Noise in biopotential recording using surface electrodes

S – 915

Noise in biopotential recording using surface electrodes

E. Huigen

November 2000

Delft Technical University

Faculty of Design, Engineering and Production

Mechanical Engineering and Marine Technology

Man Machine Systems

Title: Noise in biopotential recording using surface electrodes Author: E. Huigen Date: november 2000 Report number: S – 915 Pages: 37 Institute: University of Amsterdam Section Medical Physics Supervisors: Dr. A. Peper Prof. Dr. Ir. C.A. Grimbergen Summary: Electric events inside the human body cause potential differences on the skin. Surface electrodes can be used to measure these voltages. A surface electrode is in fact an electrochemical transducer. It converts the potential differences due to ion flows inside the human body into a measurable voltage. The measurement of bioelectric events is exposed to various sources of noise. The reactions that take place at the electrode make the electrode itself a source of noise. This paper is the result of a study of available literature on noise in biopotential recording. It focuses on the noise that originates at the electrode-skin interface. The silver-silverchloride electrode is commonly used in low-noise recording. It has a stable halfcell potential and is nonpolarizable. In a nonpolarizable electrode unhindered charge transfer is possible, making it suitable for stimulation, in a perfectly polarized electrode no charge exchange is possible. When the electrode is applied to the skin the noise level measured is higher than the level that is to be expected from thermal noise only, generated by the resistive impedance of the electrodes and intermediate tissues. The origin of the excess noise is still unclear. Localized out-of-equilibrium processes at the upper skin layers are suggested, as well as a residual EMG signal from nearby muscles. Abrading the skin lowers the excess noise. The noise level is also dependent on the type of electrode used. This suggests that processes at the electrode-electrolyte interface add to the total noise level. Examples of other noise sources are amplifier noise, capacitive and inductive interference and motion artifacts. Low-noise amplifiers are available, and measures can be taken to reduce the noise level of the latter sources to negligible value. This causes the voltages generated at the electrode-skin interface to become the dominating noise source. A capacitive electrode is a conductor coated with an insulating layer. When applied to the skin, the electrode is capacitively coupled to the skin, and its potential is therefore representative for the potential of the skin. The insulating layer has to be very thin, and is therefore vulnerable. Only few examples are mentioned in literature. Little data is available of the actual performance. In a truly capacitive electrode there is no conduction between metal and electrolyte. No halfcell potential will develop. The capacitive component of the electrode is noise-free. Therefore the capacitive electrode can be used for the study of electrode noise processes, as the interaction between electrode and electrolyte has been eliminated. The data obtained can be compared with galvanical conducting electrodes.

Titel: Noise in biopotential recording using surface electrodes Auteur: E. Huigen Datum: november 2000 Rapport nummer: S – 915 Aantal pagina’s: 37 Instelling: Universiteit van Amsterdam Sectie Medische Fysica Begeleiders: Dr. A. Peper Prof. Dr. Ir. C.A. Grimbergen Samenvatting: Elektrische verschijnselen in het menselijk lichaam veroorzaken potentiaalverschillen aan het huidoppervlak. Deze kunnen gemeten worden met behulp van oppervlakte-electroden. Een electrode is een elektrochemische transducer, die de potentiaalverschillen ten gevolge van ionenstromen in het lichaam omzet in een meetbaar voltage. Bij het meten van bioelektrische verschijnselen spelen diverse ruisbronnen een rol. Door de reacties aan het elektrodeoppervlak is de electrode zelf ook een bron van ruis. Dit verslag is het resultaat van een literatuurstudie over ruis tijdens het meten van bioelektrische verschijnselen. De nadruk ligt op de ruis die onstaat ter plaatse van de electroden. Een veelgebruikte electrode is de zilver-zilverchloride electrode. De halfcelpotentiaal hiervan is stabiel en de electrode is niet-polariseerbaar. In een niet-polariseerbare electrode is ongehinderde uitwisseling van lading mogelijk. Een gepolariseerde electrode laat geen uitwisseling van lading toe. Wanneer een electrode wordt aangebracht op de huid, is de ruis hiervan hoger dan de thermische ruis, samenhangend met de weerstand van electroden en tussenliggend weefsel. De mechanismen die ten grondslag liggen aan deze ruis zijn nog onbekend. Processen in de bovenste huidlagen en EMG signalen worden in de literatuur als mogelijke ruisbronnen gezien. Schuren van de huid verlaagt de ruis. Het ruisniveau is ook afhankelijk van het type electrode. Processen in de electrode-electrolyt overgang zijn hierdoor mogelijk ook een bron van ruis. Voorbeelden van andere ruisbronnen zijn versterkerruis, capacitieve en inductieve interferentie en bewegingsartefacten. Er zijn technieken om deze ruis uit deze bronnen te onderdrukken. Hierdoor wordt de ruis die onstaat ter plaatse van de electrode de grootste in de meetlus. Een capacitieve electrode is een geleider waarvan het actieve oppervlak is bedekt met een isolerende laag. Dit type electrode is capacitief gekoppeld met de huid. Het voltage, gemeten tussen deze electroden representeert het potentiaalverschil aan het huidoppervlak. De isolerende laag moet erg dun zijn, en is hierdoor kwetsbaar. In de literatuur zijn slecht enkele voorbeelden beschikbaar. In een capacitieve electrode is er geen galvanische geleiding tussen metaal en elektrolyt. Hierdoor ontstaat er geen halfcelpotentiaal. De capacitieve component van de electrode is ruisvrij. Omdat de interactie tussen metaal en pasta is uitgeschakeld, is deze electrode interessant voor het bestuderen van de processen die verantwoordelijk zijn voor ruis ter plaatse van de electrode.

Page i Noise in biopotential recording using surface electrodes

Assignment (in Dutch)

Opdracht literatuuronderzoek. Er zal een onderzoek uitgevoerd moeten worden naar publicaties over galvanisch geleidende en capacitieve electroden. De literatuuronderzoeken van voorgaande stages zijn erg gedateerd (K. Brinkman 1993, T. Dankers 1996). Het is goed mogelijk dat er op dit gebied in de afgelopen jaren nieuw onderzoek gepubliceerd is. Verder is het goed na te gaan of er op het gebied van de capacitieve electroden onderzoek gedaan is dat niet bekend is. Door C. Metting van Rijn zijn proeven gedaan met capacitieve elektroden, maar gegevens hierover ontbreken. Dr. A. Peper

Page ii Noise in biopotential recording using surface electrodes

Page iii Noise in biopotential recording using surface electrodes

Contents

Assignment (in Dutch) ................................................................................................................ i Contents..................................................................................................................................... iii Introduction ................................................................................................................................ 1 1 The electrode-skin interface .................................................................................................... 3

1.1 Electrodes: chemical transducers ..................................................................................... 3 1.2 Electrode impedance ........................................................................................................ 7

2 Noise in biopotential recording ............................................................................................. 13 2.1 Terminology ................................................................................................................... 13 2.2 Electrode-electrolyte noise ............................................................................................. 14 2.3 Noise at the electrolyte-skin interface ............................................................................ 14 2.4 Motion artifact ................................................................................................................ 16 2.5 Electric and magnetic field interference......................................................................... 17 2.6 Thermal noise ................................................................................................................. 18 2.7 Amplifier noise............................................................................................................... 18 2.8 Noise from additional bioelectric events........................................................................ 19 2.9 Other noise sources ........................................................................................................ 19 2.10 Overview noise sources................................................................................................ 20

3 The performance of electrodes .............................................................................................. 21 3.1 Performance of common electrode materials ................................................................. 21 3.2 A new development: the NASICON electrode .............................................................. 23

4 Capacitive electrodes............................................................................................................. 25 4.1 Some examples of capacitive electrodes ........................................................................ 25 4.2 Design criteria for capacitive electrodes ........................................................................ 27 4.3 'Wet' capacitive electrodes ............................................................................................. 28

Conclusions .............................................................................................................................. 29 Literature .................................................................................................................................. 30

