mask underetching as a function of HF concentrations. This can be explained by a more rapid interface delamination of the chromium mask. For float glass, a decrease of mask underetching was observed. This should be due to the hexafluorosilicates that 118

129 5.3 Bulk wet chemical etching of glass Fig Scanning electron microscopy pictures of float glass tips etched in a 10% HF solution at 20 for 40 minutes. A set of circular chromium masks was used for the tip shape (diameter of 140 µm, spacing of 200 µm). The tip height is 80 µm. Fig Scanning electron microscopy pictures of float glass etching results. On the left, the picture shows an underetched chromium mask prior mask detachment. Some Hexafluorosilicate precipitations still remain on the mask. On the right, top of an obtained tip with a diameter of less than one micron. precipitate at mask borders and slow down the etch rate horizontally at these locations Glass microtip results Microtips manufactured on Pyrex substrates had very different shapes depending on HF concentration. For low HF concentrations (5%), nice microtips could be achieved, but the morphology of the glass surface was rather rough (see 119

130 5 Three dimensional multi-electrode arrays Fig. 5.9, on the left). For high HF concentrations (49%), the mask underetching was too high, and resulting tips were flat, as can be seen in Fig. 5.9 (right). However, in this case, the surface morphology was smooth. Microtips manufactured in float glass showed different results. Using low HF concentrations (5% and 10%), very sharp and high tips could be realized as shown in Fig and in Fig with a smooth surface morphology. For higher HF concentrations, high rate hexafluorosilicate precipitation did not allow fabrication of usable tips. To conclude, the best glass wet chemical etching results were obtained with float glass in a 10% HF solution at 20 C, resulting in long and sharp tips with a smooth surface morphology. 5.4 FABRICATION OF 3D MULTI-ELECTRODE ARRAYS The design used for the realization of 3D MEAs is the same as that used for the 8x8 electrode matrix design described in Chapter Fabrication of tip-shaped electrode arrays is based on the previous described glass etching principle and fabrication technologies (see Chapter 2.3). The 3D MEA chips are first fabricated by micromachining techniques and then mounted on a PCB. The bottom part of a Petri dish is used to build a culture chamber sealed with DC184 silicone from Dow Corning. Fabrication process Float glass plates (thickness of 700 µm, diameter of 10 cm) were used as substrate material. First, a chromium/sc100 photoresist mask was deposited and patterned onto the glass substrate. Then the glass was etched in a 10% HF solution at 20 C until the detachment of the chromium masks (about 40 to 50 minutes) in order to obtain 60 µm high glass tips on the substrate. The next step was the deposition and patterning of AZ5214 photoresist (from Clariant) in order to define the negative of the electrode pattern. Deposition of a 50 nm titanium/150 nm platinum layer and photoresist removal in acetone completed the deposition of platinum electrodes onto the tips and the substrate. The last clean room step was the deposition and patterning of a 5 µm SU-8 epoxy insulation layer. After chip separation by substrate dicing, the MEA chips were mounted the same way as described in Chapter The fabrication process is resumed in Fig Realized structure The dimensions of the fabricated multi-electrode arrays are indicated in Table 5.2. Pictures of the obtained 3D MEAs are shown in Fig

131 5.4 Fabrication of 3D multi-electrode arrays Deposition and patterning of metallic mask for glass etching. Glass etching in HF solution. Patterning of negative photoresist for lift-off process. Titanium (50 nm)/platinum (150 nm) deposition by sputtering process. Stripping of photoresist. Patterning of an SU-8 epoxy insulation layer. Fig Fabrication process of MEAs with protruding 3D electrodes. 121

132 5 Three dimensional multi-electrode arrays Chip dimensions Microelectrode dimensions Space between electrodes 1.5 cm x 1.5 cm cone with a rectangular base of 40 µm x 40 µm and a height about 40 µm 200 µm center to center Number of electrodes matrix of 8x8 electrodes without the 4 corners Tip height about 60 µm Table 5.2 Dimensions of tip-shaped MEAs Fig SEM pictures of the workspace of a tip-shaped MEA at different magnifications. The height of the tips is about 60 µm. 5.5 ELECTRICAL PROPERTIES OF 3D ELECTRODES In order to compare planar electrodes with 3D electrodes, the electrode properties of both planar and 3D electrodes (here with a tip height of 60 µm) 122

133 5.6 Three dimensional versus planar electrodes were measured by impedance and phase shift measurements versus frequency. Measurements were done by sweeping the frequency from 100 Hz to 20 khz. The obtained results and the calculated values of the global electrode resistance and capacitance are plotted in Fig and Fig The electrode resistance and the electrode capacitance were calculated and normalized to 1 mm 2 area, considering that the resistance is inversely proportional to the area and the interface capacitance is directly proportional to it. The electrode impedance changes due to an increase of the geometrical electrode surface. A good approximation of the 3D electrode surface is given by the lateral surfaces of a pyramid (height of 40 µm and a base side of 40 µm) for a tip height of 60 µm. This approximation provides an electrode surface of about 3570 µm 2 for a 3D electrode. Comparing to a planar electrode surface of 1600 µm 2, the surface area increases by a factor The measured impedance values of the 3D electrode correspond to 48% of the planar electrode impedance, which fits roughly to the electrode surface increase. However, the phase shift gives similar values in both measurements. The resulting values of the global electrode resistance is smaller by a factor 2.14, and the global electrode capacitance is 2.08 times larger for the 3D MEA electrodes. As a results of this global resistance reduction, the electrode noise decreases roughly by a factor 2 due to the noise proportionality to R. Indeed, the measured noise level is reduced from 20 µv to 25 µv for planar electrodes to 14 µv to 17 µv for 3D electrodes. The resulting signal to noise ratio is thus also increased roughly by a factor 2. The higher capacitance of the 3D electrodes is also important for cell or tissue stimulation through the 3D electrodes because more charges can be placed onto the electrodes before irreversible Faradic reactions occur. 5.6 THREE DIMENSIONAL VERSUS PLANAR ELECTRODES Dead cell layers always appear at tissue slice borders when cut with a tissue chopper. The first layer of electrically active cells should be found about 50 µm inside the tissue slices. The idea of using protruding tip-shaped electrodes for tissue penetration, and thus get closer to active cells, should be an improvement for acute slice experimentation. To better understand the electrical characteristics of cell stimulation and recording in function of geometry, a simple 2D simulation model of planar and 3D electrodes was computed. 123

134 5 Three dimensional multi-electrode arrays Fig Results of measured impedance (top) and phase shift (bottom) versus frequency for 3D and planar MEA in a 0.9% NaCl solution. 124

135 5.6 Three dimensional versus planar electrodes Fig Calculated values of the electrode resistance (top) and capacitance (bottom) versus frequency for 3D and planar MEA. 125

136 5 Three dimensional multi-electrode arrays Planar and 3D MEA electrode recording simulations Using a conductive media model in FEMLAB, the electrical field between a cell and an electrode was simulated. The physical model used in FEMLAB was a DC conductive media model, based on a relation relating the current density J to the electrical field E through: J = σe Combining the continuity equation: (5.3) J = Q (5.4) where Q is a current source, with the definition of the electrical potential V yields the elliptic Poisson s equation defined as: Q = σ V ( ) (5.5) The two only resulting parameters are the conductivity σ and the current source Q. Dirichlet and Neumann boundary conditions were assigned to the model geometry. Three simple geometrical models were simulated to characterize the difference between planar and 3D electrodes while monitoring biological activity in tissue slices. One disadvantage of the MEA electrode/cell model is that the real electrode and cell properties are not taken correctly into account due to the use of a 2D model, which simplifies the element geometry drastically. 2D signal magnitude simulation while monitoring a fixed located cell In order to compare the measurable signal magnitude for both planar and 3D electrodes, a simulation of the electrode signal magnitude for different electrode shapes (planar and 3D electrodes with a tip height up to 80 µm) was made in the case of a biological signal source beneath the recording electrode. The signal source was modeled by a spherical cell soma, with a diameter of 20 µm, and an axon (tube with a diameter of 1 µm and an infinite length). The cell soma was locate above the electrode border, 60 µm above the substrate level. The signal applied at the cell was -60 mv at the soma and 40 mv at the axon. The electrodes were modeled by a simple planar square electrode with dimension of 40 µm x 40 µm and by a conical electrode shape for the 3D electrodes. The height of the tips varied up to 80 µm. In case of 3D electrodes, the effective electrode was located on the top of the tip (40 µm at the top of the tip). The dead cell layers in between this active cell and the electrode were assigned to be resistive homogeneous media. An electrode lead was also introduced in the model at the left of the electrode between the glass substrate and the SU-8 insulation layer. 126

137 5.6 Three dimensional versus planar electrodes Fig Plots of simulated isopotentials in presence of a single electrical source (neuron) for both a planar and a 60 µm high glass tip covered with an electrode on its top. X and Y axes represent the model geometry dimensions in µm. 127

