Design of a Peripheral Nerve Electrode for Improved Neural ... · Design of a Peripheral Nerve Electrode for Improved Neural Recording of the Cervical Vagus Nerve Bita Sadeghlo Master

Design of a Peripheral Nerve Electrode for Improved NeuralRecording of the Cervical Vagus Nerve

by

Bita Sadeghlo

A thesis submitted in conformity with the requirementsfor the degree of Master of Applied Sciences

Graduate Department of Electrical & Computer EngineeringUniversity of Toronto

c© Copyright 2013 by Bita Sadeghlo

Abstract

Design of a Peripheral Nerve Electrode for Improved Neural Recording of the Cervical Vagus Nerve

Bita Sadeghlo

Master of Applied Sciences

Graduate Department of Electrical & Computer Engineering

University of Toronto

2013

Vagus nerve stimulation (VNS) is an approved therapy for patients suffering from refractory epilepsy.

While VNS is currently an open loop system, making the system closed loop can improve the therapeutic

efficacy. Electrical recording of peripheral nerve activity using a nerve cuff electrode is a potential long-

term solution for implementing a closed-loop controlled VNS system. However, the clinical utility of

this approach is significantly limited by various factors, such as poor signal-to-noise ratio (SNR) of

the recorded electroneurogram (ENG). In this study, we investigated the effects of (1) modifying the

electrode contact dimensions, (2) implementing an external shielding layer on the nerve cuff electrode

and (3) exploring shielded bipolar nerve cuff designs on the recorded ENG. Findings from both computer

simulations and animal experiments suggest that significant improvements in peripheral nerve recordings

can be achieved.

ii

Contents

1 Introduction & Background 1

1.1 Vagus Nerve Stimulation for Treatment of Epilepsy . . . . . . . . . . . . . . . . . . . . . . 1

1.1.1 Background . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1

1.1.2 Mechanism of Action . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2

1.1.3 Different Vagal Fibers Involved in Vagus Nerve Stimulation . . . . . . . . . . . . . 3

1.2 Electroneurogram Recording and Electrode Design . . . . . . . . . . . . . . . . . . . . . . 3

2 Theory & Methods 6

2.1 Computational Study . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

2.1.1 Finite Element Model . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

2.1.2 Computational Generation of Action Potentials . . . . . . . . . . . . . . . . . . . . 7

2.2 In Vivo Study . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

2.3 Simulations/Experimental Protocol . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11

2.3.1 Effect of Electrode Dimensions on Recorded Electroneurogram . . . . . . . . . . . 11

2.3.2 Use of an External Shielding in Nerve Cuff Electrodes for Noise Reduction . . . . 12

2.3.3 Bipolar Nerve Cuff Electrode . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

2.4 Data Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

3 Results 17

3.1 Effect of Electrode Contact Dimensions on Recorded Electroneurogram . . . . . . . . . . 17

3.2 Use of an External Shielding in Nerve Cuff Electrodes for Noise Reduction . . . . . . . . . 32

3.3 Bipolar Nerve Cuff Electrode . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 41

4 Discussion 46

4.1 Middle Electrode Contact Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47

4.2 Side Electrode Contact Length . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48

4.3 Use of an External Shielding Layer in Conventional Nerve Cuff Electrodes . . . . . . . . 49

4.4 Bipolar Nerve Cuff Electrode . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 49

4.5 Conclusion and Future Approach . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 50

Bibliography 52

iii

Chapter 1

Introduction & Background

Epilepsy is a brain disorder in which patients experience spontaneous seizures due to disturbed brain

activity. Historically known as ”sudden disorderly expenditure of force” [75], epilepsy is now scientif-

ically defined as the occurrence of seizures which are accompanied with convulsive movements, loss of

consciousness, or both [12]. Seizures are considered symptoms where the causes are pathologic states

ranging from cerebral lesions, trauma and hemorrhage to water and alcohol intoxication, drug use, fever

or insulin shock and etc [12]. There are several treatments for this chronic neurologic disorder including

anti epileptic drugs (AEDs) [70, 9, 41, 33], brain surgery [69, 58, 78], and neural stimulation, such as

deep brain stimulation [64, 8, 1, 10] and vagus nerve stimulation [52, 6, 38, 18].

AEDs are the first options available to epileptic patients, however, 30% of the patients never reach

a stable seizure free state [32, 41] and many of those achieving successful treatment suffer from the side

effects of the medications [41, 26, 4, 61]. Brain surgery is one of the possible treatments available to

patients that are not responsive to AEDs, but only 25% of patients with refractory epilepsy can undergo

resective surgery. This is due to the difficulty of localizing the epileptic foci, risk of interfering with

functional regions of the brain, high risk of the surgery in certain situations, negative attitude of the

patients and their family toward the surgery, and lack of access to surgery[69]. An alternative treatment

for patients with refractory epilepsy is neural stimulation. While deep brain stimulation has emerged

as a potential option, it still remains under investigation with very few control studies having been

completed. The exact structure to be stimulated and the characteristics of the best stimuli are also still

unknown[64]. The risk and expenses of this type of treatment are such that it is considered as the last

option for those patients who are unresponsive to AEDs and are not candidates for other therapies[64].

For patients suffering from refractory epilepsy, where AEDs are not effective, vagus nerve stimulation

would be the first treatment used since it is less invasive than brain surgery or deep brain stimulation,

and side effects associated with it are minor[52].

1.1 Vagus Nerve Stimulation for Treatment of Epilepsy

1.1.1 Background

The application of electrical vagus nerve stimulation (VNS) for treatment of epilepsy can be traced

back to the late 19th century when J.L. Corning used transcutaneous VNS in conjunction with carotid

compression with the intention of reducing the cardiac output and slowing the heart beat in epileptic

1

Chapter 1. Introduction & Background 2

patients[34]. Later on, several effects of VNS on the central nervous system (CNS) in different animal

models were studied, which included cortical synchrony in cats, generation of slow cortical waves in

anesthetized primates, and thalamic responses by awake cats (cited by [16]). In 1985, the clinical

potential of VNS for treatment of epilepsy was proposed by Zabara. He hypothesized that VNS will

suppress the hyper-synchronized cortical and thalamocortical interactions that happen during the onset

of seizures [77]. Although later studies showed that VNS does not have any desynchronization or hyper-

synchronization effects on electrical activity of human brain [55, 17], its anti-convulsive effects were

supported by other studies.

First experiments on patients with intractable epilepsy were conducted during 1988 and 1989 using

the vagal nerve stimulator developed by Cyberonics Inc. According to the investigations, the results

were called ”promising” [45]. In 1994 Ben Menachem et al.[6] showed that different levels of VNS will

have different effects on the frequency of partial seizures, where patients receiving higher levels of VNS

showed a statistically significant increase in mean of seizure reduction compared to those receiving lower

levels of VNS . Common side effects of VNS were observed in the same experiment and cited by [52].

These include hoarsness, coughing, change in voice quality and muscle twitch. Chest and abdominal

pain were also observed in a few patients . However, these side effects were only felt during the VNS

pulses and were considered ”minor” [52]. Other randomized controlled trials also showed the efficacy and

safety of VNS for treatment of medically intractable seizures in which 20-40% of the patients achieved

greater than 50% reductions in frequency of seizures [15, 18]. Vagus nerve stimulation for treatment of

intractable epilepsy was finally approved by US Food and Drug Administration in 1997 and has been

widely used ever since. The current procedure of VNS for humans includes stimulating the left cervical

branch of the vagus nerve [6, 67].

1.1.2 Mechanism of Action

VNS does not have any noticeable effects on human EEG, however, it still alters EEGs during the

onset of seizures [17]. Ben-Menachem et al. [5] found a decrease in aspartate, an increase in GABA,

and an increase in ethanolamine in the cerebrospinal fluid of patients who received VNS. A study by

Henry et al. [21] measuring the brain blood flow in patients receiving VNS showed an increased blood

flow in medulla, right post central gyrus and a bilaterally increase in hypothalamus, thalamus, and

insular cortex. Bilateral decreases in hippocampus, amygdala, and posterior cingulate gyrus were also

shown. The amount of decrease or increase of the blood flow seemed to correlate with the level of the

stimulation [21]. Activation of thalamus during VNS is consistent with one of its roles in the body

which is the control of the cerebral activity [28]. Another study conducted in rats showed the Locus

Coeruleus (LC) to be a critically important structure in the circuitry necessary for the action of VNS [30].

Since LC is the main site of norepinephrine synthesis in the brain, it was interpreted that noradrenergic

neurotransmitters might enhance the effects of VNS [28]. Further studies demonstrated an intensity

dependent bilateral increase of extracellular cortical and hippocampal norepinephrine levels following

stimulation [53]. Readt et al. [49] also showed that hippocampal norepinephrine is an important factor

in the anti-convulsant effects of VNS . The increase of GABAA receptors is another factor proposed to

play a role in the procedure [36]. Mclachlan [37] showed that stimulating C fibres in anesthetized rats

have the same seizure suppression effect as heating the tail and therefore hypothesized that increasing

the activity of reticular activating system is the mechanism of action for VNS. Studies on vagus nerve

stimulation are still ongoing while the exact mechanisms by which VNS achieves seizure suppression still


remains unknown.

