(PDF) Growing Cells Atop Microelectronic Chips: Interfacing Electrogenic Cells In Vitro With CMOS-Based Microelectrode Arrays

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Growing Cells AtopMicroelectronic Chips:Interfacing Electrogenic CellsIn Vitro With CMOS-BasedMicroelectrode ArraysThis paper offers an overview of the fundamentals of bioelectronic measurements,

as well as the design, system integration, and application of CMOS-based microarrays

for recording from and for stimulating electrogenic cells from the brain and heart.

By Andreas Hierlemann, Member IEEE, Urs Frey,

Sadik Hafizovic, and Flavio Heer, Member IEEE

ABSTRACT | Complementary semiconductor–metal–oxide

(CMOS) technology is a very powerful technology that can be

more or less directly interfaced to electrogenic cells, like heart

or brain cells in vitro. To this end, the cells are cultured directly

atop the CMOS chips, which usually undergo dedicated post-

processing to obtain a reliable bidirectional interface via noble-

metal microelectrodes or high-k dielectrics. The big advantages

of using CMOS integrated circuits (ICs) include connectivity, the

possibility to address a large number of microelectrodes on a

tiny chip, and signal quality, the possibility to condition small

signals right at the spot of their generation. CMOS will be de-

monstrated to constitute an enabling technology that opens a

route to high-spatio–temporal-resolution and low-noise elec-

trophysiological recordings from a variety of biological pre-

parations, such as brain slices, or cultured cardiac and brain

cells. The recording technique is extracellular and noninvasive,

and the CMOS chips do not leak out any toxic compounds, so

that the cells remain viable for extended times. In turn, the

CMOS chips have been demonstrated to survive several months

of culturing while being fully immersed in saline solution and

being exposed to cellular metabolic products. The latter re-

quires dedicated passivation and packaging techniques as will

be shown. Fully integrated, monolithic microelectrode systems,

which feature large numbers of tightly spaced microelectrodes

and the associated circuitry units for bidirectional interaction

(stimulation and recording), will be in the focus of this review.

The respective dense microelectrode arrays (MEAs) with small

pixels enable subcellular-resolution investigation of regions of

interest in, e.g., neurobiological preparations, and, at the same

time, the large number of electrodes allows for studying the

activity of entire neuronal networks. Application areas include

neuroscience, as the devices enable fundamental neurophys-

iological insights at the cellular and circuit level, as well as

medical diagnostics and pharmacology.

KEYWORDS | Biomedical electronics; biological neural net-

works; bioelectric phenomena; complementary metal–oxide–

semiconductor (CMOS) integrated circuits; CMOS technology;

microelectrodes; mixed analog digital integrated circuits;

monolithic integrated circuits

Manuscript received December 2, 2009; accepted July 7, 2010. Date of publication

October 7, 2010; date of current version January 19, 2011. This work was supported

by ETH under the Internal Grant TH-00108, and by the Information Societies

Technology (IST) Future and Emerging Technologies program of the European

Union and the Swiss Bundesamt fur Bildung und Wissenschaft (BBW) under Contract

IST-2000-26463.

A. Hierlemann and F. Heer are with the Bio Engineering Laboratory, Department of

Biosystems Science and Engineering, ETH Zurich, CH-4058 Basel, Switzerland

(e-mail: [emailprotected]; [emailprotected]).

U. Frey was with the Bio Engineering Laboratory, Department of Biosystems Science

and Engineering, ETH Zurich, CH-4058 Basel, Switzerland. He is now with IBM Research

Zurich, 8803 Ruschlikon, Switzerland (e-mail: [emailprotected]).

S. Hafizovic was with the Bio Engineering Laboratory, Department of Biosystems

Science and Engineering, ETH Zurich, CH-4058 Basel, Switzerland. He is now with

Zurich Instruments, 8005 Zurich, Switzerland (e-mail: [emailprotected]).