138 5 Three dimensional multi-electrode arrays Fig Plot of the simulated electrode potential in case of an electrical dipole (cell at a fixed location) versus electrode height, a planar electrode corresponding to a height of 0 µm. The border condition of the electrode took into account the electrode film conductance as defined in the generalized Neuman condition. The obtained simulation results shown in Fig show that higher signal amplitudes should be measured with 3D electrodes. This is due to a geometrical advantage of the 3D electrodes. Indeed, the 3D electrode is closer to the signal source. When comparing the signal amplitudes of planar and 3D electrodes with a tip height of 60 µm, the amplitude rises with a factor close to 2.3. This result can be verified with experimental measurements on biological tissue. MEAs with 3D electrodes (tip height of 60 µm) were used for acute rat hippocampal slice culture measurements which will be described later on. Furthermore, these results suggest that higher electrodes (height of 80 µm and more) should be an even better electrode configuration than those used in this work. 2D signal magnitude simulation versus the electrode/source distance The simulation of the signal magnitude versus the electrode/source distance should show the main geometrical advantage of the 3D electrodes. This simulation was achieved by modeling the measurement electrode, i.e. a planar and a 60 µm high tip-shaped 3D electrode. Two different source configurations were simulated: 128

139 5.6 Three dimensional versus planar electrodes Fig Plots of simulated isopotentials in presence of a 100 mv dipole located at different heights above both a planar and a 60 µm high tip-shaped 3D electrode. X and Y axes represent the model geometry dimensions in µm. 129

140 5 Three dimensional multi-electrode arrays Fig Plot of the simulated electrode potential for both a planar and 3D electrode versus the distance between the electrode and the signal source, being located above the center of the electrode. For planar electrodes, the potential decreases with the square of the electrode/ source distance. the signal source was located at heights between 5 µm and 200 µm above the center of the recording electrode. the signal source was located 20 µm next to the electrode (X axis) at heights between 5 µm and 200 µm above the substrate. The signal source used in this model was modeled as a sphere divided into two parts representing an electrical dipole of 100 mv. The space in between the dipole and the electrode was assigned to be a resistive homogeneous media. This simplified the model and could also be reproduced experimentally with a glass pipette electrode as signal source in order to verify the simulation. The obtained simulation results shown in Fig and Fig show that the signal amplitude decrease, with the square of the distance between the source and the electrode (for planar electrodes). It is thus very important that an active cell in a tissue slice is located as close as possible to the recording electrode. The amplitude difference obtained between the planar and the 3D electrode is due to the increase of the electrode surface, and the resulting impedance decreases as shown in Chapter 5.5. The geometrical advantage of the protruding electrodes is not evidently shown in this configuration. 130

141 5.6 Three dimensional versus planar electrodes Fig Plots of simulated isopotentials in presence of a 100 mv dipole located 20 µm next to the electrode at different heights above the substrate level. X and Y axes represent the model geometry dimensions in µm. 131

142 5 Three dimensional multi-electrode arrays Fig Plot of the simulated electrode potential for both a planar and 3D electrode versus the distance between the substrate and the signal source in the Z axis. The signal source is located 20 µm next to the electrode (X axis). The obtained simulation results shown in Fig and Fig show that, for a cell located in a tissue slice close to an electrode location, a 3D electrode provides a better signal recording due to a shorter distance between the cell and the electrode than in a planar electrode configuration Experimental verification The effect of a tip electrode versus a planar electrode in brain slice cultures was verified using acute hippocampal slice cultures. Recordings from acute hippocampus slices with both planar and 3D electrode MEAs showed a signal amplitude increase similar to the results obtained in the simulations. However, a detailed description of these results is presented in Chapter 5.7. To verify the other two simulations, the potential variation resulting of the presence of a signal source (a glass microelectrode) was recorded for both planar and 3D electrodes in the two simulated configurations. A θ glass pipette electrode (two separated compartments) was used to generate a potential dipole in the culture chamber medium. In order to obtain values for different distances in the Z axis, the θ glass pipette was moved away from the electrode or the substrate level over a distance of 100 µm. 132

143 5.6 Three dimensional versus planar electrodes Fig Picture of a θ glass microelectrode above some planar electrodes used as signal source (potential dipole) for simulation verification. Fig Picture of the simultaneous recorded signals in case of a potential pulse applied at one electrode (black dot). Each case represents one channel (electrode spaced by 200 µm). The signals amplitude decrease versus distance in the X and Y axes is well shown in this recording. Fig shows a θ glass microelectrode above some MEA electrodes. The aperture of the pipette was about a few microns. When biphasic potential pulses (-/+) of 100 mv, 2 ms were applied, a signal could be recorded on all electrodes 133

144 5 Three dimensional multi-electrode arrays Fig Plot of signal amplitudes recorded with both a planar and 3D MEA versus the distance between the recording electrode and the stimulation microelectrode. The stimulation microelectrode was locate above the center of the recording electrode. of the MEA as shown in Fig This picture illustrates the decrease of the electrical field generated by the glass microelectrode in the solution in function of the distance in the X and Y axes. Recordings made with both some planar and 3D electrodes versus the distance in the Z axis between the signal source and the electrodes are plotted in Fig The signal magnitude ratio between the 3D and the planar electrode measurement for the same stimulation conditions decreased from 1.95 when the potential source was located direct above the electrode, to 1.5 when the potential source was located 100 µm over the electrode. This follows an intuitive physical mechanism. When the stimulation occurs far from the electrode, the effect of a tip-shaped electrode instead of a planar electrode will become negligible. Fig shows results obtained in the second simulated configuration. A glass microelectrode was placed next to the electrode and moved in the Z axis away from the substrate level. The obtained signal amplitudes recorded show the same comportment than in the simulated case (see Fig. 5.21). For planar electrodes, the distance between the signal source and the recording electrode increases and the resulting measured signal amplitude decreases. For protruding 3D electrodes, the distance between the electrode and the signal source remains 134

145 5.6 Three dimensional versus planar electrodes Fig Top: pictures of a large θ glass microelectrode located next to one recording electrode for a planar (left) and 3D electrode (right). The glass microelectrode is then moved along the Z axis away from the substrate level. Bottom: the resulting signal amplitude is plotted versus the covered distance in the Z axis. It shows a geometrical advantage of 3D electrode compared to planar electrodes. more or less constant over a vertical source displacement about 60 µm due to electrode geometry. Then the electrode/source distance also increases, resulting in a signal amplitude decrease. 135

146 5 Three dimensional multi-electrode arrays The obtained experimental results are well correlated with the simulations presented in Chapter It clearly suggests the advantages of 3D electrodes versus planar electrodes. However, the use of longer and steeper tip-shaped 3D electrodes would increase the 3D electrode improvement for tissue slice monitoring. 5.7 BIOLOGICAL EXPERIMENTS Experimentation with acute hippocampal slices from rat allows to illustrate the advantages of tip-shaped electrodes in comparison to planar electrodes. The hippocampus is a brain structure that is well known and lends itself for easy experimentation. A description of the hippocampus structure, the experimentation protocol and some results are presented in the following sections The rat s hippocampus The hippocampus or hippocampal formation is a large C-shaped structure that forms a large part of the medial wall of the rat cerebral hemispheres as shown in Fig This part of the brain plays an important role in certain aspects of learning and memory, especially the memory for facts and events called declarative memory. The hippocampus is involved in making associations between the current and previous situation that the subject has already experienced. The hippocampus is mainly composed of two types of neurons, i.e. granule neurons and pyramidal neurons, forming two lines folded in each other. The circuit considered to be the morphological substrate of the main flow of information through the hippocampus can be described as follows (see also Fig. 5.27): The information input into the hippocampus is achieved through projections from the superficial layers of the entorhinal cortex, that form excitatory synapses onto the dendritic spines of the granule neurons (the so called dentate gyrus formation). These projections are the so-called perforant path, PP. The granule neuron axons, GC, of the dentate gyrus issue numerous collaterals and the mossy fibres to the stratum radiatum of CA3 field, in which they form synapses with the large pyramidal neurons, CA3. The main axons of the large CA3 pyramidal neurons pass to the alveus, Alv. However, these axons issue coarse collateral branches, which penetrate the stratum pyramidale and pass to area CA1, where they form a compact sheet of 136

147 5.7 Biological experiments Fig Picture of a rat brain and a transverse cross-section showing the importance of hippocampal formation in rat s brain (modified from [21]). Fig Simplified representation of a transverse hippocampus slice. Abbreviations: Alv, alveus; Sch, Schaffer collaterals; Fim, fimbria; CA1, CA1 pyramidal cell layer; CA3, CA3 pyramidal cell layer; GC, granule cells in the dentate Gyrus region; PP, perforant path (modified from [22]). 137