1.1.3 Different Vagal Fibers Involved in Vagus Nerve Stimulation

The Vagus nerve consists of 4 different fibers classified based on their conductance velocity: myelinated

A-fibers, myelinated fast and slow B-fibers, and unmyelinated C-fibers. For recruiting each of the fibers,

different thresholds are needed where the largest myelinated A-fibers have the smallest threshold and

the smallest unmyelinated C-fibers have the highest thresholds [54, 72, 74]. Knowing the exact portion

of each fibre recruited during stimulation requires recording of the compound action potentials from

the nerve trunk. Woodbury & Woodbury showed that VNS in anesthetized rats only achieved seizure

suppression effects when the threshold was high enough to recruit C fibers [72]. In that study they

used tripolar recording of the electroneurogram using three hook wire electrodes attached to the nerve.

Later studies showed that chemically blocking C fibers in rats does not interfere with the anti-convulsant

effects of VNS [31]. Although some of the C-fibers including those innervating GI tract are resistant to

the chemical used for neuroblocking [7], lack of autonomic side effects in patients receiving VNS still

suggests that C-fibers do not play an important role [52, 31]. In another study, high resolution recordings

of the activity of vagal fibers was performed during VNS in canines, where the outer connective tissue

of the nerve (epineurium) was removed to achieve a higher signal to noise ratio[74]. This study also

suggests that C-fibers are not recruited at those thresholds currently used in clinical trials [74]. Based

on these findings, Yoo et al. [74] hypothesized that A fibers and a proportion of B fibers are responsible

for seizure suppression effects of VNS. On the other hand, Krahl et al have found that stimulation of

the abdominal branch of vagus will still yield seizure suppression (cited by [29]). Knowing that the

abdominal vagal afferents mainly consist of slow B fibres and unmyelinated C fibres [54, 48, 68], these

findings reveal further needs for studying the mechanism of VNS and the exact vagal fibres involved.

1.2 Electroneurogram Recording and Electrode Design

Studies have shown that on-demand vagus nerve stimulation for treatment of epilepsy is more effective

than automatic stimulation with predefined intervals [39]. Therefore, it could be argued that using a

feedback signal, from which the onset of seizures could be predicted, to make the system closed loop

would potentially improve the treatment. For example, studies have shown cardiac related cervical vagus

nerve activity as a good indicator of the onset of the seizures [19, 43]. Given the presence of cardac

fibbers within the vagus nerve trunk, it is clear that this signal could be a potential candidate as the

feedback signal for closed loop vagus nerve stimulation therapy. This reveals another aspect of the

importance of improving the signal to noise ratios of ENG signals recorded from peripheral nerves.

Improving the performance of electroneurogram (ENG) recording is not only important for VNS

therapy, but also in any peripheral nerve stimulation system with a variety of applications (e.g., reha-

bilitation [47], bladder dysfunction [27], etc). However, the current nerve recording technology is still

not capable for longterm use as a feedback signal in closed loop systems. As a result, most therapies are

implemented as open loop systems [50, 20, 44]. One of the main reasons that the ENG signals are not

reliable feedback signals is their low signal to noise ratio (SNR) [44]. The amplitude of electroneurogram

signals of the peripheral nerves are in the range of a few µV with a frequency band between 1kHz to

10kHz and a maximal power below 3kHz[44]. These nerves are surrounded by muscles that produce

electromyogram (EMG) signals of amplitudes around 1mV and a frequency band of 1Hz to 3kHz with


a peak at around 250Hz [44]. Due to this anatomical organizatory EMG signals are the most common

noise component of the ENG. Without finding a way to dramatically reduce the surrounding noise, the

signal to noise ratio of the recorded ENGs would be very small as the noise is three order of magnitudes

greater than the signal [44].

Pseudo-tripolar nerve cuff electrodes are traditionally used for recording electroneurogram signals

from peripheral nerves [59]. Such electrodes consist of a middle electrode contact that is recorded differ-

entially with respect to two symmetrically positioned side electrode contacts (connected to each other)

and an insulating cuff that is used to increase the ENG amplitude and reduce the effects of electrical

currents produced by the surrounding tissue (most importantly the EMG signal of the surrounding mus-

cles) [23, 60]. In earlier studies, it was argued that because of the short circuiting between the two lateral

electrodes, no voltage drop in the longitudinal direction can happen, therefore all the currents generated

by sources outside the cuff would flow through the short circuit connection rather than through the cuff.

Under ideal conditions this will yield zero noise in the recorded ENG [59, 50]. More importantly, later

studies showed that isolating the nerve by a relatively long and narrow cuff will linearize the field of

any signal generated by a source located outside cuff. The approximately homogeneous fluid inside the

cuff will behave like an ideal one dimensional distributed resistor whose potential varies linearly with

distance along the length of the cuff [50]. Therefore, for any electrical source located outside the nerve

cuff, the potential in the middle of the cuff would be the average of the potentials at the ends of the cuff

and short circuiting the two lateral electrodes has the same effect as averaging their potentials which

further helps this effect [2]. Therefore, noise is eliminated by differential recording. However, because

of the impedance between the tissue and electrode contacts, the potential of the lateral electrodes will

not be exactly the same. Some ionic current from outside the cuff will flow through the cuff, due to cuff

imbalance [50], and noise will not be completely eliminated by differential recording . Close proximity

of the noise source to that nerve cuff might also result in deviation from ideality [65]. Furthermore,

when the ratio of the length ande diameter of the cuff is not large enough for it to behave like an ideal

distributed resistance, departure from linearity will occur near the ends of the cuff; therefore, the length

of the cuff and the position of the lateral (side) electrodes will affect the amount of noise recorded in

the ENG [50, 2]. It was found that the effect of electrode separation on the noise is only important for

short (5-10 mm) cuffs where placing the lateral electrodes near the ends of the cuff yielded the most

noise because of the departure from linearity at these ends [50].

There have been several previous studies aimed at improving signal to noise ratio of peripheral nerve

electrodes. Some of these studies focussed on mediating cuff imbalance; for instance, adding a balancing

resistor to the side electrodes [62], or using two shorted middle electrodes placed at the centre of the cuff

instead of one [11]. Other studies focussed on avoiding the nonlinear field inside the cuff. [50] showed

that although the field inside the cuff generated by a source outside of the cuff would be linearized by

the cuff, it is still non linear at the ends of the cuff; therefore, placing the side electrodes a few mm from

the ends of the cuff would avoid this nonlinearity and improve the recorded signal. They also showed

that increasing the length of the cuff would help in decreasing the gradient of this interfering field. [51]

showed that placing two additional electrodes (connected by a wire) at the end of the cuff could improve

the signal by reducing the the amount of current flowing into the cuff.

Choosing the optimal electrode dimensions is an important factor in the fabrication of peripheral

nerve recording electrodes. However, not many studies have investigated how to find the optimum

dimensions to achieve the best SNR as the space for electrode implantation is usually limited in peripheral


nerve stimulation/recording studies.

Functional electrical stimulating (FES) systems were the first peripheral nerve stimulating/recording

systems to use information of peripheral nerve signal recordings as feedback signals in closed loop systems

[24]. Chronically implanted intrafascicular electrodes [35] have also been used for collecting sensory

feedback signals for FES systems [40, 76]. Intrafascicular recordings yield a higher SNR due to placement

of the electrode within the endoneurium [35]. Use of a Faraday cage (made from different conductive

materials) for noise reduction was also proposed in intrafascicular recording studies to further improve

the SNR [14, 40, 76].

In the current study, we investigated the feasibility of using nerve cuff electrodes as a means of

achieving closed-loop control of VNS therapy for the treatment of epilepsy. This objective was achieved

by optimizing the design of peripheral nerve cuff electrodes for improved recording of neural activity.

A computer model was implemented to test the hypotheses that (1) changing the electrode contact

dimensions and (2) implementing an external shielding layer to the nerve cuff can improve the SNR of

the recorded ENG. The results of our computer simulations were validated by complementary in vivo

experiments.

Chapter 2

Theory & Methods

2.1 Computational Study

2.1.1 Finite Element Model

A 3-dimensional computer model of a myelinated peripheral nerve was constructed using a finite ele-

ment software (Comsol Multiphysics, Comsol Inc). The model (fig 2.1) consists of a single cylindrical

nerve trunk, cylindrical electrode contacts, a cylindrical nerve cuff and surrounding saline. The overall

geometry of the surrounding saline had negligible effect on the accuracy of the model [73]. Table 2.1

summarizes the geometries and electrical conductivities of the different compartments in the model. In

order to model action potential firing of a myelinated nerve fibre, a cubical current source was used with

Ij = 100 (nA) to represent a node of Ranvier of a nerve fibre in the nerve trunk. This current source

was placed along the whole nerve at 0.2 (mm) intervals and the corresponding electrode potential of the

electrode contacts were solved by the software. The mesh size was selected in a way that an accuracy

of two decimal points was achieved. In all simulations, the external surface of the saline medium was

set as the electrical ground. The resulting electric potential values of the electrode contacts were ex-

ported (MATLAB) and used to generate single fibre action potentials (SFAPs). Electrical noise sources

that modeled adjacent muscle activity were implemented as a cubical current source placed outside the

nerve cuff. The location of the noise source and the current amplitude were adjusted according to each

simulation.