Digital Object Identifier: 10.1109/JPROC.2010.2066532

252 Proceedings of the IEEE | Vol. 99, No. 2, February 2011 0018-9219/$26.00 �2010 IEEE

I . INTRODUCTION

A high degree of connectivity and the coordinated elec-trical activity of neural cells or networks are believed to bereasons why the brain is capable of highly sophisticatedinformation processing. Likewise, the effectiveness of ananimal’s heart also largely depends on such coordinatedcell activity. To advance our understanding of these com-plex biological systems, high-spatio–temporal-resolutiontechniques to monitor cell electrical activity and an ideallyseamless interaction between cells and recording devicesat both, single-cell and network levels are desired.

Methods to directly measure the electrical activity ofcultured electrogenic cells (e.g., neurons, heart cells,retina cells, or muscle cells) include two fundamentallydifferent techniques: 1) transmembrane measurements byaccessing the cell interior or cytosol with one of the elec-trodes, the so-called Bpatch clamp[ technique [1], [2], and2) extracellular recordings, e.g., by means of externalmicrotransducers [3]–[14]. Additionally, there are indirectmethods like optical measurements using voltage-sensitiveor fluorescent dyes [15]–[17]. The patch clamp techniqueyields very accurate information on the electrophysiolog-ical properties of entire cells, or, alternatively, on currentsflowing through single ion channels. However, it is aninvasive method and is limited in the cell viability time(usually hours) and in the overall number of cells that canbe simultaneously investigated.

For extracellular recordings, the cells are cultured orplaced directly atop the sensors or electrodes [Fig. 1(a)].When electrical activity occurs in a cell, ions flow acrossthe cell membrane within milliseconds. These moving ionsgenerate an electric field, which can directly influence theopen-gate region of a field-effect transistor (FET) [18]–[22], or which can be recorded by means of metalmicroelectrodes [23]–[31]. Extracellular recordings arenoninvasive (no puncturing of the cell membrane), which

enables long-term measurements. Multisite measurements

are possible by arranging many sensors in an array. For

stimulation, voltage transients can be applied by means of

electrodes or stimulation spots, which then evoke a depo-

larization of the nearby cell membrane and solicit

subsequent electrical cell activity. A stimulation spot is a

several-micrometer-diameter spot of either metal or di-

electric materials with high dielectric constant, such as

TiO2 ð" ¼ 34Þ and HfO2 ð" ¼ 15:4Þ, connected to an elec-

trical supply, through which stimulation signals are

applied [32]–[34].The most common biological preparations studied

using microelectrode arrays (MEAs) are divided into two

main categories: acute tissue preparations (e.g., slices),

which are recorded from immediately after they have been

removed from the animal [Fig. 1(b)]; cell cultures, which

can be further divided into dissociated cell cultures and

organotypic tissue cultures. Dissociated cultures are cells

whose relative in vivo positions are no longer preserved[see, e.g., Fig. 19(b)], whereas organotypic tissue cultures

are slices that are maintained in vitro over a period of time.

These variations encompass the spectrum from short-term,

functionally preserved tissue slices to long-term organo-

typic or cell culture preparations.

Several parameters must be taken into consideration

when using MEA chips as a tool for recording from cells.

Most importantly, the cell–electrode interface must beoptimized so that tight coupling between the cells and

electrodes is established, maximizing the signal-to-noise

ratio. The high-density microelectrode arrays (HD-MEAs)

described later (Section IV-E2) are capable of recording

from as deep in the tissue as 100 �m, but the signal-to-

noise ratio drops off rapidly with distance [35]. Acute

slices, for example, have a layer of dead cells that are

damaged during the cutting of the tissue, which introducesa degree of isolation between the active cells and the

Fig. 1. (a) Schematic of a cell attached to a sensor surface. A cell featuring ion channels sits on a planar microsensor (open-gate transistor or

electrode). Moving ions in the electrode vicinity generate an electric field or voltage that is recorded by the microsensor. (b) Micrograph of an

acute cerebellar brain slice (parasagittal cut) placed on a CMOS high-density electrode chip for measurements. The white regions of the slice

constitute myelinated areas, where no electrical activity can be recorded.