Page iv Noise in biopotential recording using surface electrodes


Introduction

In this survey of literature noise during the recording of biopotentials is investigated. Various physiological processes produce electrical currents in the form of ion flows, which cause potential differences in the volume conductor of the body. The potential differences on the skin are a source of information about the bioelectrical processes that occur inside the human body. Surface electrodes convert these signals into a measurable voltage. This makes the surface potential differences available for further signal processing such as amplification and visual presentation. With the equipment currently in use, one is able to obtain a relatively noise-free recording of relatively large biopotentials such as the QRS-complex of the ECG. There are however processes in the human body that create surface potential differences that have a magnitude comparable to, or are even smaller than the noise levels commonly met in biopotential recording. Examples are EEG signals and potentials from the cardiac muscle, with a fraction of the amplitude of the QRS-complex. Examples of the latter are His bundle potentials and late potentials. The noise can, amongst others, originate from the patient, in the form of additional biopotentials, from the electrode leads and the amplifier, and from the signal transducer, the electrode. Even in laboratory conditions with high quality equipment relatively large noise potentials are recorded. There are methods available to extract signals from relatively noisy data. When a signal is periodic, averaging over a few periods will reduce noise, assuming the noise originates from random processes. This method can also be applied to bioelectric signals which are repetitive or which can be evoked. An error is made with this method because the biopotential is assumed to be stationary, while in fact it may vary in time. In this method lower noise levels means that fewer occurrences of the biopotential are necessary for averaging. This is not only time saving, but it also means a more accurate measurement of the phenomenon under study. To make a reduction of noise possible, a good understanding of the various noise sources relevant in biopotential recording is necessary. This literature study investigates these sources, and it particularly focuses at the noise processes that occur at the electrode-skin interface. Contrary to the other noise sources little is known about the source of the potential fluctuations that occur at this interface. Also literature on capacitive electrodes is studied. Capacitive electrodes are electrodes of which the active surface is coated with an insulating layer. Because the method of transducing the biopotential is different from that in 'normal' electrodes, data obtained with capacitive electrodes might give a better insight in the role the electrode plays in noise processes at the electrode-skin interface. Chapter one explains how an electrode transduces the potential differences on the skin into a measurable voltage. Physical properties of galvanic conducting electrodes are discussed. Chapter two lists the sources of noise that are likely to distort a biopotential recording. An indication of their relative importance is given. Chapter three compares the performance of various electrode types. The chapter focuses on noise data, for electrodes on which this information is available. In chapter four the capacitive electrode is discussed.


1 The electrode-skin interface

In medical science the recording of potential differences from the human body is an important source of information. Surface electrodes can be used for non-invasively measuring bioelectric activities of the heart, brain or muscles (ECG, EEG or EMG). In the human body the current flow is established through ion movement. The electrode has to convert this ion flow into a flow of electrons: the electrode is an electrochemical transducer. At the electrode-electrolyte interface ions and electrons are arranged in a specific way. This is called the electrical double layer. This arrangement of charge gives the electrode unique electrical properties. This chapter describes how surface electrodes work and models are given for the electrode-electrolyte-skin interface.

1.1 Electrodes: chemical transducers

The transfer of electrical current in living tissue is effected by the movement of ions through an electrolyte. Surface electrodes are used to measure the potential differences caused by these ion migrations. Usually the metal electrode is in direct contact with an electrolyte in the form of an electrode paste or gel. When a metal is brought in contact with an electrolyte, ions of that metal have the tendency to enter the solution, releasing one or more electrons. Ions in the solution tend to combine with the metallic surface (Geddes and Baker, 1989). As a result of these reactions a specific spatial arrangement of charge occurs. The simplest model for this was described in 1879 by Helmholtz, who stated that a layer of ions with one sign of charge is bound to the metallic surface, and a layer with the opposite sign is in the electrolyte. This is known as the electrical double layer. The charge distribution at the metal surface is given in figure 1-1. The double layer acts as a capacitor. Because of the charge arrangement at the double layer, a voltage is developed. This is called the halfcell potential. Because only the potential difference between two electrodes can be measured, a standard reference electrode has been chosen: the hydrogen electrode. The halfcell potential of a metal is defined as the voltage of that metal with respect to the hydrogen electrode. The halfcell potential of an electrode is caused by differences in ion concentration at the electrode surface and in the solution. The potential difference between two regions in a liquid with different ion concentration can be calculated with the Nernst equation. This equation can therefore be used to calculate electrode halfcell potential (Dankers, 1996):

−=

2

1lnC

C

nF

RTE (1)

where E = electrode potential F = Faraday constant T = absolute temperature n = valence of ion involved C1 = concentration of ions in the solution C2 = concentration of ions at the electrode surface


Fig. 1-1. The charge distribution at the electrode-electrolyte interface (Van Wijk van

Brievingh, 1988).

The Nernst equation can be expressed in terms of ion activities instead of ion concentrations:

γα C= (2)

where γ is the activity coefficient for a given ion. In dilute solutions γ ≈1, so α ≈ C. Potentials for commonly used electrode materials are given in Geddes and Baker (1989). Halfcell potentials can be large compared to physiological signals, so they can only be neglected when they are stable.

The silver-silver chloride electrode

In bioelectrical measurement the Ag-AgCl surface electrode is popular, because of its highly stable halfcell potential. The active surface of these electrodes consists of a combination of silver and silverchloride. At the electrode surface a reversible electrochemical reaction occurs when brought in contact with an electrolyte. When the electrode acts as the anode chloride ions are deposited on the electrode. When it acts as the cathode solid AgCl is reduced to Ag and Cl- enters the solution (Ferris, 1972). Two reversible reactions determine the electrode potential:

AgClClAg

AgAg

←

→

←

→

−+

−+

+

+ e

The Nernst equation for the first reaction is given by

++ +=Ag

0

AgAg,lnα

nF

RTEE (3)

with n = 1 for monovalent ions, +Agα is the ion activity of the silver ion, see equation 2. The

product of the ionic activities of silver and chloride ions is constant and equal to the solubility product:

sK=+ -ClAgαα (4)

with -Clα the ionic activity of the chloride ion.

Equations (3) and (4) give the Nernst equation for the Ag-AgCl electrode:


-Cl

0

AgAg,lnln α

nF

RTK

nF

RTEE s −+= + (5)

The first two terms in this equation are constant, so the electrode potential is only dependent on the activity, or concentration of the chloride ion. The Nernst equation for the Ag-AgCl electrode underlines the importance of chloridizing silver electrodes. When a pure silver electrode is placed in a solution that does not contain a significant concentration of the silver ions, there are not enough carriers in the solution to form a stable double layer. The constant terms in the Nernst equation (5) therefore are not well defined, and the halfcell potential will be unstable (Smith and Wace, 1995). The layer of AgCl on chloridized silver electrodes provides a saturated solution of silver and chloride ions at the electrode surface. In a saturated solution the solubility product is constant for a given temperature, so the electrode potential is well defined and stable. Bioelectric events cause only small and alternating currents, so the reactions at the electrode surface are continuously reversed. Therefore depletion of the silver-silverchloride layer does not occur in practice. Silver-silverchloride electrodes have also proven useful for passing larger direct currents (when used for stimulation) for shorter periods of time (Smith and Wace, 1995). When using silver-silverchloride electrodes one should be conscious of the fact that they are photosensitive (Geddes and Baker, 1989). They should either be used in subdued light or should be shaded from light. Light decomposes silver chloride so electrodes should be stored in a dark place.

Polarization

In an ideal situation the electrodes are connected to an amplifier which does not ‘load’ the transducer. In other words: no current is drawn from the electrode. Modern MOSFET amplifiers have high input impedance and therefore need negligible input current (<10-12 A). These amplifiers do not load the electrode system. A problem is encountered when large currents are passed through an electrode. This might be the case in an attempt of defibrillation, or when the measuring electrodes are also used for stimulation of the patient. When current passes through an electrode the double layer will be disturbed. Charge is built up which changes the electrode’s halfcell potential and may hinder the exchange of charge between electrolyte and electrode. This process is called polarization. Electrodes in which the flow of current through the electrode-electrolyte interface is not obstructed by this charge built-up are called ‘perfectly nonpolarizable’. Those in which no exchange of charge is possible are called ‘perfectly polarized’ (Geddes and Baker, 1989). By this definition a truly polarized electrode acts as a capacitor. A real nonpolarizable electrode does not exist. The properties of electrodes lie somewhere between perfectly nonpolarizable and perfectly polarizable. However, the silver-silverchloride electrode has proven to be extremely insensitive to polarization, and is referred to in most literature as nonpolarizable. It is clear that polarization of a galvanic conducting electrode is not desired. To describe the processes that can cause this unwanted effect, a type of polarization is explained: concentration polarization1 (Smith and Wace, 1995). Consider a pair of metal electrodes placed in an electrolyte that contains no ions of that metal. This is the case when for instance platinum or stainless steel electrodes are used in combination with organic liquid. When a potential difference is created across these

1 Another kind, ohmic polarization is described by the reaction that takes place when electrodes of a certain metal are placed in a solution that contains their own ions: M↔Mz++ze-. The polarization potential across these electrodes obeys Ohm’s law, hence the name. Because no carriers are depleted, it does not change the halfcell potential and it does not distort the recorded biopotentials. For this reason ohmic polarization is usually disregarded. Electrodes that only exhibit ohmic polarization are usually referred to as nonpolarizable.