In order to introduce noise into simulated pseudo-tripolar ENG recording, a nonideal condition of

impedance mismatch was modelled. Under ideal conditions, the electrical impedance of the nerve cuff is

symmetrical and the interfering field inside the cuff is linear. As a result, the noise recorded with such

a tripolar configuration is zero. Non-neuronal noise was therefore introduced into the model by altering

the conductivity of the saline layer - between the nerve trunk and the inner surface of the insulating

silicone layer - in only one half of the cuff. This resulted in an asymmetrical impedance along the nerve

such that the Z2/Z1 ratio (figure 2.1) of the electrode contacts deviated from unity (where 1 corresponds

to ideal conditions). In these studies, Z1 was kept constant at the same conductivity as regular saline,

while conductivity of Z2 was varied to approximate up to a 100-fold mismatch in electrical impedance.

6

Chapter 2. Theory & Methods 7

Diameter (mm) Length(mm) Electrical Conductivity(S/m)

Nerve Trunk 1 30 8.26E-02Electrode Contacts 1.025 =< 2 9.44E+06

Nerve Cuff 1.125 13 1.00E-12Saline Medium 30 60 1.05

Cubical Current Source (I = 100nA) n/a 0.1 9.44E+06

Table 2.1: Electrical Properties of Peripheral Nerve Recording Model.

Z1

Nerve CuffRecording Electrodes

Z2

Nerve Trunk

Figure 2.1: Finite element model of a peripheral nerve recording system-Modeling impedance mismatch:while the conductivity of Z1 was kept constant at the same conductivity for saline, the conductivity ofZ2 was varied in order to model a range of a 100-fold impedance mismatches

2.1.2 Computational Generation of Action Potentials

Action potentials are transmitted via saltatory conduction in myelinated nerve fibres, meaning that the

current will jump between the nodes of Ranvier. When the action potential reaches a node of Ranvier, a

potential change will happen which causes the ion channels to open and a time dependent action current

to flow through the nodal membrane. Figure 2.2 shows the action current of the membrane of each node

of Ranvier during saltatory conduction. This plot was found according to [42] with a sampling time of

17.92 µsec.

For a nerve fibre with diameter D (µm), the internodal distance is L = 100D and the conduction

velocity ν is 5.58D (m/sec) [63], although some studied use 6 as the conversion factor [25]. In the current

study the same conversion factor as [63] was used. Based on these values, the time delay between every

two adjacent nodes of Ranvier is 100D/5.58D = 17.92µsec. Meaning that if the action current reaches

the first node of Ranvier at time t, it will reach the next node of Ranvier at time (t + 17.92) µsec.

Therefore, if the amplitude of the action current of a specific node of Ranvier at time t is X1, then at

time t+ ((n−1)∗17.92) the template is Xn (figure 2.3). Furthermore, if the action current at node N is

Xn at time t, then the action current at node (N + n− 1) is X1 at the same time. Following this logic,

in order to construct a nerve fibre action potential, the electrode potentials were solved using the finite

element model, exported into MATLAB, and computationally weighted according to the time-dependent

action current template shown in figure 2.2.

Simulation of extra-neural noise was achieved in the finite element model by implementing a point

source (as described in 2.1.1) to model adjacent muscle activity. The voltage template of this source was

chosen as shown in figure 2.4 which was the laryngeal electromyogram signal recorded during one of the

in vivo experiments with a stimulus amplitude of 70 µA (refer to section 2.2).

In these studies, the electric potentials determined by the finite element software were exported into


0 20 40 60 80 100 120 140−1.5

−1

−0.5

0

0.5

1

1.5

Time (*17.92 microsec)

Cu

rren

t (n

A)

Figure 2.2: action current at each node of Ranvier during saltatory conduction

MATLAB and scaled according to the voltage template.

2.2 In Vivo Study

In order to validate the results of the computer model, in vivo experiments were conducted in sprague

dawley rats (female, weight = 250 to 450 g). The protocol was approved by the university of Toronto

division of comparative medicine. These experiments were acute non-survival experiments. Rats were

induced and subsequently anaesthetized by inhalation of 2− 5% isoflurane. The rat was then intubated

and connected to a ventilator machine. Body temperature, heart rate, and blood O2 levels were mon-

itored and maintained throughout all experiments. With the rat in the surgical plane of anesthesia,

hair was removed from ventral neck and chest. The area was washed with alcohol and iodine. The

main trunk and the recurrent laryngeal branch of the left vagus nerve were dissected and isolated. A

stimulating bipolar electrode was implanted on the recurrent laryngeal nerve and a tripolar recording

nerve cuff electrode was placed on the main trunk for recording the ENG (figure 2.5). A pair of stainless

steel wires was inserted into the laryngeal musculature to record the muscle activity. All signals were

conditioned (gain = 1000-5000, filter = 100 Hz to 10 kHz) and digitally recorded (sampling rate = 20

kHz, Powerlab 16/35). A subcutaneous needle was used as the common ground of the circuit.

The duration of each experiment was about 8 to 10 hours and the animal was euthanized at the end

of each experiment by injection of T61 into the heart.


Figure 2.3: Action current template- If the amplitude of the action current of a specific node of Ranvierat time t be X1, then at time t+ ((n− 1) ∗ 17.92) it would be Xn

0 20 40 60−1

−0.5

0

0.5

Time (*5e−5)

No

rmalized

Vo

ltag

e

Figure 2.4: Normalized laryngeal electromyogram signal used as the template signal for the noise source


Figure 2.5: Experimental setup for recording neural and muscular reponses evoked by stimulation of therecurrent laryngeal nerve.

Length (mm) Width (mm) Height (mm)

Recording Nerve Cuff Electrode (tripolar) 12 5 2Stimulating Nerve Cuff Electrode (bipolar) 6 5 2

Table 2.2: Dimensions of the nerve cuff electrodes used in in vivo study

Manufacture of Nerve Cuff Electrodes

Flat tripolar recording and bipolar stimulating nerve cuff electrodes (similar to FINE [66]) were made

as follows: Platinum electrodes were connected to lead wires using conductive silver epoxy and then

fabricated on a silicone sheet using silicone epoxy. Another layer of silicone sheet was fabricated on this

sheet in a way to cover it all except for a channel window through the length of the cuff with a width

of 1mm which was to expose the electrodes for the nerve to be placed on. Another silicone sheet was

attached on top (while having the bottom side free), in a way to be closed after implanting on the nerve,

to be able to isolate the nerve from the surrounding. The size of the manufactured bipolar and tripolar

nerve cuff electrodes are shown in table 2.2. Different sizes of platinum contacts were used according to

each study.


Length of the middle electrode contact(mm) Length of the side electrode contacts(mm)

Electrode No. 1 0.5 1Electrode No. 2 1 1Electrode No. 3 2 1

Table 2.3: Nerve cuff electrodes with three different middle electrode contact lengths (along the distanceof the nerve) used in in vivo experiments (refer to figure 2.1 for orientation of electrode).

2.3 Simulations/Experimental Protocol

2.3.1 Effect of Electrode Dimensions on Recorded Electroneurogram

Effect of Decreasing the Length of the Middle Electrode Contact-Computational Studies

In these studies, the effect of decreasing the length of the middle electrode contact was investigated. In

the first set of computational studies, the length of the middle electrode contact was varied from 2mm

to 0.25mm while keeping the side electrode contact length at 1mm. The distance between the outer

edge of the side electrode and the end of the cuff was 1.5 mm. The computer simulations generated

SFAPs for each electrode contact configuration. Additional simulations using a monopolar nerve cuff

electrode were conducted to control for the effects of changes in inter-electrode distances that resulted

from altering the length of the electrode contact.

To determine the effects of the middle electrode contact length on the recording of an external noise

source(s), simulations were conducted for a complete range of impedance mismatches (Z2/Z1= 1 to 100).

Assuming that the nerve trunk is lying on the y axis in cartesian coordinates, extending from −15mm

to +15mm, the noise source was initially positioned at (x, y, z) = (0,−10,−3) mm.

Effect of Decreasing the Length of the Middle Electrode Contact-In Vivo Experiments

Three different recording nerve cuff electrodes were used for the in vivo experiments. The side electrode

contacts were 1mm ∗ 5mm while having three different middle electrode contact sizes of 0.5mm ∗ 5mm,

1mm ∗ 5mm and 2mm ∗ 5mm (table 2.3). The recurrent laryngeal nerve was stimulated with a constant

current source using different current amplitudes changing from 5 µA to 10 mA. The neural activity

was recorded from the main trunk of the vagus nerve. The nerve electrode was connected to a band

pass amplifier of high pass frequency = 100 Hz and low pass frequency = 10 kHz, with a gain of

5000. The laryngeal muscle activity was also recorded for each stimulus amplitude. The output of the

electrodes were connected to a band pass amplifier with the same frequency band as the ENG amplifier

and with the gain of 1000. At the end of each experiment, the nerve was cut first above the stimulating

electrode to prevent stimulation of the laryngeal muscle and confirm the presence/absence of laryngeal

EMG artifact in the ENG recording. Then the main trunk was cut distal to the tripolar electrode,

where the disappearance of the neural activity during the stimulation of RLN at this point confirmed

the source of the ENG signal to be neural activity.