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electrodes. To promote the adhesion of cells or the tissuemass, the MEA chips are usually coated with cell-adhesion

substrates such as poly-D-lysine, laminin, collagen, and

fibronectin [36], [37]. Successful tactics for making

recordings from tissue sections utilize, for example,

cellulose nitrate or a coagulated plasma/thrombin clots

to stabilize the slices on the chip surface [38]. Further-

more, a perfusion system usually is necessary for the

delivery of carbogen gas (95% O2/5% CO2) and nutrients,which are required to sustain the physiologic processes of

the cells. Chip-through-holes have been realized on

circuitless chips to further improve nutrient access to the

bottom slice face on the chip and to thus enhance the

viability time [39]. This is very difficult if not impossible to

realize in high spatio–temporal resolution arrays due to

the densely packed circuitry and connection lines in the

chip as will be evidenced in Section IV.Substrate-integrated MEAs enable the study of neuro-

nal information processing [40]–[44] or the characteriza-

tion of cellular responses upon dosing biologically active

agents [25], [45]–[47]. Studying electrogenic cells in vitroprovides different experimental conditions in comparison

to investigating cells in vivo or in intact organisms: Most

animal and plant tissues comprise a variety of different cell

types, whereas a single specific cell type of interest orassemblies of selected different cell types can be grown in

culture. Moreover, the experimental parameters, such as

temperature or concentration of a chemical compound,

can be controlled much more rigorously in cell cultures as

compared to an organism. However, inferences from

in vitro-experiments to the in vivo behavior of cells have to

be experimentally verified. The investigation of acute or

organotypic slices, such as brain or other tissue slices[Fig. 1(b)] may help to bridge the gap between in vitro and

in vivo experiments to some extent [48], [49].

II . WHY USE IC OR CMOS TECHNOLOGY?

This review covers the application of complementary

semiconductor–metal–oxide (CMOS)-based, active micro-

arrays for recording from and stimulation of electrogeniccells in vitro. BActive[ array means that circuitry compo-

nents, such as amplifiers or filter units, are integrated with

the electrodes on the same substrate. The focus will be on

in vitro applications so that needle arrays carrying electrodes

and the respective in vivo applications will be omitted, as they

have been reviewed before in this journal [4], [50].

The three main advantages in using integrated circuit

(IC) or CMOS technology include most importantly1) connectivity; large numbers of transducers or electrodes

can be addressed by on-chip multiplexing architectures;

then 2) signal quality; the signal is conditioned right at the

electrode by means of dedicated circuitry units (filters,

amplifiers); and, less significantly, 3) ease of handling anduse; the devices and signals are robust (no Faraday cages,

digital signals, see below); many functions can be prog-

rammed or automated via user-friendly software anddigital interfaces that directly address digital registers

and logic or memory units on the chip side. Examples of

automation include self-identification of the chip, storage

of calibration or operation parameters, and the execution

of self-test functions. The presence of user-friendly fea-

tures is a crucial issue to encourage neuroscientists and

biologists to adopt such systems.

Traditional MEAs without multiplexers usually offer64 electrodes, with each electrode needing a connection to

external circuitry, which adds parasitic capacitance and

attenuates the small signals. CMOS-based MEAs, however,

comprise up to 16 384 electrodes with the necessary ad-

dressing circuitry on the same chip [49], [51]–[54]. The

high transducer density is an important asset, which

cannot be realized without on-chip multiplexers and ad-

dressing, simply since incorporating thousands of physicalor electrical connections on a millimeter-size chip is

impossible.

On-chip microelectronics provided by IC or CMOS

technology translate into improved system capability. For

example, signal conditioning can be performed on-chip,

ensuring that weak neural signals are faithfully recorded. It

is mandatory to not only amplify the signals but to also

limit the signal bandwidth to avoid noise aliasing in thesubsequent multiplexing. There are only very short

connection lines (a few micrometers) between electrode

and signal conditioning circuitry, and there are no long

lead lines running close to saline or the electrically active

culture under a thin layer of insulation. Therefore, any

interference with the target signals is largely obviated so

that Faraday cages and electrical shielding are not needed.