Fig. 1-2. Voltage (V1-3) between a 0.1 cm2 stainless-steel test electrode and potential-

sensing electrode in response to a 1000 ms constant-current pulse of 1 mA

(Mayer, et al., 1992).

electrodes, a current will start to flow. The positively charged sodium ions in the organic liquid will be attracted to the cathode. However, as these ions can not interact with the electrode’s metal, and deposit as solid sodium, they will simply accumulate at the surface. This buildup of charged particles will form a barrier that acts as a dielectric (the insulator in a capacitor), and the concentration of ions of one sign changes the halfcell potential of the electrode. As time elapses, the charge barrier causes a growing resistance to current flow. The resistance of the electrolyte is also increased, as the carriers are depleted from the solution. When no chemical reaction takes place at the electrode surface, changing the polarity can reverse the process of polarization. Figure 1-2 illustrates how a negative polarity, rectangular current pulse deforms the output of a polarizable electrode. The current of 1 mA through the 0.1 cm2 electrode is very high. However, the same distortion of the step input is observed at lower current densities. Because the charge barrier behaves as a capacitor, it will affect mostly the low-frequency content of signals. Most bioelectric events are in this frequency range. With this knowledge it is possible to list some measures that can be taken to prevent polarization. First, nonpolarizable electrodes, such as the Ag-AgCl electrode can be used. The electrolyte has to contain sodium ions. Second, the current density can be kept low by using electrodes with a large area. A current density less than 0.1 mA/cm2 should be used to ensure linear current transfer over a broad range of frequencies, using stainless steel electrodes (Geddes and Baker, 1989, Godin, et al., 1991). Such large current densities are not encountered in normal recording situations. The term ‘nonpolarizable’ is used for silver-silverchloride electrodes. Measurements done by Ferris and Steward (1974) on Ag-AgCl microelectrodes, however, show distortions of recorded potentials due to polarization, some much like figure 1-2. Current densities were in the range 0.01-0.1 mA/cm2. These high current densities are not unlikely when recording with small-area electrodes such as needle electrodes.

Electrolytes

In order to ensure contact between the skin and the electrode an electrode gel or electrolyte is usually applied to the skin. The skin has a relatively rough surface. Application of a gel prevents reduction of the effective electrode area caused by skin inhomogeneity, or by hair under the electrode plate. Electrode pastes are available separately, or in combination with electrodes, so-called pre-gelled electrodes. The skin has a large impedance to current flow, mainly caused by the upper layer (stratum corneum) of the epidermis, figure 1-3. It has long been known that mild abrasion of the skin with fine sandpaper can reduce the skin resistance with a factor 100 to 1000 (Geddes and Baker, 1989, Van Wijk van Brievingh, 1988). Some electrode gels contain abrasives, for better penetration of the upper layer of the skin. A condition is that the gel must not cause too


Fig. 1-3. Schematic diagram of the skin (Eggins, 1993).

much skin irritation. Chloride concentration mainly determines the conductivity and the aggressiveness of the gel (Hollander, 1982). Lewes (1965, as described in Geddes and Baker, 1989) argues that since there are now amplifiers with very high input impedance, the need for low-resistance skin-electrode contact has disappeared. Experiments done by Lewes show that a wide variety of poorly conducting substances can be used as an electrolyte. ECG’s taken with for instance toothpaste, mayonnaise, mustard, soft soap or tomato paste are indistinguishable from those taken with regular electrode paste, provided the amplifier is of high quality. The validity of Lewes’s work can be questioned when the performance of the electrode is expressed in terms of noise. Thermal noise (see paragraph 2.6) is proportional to electrode impedance. Therefore better electrolyte conductivity means less noise. This may not be visible in the recordings Lewes made of the relatively large potential differences originating from the heart. When smaller biopotentials are recorded like the EEG, noise becomes more important. The composition of the electrolyte plays a role in the formation of junction potentials. Normally the electrolyte and the organic liquids will have a dissimilar ionic composition. A junction potential develops as a result of the diffusion of these fluids. Different ions have different diffusion rates, some will move faster than others, creating a potential gradient. For instance the sodium ion is hydrated by water molecules, and therefore is much bulkier and thus slower than the unhydrated chloride ions. Potassium ions have mobility comparable to chloride ions. This is why potassium chloride is used in preference to sodium chloride in better quality electrode pastes (Smith and Wace, 1995). The mobilities of Na+, K+ and Cl- are 5.2x10-8, 7.6x10-8 and 7.9x10-8 m2s-1V-1 respectively (McCarron, 1997).

1.2 Electrode impedance

In normal biopotential recording situations, metal electrodes are connected to the skin through an electrolyte or electrode paste. In this connection two discontinuities can be observed: the electrode-electrolyte interface and the electrolyte-skin interface. When the electrode impedance is investigated, these two interfaces have to be observed separately. The electrode-electrolyte interface can be investigated by connecting two electrodes directly with an electrode paste. This interface will first be discussed. A model and experimental values for skin impedance are given next.


Electrodeterminal Electrolyte

R C


R C

Rf


R C

Rf

E

A

B

C

Fig. 1-4. (A) The Warburg circuit model for the electrode-electrolyte interface. (B)

Faradic resistance Rf in parallel with the Warburg components to account for

D.C. behavior. (C) The complete circuit model, including halfcell potential E.

Impedance of the electrode-electrolyte interface

One of the first circuit models for the electrical behavior of the double layer was derived by Warburg (1899, as described in Ragheb and Geddes, 1991). He represented the interface as a series resistance R and capacitance C, figure 1-4a. The polarization impedance for this representation is:

fC

jRZ p π2

+= (6)

Warburg found that R has the property that it is frequency dependent. This invalidates the use of the term resistor for the element R, the model is incorrect. However, most of the literature on electrode impedance uses the assumption of a frequency dependent resistor as a good approximation. Warburg stated that resistance and capacitive reactance, Xc=1/2πfC, are equal and both are proportional to f –0.5. The basis of this model was the assumption of a constant phase angle φ, given by tan(φ)=Xc/R=1 or φ=π/4 radians. In many situations this is an inaccurate model. Express the resistance and the reactance as:

α

f

AR = (7)

β

f

BX c = (8)

In 1932 Fricke proposed a constant phase model in which φ=0.5πβ. The assumption of a constant phase provides that α=β. The validity of this model depends on the electrode metal and frequency range. Ferris (1972) gives experimental values for platinum electrodes. Here α=0.5 and β ranges between 0.5 and 0.7, so the phase angle will vary with frequency. Schwan (1992) found that the constant phase model is valid for frequencies above 1 Hz. The Warburg model of figure 1-4a does not account for low frequency behavior of the electrode-electrolyte interface, as electrodes are able to transfer direct current. To model this behavior a Faradic (direct current) resistance Rf is placed in parallel with the Warburg series components, figure 1-4b (Geddes, 1997). Indeed, low-frequency measurements done by Schwan (1992) show a semi-circle in the impedance plane, implying the reactive component of the impedance approaches zero when the frequency tends towards zero (figure 1-5). The


Fig. 1-5. Impedance locus at low frequencies. Platinum electrode, area 0.0855 cm2, 0.9%

NaCl (Schwan, 1992).

Rel

Rc

Cc

Rd

CdDiffuselayer

Fig. 1-6. A circuit model for the electrode-electrolyte interface with separate electrical

components for the behavior of the compact layer (Rc and Cc) and the diffuse

layer (Rd and Cd) (van Wijk van Brievingh, 1988).

circuit model is completed by adding a halfcell potential E (figure 1-4c). The Warburg model, and models derived from it are based on available electrode impedance data, not on physical insight in the actual charge arrangement at the electrode surface. Van Wijk van Brievingh (1988) describes an accurate circuit model that accounts for the electrical properties of the double layer and the diffuse layer, separately (figure 1-6). Both are represented by a parallel resistance and capacitance, instead of a series R and C, to account for the passage of direct current. Rel is the resistance of the electrolyte. In practice the resistive and capacitive components in the circuit models described in this chapter are dependent on electrode material, area, current density, frequency, electrolyte and temperature.