Effect of Increasing the Length of the Side Electrode Contacts-Computational Studies

In this part of the study, SFAPs were generated for nerve cuff electrodes, where the length of the side

electrode contact was increased from 0.25mm to 2.0mm, while the middle electrode contact length


remained constant at 1mm.

To discriminate the effect of the electrode contact length from that of changing the interelectrode

distance, in the first part, the distance between the middle electrode contact and side electrode contacts

was kept constant while increasing the length of the side electrode. Meaning that by increasing the

length of the electrode it would approach towards the ends of the cuff. In the second part, the same set

of simulations were repeated keeping the distance of the side electrode lengths from the end of the cuff

constant at 1.5mm while increasing the length of the side electrodes from 0.25mm to 2.0mm.

The effect of increasing the length of the side electrode contacts on the noise was also determined

computationally for the complete range of impedance mismatches (Z2/Z1= 1 to 100). The length of

the side electrodes were varied in two ways, the same as mentioned above for the range of 0.25mm to

2.0mm.

2.3.2 Use of an External Shielding in Nerve Cuff Electrodes for Noise Re-

duction

Computational Studies

The effect of different configurations of an external shielding layer (Platinum with a conductivity of

9.44E + 06 S/m) on noise pick up was investigated for a nerve cuff electrode with electrode contact

lengths of 1mm. In the first set of simulations, an external shielding plate facing towards the noise

source, with the dimensions of thickness = 0.025mm, width = 1mm and the same length as the nerve

cuff (13 mm) was placed on the outside of the cuff (figure 2.6(a)).

In the second set of simulations, one external shielding ring with the diameter of 1.175mm and

coaxial with the nerve trunk was located at the mid-point along the length of the cuff (figure 2.6(b)).

The length of the shielding layer was variable ranging from 1.5mm to 13mm (covering the whole cuff).

In the third set of experiments, two shielding rings coaxial with the nerve trunk and with the diameter

of 1.175mm were located at the two ends of the cuff (figure 2.6(c)). The length of each ring was varied

from 1.5mm to 6.5mm (covering the whole cuff).

In order to include the effect of noise source proximity [65], the effect of shielding was found for close

noise sources at different locations lying on a line parallel to the nerve trunk with a distance of 3mm

from the axis of the nerve trunk. The noise source was placed on this hypothetical line at the intervals

of 5mm, starting from −15mm and ending at +15mm. In this case, a Z2/Z1 ratio of 10 was used to

model the impedance mismatch.

In the next simulation study, external noise recorded by a nerve cuff electrode with a cylindrical

shielding layer covering the entire length of the cuff (as the most effective configuration) was simulated

for a complete range of impedance mismatches (Z2/Z1 ratio = 1 to 100). The noise source was located

at (x, y, z) = (0,−10,−3) mm. For the same shielded nerve cuff electrode, the SFAP was found for the

ideal case (Z2/Z1=1) to explore the effect of shielding on the fields generated by sources located within

the nerve cuff.

The electric fields of both inside cuff and outside cuff sources were found by the finite element software

for the two conventional and shielded nerve cuff electrodes. These electric fields were plotted on the XY

plane along the nerve trunk which could be used to show the effect of the shielding layer on the resulting

electric field. The effect of impedance mismatch on the arrangement of electric field inside the cuff (i.e.,

effect on noise recording) was also further investigated by plotting the electric field of an outside cuff


(a) Having a shielding plate on the outside of the cuff

(b) having one outside ring located at the middle length of the cuff

(c) Having two outside rings located at the two ends of the cuff

Figure 2.6: Different configurations of the outside cuff shielding.


source for both ideal (Z2/Z1=1) and non-ideal (Z2/Z1=10) cases.

In Vivo Experiments

The experimental setup described in 2.2 (figure 2.5 ) was used in this part of the study. The recurrent

laryngeal nerve was stimulated with a constant current source at different amplitudes ranging from 5µA

to 10mA. First, the ENG was recorded with conventional electrode No. 2 (table 2.3). The electrode was

subsequently wrapped with an aluminum foil and the experiment was repeated for the whole range of

current amplitudes. The laryngeal muscle activity was recorded throughout the experiment. The output

of the electrodes were connected to amplifiers with the same gain and filter settings as in 2.3.1.

At the end of each experiment, the nerve was cut above the stimulating electrode to prevent stimula-

tion of the laryngeal muscle and thereby confirm the presence/absence of EMG activity in the recorded

ENG signal. Subsequently, RLN stimulation was repeated after the main trunk was cut distal to the

recording tripolar electrode. This further confirmed the source of the recorded activity obtained from

the nerve cuff electrode (e.g., stimulus artifact and neural signal).

2.3.3 Bipolar Nerve Cuff Electrode

The effects of minimizing external noise activity by tripolar nerve cuff electrodes was further investigated

in bipolar configurations. A computational model same as previous sections was implemented to generate

both SFAP and noise signals using a bipolar nerve cuff electrode. For all simulations, the noise source

was placed at (x, y, z)=(0,−10,−3) where the nerve trunk was located on the y axis extending from −15

mm to +15 mm.

In the first set of simulations, the effect of shielding on the recorded noise for a bipolar nerve cuff

electrode of the length of 6.5mm was analyzed for the range of a 100-fold impedance mismatch. These

results were compared to that of a conventional bipolar nerve cuff, conventional tripolar nerve cuff, and a

shielded tripolar nerve cuff electrode. Then, for a given 10-fold impedance mismatch, the noise recorded

with different shielded bipolar nerve cuff electrodes of different lengths was found and compared with a

conventional tripolar nerve cuff electrode with the length of 13mm.

SFAPs were also generated using different lengths of bipolar nerve cuff electrodes: l= 13mm, 10mm,

8mm and 6.5mm.

2.4 Data Analysis

To quantify the signals found in both computational and experimental studies, Vpp and RI values were

calculated as follows:

RI =

nn∑i=n1

|Vi| (2.1)

Vpp = Vmax − Vmin (2.2)

In equation 2.1, the values of n1 and nn were determined for each signal to define the time-window

during which the stimulation evoked response was expected. In the computationally generated signals,

defining a time-window was not necessary and the sum was calculated throughout the whole length


of the signal since the signal was non-zero only at times that the response occurred. However, for

the experimental data (ENG and EMG responses), defining a proper time-window was necessary to

discriminate the response from artifacts. This time-window verified visually for each stimulus response.

The responses were further confirmed by comparing the recorded signals (that included the response)

with signals recorded after the nerve was cut (i.e., when the response was disappeared but the artifact

was still recorded). The same time-window was used for equation 2.2 for determining the maximum and

minimum of the evoked signal.

The effects of changing the electrode dimensions was quantified by calculating the percent change in

signal. This was calculated as shown in 2.3, where Vref is the reference signal with respect to which the

percent difference is determined.

%difference =Vsignal − Vref

Vref∗ 100 (2.3)

To be able to compare the data found in different experiments, they were normalized with respect

to a reference signal. Equation 2.4 shows how the data was normalized in the study.

Vsignal =Vsignal −min (Vref )

max (Vref )−min (Vref )(2.4)

In the studies that required signal to noise ratio, it was calculated as follows:

SNR = 20log(signal(ENG)

Noise)[dB] (2.5)

Conduction Velocity of Myelinated Fibres

For the in vivo study, the minimum and maximum conduction velocities (vmin & vmax) of the myelinated

fibres were calculated as follows:

vmax =δl

n1 ∗ δt(2.6)

vmin =δl

nn ∗ δt(2.7)

Where δl is the distance between the stimulating electrode and the recording electrode (measured

during the rat experiments), n1 and nn refer to the start and end of the time-window of the response

respectively, and δt is the sampling time which is 1f where f is the sampling frequency which was 20

kHz throughout all of the experiments.

Recruitment Curves

In order to confirm the threshold and saturation points of ENG/EMG activation, recruitment curves were

plotted for each experiment. The RI value of the signal was calculated for each stimulus amplitude and


were plotted in the stimulus amplitude-RI plane as (x, y) = (stimulus amplitude,RI). The myelinated

fibre activity threshold was defined as the threshold for which the RI value reaches 10% of the saturation

value.

Statistical Analysis

To summarize the data from all experiments, the values are reported in the form of Mean+−Standard

Error where each were calculated as follows:

Mean =

∑ni=1 Vin

(2.8)

σ =

√∑ni=1 (Vi −Mean)

2

n(2.9)

StandardError =σ√n

(2.10)

One way analysis of variance (ANOVA1) was used to compare means of different values found with

three different electrodes (in rat experiments) to test the null hypothesis that the three values are

not statistically different (meaning that they were drawn from datasets with same mean values). The

following MATLAB function was used for ANOVA1 analysis. This was followed by a multi comparison

test that searched for statistically significant differences between every two group with the null hypotheses

that every two groups are drawn from datasets with same mean values.

Chapter 3

Results

All computer simulations in this study were based on the both noise and single fibre action potentials

(SFAPs) that were generated using pseudo-tripolar and monopolar nerve cuff configurations. As shown

in figure 3.1(a), a SFAP generated tripolarly has a smaller peak-to-peak voltage (19%reduction) and a

shorter duration(16%reduction) compared to a monopolarly generated SFAP with the same isolating

cuff and electrode contact sizes . However, the tripolar recording provides complete noise rejection in

ideal conditions, whereas monopolar recording will exhibit any outside cuff interfering signal (figure

3.1(b)).