CMOS systems provide bidirectional communication viathe electrodes as they feature stimulation and recording

units. Active circuitry allows the stimulation of cells via

single or multiple electrodes in arbitrary configurations

from the set of densely packed electrodes, all while re-

cording from other electrodes without interruption. On-

chip analog-to-digital conversion (ADC) means that the

chip produces signals that may be easily manipulated and

transferred without compromising its information content.There is no influence of mechanical and electrical con-

nections (spring contacts) that cause problems for small

analog signals. Finally, the complete system can be mono-

lithically integrated on a single chip, which features an

area on the order of 20–40 mm2. The resulting measure-

ment setups are small, easily transportable and can be used

on a lab bench, or several of the devices can be placed in an

incubator with only a flat-band cable making connection tothe rest of the setup (see Section IV-E3).

As yet, the detailed structure of the extracellular-

potential landscape of a neuron and its dynamics are still

mostly unknown, since the signal-to-noise ratio and the

spatial resolution in the recordings of extracellular poten-

tials are, in most cases, low. This prevents a clear sepa-

ration of individual signal sources and an unambiguous

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254 Proceedings of the IEEE | Vol. 99, No. 2, February 2011

reconstruction of the neuronal networks derived from fieldpotentials or multiunit recordings [55]. HD-MEAs, real-

ized in standard CMOS technology, offer the potential to

perform recordings at single-cell or even subcellular reso-

lution [53]. The devices, however, have to meet two re-

quirements. They have to provide 1) high signal-to-noiseratio in the recordings of very small signals and, at the

same time, they have to enable 2) high spatial and temporalresolution measurements. The dilemma is that the two re-quirements are diametrically opposed in a technical reali-

zation. High spatial resolution entails the use of small

electrodes featuring higher thermal noise and that only

very few and small circuitry elements for addressing and

signal amplification can be realized per electrode. The

smaller the available space for electrodes and circuitry is,

however, the larger the noise is, since the noise of a tran-

sistor scales inversely with size, and since filters or am-plifiers requiring large capacitors cannot be realized. This

means that the signal quality is compromised by the cir-

cuitry generally needed to readout such high-density

arrays. Therefore, several devices to date feature a com-

parably high noise level that prevents the revelation of

electrophysiological details. There are, however, work-

arounds as will be shown in Section IV-E2.

A disadvantage of CMOS chips is that silicon is nottransparent to visible light in contrast to standard cell

culture substrates used in biology. Additionally, the chip or

its components may corrode upon operation and long-term

exposure. This holds particularly true as CMOS chips are

usually not made for being immersed in salt water over

extended times (weeks and months). Moreover, as soon as

stimulation features are implemented on-chip, severe

problems may arise from electrochemistry upon deliveringcurrent or voltage pulses and from the different redox

potentials of the variety of exposed materials, which may

generate local electrochemical elements and cause massive

corrosion (e.g., complete dissolution of CMOS metal alu-

minum). Therefore, a good packaging solution is needed,

on the one hand, to protect the chip against metabolic

products and chemicals of the cell culture, and, on the

other hand, to prevent the cells from being poisoned ordisturbed by toxic materials released by the chip, such as

the CMOS metal aluminum (also copper in modern pro-

cesses) that dissolves in saline solution. Viable solutions

will be presented in Section IV.

In the following, this review will first provide a short

introduction to the fundamentals of bioelectronics and

bioelectronic measurements (Section III), and then, an

overview on the design and application of CMOS-basedMEAs for interacting with electrogenic cells in vitro(Section IV). All systems feature the sensors or transducers

and the circuitry components on the same chip (monolithic

design). CMOS circuitry-only chips featuring recording

and/or stimulation units (see, e.g., [56] and [57]), which

are then connected to separate circuitless electrode array

chips (mostly on glass substrates), will not be covered here.