Impedance of electrodes on the skin

In skin impedance measurements, the electrode impedance is always included. However, when using low-impedance electrodes, the skin impedance is dominating. The layer of dead cells in the epidermis, the stratum corneum causes the largest impedance to current flow. The total impedance can be up to 5 MΩ. Eggins (1993) gives a circuit model for the electrode-skin interface, figure 1-7. In his work he argues that the performance standards drawn up by the Association for the Advancement of Medical Instrumentation (AAMI) for pre-gelled electrodes, used by manufacturers and distributors, are largely irrelevant. The AAMI tests electrode pairs connected face to face (gel-to-gel), at only one frequency, 10 Hz. According to Eggins a broader frequency spectrum must be used, and more important, as the impedance of the epidermis is much higher than the other impedances in figure 1-7, the electrodes have to be tested on the skin. Gel-to-gel impedance measurements do not correlate well with measurements on unprepared human skin, however a strong correlation has been found between gel-to-gel measurements and those on abraded skin, with the stratum corneum and the barrier layer removed.


Fig. 1-7. Simple equivalent circuit model of the electrode-skin impedance (Eggins, 1993).

The impedance of the epidermis dominates the others, and has the form:

( )αωTj

RZ

p

skin+

=1

(9)

The element Cp therefore is not purely capacitive, but has an impedance

( ) αω −= jKZpC (10)

with

α

T

RK

p= (11)

In Eggins work α ranges between 0.8 and 1.0, and remains constant in time after application on the skin. The resistive and reactive components of the impedance can differ strongly when electrodes are applied to different persons. Rosell, et al. (1988) shows a two-decade spread (from 10 kΩ to 1 MΩ) over various 1 Hz impedance measurements. As frequency increases, the impedance decreases, and so does the variance. Figure 1-8 shows impedance measurements from various sources. All plots have a semi-circular form, only the magnitudes differ. The differences in impedance between different electrodes and the variations in time are caused by various factors: • Electrode area. The impedance is inversely proportional to electrode area, or more

specific: the area covered by electrode gel (Brinkman, 1993). • Time of application. As time increases, the semi-circular arcs in the impedance plane

become smaller (Eggins, 1993). Variation over time can be explained by initial contraction of pores due to the cold gel, penetration of electrolyte through the skin, spreading of electrolyte, diffusion with sweat, skin recovery after abrasion. Variations can occur in time scales as short as a second.

• Condition of the skin. Abrasion lowers impedance. Skin impedance differs between different subjects, and between different places on the body.

• Electrolyte composition. For instance, curve A in figure 1-8 is recorded with a ‘dry’ electrode, i.e. no gel is used. Body fluids serve as an electrode gel.

When all these factors are taken into account it is extremely difficult to predict the impedance. This emphasizes the importance of high input impedance amplifiers, to prevent the recorded biopotentials being altered by impedance variations. In chapter 2 it will be explained that lowering the impedance results in less noise. This makes skin abrasion and the use of an electrode paste necessary in low-noise biopotential recording.


Fig. 1-8. Electrode impedance measurements in the impedance plane. (A) NASICON dry

electrode, 10 minutes after application, electrode area 0,90 cm2 (Gondran, et al.,

1995). (B) Marquette wet electrode, 6 minutes after application (Eggins, 1993).

(C) Hydrogel electrode HP 13951C, 4.5 minutes after application (Eggins,

1993). (D) Ag-AgCl electrode (Searle and Kirkup, 1999).


2 Noise in biopotential recording

As surface biopotential recording involves the measurement of extremely small potential differences, noise is likely to play an important role. This chapter gives an overview of noise sources that may be encountered during these measurements. The ECG has a typical amplitude of 1 mV. One per cent of this value is generally accepted as the upper level of noise voltage in ECG recording (Metting van Rijn, et al., 1990). In the preceding chapter the complex reactions that occur at the electrode surface were described. These reactions cause a certain arrangement of charged particles at the metal surface. At the same time reactions take place at the electrolyte-skin interface. These processes give the interface unique physical properties. It implies that the interface itself is an unstable voltage source, which is added to the recorded biopotential. In the case of silver-silverchloride electrodes the halfcell potential has proven to be stable compared to other electrode materials. However, silver-silverchloride electrodes still exhibit a significant amount of noise. In this chapter the noise at the electrode-skin interface is described, and an inventory is given of other noise sources which may affect a biopotential recording. When available, methods for the reduction or suppression of noise are portrayed. At the end of this chapter a quantitative comparison of the various noise sources is made. The terminology used in this paper is discussed in the first section.

2.1 Terminology

The output voltage of an amplifier used to measure bioelectric events is not always an accurate representation of the event under examination. It is composed of the desired voltage, and a number of unwanted voltages. The desired voltage is called the signal. All unwanted voltages are referred to as noise. When the term noise is used, it does not mean that the origin of the fluctuating voltage is unknown. For example, in some cases the ECG is the signal under examination. In other cases the ECG is classified as noise, when activities of muscles, other than the cardiac, are measured. The ECG noise can be easily distinguished from the signal by its repetitive nature. The recorded noise may originate from various sources. For this reason noise is divided in some categories, depending on the nature of the noise. The terminology is as follows: In this paper the term noise is applied to random fluctuations of voltage with a more or less Gaussian distribution of instantaneous amplitudes. An example is thermal noise. The American Heritage Dictionary (1996) describes the term artifact as “an object produced or shaped by human craft (…) a typical product or result (…) a structure or substance not normally present but produced by an external agent or action”. In this paper artifact is used for fluctuations of voltage with a non-recurring nature. For instance, movement of the electrode causes a single peak in the time domain. Interference will be used to describe unwanted voltages that are continuous in the time domain and have a more deterministic repetitive nature. An example is 50 Hz interference caused by capacitive coupling of the measurement setup with power lines. This interference will cause a narrow band peak in the frequency domain. Another example is the ECG, when interest goes out to other signals. Of course there is some overlap between these terms. For instance a movement artifact with a recurring nature may be encountered when an electrode is positioned on a pulsating vein. Note that another term is used in this paper, which does not classify as a synonym for noise. Distortion is used to describe the alteration in waveform of a signal. This is the case when the signal is sent through some kind of filter, for instance an RC-network, or skin acting like one.


In addition to this linear distortion non-linear distortion can occur. In this case the original waveform might be difficult to recover.

2.2 Electrode-electrolyte noise

The stability of the halfcell potential of an electrode depends on the used material and electrode preparation. The material that is mostly used for low-noise recording of biopotentials is silver-silverchloride, as its halfcell potential is relatively stable. The noise introduced by the electrode-electrolyte interface can be recorded by connecting electrodes gel-to-gel. This means the electrode metals are directly connected through a layer of electrolyte. Geddes (1972) measured fluctuations in halfcell potential of more than 200 µV peak-to-peak on pure silver electrodes, and noise levels up to 10 µV peak-to-peak after chloriding. The electrodes were immersed in a 0.9% NaCl solution. More recent measurements done by Gondran, et al. (1996) show peak-to-peak voltages as low as 0.25 µV for 3M Red Dot electrodes. This extremely low noise voltage can not be verified. No frequency range on which the peak-to-peak voltages were determined is given in Gondran’s article. The noise level of the electrode-electrolyte interface varies proportional to f -n, according to different publications 0.5<n<1.5. Dankers (1996) took the frequency at which the electrode noise level dropped below the amplifier noise as a measure of performance of different electrodes. This is called the corner frequency. The corner frequencies Dankers found ranged between 3 Hz for Brain Wave Ag-AgCl electrodes gel-to-gel and 31 Hz for gold electrodes. The composition of the electrode paste also has influence on the point of intersection. However, no clear relationship has been established. The literature is not unequivocal about the origin of the spontaneous fluctuations in halfcell potential. The noise level is higher than thermal noise (section 2.6). Gondran, et al. (1996) state that, as the noise of a redox system in equilibrium is equal to thermal noise, the electrodes possibly cannot reach a state of equilibrium under their operating conditions. Brinkman (1993) has found two possible explanations for the fluctuations in literature. The first states that the resistance of the fluid film is unstable, under constant current this creates potential fluctuations. The second describes random microscopic ionic discharges at the electrode surface due to irregularities and impurities of the electrode material. Methods for stabilizing silver-silverchloride electrodes are applying the electrolyte some time before usage (pre-gelled electrodes) (Dankers, 1996) or storage of the electrodes in a saline solution and connecting a short circuit between them (Ferris, 1972). The noise level of gel-to-gel electrodes was found to be inversely proportional to electrode diameter (Fernández and Pallás-Areny, 2000).