3.1 Effect of Electrode Contact Dimensions on Recorded Elec-

troneurogram


Effect on the signal

The effect of electrode dimensions was investigated for the following: peak-to-peak voltage (Vpp), rectified

integrated (RI) value and the shape of the SFAPs.

Figure 3.2 shows the effect of the middle electrode contact length on the Vpp of the SFAP. The Vpp of

the SFAP increased as the middle electrode contact length is decreased from 2mm to 0.25mm (with the

same side electrode contact length of 1.0mm for all of the electrodes). The solid line in figure 3.2 shows

the percent increase of the Vpp for a monopolar electrode and the dotted solid line shows the effects in a

tripolar electrode. Percent change in values were calculated with respect to the Vpp of the SFAP for the

electrode with middle electrode contact length of 2.0mm (and side electrode contacts of 1.0mm). The

effect of changing the middle electrode contact length on the RI value of the SFAPs was found to be less

than that for Vpp (less than 3%).

Figure 3.3 shows the effect of middle electrode contact length on the shape of the SFAP. By decreasing

the length of the middle electrode contact (tripolar configuration), the negative phase of the SFAP was

primarily affected (increased); whereas the change in Vpp of the positive phase was negligible (figure

3.3(a)). Changing the middle electrode contact length had minimal effect on the duration of the SFAP.

By increasing the length of the middle electrode contact, the duration of the negative phase increased

slightly whereas the duration of the positive phase slightly decreased (figure 3.3(b)).

17

Chapter 3. Results 18

0 50 100 150 200−3

−2.5

−2

−1.5

−1

−0.5

0

0.5

1x 10

−6


Sin

gle

Fib

re A

cti

on

Po

ten

tial

Tripolar Recording

Monopolar Recording

(a)

0 10 20 30 40 50 60−1.5

−1

−0.5

0

0.5

1x 10

−6

Time (*5e−5)

Re

co

rde

d N

ois

e (

V)

Tripolar Recording

Monopolar Recording

(b)

Figure 3.1: Monopolar vs tripolar recordings. (a)Computationally generated single fibre action poten-tial.(b)Computationally generated noise


0 0.5 1 1.5 20

1

2

3

4

5

6

7

Middle Electrode Contact Length(mm)

Peak−

to−

Peak P

erc

en

t In

cre

ase

Monopolar Recording

Tripolar Recording

Figure 3.2: Effect of middle electrode contact length on the peak-to-peak voltage of the single fibre actionpotential-Solid line: Percent increase of the Vpp of the single fibre action potential for a monopolar nervecuff electrode -Solid plus cross line: Percent increase of the Vpp of the single fibre action potential for atripolar nerve cuff electrode.


0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 20

1

2

3

4

5

6

7

Length of the Middle electrode(mm)

Pe

rcen

t In

cre

as

e o

f th

e P

ea

k−

to−

Pea

k V

olt

ag

e

% Increase of −ve part

% Increase of +ve part

(a)

0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2400

450

500

550

600

650

700

Length of the Middle electrode(mm)

du

rati

on

of

the

Sin

gle

Fib

re A

cti

on

Po

ten

tia

l(m

icro

se

c)

Duration of the +ve and −ve Phases of the SFAP recorded by different Electrodes

Duration of the +ve phase of SFAP

Duration of the −ve phase of SFAP

(b)

Figure 3.3: Effect of the middle electrode contact length on the shape of the single fibre action potentialfor a tripolar nerve cuff electrode.(a)Effect of middle electrode contact length on the peak-to-peak voltageof the negative and positive parts of the single fibre action potential.(b)Effect of middle electrode contactlength on the duration of the negative and positive parts of the single fibre action potential.


0 0.5 1 1.5 2−5

−4

−3

−2

−1

0

Side Electrode Contact Length (mm)

Pe

rce

nt

Dif

fere

nc

e V

alu

e

Peak to Peak Value

Rectified Integrated Value

Figure 3.4: Effect of side electrode contact length on Vpp and RI values of the single fibre action potential.The interelectrode distance between the middle and side electrodes was kept constant at 3.5 mm

Figure 3.4 summarizes the effect of changing the length of the side electrode contacts on the RI value

of the SFAP. Percent difference of the values were found with respect to the RI value of the SFAP for

the electrode with middle electrode contact size of 1mm and side electrode contact sizes of 2mm. As

shown in the figure, by decreasing the length of the side electrode contacts, the RI value of the SFAP

was modestly decreased. The effect of the side electrode contact sizes on the Vpp of the SFAP was found

to be minor (less than 1% while changing the length from 2mm to 0.25mm.). In these simulations, the

inter-electrode distance between the side electrodes and the middle electrode contacts was kept constant

at 3.5mm .

Decreasing the length of the side electrode contact increased the Vpp of the positive phase (less than

3%) while decreasing the Vpp of the negative phase (approximately 2%) (figure 3.5). The side electrode

contact length had negligible effect on the duration of the SFAPs.

Simulations (increasing side electrode length from 0.25mm to 2.0mm while keeping the middle elec-

trode contact length constant at 1mm) were repeated keeping the distance between the end of the side

electrodes and the end of the cuff constant at 1.5mm. Figure 3.6 shows three different SFAPs generated

using three different side electrode contact lengths of 2mm, 1mm and 0.25mm. As shown in the figure,

not only was the Vpp significantly affected, but the duration of the SFAP was also significantly increased.

Figure 3.7 shows the RI and Vpp values and the corresponding percent increases with respect to values

found with the tripolar nerve cuff electrode with the side electrode contact lengths of 2mm.

Effect on noise

Figure 3.8 summarizes the effect of middle electrode contact length on the recorded noise for the complete

range of impedance mismatches (Z2/Z1=1 to 100). In these simulations a monopolar noise source located

at (x, y, z)=(0,−10,−3) was used while the nerve trunk was lying on the y axis extending from −15

mm to +15 mm.The results showed an approximately 6% increase in Vpp of the noise while decreasing

the middle electrode contact length from 2mm to 0.25mm. The SNR(Vpp) remained constant and the

SNR(RI) decreased by 2% while changing the length of the middle electrode contact.


0 0.5 1 1.5 2−3

−2

−1

0

1

2

3

Length of Side electrode (mm)

Pe

rce

nt

Inc

rea

se

of

the

Pe

ak

−to

−P

ea

k V

ota

ge

% Difference of +ve part% Difference of −ve part

Figure 3.5: Effect of side electrode contact length on the peak-to-peak voltage of the negative andpositive parts of the single fibre action potential. The interelectrode distance between the middle andside electrodes was kept constant at 3.5 mm

Changing the length of the side electrode contacts had a completely different effect on noise depending

on whether it is increased towards the ends of the cuff or towards the centre. Figure 3.9(a) shows the

effect of decreasing the length of the side electrode contact on noise while the distance between the side

electrodes and the middle electrode was kept constant.

The noise significantly decreased as the side electrode contact length was decreased ( 40% decrease,

from 2 to 0.25 mm). This yielded a 25% increase in the SNR(RI) value and 27% increase in the

SNR(Vpp) value when changing the side electrode contact length from 2mm to 0.25mm.

However, the trend was the opposite if the distance between the end of side electrodes and the end of

the cuff was kept constant while deccreasing the electrode contact length. Figure 3.9(b)shows the results

while changing the side electrode contact length from 2 to 0.25 mm. In these simulations the distance

between the end of each side electrode contact and the cuff end was kept constant at 1.5 mm. As

shown, decreasing the side electrode contact length yields greater noise (6-7%increase in Vpp of the noise

while decreasing from 2mm to 0.25mm). Combined with the effects on the simulated SFAP (figure 3.7),

the overall effect of changing the dimension of the side electrode yielded a 60% increase in SNR(Vpp)

(changing from 2mm to 0.25mm). The same trend was also found when replacing the mono polar noise

source with a bipolar noise source while having the two sources on the ends of a hypothetical line parallel

to the nerve trunk and at a distance of 3 mm from it’s axis.

In Vivo Studies

The focus of this study was on myelinated fibre activity recordings. For each stimulus amplitude, the

signal was averaged over ten pulses to eliminate the random noise. The threshold of A fibre activity

(recorded with conventional electrode No. 2) was found to be 25.7 +− 4.45 µA (mean +

− standard error)


60 80 100 120

−25

−20

−15

−10

−5

0

5

x 10−6


Sin

gle

Fib

re A

cti

on

Po

ten

tial

(V)

Side Electrode Contact Length=2mm

Side Electrode Contact Length=1mm

Side Electrode Contact Length=0.25mm

Figure 3.6: Effect of side electrode contact length on the shape of the single fibre action potential, wherethe distance between the outer edge of the side electrode and the end of the nerve cuff was kept constantat 1.5 mm. The middle electrode contact length was 1mm for all of the three electrodes with-Red Plot:side electrode contact lengths = 2mm.-Bue Plot: side electrode contact lengths = 1mm.-Green Plot:sideelectrode contact lengths = 0.25mm.