III . FUNDAMENTALS OF RECORDING OFELECTRICAL CELL ACTIVITY

A. Electrogenic CellsMany types of cells in the body have the ability to

undergo a transient electrical depolarization and repo-

larization that is either triggered by external mechanisms

(e.g., motor-nerve stimulation of skeletal muscle, or cell-

to-cell depolarization in the heart) or by intracellular,spontaneous mechanisms (e.g., cardiac pacemaker cells)

[58]. Cells that exhibit the ability to generate electrical

signals are called electrogenic cells, the most prominent of

which include brain cells or neurons and heart cells or

cardiomyocytes. Only a brief description of the biology of

electrogenic cells will be given in this section; for more

details the reader is directed to various books on this topic

[58]–[60].A huge number of neuronal cells form the nervous sys-

tem that regulates all aspects of body functions. The hu-

man brain contains about 1012 neurons (nerve cells) with

each neuron forming thousands of connections to other

neurons so that a large network of electrical connections

with massively parallel information processing character-

istics results. The output of a nervous system is the result

of its inputs and its circuit properties, that is, of the wiringor interconnections (synapses) between the single neu-

rons, and of the strength of these interconnections. The

synaptic connections between neurons can be reorganized,

which is known as synaptic plasticity, and is believed to be

a mechanism of learning in our brain. Different types of

neurons can be distinguished according to their physiology

and function in the body [58], [60].

Neurons exhibit four distinct regions with differentfunctions: the cell body, the dendrites, the axon, and the

axon terminals (Fig. 2). The cell body or soma with a dia-

meter of 10–30 �m contains the nucleus and is the pro-

duction site of most neuronal proteins. The majority of

neurons have a single axon, whose diameter varies from a

micrometer in the human brain to a millimeter in the giant

squid. Axons are specialized for the conduction of elec-

trical pulses, termed action potentials, away from the cellbody towards the axon terminals. Most of the neurons are

polarized: They receive synaptic input at the dendritic end

and provide output through synapses at the axonal end

(Fig. 2) [58].

Fig. 2. Schematic view of a neuron. Modified from [58].

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Neurons communicate with other neurons through

specialized contact zones, which are called synapses. Sy-napses generally transmit signals in only one direction,

from the axon terminal of the presynaptic cell to the den-

drites of the postsynaptic cell. Synapses can be either

electrical or chemical [58], [60]. Electrical synapses are

normally located at the gap junctions, which consist of

specialized proteins that form channels bridging the in-

teriors of two neurons and allow current flow from one

neuron to the other. In a chemical synapse, action po-tentials trigger the secretion of neurotransmitters in the

presynaptic neuron. The neurotransmitters then diffuse

across the synaptic cleft (20–25-nm width), and bind to a

receptor site on the postsynaptic membrane and cause

electrical signals there. The time required for this process,

the so-called synaptic delay, is typically 0.3–0.5 ms. The

synapse is an important feature for neural networks to

learn and to perform computation.Different ion concentrations inside and outside the

neuronal cell lead to an electric potential across the plasma

membrane (Fig. 3). In virtually all cases, the inside of the

cell membrane is negatively charged relative to the outside

[58], [60], i.e., a concentration gradient of ions between

the extracellular and the intracellular space (cytosol)

exists. Ion channels are embedded transmembrane proteins

that allow for the selective passage of certain ions. Ionchannels are termed gated if they can be opened or closed.

There are three types of gated ion channels: ligand-gated,

mechanically gated, and voltage-gated channels. The con-

centration of Kþ ions inside typical metazoan cells is about

ten times larger than that in the extracellular fluid, where-

as the concentrations of Naþ and Cl� ions are much higher

outside the cell than inside (Fig. 3). These concentration

gradients are maintained by membrane proteins that are

capable of pumping ions from one side of the membrane tothe other, often against their concentration gradients,

which consumes energy. There are additional types of ion

channels involved, but the membrane potential of most

electrically active cells is primarily affected by the opening

and closing of ion channels for Kþ, Naþ, Cl�, and

sometimes Ca2þ [58], [60].