148 5 Three dimensional multi-electrode arrays Fig Typical experiment driven onto hippocampus slices. Electrical stimulation of Schaffer collaterals (1) will induce evoked responses at CA1 pyramidal neurons (2). fibres. Within the CA1 region, these so-called Schaffer collaterals, Sch, issue short branches which enter into synaptic contact (excitatory action) with the dendrites of pyramidal and probably of non-pyramidal neurons, CA1. The main axons of the large CA1 pyramidal neurons also pass to the alveus, Alv. The CA1 pyramidal neurons project heavily to the subicular complex, which makes projections to the superficial layers of the cortical area. This simplified explanation highlights the unidirectional aspect of hippocampus connectivity and data processing. Because of the highly regular organization of the hippocampus, it provides high field potential recordings. Furthermore, the interesting functions treated by the hippocampus makes it one of the most studied parts of the brain. Several types of experiments can be performed on the information path in the hippocampus. One of the most used experiments is the electrical stimulation at CA3/CA1 border of Schaffer collaterals and the simultaneous recording of CA1 pyramidal neuron evoked responses. Fig illustrates the measurement of electrical activity using a multielectrode array for stimulation and recording from hippocampus slices. Schaffer collaterals can be stimulated electrically through an electrode locate underneath (1). Recording of evoked responses at pyramidal neurons in CA1 region can be monitored through other electrodes (2). Moreover, electrodes cover all of the hippocampus structure and record its activity at all locations simultaneously. More information about the hippocampus can be found in literature [21, 23, 24]. 138

149 5.7 Biological experiments Sodium chloride (NaCl) Potassium chloride (KCl) Sodium hydrogen carbonate (NaHCO 3 ) Magnesium chloride (MgCl 2 ) Glucose (C 6 H 12 O 6 ) Calcium chloride (CaCl 2 ) Oxicarbonate gas (95% O 2, 5% CO 2 ) 135 mm 5 mm 15 mm 1 mm 10 mm 2 mm Table 5.3 Composition of Artificial Cerebro-Spinal Fluid (ACSF) used for experimentation with hippocampal brain slices. The ACSF is saturated with oxicarbonate gas Preparation of acute hippocampal slices The use of acute slices for in-vitro experimentation provides the closest invivo model. Due to the short duration between animal death and slice stimulation and recording, the neuronal connectivity in the tissue preparation is preserved. This point is very important, because the obtained experimental results are close to results that would be obtained when measured under in-vivo conditions. The slice preparation follows mainly the guide to multi-channel-recording of acute hippocampal slices from [25]. Animal handling and preparation was carried out according to the animal rights Swiss ethical rules. Experiments were carried out with young male rats (ca. 4 weeks old, 120 g), that were sacrificed by decapitation. Anesthesia was avoided, since most anesthetics do obviously affect the brain and its efficacy depends strongly on the weight of the animal. The brain was quickly removed and put into ice-cold (4 C) saturated with 95% O 2 /5% CO 2 artificial cerebro-spinal fluid (ACSF) without calcium. Removal of calcium ions removes synaptic propagation of action potentials during preparation, reducing cell damage. Transverse hippocampal slices were cut at 350 µm thickness with a vibrating tissue chopper. The slices were then preincubate in normal gazed ACSF solution with 2 mm CaCl 2 at room temperature for at least 1 hour for recovery. The used electrode arrays were pre-coated with 0.1% polyethylenimine (PEI) for 2 hours, extensively rinsed with distilled water and dried. Slices were then placed into pre-coated and dry planar or 3D MEA and positioned to cover the electrode matrix. After the positioning of the slice, the surrounding solution was 139

150 5 Three dimensional multi-electrode arrays Fig Top: typical acute hippocampal slice culture on a multi-electrode array. Bottom: recording obtained when stimulation (current pulses) was applied at dentate gyrus perforant path (channel with black dot). 140

151 5.7 Biological experiments Fig Recordings obtained when a stimulation (current pulses) was applied at Schaffer colatterals in CA1 region (channels with black dot). Obtained signal amplitudes are larger than in the CA3 region. 141

152 5 Three dimensional multi-electrode arrays removed with a pipette. Small filter papers were then used to absorb the solution remaining between the slice and the MEA surface. Fresh Ringer solution was quickly introduced in the culture chamber on top of the slice. The MEA was then transferred to the MEA1060 amplifier interface from Multi Channel Systems and a continuous perfusion rate of 2-3 ml/min of ACSF was maintained during the recording session Electrophysiological stimulation and recording method During measurements, the slices were stimulated with biphasic current pulses through some electrodes and evoked biological responses were measured at the other remaining 59 electrodes. To construct an input/output (I/O) curves with the STG stimulator (Multi Channel Systems), we used biphasic current pulses (-/+ values, 120 µs duration at each level) ranging from 0 to 360 µa in 30 µa steps, i.e. we ended up with 12 values on the I/O curve. It is important to start the pulse with the negative flank first, since it was found to be best for stimulation in terms of minimizing the stimulation artifacts. For each stimulation step three signal responses with an inter-spike interval of two seconds were recorded. The interval between two different stimulation steps was 10 seconds. Moreover, the I/O curve was measured over three times every 10 minutes, i.e. a complete experiment lasted not more than 1 hour. The first 15 minutes were reserved for selecting a stimulation electrode and establishing a steady-state condition between the electrodes and the slice. Then, every 10 minutes the STG stimulator initiated the I/O curve. We attempted to select only one site at the CA3/CA1 border for stimulation. An electrode just below the cell layer was used to stimulate Schaffer collateral fibres in order to get evoked responses at pyramidal neurons in the CA1 region. When no clear signals were present in the CA1 region, the end of the inner part of the dentate gyrus, which generates evoked responses in CA3 region via the mossy fibre path, was stimulated Recordings from rat s hippocampus The Fig and Fig show typical evoked responses in the hippocampal formation and illustrate the schematic representation of information paths shown in Fig A ~1 ms long stimulation artifact occurred first on all measurement electrodes followed by eventual evoked biological responses, depending on the tissue organization and stimulation location. The recordings shown in Fig show typical Schaffer collateral fibre stimulations (electrodes marked with a black dot) and evoked population spike 142

153 5.7 Biological experiments Fig Biological responses to electrical stimulation (on the left) can be eliminated by application of a low calcium solution perfusion in the culture chamber due to synaptic transmission inhibition. The traces shown in the middle column correspond to the remaining electrode stimulation artifacts at different locations after 4 minutes of low calcium solution perfusion. This experiment is reversible as shown on the right traces measured 10 minutes after low calcium solution wash out. responses in the CA1 pyramidal neuron layer (electrodes situated one row over the stimulation electrodes). In addition, the recording shown in Fig shows a typical stimulation of granule neurons located in the dentate gyrus (electrode marked with a black dot) and population spike responses in the CA3 region (electrodes situated in the two rows under the stimulation electrode) evoked via the mossy fibre path. However, the signal magnitude varies dramatically with stimulation and recording location due to the defined cell connectivity in the hippocampus. In the CA1 region, for example, the highest signal magnitudes are obtained when the recording electrodes are located direct under the pyramidal neuron layer. The application of a low calcium ACSF solution (0 mm CaCl 2 ) is one possible reversible experiment allowing identification of the part of the signal that belongs to the stimulation artifact and to biological activity. The absence of calcium eliminates synaptic transmission and should leave only the stimulation artifact and eventually presynaptic action potentials. The Fig shows the results of a low calcium experiment. The first column of signals (on the left) 143

154 5 Three dimensional multi-electrode arrays shows evoked responses on four different electrodes before application of a low calcium solution in the culture chamber. The second column of signals (in the middle) shows responses from the same electrodes but after four minutes of low calcium solution perfusion in the culture chamber. The remaining signals are the electrode stimulation artifact or presynaptic action potentials (small signal after the stimulation artefact in the lowest row). The third column of signals (on the right) shows recovery of biological responses after a 10 minute wash out of the low calcium solution by standard ACSF solution. A low concentration solution (1 µm) of tetrodotoxine (TTX) can be used to eliminate all biological responses.tetrodotoxine is a poison found in the sex organs and liver of marine puffer fish and other species of the Tetredontiformes, but also in the blue-ringed octopus and some salamanders. It blocks the membrane sodium channels, thus blocking all neural activity. Experiments of TTX application showed similar responses than shown in Fig. 5.31, but the duration of response recovery was much longer (30 minutes to one hour) Stability of recorded signals Long-term stimulation experiments allowed the estimation of acute slice survival under measurement conditions. This is important for experiments lasting more than one hour in order to evaluate if response variation are due to normal activity of are due to tissue slice degradation. Fig shows the results of a long term stimulation experiment. A 200 µa stimulation pulse lasting 120 µs was applied to an electrode once every 30 s for more than one hour, i.e. we ended with more than 120 values. This was made using both a planar and a 3D MEA for the acute slice culture. The results obtained show that the acute slices remained stable during the experiments. As also shown in Fig (top), the signal magnitude for both MEA configurations did not vary more than 15% over the overall experiment. As shown in Fig (middle and bottom), where all measured signals from the electrode providing the signal with the highest magnitude were superposed, the shape and the amplitude of the evoked responses did not change during the experiment. The slope of the obtained population spikes reflects the signal integration of neurons lying more or less nearby the recording electrodes. When the signal slope is steep, like for glass pipette microelectrodes, the recording reflects the responses of neurons located in close vicinity of the electrodes, and the recorded responses are more local. When the slope is less steep, responses from neurons located more far away are taken into account in the population spike, suggesting a response measured from a larger area around the electrode. The slopes of the population spikes obtained in the stability experience were calculated for the 144