−0.5 0 0.5 1 1.5 2 2.50

1

2

3

4

5

6

7x 10

−5

Re

cti

fie

d I

nte

gra

ted

Valu

e (

V)

70%

increase

42%

increase

(a)

−0.5 0 0.5 1 1.5 2 2.50

0.5

1

1.5

2

2.5

3

3.5x 10

−6

Side Electrode Contact Length (mm)

Pe

ak−

to−

Pea

k V

alu

e (

V)

13%

Increase

21%

Increase

(b)

Figure 3.7: Effect of side electrode contact length on the single fibre action potential, the middle electrodecontact length was kept constant at 1mm. (a) Effect on the rectified integrated value. (b) Effect on thepeak-to-peak voltage


0 10 20 30 40 50 60 70 80 90 1000

0.5

1

1.5

2

2.5

3

3.5

4

4.5x 10

−6

Z2/Z1 Ratio

No

ise

(V

)

midelec=2mm

midelec=1mm

midelec=0.25mm

Figure 3.8: Effect of middle electrode contact length on the recorded noise for the whole range ofimpedance mismatches

Probe > F

Threshold of A Fibre Activity 0.8009Minimum Conduction Velocity of A Fibres 0.4971Maximum Conduction Velocity of A Fibres 0.2157

Table 3.1: Threshold of A fibre activity, minimum velocities and maximum velocities of A fibres foundwith the three different nerve cuff electrodes were not significantly different with p values mentioned inthe table (N=5 rats)

with a minimum conduction velocity of 9.7 +− 1.5 m/s and a maximum conduction velocity of 31.6 +

−2.63 m/s for N = 7 rats. Figure 3.10 shows the recruitment curve of myelinated fibres found with

electrode No. 2 (length=1 mm). The thresholds and conduction velocities found with three different

electrodes (table 2.3) for five experiments were not significantly different. Table (3.1) summarizes the p

values found with one way analysis of variance (ANOVA1) for each variable.

Figure 3.11 shows the ENG signals recorded from the main trunk of the left vagus nerve while

stimulating the left recurrent laryngeal nerve with a stimulus amplitude of 40µA. The stimulus artifacts

and ENG signals are labeled in the figure. Top trace is the ENG signal recorded with electrode No. 1

(length=2 mm), middle trace is the same signal for electrode No. 2 and the bottom trace shows ENG

signal recorded with electrode No. 3 (length=3 mm). An increase in V pp of the signals recorded with

the nerve cuff electrodes having smaller middle electrode contacts was observed in figure 3.11. The

normalized V pp of the signals (with the corresponding standard errors for N = 5 rats) recorded with

the three different electrodes while stimulating the RLN with a stimulus amplitude of 70 µA (which was

on the saturation interval of the recruitment curve and therefore all the myelinated fibres were recruited

at this amplitude) are shown in figure 3.12. Data are normalized with respect to Vpp of the signal

recorded with electrode No. 3. Figure 3.13 and table 3.2 summarize the ANOVA1 analysis results for

V pp values showed in figure 3.12. Multiple comparison of the mean values is shown in figure 3.14, which


0 20 40 60 80 1000

1

2

3

4

5x 10

−6

Z2/Z1 Ratio

No

ise (

V)

sidelec=2mm

sidelec=1mm

sidelec=0.25mm

(a)

0 20 40 60 80 1000

1

2

3

4

5x 10

−6

Z2/Z1 Ratio

No

ise (

V)

sidelec=2mm

sidelec=1mm

sidelec=0.25mm

(b)

Figure 3.9: Effect of side electrode contact length on the recorded noise for the range of a 100-foldimpedance mismatch.(a)The length of the side electrodes were change in a way to keep the distancebetween the side electrode and middle electrode constant.(b)The length of the side electrodes werechange in a way to keep the distance between the side electrode end the end cuff constant at 1.5mm.


0 0.02 0.04 0.06 0.08 0.1 0.120

20

40

60

80

100

120

Stimulus Amplitude (mA)

Re

cti

fie

d I

nte

gra

ted

Va

lue

(M

icro

V)

Figure 3.10: Recruitment curve of myelinated fibres. ENG signals were recorded from the main trunkof the vagus nerve using electrode No. 2

Source SS df MS F Probe > F

Columns 3.13711 2 1.56856 5.73 0.0179Error 3.28488 12 0.27374Total 6.422 14

Table 3.2: Summary of ANOVA1 analysis of the Vpp of the ENG signals recorded with three differentelectrodes-N = 5 rats

indicated significant differences between the Vpp of the signals recorded with electrodes No. 1 & 3, and

that between electrodes No. 2 & 3.

Rectified integrated (RI) values were also calculated for the recorded ENGs. For N = 5 rats, the

rectified integrated values for the stimulus amplitude of 70µA is shown in figure 3.15 . Data were

normalized with respect to the RI value of the signals recorded with electrode No. 3. Table 3.3 and

figure 3.16 summarize the ANOVA1 analysis of the data. Multiple comparison of the means is shown in

figure 3.17 which showed significant difference between means of RI value of the signals recorded with

electrodes No. 1 & 3, and also between those of electrodes No. 2 & 3.


0 2 4 6 8 10 12−0.05

0

0.05

EN

G

(mV

)

0 2 4 6 8 10 12−0.05

0

0.05

EN

G

(mV

)

0 2 4 6 8 10 12−0.05

0

0.05

Time (ms)

EN

G

(mV

)

ENG

ENG

ENGStimulus Artifact

Stimulus Artifact

Stimulus Artifact

Figure 3.11: Recorded ENG from the main trunk of the left vagus nerve while stimulating the left RLNwith I = 40 µA -Top trace: recorded ENG with electrode No. 1. -Middle trace: recorded ENG withelectrode No. 2.-Bottom trace: recorded ENG with electrode No. 3.

Source SS df MS F Probe > F

Columns 2.08226 2 1.04113 6.56 0.0119Error 1.90413 12 0.15868Total 3.98639 14

Table 3.3: Summary of ANOVA1 analysis of the RI values of the ENG signals recorded with threedifferent electrodes-N = 5 rats


0

0.5

1

1.5

2

Middle Electrode Size

No

rm

alized

Peak−

to−

Peak V

alu

e

Electrode No. 1

Electrode No. 2

Electrode No. 3

Figure 3.12: Normalized peak-to-peak voltage of the ENG signals recorded with three different electrodes.Values are shown as : Mean+

−Standard Error for N = 5 rats

0.5 1 2

1

1.2

1.4

1.6

1.8

2

2.2

2.4

2.6

2.8

3

Figure 3.13: Anova1 analysis on normalized peak-to-peak value. Data were normalized with respect toVpp of the signal recorded with electrode No. 3-N = 5 rats


0.5 1 1.5 2 2.5

2

1

0.5

Click on the group you want to test

The means of groups 1 & 2, and 0.5 & 2 are significantly different

Figure 3.14: Multiple comparison analysis on normalized peak-to-peak values. Data were normalizedwith respect to Vpp of the signal recorded with electrode No. 3.

0

0.2

0.4

0.6

0.8

1

1.2

1.4

1.6

1.8

2

Middle Electrode Size

No

rma

lize

d R

ec

tifi

ed

In

teg

rate

d V

alu

e

Electrode No. 1

Electrode No. 2

Electrode No. 3

Figure 3.15: Normalized rectified integrated value of the ENG signals recorded with three differentelectrodes. Values are shown as : Mean+

−Standard Error for N = 5 rats


0.5 1 2

1

1.2

1.4

1.6

1.8

2

2.2

2.4

2.6

Figure 3.16: Anova1 analysis on normalized RI values. Data were normalized with respect to the RIvalue of the signal recorded with electrode No. 3

0.8 1 1.2 1.4 1.6 1.8 2 2.2

2

1

0.5

Click on the group you want to test

The means of groups 0.5 & 2, and 1 & 2 are significantly different

Figure 3.17: Multiple comparison analysis on normalized RI values. Data were normalized with respectto the RI value of the signal recorded with electrode No. 3


−0.02 −0.015 −0.01 −0.005 0 0.005 0.01 0.015 0.02−60

−50

−40

−30

−20

−10

0

location of Noise Source on (0,y,−0.003) (m)

Perc

en

t N

ois

e R

ed

ucti

on

Figure 3.18: Percent difference of noise recorded tripolarly by a cuff having an outside shielding platefacing the noise source from noise recorded with a conventional electrode. The X abscissa shows thelocation of the noise source which was moved parallel to the nerve cuff electrode, at a distance of 3 mmfrom the centre of the nerve trunk.

3.2 Use of an External Shielding in Nerve Cuff Electrodes for

Noise Reduction


Impedance mismatch was introduced to the system in order to deviate from ideal conditions and produce

noise in the tripolar nerve recording configuration. While the system had zero noise in the ideal case

(Z2/Z1=1), the peak-to-peak voltage of the recorded noise was increased by increasing Z2/Z1 ratio (solid

line in figure 3.21). For the case of Z2/Z1= 10, three different configurations of outside cuff shielding

layer was tested. In the first set of simulations, a shielding plate (figure 2.6(a)) was placed on the outside

of the cuff facing towards the noise source. Figure 3.18 shows the percent reduction in the noise signal

while changing the location of the point source along a hypothetical line, facing the shielding plate and

lying parallel to the nerve trunk at a distance of 3 mm from it, at 5 mm intervals starting from −15

mm to +15mm. As shown in the figure, the amount of noise reduction is dependent on the location of

the noise source; however, significant noise reduction (20% to 60%)was achieved for all locations. All

percent change values were calculated with respect to the noise recorded with a conventional tripolar

nerve cuff.