The concentration gradient across the cell membrane

leads to an electrical potential, which is called restingpotential. The resting potential is mainly a consequence of

the membrane permeability to Kþ ions. The plasma mem-

brane contains resting Kþ ion channels, which are open in

the resting state and allow the passage of only Kþ ions. The

Kþ ions move across the membrane down their concen-

tration gradient, which leaves an excess negative charge on

the cytosolic face and deposits positive charge on the

exoplasmic face

Ex ¼RT

zxFln

cox

� �

dcixe: (1)

A variant of the Nernst equation (1) allows us to calcu-

late the maximum potential difference across a membrane

Ex (electromotive force), for any ion x at concentration cx.

Ex is given by the product of the thermal voltage RT=zxFand the concentration difference. R is the molar gas

constant, T is the temperature in Kelvin, zx is the number

of exchanged charges (1 for Kþ, Naþ, and Cl�), and F is theFaraday constant. The superscripts Bi[ and Bo[ denote the

ion concentration inside and outside the cell, respectively

[see (1)]

VM ¼RT

F� ln ½Ko�PK þ ½Nao�PNa þ ½Cli�PCl

½Ki�PK þ ½Nai�PNa þ ½Clo�PCl: (2)

The electric potential VM across a cell membrane can

be calculated by the Goldman–Hodgkin–Katz voltageequation (2), a more complex variant of the Nernst

equation, in which the concentrations of the ions are

weighted in proportion to the magnitudes of their trans-

membrane permeability constants P. P is a measure of the

ease with which an ion can cross a unit area of a membrane

driven by a one-molar difference in concentration; it is

proportional to the number of open ion channels and the

number of ions that each channel can conduct per second.Thus, PNa, PK, and PCl are measures of the Bleakiness[ of a

unit area of the membrane to these ions. The permeability

of the membrane to Naþ and Cl� ions is about one tenth

of that for Kþ, since there are about ten times more open

Kþ channels than open channels for Naþ or Cl�. In

theory, the resting membrane potential may be anywhere

between EK (�90 mV), the potassium equilibrium

Fig. 3. Intra- and extracellular ion concentrations for the species

primarily involved in the electrogenic properties of the neuronal

membrane, Naþ, Kþ, and Cl�. Redrawn according to [58]. Note that the

outside cell membrane is considered to be positively charged.

A� denotes all nonchloride anions (phosphates, etc.).

Hierlemann et al.: Growing Cells Atop Microelectronic Chips: Interfacing Electrogenic Cells In Vitro

256 Proceedings of the IEEE | Vol. 99, No. 2, February 2011

potential, and ENa (55 mV), the sodium equilibrium po-

tential. By inserting the typical ion concentrations into (2),

the potential of the resting membrane amounts to, e.g.,

�62 mV for the squid giant axon. Most neurons feature

resting potentials around �60 mV.

An action potential event is largely the result of

voltage-gated ion-channel conductances that cause a depo-

larization and repolarization of the neuronal plasma mem-brane within 2–3 ms [Fig. 4(a)]. The resting potential

across the membrane is, as already mentioned, maintained

by large transmembrane proteins that pump ions, against

their electrochemical gradient, in and out of the cell. The

action potential is triggered when the membrane is depo-

larized to a threshold value causing voltage-gated ion

channels to open. Naþ ions then flow down their elec-

trochemical gradient across the cell membrane (inwardcurrent, INa), further depolarizing the membrane. Depo-

larization causes Kþ ion channels to open, causing a cur-

rent flow in the opposite direction (outward current, IK)

that repolarizes the membrane [see Fig. 4(a)]. Finally, all

ion channels relax to their original state, and the mem-

brane returns, often after a brief hyperpolarization (mem-

brane potential more negative than resting potential), to

its resting potential. It has to be noted that the occurrenceof an action potential is an electrical event generated by a

change in the distribution of charge across the membrane

and not by a marked change in the overall intracellular or

extracellular concentration of Naþ or Kþ ions (ion con-

centration change of less than 1 per mill inside the cell).