155 5.7 Biological experiments Fig Stability of an acute slice preparation over time. Top: Plot of the signal magnitude for the overall experiment for both MEA configurations. The magnitude variability did not exceed 15%. Bottom: Superposed traces of a response to an 200 µa, 120 µs biphasic stimulation pulse applied once every 30 s over more than one hour using both a planar MEA and a 3D MEA configuration. 145

156 5 Three dimensional multi-electrode arrays Fig Extracellular field potential recorded from CA1 pyramidal neurons using a glass microelectrode. The stimulation artifact (arrow) is also shown (from [26]). planar (228 ± 22 µv/ms) and the 3D (272 ± 29 µv/ms) electrodes. The obtained results suggest that the recordings made with 3D electrodes were more local (the neurons were located closer to the recording electrodes) than with planar electrodes. Fig shows an extracellular field potential recording made using a glass microelectrode. The signals recorded with the 3D MEA electrode and the glass microelectrode have a similar shape and duration D versus planar electrode arrays In order to compare the responses that can be obtained from planar and 3D MEAs, I/O curves were measured (n=10) for both electrode types. Data from one 3D electrode experiment is illustrated in Fig and Fig In this experiment, the hippocampus was stimulated at CA1/CA3 border in order to stimulate Schaffer collateral fibres projecting to the CA1 pyramidal cell layer (dotted electrode in Fig. 5.34). The overall hippocampus slice response to a 360 µa, 120 µs biphasic stimulation pulse is shown in Fig (bottom). The high magnitude signals obtained at the electrodes one row above the black dotted electrode location are the responses from the CA1 pyramidal neurons. The magnitude of these evoked responses was measured in the millivolt range in this experiment. The traces plotted in Fig show three signals recorded from the electrode with the largest signal amplitude (electrode one row above the stimulation electrode) for a 90 µa, 180 µa, and 360 µa stimulation pulse. Similar responses as those shown in Fig and Fig were also obtained with planar MEA recordings. However, the obtained signal amplitudes 146

157 i 5.7 Biological experiments Fig Illustration of a I/O curve measurement. Top: picture of an acute hippocampus slice on a 3D MEA. Schaffer collaterals were stimulates (white dot) using current pulses. Bottom: picture of a recording corresponding to a 360 µa, 120 µs biphasic stimulation current pulse (electrode with black dot). The largest signal amplitudes (about 1.1 mv) were measured at CA1 pyramidal neurons located one row above the stimulation electrode. 147

158 5 Three dimensional multi-electrode arrays Fig Signals of the largest evoked response amplitude obtained in the experiment shown in Fig at different stimulation pulse amplitudes. were smaller than those obtained with 3D MEAs as illustrated in Fig (top). These results illustrate that a minimum stimulation current of about 50 µa is needed in both configurations to obtain a detectable biological response. Remember that in order to generate an action potential, a cellular stimulation threshold has to be reached. However, when this threshold is reached, the response behavior is different in both configurations. The 3D MEA configuration responses, especially at low current stimulation amplitudes, increased with a steeper slope than in the planar MEA configuration. This can be attributed to better tissue stimulation characteristics due to the three dimensional electrode shape, and thus more local stimulation obtained with the 3D electrodes. In these measurement series, responses to the largest stimulation pulse (biphasic stimulation pulse of 360 µa, 120 µs) showed that larger signal amplitudes can be obtained with 3D electrodes (factor around 2.2) than with planar MEA electrodes. This significant signal amplitude increase can be explained by the better electrical properties of the 3D MEAs (see Chapter 5.5) due to a larger electrode area, however, the geometrical measurement advantage of tip-shaped electrodes versus planar electrodes (see Chapter 5.6), and a shorter distance between the electrodes and the active cells due to tissue penetration cannot be clearly seen. The advantage of the electrode geometry can be seen when representing the same data against 100% of the signal amplitude (see Fig bottom). In this case, a shift to the left of the input/output relation for the 3D electrodes can be seen demonstrating that a) a more local stimulation occurred due to the tip geometry and that b) shorter distance between the cells and the electrodes. About 25% less current is necessary to obtain 50% of the maximum signal amplitude when using 3D instead of planar stimulation and recording electrodes. 148

159 5.7 Biological experiments Fig Top: plot of the highest signal amplitude mean value obtained for both planar (n=10) and 3D (n=10) MEA measurements versus current stimulation. The measurements made with 3D MEAs show signal amplitudes up to 2.25 times those measured with planar MEAs. Bottom: 149

160 5 Three dimensional multi-electrode arrays Fig Distribution of the 95 highest signal amplitudes of both planar (n=10) and 3D (n=10) MEA measurements. There is a clear shift to higher amplitude values for 3D MEA measurements. In addition, the distribution of the largest signals obtained (n=95) in these experiments for both planar (n=10) and 3D (n=10) MEA electrodes shows a shift to larger amplitudes (Fig. 5.37), suggesting a global effect due to the different electrical and geometrical characteristics of the used planar and 3D electrodes. 5.8 CONCLUSIONS Fabrication of tip-shaped electrodes on glass substrates by wet chemical etching in hydro-fluoric acid allows realization of transparent 3D multi-electrode arrays. This process is easy to control and do not give rise to high additional manufacturing costs. It has been shown in this chapter that the integration of three dimensional electrodes based on glass microtips has been an improvement for biological activity recording. Indeed, simulations and experimental verification showed that three dimensional electrodes shows properties superior to planar electrodes. Recordings of evoked population spike responses from acute rat hippocampus slices showed that the measured signal amplitude is more than doubled when using 3D electrodes instead of planar electrodes. Moreover, better tissue 150

161 5.9 References stimulation is obtained with the 3D electrodes due to a more local stimulation in the tissue and a shorter distance between the electrode and the active cells in the tissue slice. 5.9 REFERENCES [1] M. M. Farooqui and A. G. R. Evans, Silicon sensors with integral tips for atomic force microscopy: a novel single-mask fabrication process, Journal of Micromechanics and Microengineering, Vol. 3, pp. 8-12, 1993 [2] R. C. Davis, C. C. Williams and P. Neuzil, A micromachined submicrometer photodiode for scanning probe microscopy, Proc. of the International Conference on Solid-State Sensors and Actuators (Transducers '95) and Eurosensors IX, Vol. II, Stockholm, Sweden, June 25-29, pp , 1995 [3] P.-F. Indermühle, C. Linder, et al., Design and fabrication of an overhanging xy-microactuator with integrated tip for scanning surface profiling, Sensors & Actuators: A. Physical, Vol. 43, pp , 1994 [4] J. Brugger, R. A. Buser and N. F. de Rooij, Silicon cantilevers and tips for scanning force microscopy, Sensors & Actuators: A. Physical, Vol. 34, pp , 1992 [5] W. Trimmer, P. Ling, et al., Injection of DNA into plant and animal tissues with micromechanical piercing structures, Proc. of the 8th IEEE Int'l Workshop on Micro Electro Mechanical Systems, MEMS '95, Amsterdam, the Netherlands, January 29 - February 2, pp , 1995 [6] H. Jansen, M. d. Boer, et al., The black silicon method IV: the fabrication of trhee-dimensional structures in silicon with high aspect ratios for scanning probe microscopy and other applications, Proc. of the 8th IEEE Int'l Workshop on Micro Electro Mechanical Systems, MEMS '95, Amsterdam, the Netherlands, January 29 - February 2, pp , 1995 [7] S. Henry, D. V. McAllister, et al., Micromachined needles for the transdermal delivery of drugs, Proc. of the 11th IEEE Annual International Workshop on Micro Electro Mechanical Systems, MEMS '98, Heidelberg, Germany, January 25-29, pp , 1998 [8] M. Ishida, K. Sogawa, et al., Selective growth of Si wires for intelligent nerve potential sensors using vapor-liquid-solid growth, Proc. of the International Conference on Solid-State Sensors and Actuators, Transducers '99, Vol. 2, Sendai, Japan, June 7-10, pp , 1999 [9] D. V. McAllister, F. Cros, et al., Three-dimensional hollow microneedle and microtube arrays, Proc. of the International Conference on Solid-State 151