In the next set of simulations, the same method was applied while having one external shielding ring

located in the middle of the cuff (figure 2.6(b)) with the diameter of d = 1.175 mm. Figure 3.19 shows

the percent noise reduction for shielding rings with different lengths along the nerve. The X axis shows

the location of the noise source for each group of bars. For each length of the shielding ring, colour

coded in figure 2.6(b), the percent noise reduction is shown with the noise source placed at the following

locations: (0, y, 3) and y={-15,-10,-5,0,5,10,15} (values are shown in mm). Percent difference values

were calculated with respect to the noise recorded with a conventional tripolar nerve cuff.

The effect of having two external shielding rings (d = 1.175 mm) at the two ends of the cuff as shown

in 2.6(c) was also investigated. Figure 3.20 shows the percent noise difference achieved with different

lengths of the shielding rings. The length of the rings is colour coded in the figure. For instance, green

indicates the percent noise reduction that results from having two rings with lengths of 4.5mm (9mm


−0.02 −0.015 −0.01 −0.005 0 0.005 0.01 0.015 0.02−90

−80

−70

−60

−50

−40

−30

−20

−10

0

10

location of Noise Source on (0,y,−0.003) (m)

Perc

en

t N

ois

e R

ed

ucti

on

1.5 mm

3 mm

4.5 mm

6 mm

10 mm

whole cuff covered

Figure 3.19: Percent difference of noise recorded tripolarly by cuff having an external shielding ringlocated at the middle length of the cuff from noise recorded with a conventional nerve cuff electrode.The X abscissa shows the location of the noise source which was moved parallel to the nerve cuff electrode,at a distance of 3 mm from the centre of the nerve trunk. Different bars show the percent reduction fordifferent lengths of the shielding layer with the colour code showed in the legend.

of the cuff was covered by the two rings). Noise reduction was investigated for noise sources located at

different positions as mentioned above: (0, y, 3) and y={-15,-10,-5,0,5,10,15} (values are shown in mm).

Percent difference values were calculated with respect to the noise recorded with a conventional tripolar

nerve cuff.

The effect of having an external shielding layer completely covering the cuff (cylinder, coaxial with

the nerve trunk, d = 1.175mm) was also found for a range of a 100-fold impedance mismatches. In this

study, the noise source was located 3 mm from the centre of the nerve and 10 mm offset from the middle

of the nerve cuff electrode (at (x, y, z) = (0,−10,−3) mm). Our findings showed an approximately

consistent 70% noise reduction throughout the mismatch range. Results are shown in figure 3.21 (dotted

solid line).

Finally, the effects of the external shielding layer on the recorded SFAP (and noise) was tested by

comparing the electric potential generated at the electrode contacts by sources located both inside and

outside the cuff. The shielding layer used in this study was cylindrical, coaxial with the nerve trunk

with d = 1.175mm. Figure 3.22 shows the effect of shielding layer on the electric potential field inside

the cuff generated by an active node of Ranvier. Figure 3.23 shows the same effect for a node of Ranvier

closer to the end of the cuff. The external shielding layer had no effect on the electric field inside the

cuff i.e., SFAPs found with a shielded cuff were the same as those recorded with a conventional cuff.

In both figures (3.23,3.22),the top trace shows the electric field for a conventional nerve cuff electrode

where the bottom trace shows the same field for a shielded nerve cuff electrode.

Figure 3.24 shows the effect of an external shielding layer on the electric field of noise sources located

outside the cuff. In these simulations, the noise source was located at (0,−10,−3) mm where the centre


−0.02 −0.015 −0.01 −0.005 0 0.005 0.01 0.015 0.02−100

−80

−60

−40

−20

0

20

location of Noise Source on (0,y,−0.003)

Perc

ent N

ois

e R

eduction

1.5 mm

3 mm

4.5 mm

whole cuff covered

Figure 3.20: Percent difference of noise recorded tripolarly by cuff having two external shielding ringsat the two ends of the cuff from noise recorded with no Outside bar. The X abscissa shows the locationof the noise source which was moved parallel to the nerve cuff electrode, at a distance of 3 mm from thecentre of the nerve trunk.

0 20 40 60 80 1000

1

2

3

4x 10

−6

Z2/Z1 Ratio

No

ise (

V)

Noise of the conventional tripole cuff

Noise of the shielded tripole cuff

70%Noise

Reduction

Figure 3.21: Effect of impedance mismatch on the peak-to-peak voltage of a tripolarly recorded noise.-Solid line: Peak-to-peak voltage of the noise for a conventional nerve cuff. Dotted solid line: Peak-to-peak voltage of the noise for a completely shielded nerve cuff.


(a)

(b)

Figure 3.22: Effect of shielding on the electric potential field of a current source located in the middleof the cuff modelling a firing node of ranvier-(a)Electric potential for a conventional nerve cuff elec-trode.(b)Electric potential for a shielded nerve cuff electrode


(a)

(b)

Figure 3.23: Effect of shielding on the electric potential field of a current source located inside the cuffclose to the ends of the cuff modelling a firing node of ranvier-(a)Electric potential for a conventionalnerve cuff electrode.(b)Electric potential for a shielded nerve cuff electrode


of the nerve trunk was located at (0,0,0). The top trace shows the electric field for a conventional nerve

cuff where the bottom trace shows the same field for a shielded nerve cuff electrode.

It was shown in previous sections that impedance mismatch introduces noise into the ideal tripolar

recording system. Similarly, figure 3.25 shows the effect of impedance mismatch on the electric field of

an outside cuff noise source. In the top trace, the electric potential is shown for an ideally symmetrical

nerve cuff electrode where the bottom trace shows the same field for a nerve cuff electrode with an

impedance mismatch of 10-fold.

In Vivo Study

The effect of wrapping an electrical shielding layer around a conventional nerve cuff electrode was

explored. Figure 3.26 shows the ENG signals recorded from the main trunk of the left vagus nerve

while stimulating the corresponding left recurrent laryngeal branch with a stimulus amplitude of 40µA.

The evoked EMG signal recorded from the laryngeal musculature is also shown (bottom trace). Top

trace in figure 3.26 shows the ENG signal recorded with a conventional nerve cuff electrode, which

includes stimulus artifact, EMG artifact and ENG contaminated with EMG artifact. In the second

trace, the signal recorded with the shielded electrode is shown. With the shielded electrode, not only

was the stimulus artifact reduced, but contamination of the ENG by EMG-derived noise was also notably

reduced. Percent reduction values between the two shielded and non-shielded nerve cuff electrodes were

found for two different variables: (1)Vpp of the stimulus artifact. (2)RI value of the ENG signal. The

amount of reduction in RI value of the ENG signal shows the amount of EMG artifact contamination

of the signal. These two values were found for two different stimulus amplitudes of 40µA and 100µA

which are shown in figure 3.27 for two experiments (N = 2). For nine different stimulus amplitudes in

the saturation interval of the recruitment curve (i.e. 40µA to 120µA ), a 59.36+− 3.11 percent reduction

in the RI value of the EMG artifact was found.

In the third trace of figure 3.26 the recorded signal from the main trunk of the vagus is shown while

stimulating the RLN but having the nerve cut below the recording nerve cuff electrode. As shown in

the figure, while the ENG response is disappeared, the signal still contains EMG artifact.


(a)

(b)

Figure 3.24: Effect of shielding on the electric potential field of a noise source located outside the cuff at(0,-10,-3) mm where the centre of the nerve trunk was located at (0,0,0) mm-(a)Electric potential for aconventional nerve cuff electrode.(b)Electric potential for a shielded nerve cuff electrode


(a)

(b)

Figure 3.25: Effect of impedance mismatch on the electric potential field of a noise source located outsidethe cuff at (0,-10,-3) mm where the centre of the nerve trunk was located at (0,0,0) mm-(a)Electricpotential for an ideally symmetrical nerve cuff electrode.(b)Electric potential for a nerve cuff electrodewith an impedance mismatch of 10-fold


0 50 100 150 200 250

−0.05

0

0.05

EN

G(m

V)

0 50 100 150 200 250

−0.05

0

0.05

EN

G

(mV

)

0 50 100 150 200 250

−0.05

0

0.05

EN

G(m

V)

0 50 100 150 200 250−0.5

0

0.5

EM

G(m

V)

EMGArtifact

EMGResponse

Time (*5e−5)

EMGArtifact

ContaminatedENG

ENGResponse

Stimulus Artifact

Stimulus Artifact

Stimulus Artifact

Stimulus Artifact

Figure 3.26: Effect of shielding on ENG and noise reduction -Top trace: ENG signal recorded with aconventional nerve cuff electrode. Second trace: ENG signal recorded with a shielded nerve cuff electrode.Third trace: Signal recorded from the main trunk while the nerve was cut below the recording nerve cuffelectrode that shows the disappearance of ENG response while including EMG artifact. Bottom trace:EMG signal recorded from the laryngeal musculature


1 2−70

−60

−50

−40

−30

−20

−10

0

Pe

rce

nt

Re

du

cti

on

Va

lue

Stimulus Artifact

EMG Artifact

40 Micro A

Stimulus Amplitude

100 Micro A

Figure 3.27: Amounts of stimulus artifact and EMG artifact reductions for an electrically shielded nervecuff electrode. Values are shown for two different stimulus thresholds of 40µA and 100µA.