Action potentials propagate unidirectional from the

axon hillock down the axon to the axon terminals (Fig. 2).

This is due to the fact that the voltage-gated Naþ channels

remain inactive for several milliseconds after opening, a

time span called the refractory period [58], [60]. The

inability of Naþ channels to reopen during the refractory

period also limits the number of action potentials per

second that a neuron can conduct.

Cardiac muscle cells, also denoted cardiac myocytes (or

cardiomyocytes), share similarities with skeletal muscle

cells with regard to their striated appearance and contrac-tion. A coordinated and synchronous activity of the cardio-

myocytes is necessary to effectively pump blood through

the vessels. Therefore, there is a direct connection be-

tween the cytoplasms of cardiac cells, called gap junctions,

which allows various molecules and ions to pass freely and

without attenuation. Accordingly, a depolarization wave

rapidly spreads through those direct connections without

characteristic temporal delay, which is observed forneuronal chemical synapses [58]. This direct cell-to-cell

connection means that one usually does not observe single-

cell action potentials, but so-called Bfield potentials[ upon

using extracellular electrodes to record from cell ensem-

bles. Field potentials are comparably large-amplitude

signals that result from simultaneous, collective electrical

activity of patches of electrogenic cells.

There are two general types of cardiac action poten-tials: 1) nonpacemaker action potentials, which are found

throughout the heart, and 2) spontaneous action potentials

of pacemaker cells. The latter are found in the sinoatrial

and atrioventricular nodes of the heart [6], [58]. Both

types of action potentials in the heart differ considerably

from those of neuronal cells in the duration of the action

potentials [neuronal cell about 1 ms; see Fig. 4(a); cardiac

action potential from 200 to 400 ms; see Fig. 4(b) and (c)]

Fig. 4. Schematic, time course, and current contributions of action potentials of different electrogenic cells as measured intracellularly.

Arrows pointing to the right represent ion currents flowing into the cell, arrows pointing to the left represent outward currents. (a) Neuron: Onset

of the action potential at 0.5 ms; the membrane potential becomes more positive (depolarization) through Naþ influx. After 1 ms return via

hyperpolarization state (potential more negative than resting potential) to resting potential through opening of Kþ channels: Kþ efflux.

Adapted from [58]. (b) Cardiomyocytes (ventricular myocyte): The numbers 0–4 mark the initial rapid depolarization phase (0), an initial

repolarization phase (1), a plateau phase (2), a rapid repolarization phase (3), and the membrane resting potential (4); for more information,

see text. (c) Action potential of a cardiac pacemaker cell: The numbers here mark the depolarization phase of the action potential (0),

a repolarization phase (3), and a spontaneous depolarization phase (4); for more information, see text.

Hierlemann et al. : Growing Cells Atop Microelectronic Chips: Interfacing Electrogenic Cells In Vitro

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and the involved ion channels. In neuronal and nonpace-maker cardiac cells, the depolarization phase of the action

potential is caused by an opening of sodium channels,

whereas in cardiac pacemaker cells, calcium ions are in-

volved in the initial depolarization.

Nonpacemaker cardiac cells have a true resting mem-

brane potential [phase 4, Fig. 4(b)] that remains near the

equilibrium potential for Kþ ðEKÞ. When these cells are

rapidly depolarized to a threshold voltage of about�70 mVby another conducted action potential (phase 0), a fast

Naþ-current INa flows into the cell. At the same time, the

outward Kþ current IK is interrupted. These two changes

in ionic currents move the membrane potential farther

away from the resting potential and closer toward the

sodium potential. Phase 1 then represents an initial repo-

larization caused by the opening of a special type of Kþ

channel producing a transient outward Kþ current, IKðtoÞ.However, due to the large increase in the slow inward

Ca2þ current ICaðlÞ, the repolarization is delayed, and there

is a plateau phase in the action potential (phase 2). This

inward calcium movement is through long-lasting (l-type)

calcium channels that open up when the membrane

potential depolarizes to about �40 mV. Repolarization

(phase 3) occurs, when the outward Kþ current IK in-

creases again, and the Ca2þ current ICaðlÞ decreases.Therefore, the action potential in nonpacemaker cells is

primarily determined by changes in fast Naþ, slow Ca2þ,

and Kþ currents. During phases 0, 1, 2, and a part of

phase 3, the cell is refractory to the initiation of new action

potentials, i.e., the cell does not show any electrical

response.