162 5 Three dimensional multi-electrode arrays Sensors and Actuators, Transducers '99, Vol. 2, Sendai, Japan, June 7-10, pp , 1999 [10] A. Sayah, C. Philipona, et al., Fiber tips for scanning near-field optical microscopy fabricated by normal and reverse etching, Ultramicroscopy, Vol. 71, pp , 1998 [11] A. Ruf, J. Diebel, et al., Ultra-long glass tips for atomic force microscopy, Journal of Micromechanics and Microengineering, Vol. 6, pp , 1996 [12] P. Thiébaud, N. F. de Rooij, et al., Microelectrode arrays for electrophysiological monitoring of hippocampal organotypic slice cultures, IEEE Transactions on Biomedical Engineering, Vol. 44, N 11, pp , 1997 [13] P. Thiébaud, C. Beuret, et al., An array of Pt-tip microelecrodes for extracellular monitoring of activity of brain slices, Biosensors & Bioelectronics, Vol. 14, pp , 1999 [14] P. Thiébaud, Fabrication of microelectrode arrays for electrophysiological monitoring of hippocampal organotypic slice cultures by interface, PhD. dissertation, University of Neuchâtel, Neuchâtel, Switzerland, 1999 [15] H. Scholze, Glass: nature, structure, and properties, Springer Verlag, New York, 1991 [16] G. W. McLellan and E. B. Shand, Glass engineering handbook, McGraw- Hill, New York, 1984 [17] J. Suire, Réactions entre le verre et l'acide fluorhydrique, Silicates Industriels, Vol. 36, pp and pp , 1971 [18] D. J. Monk, D. S. Soane and R. T. Howe, Determination of the etching kinetics for the hydrofluoric acid / silicon dioxide system, Journal of the Electrochemical Society, Vol. 140, N 8, pp , 1993 [19] D.-T. Liang and D. W. Readey, Dissolution kinetics of crystalline and amorphous silica in hydrofluoric-hydrochloric acid mixtures, Journal of the American Ceramic Society, Vol. 70, N 8, pp , 1987 [20] G. A. C. M. Spierings, Wet chemical etching of silicate glasses in hydrofluoric acid based solutions, Journal of Materials Science, Vol. 28, pp , 1993 [21] M. F. Bear, B. W. Connors and M. A. Paradiso, Neuroscience: exploring the brain, Williams & Wilkins, USA, 1996 [22] I. A. Langmoen and P. Andersen, The hippocampal slice in vitro: a description of the technique and some examples of the opportunities it offers, in G. A. Kerkut and H. V. Wheal, Electrophysiology of isolated mammalian CSN preparations, Academic Press, London,

163 5.9 References [23] D. G. Amaral and M. P. Witter, Hippocampal formation, in G. Paxinos, The rat nervous system, Academic Press, Inc, 1995 [24] R. Nieuwenhuys, H. J. t. Donkelaar and C. Nicholson, The central nervous system of vertebrates, Springer Verlag, Berlin, 1998 [25] [26] M. Raggenbass, J. P. Wuarin, et al., Opposing effects of oxytocin and of a µ-receptor agonistic opiod peptide on the same class of non-pyramidal neurones in rat hippocampus, Brain Research, Vol. 344, pp ,

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165 6 MICROCHANNEL FABRICATION TECHNIQUES One important improvement of actual multi-electrode arrays would be the integration of a perfusion system to locally apply chemical compounds. This new feature would allow investigation of local chemical effects on cell activity (stimulation or inhibition) in tissue preparations, which can have an impact on drug discovery applications. As we show in this chapter, such a perfusion system could be fabricated at the MEA chip surface by a PECVD silicon nitride deposition onto a sacrificial photoresist layer or a polymeric multi-layer fabrication technique. However, the fluidic connection to these integrated microchannels can generate a serious problem without some MEA packaging modifications. Other applications like white blood cell (leukocytes) migration following growing chemical gradients, also called chemotaxis, can also be addressed using microchannel devices. A short introduction to the first microchannel device realized for this application will be presented at the end of this chapter. 6.1 MICROFLUIDIC COMPONENTS FOR MEAS Active cells in culture over a multi-electrode array can be stimulated by electrical and or by chemical provocation. Local electrical stimulation can be achieved through the electrodes by applying biphasic current or voltage pulses affecting only the cells located close to the electrodes as shown in Chapter 5.7. Chemical stimulation is very important in electrophysiology investigations because it opens many research and drug discovery applications. Unfortunately, local chemical stimulation is more difficult to achieve due to the surrounding culture solution. It occurs by replacing the culture solution by a new test solution. This can also be done more locally using a micropipette, however, the whole cell/ tissue preparation is stimulated due to fast diffusion of the delivered compounds in the whole culture solution. In this case, the chemical investigations are limited to global compound effects, but do not allow investigation of local effects on tissue structures. To overcome this problem, we proposed to integrate fluidic channels in the MEA chips on the substrate surface as shown in Fig. 6.1 (right). The stimulation 155

166 6 Microchannel fabrication techniques Fig. 6.1 On the left, schematic of chemical compound top perfusion in the culture chamber. The whole tissue will be exposed to the perfusion due to diffusion in the culture medium. On the right, schematic of an integrated microchannel perfusion system. The cells that will be exposed to chemical stimulation are located adjacent to the microchannel end. will occur only around the microchannel end, however, the perfusion flow should be as low as possible in order to avoid the detachment of cells or tissue slices from the device substrate. 6.2 MICROCHANNELS Microchannel fabrication techniques were developed for microelectronic component cooling [1], electrophoresis [2], gas or liquid chromatography [3], and cell properties characterization [4] applications. Various materials such as silicon, glass, metals (copper, gold, etc.), and polymers were used for surface and embedded microchannel realization. The fabrication techniques that have been employed, were wafer bonding sealing of bulk etched channels [2, 3, 5], laser machining and anisotropic silicon etching [5], sacrificial layer etching and deposition sealing [6], metal electroplating [1], and polymer and metal microstructuring [4, 7, 8]. However, a compatible fabrication technology is required in order to integrate surface microchannels on MEA devices. This excludes silicon and glass wafer bonding-based fabrication techniques. Another point is that the channels should be made of an insulating material, which excludes metal electroplating sealing techniques. Furthermore, the used materials should be biologically inert, 156

167 6.3 Silicon nitride microchannels and transparent. This would suggest polymers or other insulation-layer materials such as silicon nitride. The two next sections will present microchannel fabrication technologies based on silicon nitride and SU-8 photoepoxy materials. 6.3 SILICON NITRIDE MICROCHANNELS Silicon nitride has been used extensively in semiconductor technology for device passivation, interlayer insulation and mechanical protection. The conventional silicon nitride deposition method is the use of chemical vapor deposition (CVD), a chemical reaction that is activated by thermal energy. The main disadvantage is the high deposition temperature, which ranges usually from 700 to 900 C. However, the use of plasma-enhanced chemical vapor deposition (PECVD) has the advantage that the substrate can be held at low temperatures (below 400 C). In PECVD, plasma activation provides the radicals that result in the deposited films, and ion bombardment of the substrate provides the energy required to arrive at the stable desired end-products. The operational temperatures are lower, as part of the activation energy needed for deposition comes from the plasma [9-11]. Fabrication of surface microchannels by a method based on a PECVD silicon oxynitride (SiON) deposition on a sacrificial photoresist layer was reported [12]. Microchannels could be obtained due to the room temperature PECVD deposition, since heating of the sacrificial photoresist layer was avoided. Consequently, the photoresist could easily be removed in a subsequent process step. In order to produce silicon nitride surface microchannels for local cell stimulation applications, we adopted the above procedure. It was possible to obtain silicon nitride layer thicknesses in the micrometer range. The depth of the microchannels were only a few tens of microns for channel lengths of several millimeters Process flow of silicon nitride microchannels The microchannels were fabricated in four steps: first a sacrificial photoresist layer was deposited on a glass substrate and structured by photolithography. Then, a silicon nitride layer was deposited by a room temperature PECVD process, defining the future channels. Further, an SU-8 epoxy insulation layer was deposited and patterned above the deposited silicon nitride layer. This step defined the channel openings. Finally, the channel extremities and the channels 157