2.3 Noise at the electrolyte-skin interface

When measuring the noise spectrum of the skin with surface electrodes the noise level is higher than the thermal noise associated with the resistance of the skin and electrodes. The source of the excess noise is uncertain. Measurement of only the noise signal produced by the skin, is a difficult task, as EMG signals and other noise sources will always contaminate the recording. Noise measurements are frequently carried out on the inner forearm, the muscles in this area are easily controlled. A typical example of a noise spectrum obtained on the inner forearm is given in figure 2-1A. The figure is based on the numerical results found by Brinkman (1993). At low frequencies (<10 Hz), 1/f noise is observed, followed by a horizontal line (approximately 10-100 Hz). At high frequencies (>100 Hz), the spectrum decreases again because of the limited bandwidth of the amplifier. The spectrum for the high frequencies is in correspondence with the thermal noise generated by the impedance of the skin and the electrodes. For low frequencies the


Fig 2-1. Noise spectra of electrodes on the inner forearm. (A) Normal spectrum. (B)

Abraded skin. (C) Muscles contracted. (D) Muscles contracted and skin

abraded.

noise spectrum can be 100 times higher than is to be expected from thermal noise only (Brinkman, 1993). The values in the graph may differ per measurement. For example Gondran, et al. (1996) found spectra in agreement with thermal noise for frequencies over 10 Hz. The peak-to-peak values were in the order of 10-60 µV. Skin abrasion lowers the low-frequency noise level (figure 2-1B). A logical conclusion is that the upper skin layers are responsible for low-frequency noise. Frequencies over 10 Hz are not affected by skin abrasion. The residual noise signal cannot fully be associated with thermal noise. It is assumed that it originates from a residual EMG signal (Godin, et al., 1991, Gondran, et al., 1996, Brinkman, 1993). Contracting the muscles in the forearm (for instance, bending of the fingers) causes a substantial rise in the spectrum over 10 Hz (figure 2-1C). Contraction of the muscles when measuring on abraded skin shows the EMG also affects frequencies below 10 Hz (figure 2-1D). Brinkman (1993) suggests from these observations that excess noise measured in relaxed state originates from a residual EMG signal, to which a low-frequency 1/f noise is added by the upper skin layers. It can be questioned however whether the frequency content of the noise spectrum of a relaxed muscle has a strong correlation with the frequency content of a tense muscle. Additional research is necessary. Gondran, et al. (1996) examined the noise spectra of measurements done with different electrode diameters. In general, increasing the diameter reduces the excess noise. They found that the reduction in excess noise with increasing electrode diameter has a limiting value. This might be explained by assuming that part of the noise originates at some distance from the electrode. A relatively distant signal originating from muscles is not likely to be affected by electrode diameter. According to his measurements this residual EMG signal has a peak-to-peak value of 5 µV. Godin, et al. (1991) found a smaller slope in the diameter-RMS noise voltage graph than was to be expected from their calculations, and stated also that this was due to incomplete relaxation. Gondran, et al. (1996) give two possible explanations for the low-frequency 1/f noise generated on unabraded skin. First, it could be due to an out-of-equilibrium process, for instance the slow drift of electrode voltage. Secondly, it could be caused by the ionic nature of the skin, for instance fluctuations in ion concentration or mobility. This could explain the decrease in excess noise with increasing electrode diameter. Brinkman (1993) examined changes of the noise spectrum in time. He found that intact skin tends to become noisier after some time, but abraded skin becomes less noisy. This phenomenon remains unexplained, as little is known about the interaction between skin and electrolyte.


Gel soaked sponge

Silver

Silver chloride layer

External snap (stainless steel)Plastic cup

Adhesive

Fig. 2-2. Cross-section of an electrode in which the active surface in recessed in a plastic

cup to reduce motion artifact due to deformation of the electrode-electrolyte

interface.

Dankers (1996) found that the noise produced by the skin is not independent of the electrode metal used. He found different noise spectra for different metals, using the same electrolyte. This was not to be expected, as the noise level produced by electrodes placed gel-to-gel is much lower than noise produced by the skin and residual EMG. No explanation is given.

2.4 Motion artifact

Movement can cause changes in the potentials that are created when an electrode is applied to the skin. Normally, when the patient is relaxed, and high quality electrodes are used, the recording is not distorted by motion artifact. Brinkman (1993) has found no significant correlation between the intentional movement of the arm and the noise signal. The mechanisms that can cause motion artifact are described next.

Liquid junction potential variations by electrode movement

The various phase junctions in the electrode-electrolyte-skin interface all cause junction potentials, sensitive to motion artifact. The skin-electrolyte interface can cause artifacts of 400-600 µV when the electrode is moved parallel to the skin surface (Smith and Wace, 1995). When the electrode is moved perpendicular to the skin the potential changes can be up to 900 µV. Firm attachment to the skin can reduce the potential changes. The electrode-electrolyte interface also produces artifacts when mechanically disturbed. Gatzke (1974, as described in Webster, 1984) measured a 15 mV potential change when a pure silver electrode is moved in electrolyte. Coating with silver chloride, thus creating a stable double layer, produces a 10-fold reduction of the artifact. Further reduction (up to negligible value) can be achieved by recessing the electrode-electrolyte interface in a protective cup, in which a sponge soaked in gel is placed (figure 2-2).

Skin potential changes

Earlier, the stratum corneum and the barrier layer have been identified as the major sources of the impedance of the skin. Webster (1984) has also observed a potential difference between the inside and the outside of the barrier layer. Van Wijk van Brievingh (1988) however states that this skin potential is a liquid junction potential between deeper skin layers and the electrolyte. The skin potential has a typical value of 30 mV at the thorax. Stretching of the skin decreases the skin potential to about 25 mV. This artifact can be reduced to negligible value by abrading the skin. Webster gives 20 strokes with fine sandpaper as an indication. Shackel (1959, as described in Geddes and Baker, 1989) developed a method for shorting the skin potential, called the skin-drilling technique. The epidermis is eroded using a dental burr. The capillaries remain undisturbed, so no blood is drawn. Unfortunately, the epidermis is also the layer that protects the skin from irritation. A mild electrode gel has to be used to prevent unwanted effects.


Ödman (1981, as described by Smith and Wace, 1995) noticed that due to visco-elastic properties of the skin, the skin potential does not change linear with skin deformation. As the skin stretches the potential rises with a speed greater than 1 mV/s, causing significant artifact, however when the skin is released the 90-10% fall time is about 40 seconds. This causes very low frequency noise, which usually has no effect on the interpretation of the recorded ECG or EMG.

Cable movement

A source of artifact, only measured in moving cables (e.g. during exercise) is triboelectric noise (Webster, 1984). Cable insulation can act as a movement transducer when it is deformed. Special low-noise cable materials have been developed. Movement of the electrode cables inside a magnetic field may cause an inductive signal in the wires. An example in which cable movement causes artifact is in the strong magnetic field of a MRI-scanner. Inductive coupling to (alternating) magnetic fields is discussed in the next section. Amplifying the signal before it enters the cable can also reduce all forms of artifact produced by the wires. Various electrodes with amplifier circuits attached to the electrode are described in literature (Webster, 1984, Prutchi and Sagi-Dolev, 1993, Metting van Rijn, et al., 1996).

2.5 Electric and magnetic field interference

Capacitive coupling

The patient, the electrodes and the electrode cables are capacitively coupled to nearby electrical fields, for instance power lines. In theory, the voltages induced by this capacitive coupling will appear as common mode voltages over the electrode leads, assuming that both electrodes and electrode leads are equally affected by the interference. When a differential amplifier is used the common mode voltage is rejected. A common-mode rejection ratio of 80-90 dB is customary in clinically used differential amplifiers. Interference appears in the non-ideal situation in which the impedance of electrode or cable pairs differs. The common-mode voltage is then converted into a differential signal. This is known as the potential divider effect. A difference in cable impedance may for instance be caused by the usage of cables of different length. Difference in electrode impedance may lead to an interference potential of 200 µV peak-to-peak (Metting van Rijn, et al., 1990). Several methods are available for the reduction of interference caused by electrical fields. Physical separation is the most obvious method, but not always possible. Skin abrasion lowers the difference between electrode impedances, and also reduces the common-mode voltage. Cables are usually shielded. This lowers the interference level in the wires. Some amplifiers are equipped with a circuit that drives the shields to the signal of the inner wire. This technique is called guarding. It results in smaller cable capacitance (Metting van Rijn, et al., 1990). Because this is a form of feedback, there are limits to its stability. Another way of lowering the common-mode voltage is the use of a driven right leg circuit, figure 2-3. The common-mode signal is fed back to the patient through an extra amplifier. In this way the patient is virtually connected to the earth.

Inductive coupling

Because the measurement setup forms a closed loop, a magnetic field can cause an inductive current to flow in the loop. The simplest way to reduce interference due to inductive coupling is twisting of cable pairs (Bentley, 1983). The sign of the induced voltage is dependent on the orientation of the two cables. When twisting the cables, multiple loops are created, with opposing orientation. In every loop the induced voltage opposes the voltage in the preceding


Fig. 2-3. Driven right leg circuit (Webster, 1984).

loop, so they cancel each other out. As twisting of the cables is not always possible in complex ECG measurement setups, shielding of the magnetic source may be necessary (Metting van Rijn, et al., 1990).