Pacemaker cells are characterized as having no true

resting potential, but instead generate regular, spontane-ous action potentials [58]. There are, in fact, no fast Naþ

currents operating in those cells [Fig. 4(c)]. At phase 0, the

depolarization is primarily due to increased inward Ca2þ

currents ICaðlÞ. As the movement (or conductance) of Ca2þ

through the channels is not rapid, the rate of depolariza-

tion (slope of phase 0) is much slower than in other cardiac

cells. Repolarization occurs (phase 3) as the outward Kþ

current IK increases and the Ca2þ current ICaðlÞ decreases.Spontaneous depolarization (phase 4) is due to a fall in the

Kþ current IK and a small increase in the Ca2þ current

ICaðtÞ. A slow inward Naþ current also contributes to

phase 4, and is thought to be responsible for what is

termed the pacemaker or Bfunny current[ If . Once this

spontaneous depolarization reaches threshold (approxi-

mately �40 mV), a new action potential is triggered.

Although pacemaker activity is spontaneously generated,the rate of this activity can be modified significantly by

external factors such as autonomic nerves, hormones,

ions, and ischemia/hypoxia.

The rhythm of the heart is normally generated and

regulated by pacemaker cells within the sinoatrial (SA)

node, which is located within the wall of the right atrium.

SA nodal pacemaker activity normally governs the rhythm

of the atria and ventricles [6]. When this rhythm becomesirregular, too fast (tachycardia) or too slow (bradycardia),

or the frequency of the atrial and ventricular beats are

different, then this is called an arrhythmia [6]. Arrhyth-

mias can develop from either altered impulse formation or

altered impulse conduction. The former concerns changes

in rhythm that are caused by changes in the automaticity of

the pacemaker cells or by abnormal generation of action

potentials at sites other than the SA node. Altered impulseconduction is usually associated with a complete or partial

block of electrical conduction within the heart and

commonly results in reentry, which can lead to tachyar-

rhythmias. Many origins of arrhythmia, e.g., the so-called

QT-prolongation (Q and T denote specific features in the

waveform of an electrocardiogram), a delayed repolariza-

tion following the depolarization or excitation of the heart

[47], can be also observed in cell cultures, which opens alarge field of applications for MEAs in the pharmaceutical

industry (see also Section IV-E4) [25], [61].

B. Electrical Interface and Signal RecordingThe electrical interface includes the cell membrane in

physiological solution on the one hand and the transducer

on the chip on the other hand. The electrical properties of

the cell membrane are usually described by the gate model

of Hodgkin and Huxley [62]. The gate model has been

developed studying the squid axon and proposes that ion

currents are flowing through transmembrane channelsdown their concentration gradients. The channels feature

voltage-sensitive gating or gating particles. Fig. 5 shows the

equivalent circuit of the neuronal membrane according to

the gate model of Hodgkin and Huxley. The equivalent

circuit consists of parallel branches for the different ionic

contributions and the membrane capacitance. The tunable

resistors represent voltage-gated ion channels, and gx and

Ix denote the specific voltage- and time-dependent con-ductance for ion x and the current carried by ion x (e.g.,

Naþ, Kþ, and Cl�). The voltage sources arise from ionic

concentration gradients across the cell membrane, and the

Nernst potential Ex represents the equilibrium potential of

ion x according to (1). The main ionic currents include the

potassium current IK, the sodium current INa, and some

smaller currents summarized in a leakage current IL:. The

resulting total membrane current is then described by

(PDF) Growing Cells Atop Microelectronic Chips: Interfacing Electrogenic Cells In Vitro With CMOS-Based Microelectrode Arrays - DOKUMEN.TIPS (2024)

References