168 6 Microchannel fabrication techniques Deposition and patterning of SJR5740 photoresist. The resist was reflown at 110 C during 90 s. Room temperature PECVD deposition of a 2 µm thick silicon nitride layer. Deposition and patterning of a 5 µm thick SU-8 layer defining the channel ends. The channels were released by silicon nitride etching in HF solution and photoresist dissolution. Table 6.1 Process flow of silicon nitride surface microchannel fabrication. were released by wet chemical etching of the silicon nitride layer and the dissolution of the sacrificial photoresist layer. The photoresist chosen for sacrificial layer was the Microposit SJR 5740 resist from Shipley. First, layers with a thickness between 10 µm and 30 µm were spun on a standard spin coater. A 17 µm thick layer could be obtained by spinning at 2000 rpm during 10 s. Resist exposure to UV light at 365 nm wavelength, dose of 200 to 400 mj/cm 2, was done on a MA6 mask aligner (Karl Süss). The resist was then developed during 1 min in Microposit 2401 Developer (from Shipley) diluted with deionized water 5:1. Finally, a postbake at 110 C during 90 seconds was necessary to round the profile of the obtained photoresist pattern. The silicon nitride layer was then deposited at room temperature by PECVD (Alcatel 601D PECVD system) using a mixture of SiH 4 and N 2 /NH 3 (1:4). A 3 µm thick silicon nitride layer was deposited in a 30 minutes step at a substrate temperature below 35 C. A 5 µm thick SU-8 epoxy insulation layer was then deposited at 5000 rpm during 40 seconds on a spin coater. A prebake was done at 80 C during 2 hours. 158

169 6.3 Silicon nitride microchannels SU-8 exposure to UV light (365 nm), with a dose of 300 mj/cm 2, defined the future channel openings. Polymerization of the exposed parts was done by a few hours bake in an oven at 60 C. SU-8 development was finally done using polyglycol-methyl-ether-acetate (PGMEA) during 1 minute. To open the channel extremities, the silicon nitride layer was etched in a 10% HF solution during a few seconds. The release of the sacrificial photoresist layer was done in a Microposit Remover 1165 solution (from Shipley) during 2 hours at 80 C Fabrication process discussion Some processing steps were modified in comparison to conventional processing in order to obtain good results. This was specially the case of the SJR5740 and the SU-8 photoresists, which define the channel shape and final cover. The SJR5740 photoresist can be considered as a large thickness photoresist. Layer thicknesses up to 30 µm can be deposited in one single step, which is more than necessary for our application. For high resist thicknesses, it is very important to ensure complete solvent evaporation during softbake. A longer but lower temperature softbake will improve solvent evaporation, thus avoid possible deterioration in the uniformity of the layer and improve the adherence to the substrate. After resist patterning, a high temperature postbake was performed to melt and reflew the resist without burning it [13-15]. The postbake duration corresponds to the minimum time of solvent evaporation needed to evaporate the remaining solvents of the photoresist. It is very important to make this postbake before silicon nitride deposition because the resist profile is modified (no edges left as shown in Fig. 6.2), allowing a homogeneous silicon nitride layer deposition as shown in Fig Moreover, the defined resist pattern will be damaged during the PECVD silicon nitride deposition process due to remaining solvent in the resist layer as shown in Fig. 6.4 (bottom left). At room temperature PECVD deposition, a large concentration of hydrogen atoms remain in the obtained layer. Thus, the silicon nitride layer does not correspond to layers obtained at high temperatures and should be expressed as Si x N y H z. For good silicon nitride layer continuity at the substrate, the contact angle between the substrate and the sacrificial photoresist layer has to be higher than 90 (photoresist shape less than hemispherical as in Fig. 6.2 left). To achieve this angle, the resist profile should respect the following relation between the width and the height of the photoresist line: 159

170 6 Microchannel fabrication techniques Fig. 6.2 Scanning electron microscopy pictures of SJR5740 photoresist melted lines on a glass substrate. The thickness of the deposited resist layer was about 30 µm, and the channel widths shown are 100 µm (on the left) and 40 µm (on the right). Fig. 6.3 Scanning electron microscopy pictures of silicon nitride channels on a glass substrate. Channel aspect ratio of 0.2 for a width of 100 µm (on the left) and aspect ratio of 0.5 for a 40 µm width (on the right). The channel height is 18 µm. width > 4 (6.1) height which corresponds to an aspect ratio that is smaller than As shown in Fig. 6.4 (on the top right), the PECVD deposited silicon nitride follows a columnar growth structure. This is a problem if the contact angle between the substrate and the resist pattern is lower than 90 (photoresist shape greater than hemispherical as shown in Fig. 6.4 top left). In the case of columnar growth, the silicon nitride layer grows perpendicularly from the substrate and 160

171 6.3 Silicon nitride microchannels Fig. 6.4 Scanning electron microscopy pictures of silicon nitride surface microchannels. Top left: when the aspect ratio of the channel is larger than 0.25, the channel walls can easily separate from the substrate. Top right: bubbles can appear at the channel surface due to gaseous hydrogen accumulation in the silicon nitride layer. Bottom left: channel destroyed by resist out-gassing during silicon nitride deposition. Bottom right: cross section of a finished channel embedded in a SU-8 layer. The surface is leveled around the channel by the SU-8 layer. cannot intersect with the channel cover. If two silicon nitride growth fronts form an intersection, the deposition stops and there is a poor adhesion in between. The resulting channel will have poor adhesion to the substrate and will be removed easily when a force is applied (Fig. 6.4 top left). A slower silicon nitride deposition rate would change the growth structure and solve this problem. The deposition of the PECVD silicon nitride layer has two shortcomings: a high hydrogen content (in the range of 20 to 30 atomic percent) and high stress. It is well known that hydrogen is widely present in low temperature silicon nitride layers [9, 16]. The high hydrogen concentration in the silicon nitride layer had 161

172 6 Microchannel fabrication techniques Fig. 6.5 Picture of a silicon nitride microchannel (width of 40 µm, length of 4 mm) filling due to capillary forces by deposition of a solution drop at one channel entrance. Time duration between the two pictures is 1/6 second. This first experiment shows that these channels are watertight. created bubbles at the channel top as shown in Fig. 6.4 (top right). The stress in these thin films varies with the film thickness as well as its hydrogen content. Poor adhesion on glass substrates was also observed without adhesion promotion. An oxygen plasma prior to PECVD silicon nitride deposition was an effective technique to promote the adhesion of the silicon nitride layer onto the glass substrate. A 5 µm thick layer of SU-8 epoxy was deposited as insulation layer. In the developed process, the temperatures of the softbake and polymerization bake were decreased to be compatible with the underlying sacrificial resist and the silicon nitride layer. Decreasing the bake temperature down to 60 C instead of 95 C implied that the corresponding bake duration was increased from 15 minutes to 2 hours. However, the SU-8 insulation layer flattened the surface of the device (see Fig. 6.4 bottom right) and ensured a good mechanical stability of the resulting open channels. Initial channel watertightness and filling tests showed that these surface microchannels were suited for local solution delivery. The test channels remained watertight and could be filled in less than a second due to capillary forces as shown in Fig

173 6.4 SU-8 photoepoxy microchannels Compatibility with MEA fabrication technology The compatibility of the different process steps was demonstrated. Fig. 6.4 (bottom right) illustrates a cross-section of an embedded microchannel. Furthermore, this fabrication process is compatible with the overall MEA fabrication process. After MEA electrode deposition, silicon nitride surface channels can be fabricated instead of a single SU-8 insulation layer. However, no MEAs with integrated surface microchannels have been developed yet, but this is planned for future work. Thus, the solution delivery conditions are still not known in case of such a surface channel ending under a tissue slice. 6.4 SU-8 PHOTOEPOXY MICROCHANNELS The SU-8 photoepoxy was introduced as thick layer photoresist for micromachining applications at IBM in 1995 [17, 18]. It was then specially used as mold material for nickel electroplating [19]. The micromachining of the SU-8 resist was well studied at the EPFL in order to optimize the fabrication of electroplating molds [20-23]. On the other hand, first attempts to produce microchannels made of SU-8 were done [24-26]. Results of this work leaded to the following optimized multi-layer microchannel fabrication process, which is specially adapted for channel dimensions below 50 µm x 50 µm Process flow of SU-8 photoepoxy microchannels The realization of SU-8 surface microchannels is based on the use of a multilayer technique. A first SU-8 layer is deposited and patterned in order to define the channel walls. Then a dry film SU-8 layer is laminated over the first SU-8 layer. This second SU-8 layer is then exposed to UV light, polymerized, and developed in order to obtain a finished channel. Preparation of SU-8 dry films SU-8 dry films can be obtained by deposition and solvent evaporation onto a Mylar support substrate. One way of SU-8 dry film fabrication is the following: A Riston 4700 Series sheet (Du Pont de Nemours International SA) is laminated onto a glass or silicon wafer. Riston is a photopolymer dry film resist that has been developed to perform primary imaging operations required in the production of printed circuit boards as well as chemical milling imaging applications. However, the Mylar protection sheet of the Riston dry film is used as SU-8 support. 163