2.6 Thermal noise

Every resistor will produce a certain amount of noise above 0° K. This is due to the fact that heat is the random movement of elementary particles. In theory the movements are in all directions, and should cancel each other, however some imbalance remains (Fitch, 1973). The random movement of charged particles causes thermal noise, and it has a root mean square value of

kTBRVth 4= (12)

where k is Boltzman’s constant (1.38.10-23 J/K) T is the temperature in Kelvin B is the bandwidth of the voltage measuring device in Hertz R is the resistance in Ohms

When a non-negligible capacitance parallel to the resistor is present, the thermal noise is filtered:

222

2

1

4

RC

kTBRVth ω+

= (13)

2.7 Amplifier noise

The noise introduced by the amplifier can be modeled as two uncorrelated noise sources, Vn and In in figure 2-4 (Fitch, 1973). Ze is the electrode-electrolyte impedance and Ve is a voltage noise source related to electrode processes. As the noise sources are independent, the total amount of noise from the amplifier is given by

( )222nena IZVV += (14)


Noiselessamplifier

In

Vn

Ve

Ze

Output

Fig. 2-4. Amplifier noise source model. Vn and In are uncorrelated voltage and current

noise sources, Ze and Ve are the impedance and the voltage noise source created

by the electrode and electrolyte.

The voltage noise of the amplifier can be measured by shorting the inputs. At low frequencies the voltage noise is 1/f noise (typically < 1 Hz), followed by a white noise characteristic for the rest of the bandwidth of the amplifier. In state of the art amplifiers the white noise spectrum can be as low as 10 nV/√Hz, for frequencies over approximately 1 Hz (Brinkman, 1993). The voltage created by the amplifier current noise is proportional to the impedance of the signal source. As the impedance of the skin can be high, current noise can become dominant. Godin, et al. (1991) found dominating current noise for loads over 30 kΩ. Brinkman (1993) states that a design criterion for amplifiers for bioelectric events could be that the current noise is smaller than the voltage noise for all impedances of signal sources likely to be connected to the amplifier. In this way amplifier noise is independent of electrode impedance.

2.8 Noise from additional bioelectric events

Unstable liquid junction potentials created by the skin have already been discussed, as well as EMG signals due to incomplete muscle relaxation. The EMG of active muscles can be a major disturbance in measuring biopotentials. Some examples are noise in the ECG due to muscle activity in the thorax while breathing, noise in the EEG originating from the EOG (electro-occulogram), and major EMG signals while recording the ECG during exercise. In order to prevent contamination of the recorded biopotential muscle activity must be avoided. When this is impossible, body surface areas with minimal EMG levels have to be sought. The highest level of EMG comes from the motor point of the muscle (Inbar and Noujaim, 1984). This is the point used for recording the EMG, but has to be avoided when other sources are monitored. It can be found by electrical stimulation of the muscle. The motor point is the location that gives maximum contraction. Webster (1984) describes electrode locations on the body surface with minimal contamination for ECG monitoring during exercise and transport in an ambulance. In both cases, of course, motion artifacts also play an important role. Respiratory signals can be recorded simultaneously with the ECG. Afterwards, the ECG can be 'filtered' by compensating for the respiratory signal.

2.9 Other noise sources

In the preceding, most of the noise sources, observed in clinical practice are described. However, noise may come from unexpected sources. Webster (1984) describes how static electricity originating from synthetic clothing causes noise. During low humidity the friction induced currents flowing through the patient to common were up to 10 µA. The noise voltage can be reduced by lowering the skin impedance. He also describes radio-frequency interference caused by electrosurgical units used to cut and coagulate tissue. Shielding can


cure this problem. Judicious building of hospitals can provide adequate distance between these sources of interference and the recording equipment. Brinkman (1996) discusses noise due to discretisation of the signal in the AD-converter, and noise due to turning on the high-pass filter of the preamplifier used in his research.

2.10 Overview noise sources

Figure 2-5 gives an overview of the discussed noise sources, with indications of the peak-to-peak voltages that can be created. Typical amplitudes of interest of ECG, EEG and EMG are drawn for comparison. Figure 2-5 is based on numerical values found in the various literature sources mentioned in this chapter. The 10 µV line (1 per cent of ECG) is the maximum peak-to-peak noise level acceptable in ECG recording, as stated in the beginning of this chapter (Metting van Rijn, 1990). The horizontal lines in figure 2-5 indicate upper and lower values of noise peak-to-peak voltages. With some effort (the noise reducing measures described in this chapter) all noise voltages can be reduced to their lower values. The figure shows that movement artifact and capacitive coupling with electric fields can cause intolerable noise levels in ECG recording, but when the right measures are taken the noise level can also be reduced to negligible value. The other four noise sources in the figure have nonzero lower values. For instance thermal noise will always be present, as the resistance of tissue material is nonzero. Skin-gel interfacial noise has its lower value at 10 µV. In low noise recording this interface is an important noise source. Literature does not provide a clear description of the processes that may be responsible for the voltage fluctuations. Further research is necessary. Amplitudes of interest in EEG recording range around 100 µV. One per cent of this value, 1 µV can be chosen as a maximum tolerable noise level in EEG recording. In figure 2-5 it can be seen that amplifier noise, thermal noise and electrode-electrolyte interfacial noise can all reach peak-to-peak levels above 1 µV. When EEG signals or other small biopotentials are recorded, noise will always be of importance.

Fig. 2-5. Amplitudes of interest in ECG, EEG and EMG recordings, and typical peak-to-

peak voltages of different noise sources.


3 The performance of electrodes

In studying literature, it is remarkable how little comparative research can be found of commonly used electrode materials, of different types of Ag-AgCl electrodes and of electrodes in combination with different electrolytes. Extensive research has been done in trying to model electrode processes. However, in many studies the interaction with the human skin is neglected. This gives little information on the situation when electrodes are placed on the skin, because it is the whole interface electrode-electrolyte-skin that determines the final performance of the electrode (Eggins, 1993). It can be assumed that electrode manufacturers are doing research but, when so, it stays behind closed doors. This chapter gives an overview of some performance characteristics that can be found in literature. When available, quantitative noise data is given, to give an indication of the minimal signal amplitude that can be transduced by the electrodes.

3.1 Performance of common electrode materials

Ag-AgCl electrodes

Silver-silverchloride has a superior quality over other electrode materials with respect to the stability of the half-cell potential. This is the reason that there are now numerous different types of Ag-AgCl electrodes available. Because of this extensive supply it is difficult to compare the performance of ‘the typical Ag-AgCl electrode’ with other electrode materials. Various subdivisions can be made: • Type of bioelectric event

- ECG - EEG - EMG - ……

• Application period - Short term electrodes - Long term electrodes

• Monitoring situation - ICU monitoring - Holter electrodes (used for 24-hour monitoring for heart rhythm during daily activities) - Stress test electrodes (used during exercise on a treadmill) - Electrodes for burn patients - Pediatric electrodes - ……

Each of these variables puts certain demands on the electrode: • Acceptable noise level • Offset potential and its drift • Recovery time after defibrillation (high currents will temporarily polarize the electrode) • Sensitivity to motion artifact • Ease of application (for instance pre-gelled or not) • Aggressiveness of the electrolyte in skin penetration • Time in which an electrolyte evaporates or a gel crystallizes • Smearing of electrolyte (smearing may short closely spaced electrodes) • Strength of the adhesive and its effect on the skin • Disposable or reusable


It is clear that there is overlap between the different fields of application. It is up to the manufacturer to define in which area an electrode can be used, which has resulted in various definitions for for instance 'long-term' or 'disposable'. Standards have been defined by the AAMI (as described in Hollander, 1982), which state a minimal required level of performance with respect to offset potential and recovery time. Hollander has done comparative study of surface ECG-electrodes. He compared 24 electrodes that the manufacturers considered to be long-term. Some of his findings were: Gel-to-gel measurements • The DC-offset 1 minute after application ranged from 0.1 to 218 mV for different makes,

and up to 250 mV after 24 hours. The AAMI standard is a maximum of 100 mV after 1 minute.

• The offset drift over 30 seconds ranged from 0 to 7 mV, and over 300 seconds from 0 to 35 mV for different makes. The course of the offset during the time interval is not specified, only the absolute difference between beginning and end.

• The mean value of the 10 Hz impedance measured over a one week ranged from 23 to 3000 Ω for different makes. The AAMI standard is 2 kΩ. Per make the impedance varied only a few percent over the whole week.

Measurements on human skin • The short-term impedance generally decreases over time (4 hours) and ranged from 10 to

100 kΩ for different makes. The impedance can decrease up to 50% in one hour. • The offset ranges from 2 to over 100 mV for different makes. Generally electrodes with a

large offset also showed the largest variations in offset potential between the measurement periods (1 hour).

• The 'dynamic offset' is also measured. It is defined as the offset drift after a load of 5 N is released from the electrode. It ranged from 1 mV for the best electrode after one day of application, to over 125 mV (the limit of the recording equipment) for the worst.

Other aspects of Hollander's research include defibrillation tests, adhesion period and skin reaction. The course of the offset potential was measured as an indication of the baseline drift variations of the ECG equipment. Unfortunately, noise measurements were not carried out. The noise level determines the minimum amplitude of bioelectric signals that can be measured. Some information regarding noise can be gathered from Hollander's research: the worst performing electrodes gel-to-gel have an offset variation of 7 mV over 30 seconds, i.e. a mean of 0.2 mVs-1. This drift can elevate the low-frequency signal spectrum. Dankers (1996), compared three different Ag-AgCl electrodes, and found large differences in the noise levels. An indication: measurements on the inner forearm gave a noise rms value of 3.4±0.7 µV for one electrode, 18±9 µV for the other. Corner frequencies of 3 and 9 Hz were measured for two different electrodes. He also found that next to electrode material, the electrolyte plays an important role in the resulting noise level. No clear relationship has yet been found between electrolyte composition and noise. The peak-to-peak noise levels reported by Gondran (1996) for pregelled Ag-AgCl electrodes are 0.25 µV gel-to-gel (no bandwidth is mentioned for this exceptionally low value), and 10-60 µV on the inner forearm. He used only one make of Ag-AgCl electrodes. Concluding, it can be said that in the literature some clinical aspects regarding Ag-AgCl electrodes, such as skin reaction and adhesion as well as some technical aspects such as offset potential, are well described. However, a comparative study of the noise levels cannot be found.


Other electrode materials

There are situations where the use of Ag-AgCl electrodes is not desired. Silver may be toxic when used inside the human body. Sometimes the use of electrode paste is undesirable, for instance in myoelectric control of protheses or when there's danger of bridging in closely spaced electrodes. Silver is not transparent to X-rays. Silver-silverchloride electrodes have also proven to be mechanically vulnerable. For this reason, some other electrode materials are in use, such as stainless steel, carbon and gold. Godin, et al. (1991) investigated the noise of stainless-steel electrodes, applied on the skin without electrode paste. They found the noise level inversely proportional to electrode diameter. No excess noise above amplifier noise was found, except for what might be some residual EMG signal. The amplifier noise was 60 µV/√Hz, for loads under 40 kΩ. Poor amplifier choice is the reason why they found no excess noise. Because of the high impedance of the pasteless electrodes, in many measurements the amplifiers current noise dominated the voltage noise. The bandwidth of the measurements was 8 Hz-10 kHz. It is suggested that under 8 Hz excess noise may be present. Other research has shown a considerable amount of excess noise. Dankers (1996) found that gold electrodes in Elefix paste gel-to-gel had rms noise levels of 6 µV, comparable to Ag-AgCl electrodes. In Redux paste the noise was much higher (39 µV). On the skin, the difference between the two pastes was less pronounced. He showed that although various materials can be used to record biopotentials, Ag-AgCl electrodes have the best properties with regard to the generated noise level.

3.2 A new development: the NASICON electrode

Silver-silverchloride surface electrodes are 'compatible' with organic fluids, because at the electrode surface an exchange of chloride ions is possible. This is what makes the Ag-AgCl electrode insensitive to polarization. Gondran, et al. (1995) has proposed a transducer which can interact with the sodium ion, another important charge carrier in organic fluid. The electrode is based on a ceramic material, NASICON, formula Na3Zr2Si2PO12. This material is a good conductor for Na+ ions. When brought in contact with the skin, an ion exchange will occur:

++←

→NASICONNaNa skin

In his work, Gondran shows that, when in contact with the skin, the inner potential of the NASICON ceramic is representative of the skin potential, without electrode paste applied. The electrode is shown in figure 3-1. The signal-transducing element is an Ag-AgCl wire, in contact with the NASICON ceramic via an electrolyte. The electrode behaves much like standard Ag-AgCl electrodes. The impedance is somewhat higher (figure 1-8). This is to be expected from a dry electrode. The noise level was higher than Ag-AgCl electrodes, of course with the higher impedance the corresponding thermal noise will be greater. Gondran, et al. (1996) found the 0.5-100 Hz peak-to-peak noise level ranged from 10-60 µV for different patients. The yet unexplained excess noise, as observed with the Ag-AgCl electrodes, is also witnessed using the NASICON electrode. Impedance is inversely proportional to the electrode area. It is remarkable, that the inverse proportional relationship between area and excess noise level, which is observed in Ag-AgCl electrodes is much less pronounced in NASICON electrodes. However, these results are based on limited data. Only a few curves have been recorded, three different electrode diameters on two patients.


Fig. 3-1. The construction of the NASICON electrode


4 Capacitive electrodes

It is stated earlier that the application of electrode paste is not always desired. The use of pregelled electrodes can save time, however the subject still is in contact with electrode gel, which may cause skin irritation. Smearing of the electrolyte can short circuit closely spaced electrodes. The resistive path provided by the electrolyte also tends to deteriorate in long-term measurements. According to literature, recording of biopotentials with ‘dry’ electrodes has proven to be successful over the last decades (Gondran, 1995, Godin, et al., 1991). It is possible because modern high-input-impedance amplifier design has made a low-resistance path between skin and electrode unnecessary. In these articles it is not mentioned that high electrode resistance can elevate thermal noise voltages well within the range of interest in biopotential recording (figure 2-5). Note that a truly dry electrode does not exist, as sweat serves as an electrolyte. Dry galvanic conducting electrodes are in use for some applications, however the qualities of silver-silverchloride electrodes can not be matched. A different approach is the use of insulated or capacitive electrodes. In this type, the insulation provides a very high resistance between subject and electrode, creating a capacitive coupling between them. To create a large capacitance, dielectric has to be thin. This makes it vulnerable to mechanical disturbance, and there is danger of electrostatic breakdown. The use of a capacitive electrode also requires some modifications in amplifier design. Various capacitive electrodes have been designed and tested. However, there are no reports of extensive use of capacitive electrodes in practice. The next section describes some realizations of the idea, followed by some possible design criteria that may be used in determining the parameters of a capacitive electrode. After that, another implementation is suggested.

4.1 Some examples of capacitive electrodes

One of the first insulated electrodes is described by Richardson, et al., and Lopez, et al. (from Geddes, 1972). The transducer in this electrode was a plate of anodized aluminum, the dielectric was provided by aluminum oxide. Richardson reported the resistance between subject and electrode plate to be between 1000 and 30,000 MΩ. This proved the capacitive coupling according to Richardson, however no data on thickness of the dielectric or the capacitance it provided is given. The aluminum plate is attached to a field-effect transistor, connected as a source follower (figure 4-1). A metal ring connected to the ground acts as an electrostatic shield. According to literature, ECG's recorded between two of these electrodes were indistinguishable from those obtained with regular Ag-AgCl electrodes. Lagow (as described by Geddes, 1972) investigated the influence of saline solution on Richardson's electrode, and found that saline breaks down aluminum oxide, creating a resistive path for current flow. He argues that as saline is a component of human sweat, this makes the electrodes useless for prolonged use. Lagow has suggested the use of oxidized tantalum, and proved it to be saline resistant. It is not mentioned by Lagow, but very well possible that next to saline other substances in perspiration may also cause breakdown of the insulating layer. Prutchi and Sagi-Dolev (1993) describe a more recent development. They were concerned with developing a dry electrode that could be integrated into flight suits and helmets. The metal used in this design is again aluminum, however the insulating oxide layer is now provided by a new technique, they call hard anodizing "Super". The process creates an aluminum oxide layer which is mechanically strong, resistant to corrosion of saline, and has stable dielectric properties. The electrode includes a buffer amplifier which lowers the impedance to minimize the sensitivity to noise (figure 4-2). The plate is additionally protected from interference by a driven shield. The use of a conducting foam as part of the headphone is suggested as a possible reference electrode. Again, this electrode records biopotentials with


Fig 4-1. Capacitive electrode by P. Richardson (from Geddes, 1972).