174 6 Microchannel fabrication techniques Deposition and patterning of a SU-8 layer defining the channel walls. Lamination of a SU-8 film onto the substrate. It will build the channel cover. Exposure and polymerization of the second SU-8 layer to define the channel ends. Development of the second SU-8 layer releasing the microchannels. Table 6.2 process flow of SU-8 multi-layer microchannel fabrication. SU-8 resist is then spun onto the Mylar sheet with thicknesses depending on SU-8 viscosity and deposition conditions (from a few microns up to 1 mm). A prebake at 95 C is then made in order to evaporate the solvent in the SU-8 layer. The duration of this step depends on the SU-8 layer thickness and can last more than 12 hours for 500 µm thick SU-8 layers. The SU-8/Mylar sheet can be released easily from the Riston layer, due to mainly electrostatic adhesion between these two films. On the other hand, the SU-8 dry film can also be released from the Mylar sheet by peeling off. Process flow The first SU-8 layer is spun onto a substrate (glass, silicon, polymer), the film thickness being dependent on SU-8 viscosity and the spinning speed. The solvent 164

175 6.4 SU-8 photoepoxy microchannels Fig. 6.6 Scanning electron microscopy pictures of a channel entrance on a glass substrate (on the left) and a cross section of a channel made in three SU-8 layers (on the right). Channel dimensions are 20 µm x 25 µm. The channel length is several millimeters. in the SU-8 is then evaporated by a bake at 95 C. Here, the duration also depends on the SU-8 thickness. An exposure to UV light (365 nm), a bake at 95 C for 15 minutes for polymerization, and development in poly-glycol-methyl-etheracetate (PGMEA) defines the future channel walls. Next, a previously prepared SU-8 dry film is laminated onto the structure at 0.2 m/minute at 60 C. The laminated SU-8 film is then exposed to UV light and polymerized at only 60 C for 2 hours in order to avoid filling of the channel by the SU-8. Finally, the development of the second SU-8 layer in PGMEA opens the channel Fabrication process discussion The deposition and patterning of the first SU-8 layer defining the channel walls is done following the standard SU-8 photoepoxy photolithography parameters. Further, the preparation of SU-8 dry films is very simple and uses no special technological features. However, the lamination process is a very delicate step. First the Mylar sheet with the SU-8 film is peeled off from the substrate. This is a delicate step, specially when the upper laying SU-8 layer if thicker than 100 µm. If the SU-8 layer thickness is more than 500 µm, is it impossible to release the film without breaking it. Thus, SU-8 dry films with thicknesses over 500 µm should be avoided. 165

176 6 Microchannel fabrication techniques Lamination of SU-8 film occurs after putting it on the patterned substrate. The lamination is done without pressure in order to avoid that the SU-8 film is pressed into the previously defined channel. The heating is necessary for good adhesion of the laminated SU-8 layer. However, exposure to an oxygen plasma prior to the lamination improves the adhesion between the two SU-8 layers. Moreover, to avoid destruction of the channel during lamination, the thickness of the laminated SU-8 film should not be larger than twice the height of the final channel. When the laminated SU-8 layer thickness is larger, the channel is filled during lamination. After lamination, the Mylar sheet is delicately peeled off the laminated SU-8 layer. However, when the laminated SU-8 layer is very thin, some parts of the this film are also peeled off from the underlying SU-8 layer, generating holes in the laminated layer. This phenomenon can be avoided by having a minimum SU-8 layer thickness of 20 µm. The polymerization of the laminated SU-8 layer is done at a lower temperature than in a standard process. This is done to avoid flowing the SU-8 channel cover into the channel. The best way to insure that the channel remains open is to put another Mylar sheet on the SU-8, and to make the polarization step with the device up side-down. The final development of the SU-8 opens the channel. However, the device should be rinsed thoroughly in isopropanol to insure that all the developer is eliminated inside the channels. One interesting advantage is the possibility to laminate several SU-8 layers on a substrate in order to realize more complex channel systems (channels at different levels). Then the lamination process has to be repeated for each new SU-8 layer Compatibility with MEA fabrication technology The SU-8 multi-layer fabrication technique presented is compatible with the MEA fabrication process. However, as mentioned previously, the minimum layer thickness of SU-8 films that can be laminated is about 20 µm. Due to this large film thickness, wells will be formed at the electrode sites on the MEA surface, which should be avoided because of low electrode sealing when cells are present on the MEA. Thus, the SU-8 would be too thick for adequate measurement conditions. On the other hand, a polymer lamination process should be avoided on 3D MEAs because of possible tip damage during lamination. Unfortunately, no tests were made yet to integrate such SU-8 surface microchannels onto MEAs. Thus, the real measurement conditions are not known. However, this simple fabrication technique should be tried in close future 166

177 6.5 Microchannel connection in order to evaluate the stimulation and measurement conditions with such SU-8 microchannels. Besides the integrated MEA solution delivery, other applications can be considered with those SU-8 microchannels. It should be suitable for electrophoresis, cell counting and sorting, cell migration and rheology applications. The first leukocyte migration device composed of a SU-8 microchannel onto a glass substrate, in which measurement electrodes where integrated, was fabricated and is described in the last section of this chapter. 6.5 MICROCHANNEL CONNECTION Two fabrication techniques of small surface microchannels were described using two different materials and its manufacturing techniques. As shown previously, small channels with micron-scale dimensions can be realized over a length of several millimeters. However, connection between an external fluidic system and these integrated channels remains a problem. One possibility is the gluing of a tube connector onto the substrate. This solution works fine, but the overall dimensions of the connectors are of several millimeters, which is too large for easy integration onto MEAs without modifying MEA chip dimensions. Another possible connection to the entrance of a channel can be achieved using an O-ring or a polydimethylsiloxane (PDMS) ring. The PDMS forms a sealed connection, when put in contact with another material. The fabrication of a tube with an O-ring or a PDMS extremity could be a solution for a removable connection between an integrated channel and external fluidics by simple pressing on the MEA substrate at the channel entrance. Here, the main disadvantage would be the mechanical instability of the connection. 6.6 LEUKOCYTE CHEMOTAXIS APPLICATION The ability for active and oriented migration of certain blood cells represents a decisive protection mechanism for our body in the case of infection. These specialized cells, the leukocytes, have the unique capability of detecting with a high sensitivity a chemical signal and then to initiate a directed migration towards its source. A first step in the response to infection is the release by infected cells of chemoattractants that are detected by leukocytes. These cells promptly migrate [27] from the blood toward the infected region where they phagocyte the altered tissue. 167

178 6 Microchannel fabrication techniques Leukocyte migration therefore constitutes an important and, sometimes limiting factor, in the protection of our body from external aggressions. The determining nature of the leukocyte migration is further illustrated in some diseases that are characterized by leukocyte hypermotility whereas other pathologies are related to their hypomotility. Although of clinical relevance, determination of the integrity of leukocyte migration still remains cumbersome. The engineering of a new device allowing rapid characterization of the leukocyte motility would therefore a) be suited for leukocyte receptor characterization and b) represent an additional tool for the clinician to quantify a crucial element of the immune response. We have developed a first chemotaxis microchamber suited for use in the hospital routine, using a microchannel-based device. Since its main usage will be the evaluation of leukocyte chemotaxis in patient care and in clinical studies, the planned chamber should have the following properties: a) the requirement for only small volumes of blood (5µl), b) a low price about 10 $/chamber, c) simple use for bed-side applications through non-specialized personal, and d) reproducible and operator-independent results. The current methods employed to characterize leukocyte motility are the socalled Boyden chamber and agarose gel described below Current measurement methods of leukocyte migration The classical Boyden chamber [28] consists in a semi-permeable filter that separates two compartments (Fig. 6.7 left). Purified blood leukocytes are placed in the upper compartment at approximately 1x10 6 cells/cm 2 and the lower chamber is filled with medium containing a chemoattractant. The stimulated cells transit trough the micropores of the filter. Using this method, the number of cells which have crossed the filter in a given time period can be measured. Thus, cell motion cannot be directly quantified but only indirectly estimated from the cell count. Furthermore, the Boyden chamber technique needs a relatively large quantity of purified leukocytes precluding its use either in experimental setting with small animals or when patient s blood is repeatedly examined. The linear agarose assay [29, 30] is a modification of the under-agarose assay originally proposed by Cutler and Nelson [31, 32]. One percent agarose mixed with medium solution is poured on a glass coverslip and allowed to solidify. Three parallel rectangular wells are cut in the agarose using a template. The middle well is filled with leukocytes. Another one is previously filled with the desired chemoattractant, and adequate time was given to allow the establishment of a chemical gradient by diffusion. The third one is a negative control. The device is placed into the incubator were temperature and humidity are 168