Fig. 4-2. Capacitive electrode by Prutchi and Sagi-Dolev (1993).

the same quality as Ag-AgCl electrodes. Some of the design criteria used were noise levels under 0.7 µV rms over a bandwidth 0.5-500 Hz, when recording biopotentials a noise level under 1 µV rms, electrostatic discharge protection, and an output impedance low enough to drive a distant bioamplifier. Unfortunately, the article by Prutchi and Sagi-Dolev gives no data how these criteria were accomplished, such as the thickness of the aluminum oxide layer, the properties of the electronic components used, or skin preparation. A second recent development of capacitive electrodes is given by Taheri, et al., (1994). The goal was to create an EEG electrode that could be applied without paste, to prevent bridging when the electrodes are used in a closely spaced array. The sensing element used in this design is a stainless steel plate with a 200 nm thick silicon nitride coating, connected to a buffer amplifier with an input impedance of 1015 Ω. The amplifier is battery powered, which was mounted on the electrode. The active surface is small enough for good contact with the scalp. The electrode can be connected to standard EEG instrumentation. Biopotentials were recorded with use of an Ag-AgCl reference electrode. Again a performance comparable to standard Ag-AgCl electrodes was achieved. The only drawback was excessive low-frequency motion artifact, due to bulkiness of the prototype. Plans for a smaller design have been made.


Fig. 4-3. The capacitive electrode and the input impedance of the amplifier form a high-

pass filter. ττττ=RC.

The article by Taheri, et al. states that a capacitance of 150-300 pF is desired in order to dominate skin impedance. No data is given on the actual capacitance the authors achieved. In fact, the numerical data given in the article is doubtful. For instance, a design criterion stated by Taheri, et al. is that the electrode must be able to pick up frequencies as low as 0.01 Hz. It is unclear why; frequencies of interest in EEG recording start at 0.5 Hz. But, using this value, in the article it is argued that the desired 150-300 pF imposes an input impedance of at least 6.7.1014-3.3.1014 Ω on the amplifier in order to reach a cutoff frequency of 0.01 Hz (see next paragraph). This RC combination will result in a cutoff frequency of approximately 10-6 Hz, much smaller than 0.01 Hz. The objectives of the authors do not match the actually implemented numerical values. This will be argued further in the next section.

4.2 Design criteria for capacitive electrodes

When designing a capacitive electrode, several factors have to be taken into account. The insulating layer has to be strong enough mechanically, resistant to electrostatic breakdown and corrosion due to perspiration. The relation between capacitance and thickness of the dielectric is:

d

AC rεε 0= (15)

Another criterion that has to be addressed is that the capacitive electrode and the input impedance of the amplifier form a high-pass filter (figure 4-3). The cutoff frequency of the filter is 1/2πRC, and has to match the lowest frequency of interest of the measured biopotential. This matter is complicated by the requirement that the amplifier has to be galvanically connected to the circuit, to provide a path for the input bias current of typically 1-10 pA. When using regular bioelectrodes this path is provided by the galvanic conduction of the electrodes. Using capacitive electrodes this path might be created by connecting a resistor between amplifier input and common. The value of the resistor is limited by the voltage drop it creates. In practice this value will be much lower than the amplifier input resistance, raising the cutoff frequency. None of the articles on capacitive electrodes discussed in the preceding paragraph mentions a provision for the input bias current of the amplifier, except for Prutchi and Sagi-Dolev (1993). For unknown reason the path is not drawn in figure 4-2. The claim of the other authors that biopotentials were successfully recorded using capacitive electrodes might be because the insulating layer was punctured, invalidating the term "capacitive". The acceptable noise level is another criterion that can be used in design. When using galvanic conducting electrodes the amount of thermal noise is determined by the resistance of electrodes and tissues. When using capacitive electrodes, the input resistance of the amplifier becomes dominant, and will account for the thermal noise. Together with the electrodes capacitance it will create a low-pass filter that will weaken noise above the cutoff frequency.


Taheri, et al., (1994) have used the following criterion for determining the electrode's desired impedance: "The skin impedance at 10 kHz is approximately 1 kΩ which results in a capacitance value of 15 nF. (...) The electrode capacitance should be at least two orders of magnitude smaller than 15 nF at 10 kHz to become the dominant frequency setting component in the system. Thus, values of electrode capacitance between 150-300 pF are adequate to dominate the frequency characteristics of the electrode". The authors used this reasoning to minimize the effect the skin impedance has on the recording. However, at the beginning of this paragraph it is argued that a path is needed for the amplifier's input bias current, for example a resistor of 10 MΩ, connected from amplifier input to common. The high-pass filter that is formed then has a cutoff frequency of 53 Hz, much higher than the 0.01 Hz aimed at by the authors. The product of R and C has to be increased to lower the cutoff frequency. Considering these poorly founded numerical values, the validity of the article can be questioned.

4.3 'Wet' capacitive electrodes

The purpose of using capacitive electrodes is generally to obtain a biopotential recording of the same quality as those recorded with Ag-AgCl electrodes, without the use of an electrolyte. A problem with dry electrodes is that poor contact is made with the skin. When using dry capacitive electrodes, the second capacitor plate is defined by the skin, an inhomogeneous layer (figure 4-4). It is clear that the capacitor is not well defined, it depends on the skin structure and the amount of sweat secretion, which changes in time. The dielectric is not only defined by the insulating layer, but also by gaps between skin and electrode. When an electrode gel is applied, the second plate is defined by the homogenous surface of the gel. The dielectric is well defined as the insulating layer. When electrodes are applied with paste, the advantage of a dry electrode is lost. However, another advantage is gained over galvanic conducting electrodes. When a capacitive electrode is used, there is no contact between metal and electrolyte. This means no halfcell potential will develop, and therefore a possible source of noise is eliminated. For this reason capacitive electrodes may prove an interesting tool in the investigation of the noise processes at the electrode surface. Comparison of noise measurements obtained from wet capacitive electrodes and galvanic conducting electrodes might give insight in the role the metal-paste interface plays in electrode noise processes. When the metal-electrolyte interface plays an important role in the generation of noise, wet capacitive electrodes can provide less noisy measurements. No literature can be found in which a truly wet capacitive electrode is tested. Research has to be done to test its performance.

Skin surfaceSweatHair

Metal plate

Dielectric?

Second plate?

First plate

Electrode lead

Insulating layer

Fig. 4-4. Schematic cross-section of the interface between a ‘dry’ capacitive electrode

and the skin.


Conclusions

Measurement of bioelectric events is subject to noise from various sources. Part of this noise originates from the electrode-skin interface. When the amplifier is of high quality, and noise and interference reducing measures have been taken (such as preparation of the skin, driven right leg circuit and guarding circuit), the noise level at the electrode-skin interface can become dominating. Peak-to-peak values of the voltages created can be in the same order of magnitude, or larger than some bioelectric events (for example evoked potential EEG and ventricular late potentials). Above 0 °K every resistor exhibits a certain amount of thermal noise. This is a fundamental limitation in biopotential recording. However, the noise originating from the electrode-skin interface is higher than the thermal noise corresponding with the resistance of electrodes and tissues. Little is known about the source of this excess noise. Abrading the skin lowers the low-frequency excess noise, indicating it may originate from out-of-equilibrium processes in the upper skin layers. The interaction between body fluids and electrode gel is generally assumed to be the cause of part of the noise. The noise level is dependent on the electrode material, so processes at the electrode-electrolyte interface may contribute to excess noise. Another explanation mentioned in literature is that residual EMG signals due to incomplete muscle relaxation might contribute to the excess noise. The broad field of application for surface electrodes has resulted in an extensive supply of electrodes. Aspects such as offset potential and its drift, impedance, skin reaction and adhesiveness are well documented for silver-silverchloride electrodes. A large comparative study on the noise of different electrodes cannot be found in literature. The need for dry electrodes has resulted in the idea of a capacitive electrode. In this type the electrode's conducting surface is coated with an insulating layer, and is therefore capacitively coupled to the skin. The insulating layer has to be very thin, which has the consequence that it is mechanically vulnerable. It also has to be resistant to electrostatic breakdown. Puncture of the insulating layer causes the electrode to become galvanically conducting. Some examples of capacitive electrodes can be found in literature. However, there is very little information available on the actual performance of the electrodes. When a capacitive electrode is applied to the skin with the use of an electrode paste, there is no interaction between metal and electrolyte. No halfcell potential can develop, in contrast to galvanic conducting electrodes. When the coupling between electrode metal and electrolyte is purely capacitive, it can be assumed not to contribute to the noise. For this reason the 'wet' capacitive electrode can prove to be a useful tool for determining the contribution of the various transitions in the electrode-electrolyte-skin interface in the total amount of noise produced. There was no documented research found to include in this paper on capacitive electrodes applied with paste.


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Noise in biopotential recording using surface electrodes

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