179 6.6 Leukocyte chemotaxis application Fig. 6.7 On the left, diagram of a classical Boyden chamber. Cells are placed in the upper compartment and migrate through a semi-permeable filter toward the lower compartment filled with a chemoattractant. Migration is quantified by counting transmigrated cells in the lower compartment. On the right, picture of an under-agarose migration of leukocytes: on top cells migrating toward a chemoattractant gradient, middle (black) well filled with cells (whole blood), and bottom cells migrating following random migration pattern. maintained. Leukocytes migrating under the agarose (Fig. 6.7 right) from the well toward chemoattractant gradient can be quantified by optical measurements. In particular, the use of a digital camera allows both cheap and fast electronic treatments necessary for the cell quantification using the standard optical analysis package NIH-image (National Institute of Health). Image analysis carried out on the chemoattractant side reveals the typical cell migration toward the compound of interest. In contrast, on the opposite side of the well leukocytes display only their typical random migration pattern. Comparison of the cell speed on the two sides allows further characterization of the compound s effects. Cell migration results from the projection of cell membrane processes (lamellipodium) by polymerization of actin extension and contraction of the remaining cell body. Although the under-agarose method linked with computed image analysis is a powerful tool to explore cell locomotion, this method still remains time consuming and cannot be readily applied clinically. 169

180 6 Microchannel fabrication techniques Fig. 6.8 Schematic of the chemotaxis measurement device. It is composed of two compartments connected by a microchannel allowing leukocyte passage in a single file. Integrated electrodes in the microchannel should detect the passage of a leukocyte by impedance variation measurement between the electrodes Proposed devices Due to new microchannel fabrication techniques, microchannels that have dimensions comparable to those of a single cell can be realized. We designed a dual-compartment chamber linked by a microchannel [4, 33]. The two compartments or reservoirs designed to receive on one hand the leukocytes and on the other hand a defined chemoattractant solution are connected by the elongated microchannel permits passage of leukocytes in single file between the two compartments (Fig. 6.8 left). Gold electrodes were placed at different positions along this fine tunnel to detect by impedance variations measurement the leukocyte s passage (Fig. 6.8 right). Cell detection The cell detection system in the microchannel will be determining for the simplicity of the method of cell migration evaluation. In a first step, optical tracking of the cells will be necessary to assess and verify the quality of the leukocyte migration. Sequences of digital images taken at regular interval will allow the production of short movies in which cell migration within the channel can be followed with high precision. Moreover, this technique will be employed to verify the adequacy of our system in different experimental conditions. Although allowing an excellent cell detection and even pattern recognition for the different cell subtypes the use of an optical method may be difficult to establish for the bed side measurements. We shall therefore develop in parallel an 170

181 6.6 Leukocyte chemotaxis application alternative electrical method based on variation of the channel resistance. An electrical detection of the cells in the channel presents multiple advantages versus an optical method. First, it is cheaper and simpler to implement, and second it does not require optical analysis to be performed. The principle of an electrical detection of the cells is based on an impedance variation in the channel between two electrodes. Since it is known that cell impedance is higher than the medium it follows that the presence of a cell within a given channel section should result in a change of electrical impedance. Providing that this signal is large enough, it should be possible to detect with a rather simple electronic device. Geometry and size of the channel will condition the possible impedance changes and therefore the signal to noise ratio. Thus, although initial chip design is made on the basis of a linear microchannel, other structure may also be evaluated in the future. Providing that cells are crawling through the channel in a single file, the design of multiple electrodes should further allow to quantify both the number of cells passing through the channel as well as their individual speed. The electrical detection necessitate pairs of metallic electrodes deposited on the bottom of the channel. In a first step, the material and the dimensions needed for an optimal detection have to be established by experiments conducted using, for instance, microbeeds as cell simulation. Once established the detection method will then be challenged with leukocytes obtained by purification of a blood sample. Chemoattractant gradient The chemoattractant gradient plays the key role in cell migration. In fact, cells are able to detect a variation of chemoattractant concentration between their front and rear by the number of bindings of chemoattractants to cell receptors of the cell membrane [27]. The detection threshold of chemoattractant concentration variation is about 1% on the length of the cell (about 10 µm) and maximal cell response (migration) was obtained for a concentration gradient of about 10% [34]. A previous model [35] studied diffusion of solutions in a channel between two wells. Expectations showed that the concentration gradient in a channel becomes linear after a short transition period. The main hypothesis of this model is the fact that the volume of the channel is much smaller than the volumes of the reservoirs (more than 500 times smaller) [36]. The persistence of the gradient depends on the viscosity of the solutions. For leukocyte chemoattractant FMLP the gradient remains linear between 5 minutes and 15 hours after the introduction of the chemoattractant in the well. The magnitude of the gradient depends on the channel length (a length of one millimeter is needed to ensure a concentration variation of more than 1% over a length of 10 µm in the channel). 171

182 6 Microchannel fabrication techniques Fabrication process Construction of such a cell counter chamber was realized using the following procedure. First, a 100 µm thick SU-8 dry film was prepared on a Mylar sheet. It is then exposed to UV light and polymerized (bake at 95 C lasting 15 min) in order to form the bottom part of the devices. Then, a chromium (50 nm)/gold (150 nm) layer was evaporated onto the SU-8. A positive photolithography process (S1813 from Shipley) and wet chemical etching defines the electrode pattern. A 10 µm thick SU-8 epoxy layer was spun on the substrate to define the bottom part of the two compartments and the microchannel walls (microchannel width and height of 10 µm). The lamination of a 20 µm thick SU-8 dry film onto the structure allows to generate a covered channel between the two compartments. Finally, the lamination of two 400 µm thick SU-8 dry film layers, UV exposure, polymerization and development builds liquid reservoirs with a volume of approximately 4 µl. To release the chips, the Mylar sheet is peeled off the underlying Riston film and the chips are separated from the Mylar. Obtained results Functional devices were obtained by this multi-layer SU-8 fabrication technique. However, poor adhesion between the SU-8 layers appeared in several devices due to internal stress in the SU-8 polymer, specially for thick layers like the reservoirs. It resulted that many channels were leaking due to delamination. The Table 6.3 gives the dimensions of the devices and Fig. 6.9 (top) shows a whole device and an open microchannel with integrated pairs of gold electrodes. The height of the reservoirs of this chip is 800 µm made by two 400 µm thick SU-8 layers, generating a volume of 4 µl. The first test made on these devices was the control that the channel is open. This was done by applying a solution in one reservoir, i.e. the channel is filled by capillary forces in presence of a solution (see Fig. 6.9 bottom left). First experiments showed that white blood cells are able to flow through the fabricated microchannels. The solution volume disequilibrium between the reservoirs generated a flow through the microchannel for volume compensation. When leukocytes were placed in the reservoir with the larger solution volume, some cells flew due to the pressure drop through the channel. Furthermore, chemoattractant gradients were established along the microchannels as demonstrated by the accumulation of cells close to the entrance of the channels (see Fig. 6.9 bottom right). A global migration in the bottom of the well in the direction of the channel entrance was also observed in these experiments. Although the cells accumulate at the channel entrance, no cells migrated inside the channel. Thus, no cell detection and counting could be achieved with 172

183 6.6 Leukocyte chemotaxis application Part of the device Chemotaxis device One compartment Dimensions 7 mm x 7 mm diameter of 2.5 mm, height of 800 µm It results a volume of ~4 microliters Microchannel 10 µm x 10 µm, length of 1 mm It results a volume of 100 picoliters Electrodes 20 µm x 10 µm, interspace of 20 µm Table 6.3 Dimensions of the device Fig. 6.9 Top: pictures of the whole chemotaxis device (on the left) and the channel (on the right) realized using a multi-layer SU-8 fabrication method. Bottom: Microchannel filling test (on the left) and leukocytes chemotaxis experiment (on the right). these devices yet. This stopping of cell migration should be due to geometrical effects, due to the fact that many cells migrate to only one single point (the 173

184 6 Microchannel fabrication techniques Fig Migration of leukocytes in a microchannel network (channel width of 30 µm, separation between two channels 30 µm). Leukocytes migration toward a chemoattractant gradient is promoted in this configuration due to many possible paths. channel entrance). It seems that when too many cells migrate to the same point, they stop migrating because of a cell accumulation eliminating the chemoattractant gradient seen by the cells. One way to overcome this problem should be the use of multiple channels allowing migration through a larger surface instead of one single channel Considerations for future devices To test if multiple microchannels placed side by side will promote cell migration, arrays of microchannels with different widths were realized (without measurement electrodes). First experiments driven with leukocytes showed that the cells enter the channels and migrate following the channel walls toward an established chemoattractant gradient. However, in this test, the gradient was not present in the solution but was previously fixed into an agarose gel layer placed onto the channels as cover. This migration can thus be considered like an under-agarose gel migration. The cells prefer the microchannel space for migration but some cells can also be found at places in between the channels as shown in Fig The combination of multiple microchannels with an agarose gel cover seems to be promising for future functioning leukocytes chemotaxis devices. However, new devices integrating multiple microchannels combined with measurement 174