3.3 Bipolar Nerve Cuff Electrode

This section reports the preliminary findings on using shielded bipolar nerve cuff electrodes. Figure

3.28 compares the noise of a bipolar nerve cuff electrode with the length of l = 6.5mm with the noise

of a tripolar nerve cuff electrode with the length of l = 13 mm(double the length of the bipolar cuff).

As shown in the figure, electric shielding will yield a 75% noise reduction for a bipolar cuff versus a

conventional bipolar cuff. A conventional bipolar cuff has significantly larger noise than a conventional

tripolar nerve cuff. However, for impedance mismatches higher than 3-fold, a shielded bipolar cuff

exhibits smaller noise compared to the conventional tripolar cuff, but still larger noise than a shielded

tripolar nerve cuff. Figure 3.29 shows the Vpp of the noise recorded with shielded bipolar nerve cuff

electrodes versus a conventional tripolar nerve cuff electrode for an impedance mismatch of 10-fold.

Compared to a conventional tripolar nerve cuff electrode with a length of l = 13mm, the shielded

bipolar nerve cuff electrodes with lengths smaller than 13mm yield smaller noise recordings.

Figures 3.30 and 3.31 show the RI and Vpp values of the SFAP potentials generated with bipolar

nerve cuff electrodes of different lengths and a tripolar nerve cuff electrode of the length of 13mm. As

shown in the figures, these two values decrease by decreasing the length of the bipolar cuff.

The signal to noise ratios (SNRs) were calculated for different lengths of shielded bipolar nerve

cuff electrodes and compared with a conventional tripolar nerve cuff electrode (with 10-fold impedance

mismatch). The SNR values were found for both RI values (figure 3.32) and Vpp values (figure 3.33).


0 20 40 60 80 1000

1

2

3

4

5x 10

−6

Z2/Z1 Ratio

No

ise (

V)

Conventional tripolar cuff (l=13 mm)

Shielded tripolar cuff (l=13mm)

Conventional bipolar cuff (l=6.5mm)

Shielded bipolar cuff (l=6.5mm)

Figure 3.28: Comparing the noise of a shielded bipolar nerve cuff electrode with a conventional bipolarand tripolar nerve cuff electrodes


V

Noise

Figure 3.29: Vpp of the noise recoded with shielded bipolar cuffs of different lengths versus a conventionaltripolar cuff with a 10-fold impedance mismatch


Rectified Integrated Value

V

Figure 3.30: Rectified integrated values of single fibre action potentials generated using bipolar nerve cuffelectrodes of different lengths and the tripolar nerve cuff electrode used in previous sections (l = 13mm)

V

Peak−to−Peak Voltage

Figure 3.31: Vpp values of single fibre action potentials generated using bipolar nerve cuff electrodes ofdifferent lengths and the tripolar nerve cuff electrode used in previous sections (l = 13mm)


Signal To Noise Ratio

dB

Figure 3.32: Signal to noise ratios (for RI value) bipolar nerve cuff electrodes of different lengths andthe tripolar nerve cuff electrode used in previous sections (l = 13mm)

Signal To Noise Ratio

dB

Figure 3.33: Signal to noise ratios (for Vpp value) bipolar nerve cuff electrodes of different lengths andthe tripolar nerve cuff electrode used in previous sections (l = 13mm)

Chapter 4

Discussion

In the current study, a computational model of a peripheral nerve for neural recording with nerve cuff

electrodes was implemented and tested. Using this model, the hypotheses that changes in the electrode

dimensions can have significant effect on the recorded neural signal was investigated. Furthermore,

the idea of applying an electrically insulating layer onto nerve cuff electrodes was tested in which a

highly conductive material was applied to the external surface of the nerve cuff to suppress noise signals

generated by external sources (e.g., muscles). Moreover, the feasibility of applying these techniques for

designing nerve cuff electrodes with a significantly smaller physical footprint, such as bipolar nerve cuff

electrodes, was also tested. Finally, these computational results were validated in animal experiments

that used the cervical vagus nerve of anesthetized rats as the physiological test-bed.

The computational model was constructed based on the physiological mechanisms of saltatory con-

duction of action potentials in myelinated nerve fibres. Similar to previous works that have implemented

mathematical and computational models for the same purpose ([63, 46, 73]), a pre-determined template

of the transmembrane current from a single node of Ranvier was used for constructing SFAPs from

electrode potentials computed by the finite element model. The SFAPs generated using monopolar and

tripolar nerve cuff electrodes agreed with those computationally found by [46] and [73] and experimen-

tally found by [59], with the difference that SFAPs in the current study were biphasic instead of triphasic.

This was due to the fact that our action current template approximation did not include the third phase

of the signal [46]. A working model of asymmetrical electrode impedances was also implemented to cre-

ate impedance mismatches that would result in a lower signal to noise ratio (i.e., increased noise levels).

Computational findings were further validated by in vivo experiments. The focus of the in vivo study

was on myelinated fibre activity. The threshold and conduction velocities of myelinated fibres found in

our rat experiments agreed with [71, 56, 22].

Previous studies have reported the effects of varying the cuff length ([50, 2]) or electrode configuration

([11, 3]) on the performance of nerve recording electrodes. However, there has been little work examining

the effects of the electrode contact length or the presence of an external shielding mechanism on the

recorded ENG in nerve cuff electrodes. This study focused on the effects of these variables using a

pseudo-tripolar nerve cuff electrode. In both computer simulations and in vivo experiments, the results

of this study suggest that reducing the size of the middle electrode contact increases the signal component

of the ENG, while implementing an electrical shield around the nerve cuff effectively reduces the noise

component of the ENG recording.

46

Chapter 4. Discussion 47

4.1 Middle Electrode Contact Length

Our computational and experimental studies on middle electrode contact length showed that decreasing

the length of this contact would increase the Vpp value of the SFAP. However, when the length of

the middle electrode contact is changed in a tripolar nerve cuff electrode, the inter-electrode distance

between middle contact and side contacts also changes. To be able to distinguish between these two

variables (i.e., middle electrode contact length and interelectrode distance), the idea of changing the

length of the electrode contact was also tested for a mono polar nerve cuff electrode with the same cuff

length. The same results were found for mono polar recording which was larger Vpp values for smaller

electrode contact lengths. The percent increase found in tripolar recording was only slightly (around

1%) greater than that of mono polar recording, which showed that the change in inter-electrode distance

has negligible effect on the recorded signal. One possible reason behind this is that when the middle

electrode contact length was changed δd mm, the inter-electrode distance was changed δd/2 mm. While

the inter-electrode distance was 3.875 mm for middle electrode contact length of 0.25 mm, it was 3 mm

when having a middle electrode contact length of 2 mm. Meaning that changing the middle electrode

contact length from 2 mm to 0.25 mm caused an approximately 20% change in the inter-electrode

distance which had negligible effect on the SFAP. Moreover, changing the length of the middle electrode

contact did not change the tripolar length (i.e., distance between the two side electrodes) and therefore

the overall SFAPs were not affected significantly.

Furthermore, it was found that decreasing the middle electrode contact length not only increases the

Vpp of the SFAP, but it also increases the Vpp of the noise. According to this, it could be argued that

whatever signal we are recording, using a smaller electrode contact gives a larger peak-to-peak value.

One possible reason could be the impedance of the electrode contact which is larger for smaller contacts

(r =ρ l/A [57]) and therefore yields larger peak-to-peak readings with smaller surfaces.

While changing the length of the middle electrode contact from 2 mm to 0.25 mm changed the Vpp

value around 6%, it only changed the RI value around 3%. It could be argued that the dominant effect

of changing the length of the middle electrode contact would be noticed for the Vpp value and not the

RI value.

It was noted that the effects of reducing the longitudinal dimension of the middle electrode contact

resulted in significantly larger effects in the animal studies than what was observed in the computationally

generated SFAPs. This was likely due to the additive effects of multiple axons that contributed to the

stimulation-evoked compound nerve action potentials (CNAP). Moreover, although the source of the

ENG signals was confirmed at the end of each experiment, because of the overlap between the ENG

and EMG responses , it wasn’t possible to fully discriminate between the pure ENG component and

EMG artifact component of the recorded ENG signal. However, according to the computer simulations,

at least half of the increase in Vpp of the recorded ENG is specifically related to the ENG component

and not the EMG artifact component of the recorded signal. The noted increase in the peak-to-peak

amplitude of the ENG signals in both computer and animal studies were also comparable to previous

work involving percutaneous monopolar recording of motor unit potentials [13].

Statistical analysis of the experimental data showed that both Vpp and RI values of the ENG signals

recorded with electrode No. 3 (longer middle electrode contact) were smaller than those values of the

ENG signals recorded with electrodes No. 1 and 2 (shorter middle electrode contacts). Using ANOVA1

and Multi Comparison tests, it was found that the difference in both Vpp values and RI values were

statistically significant with p values smaller than 0.05.


4.2 Side Electrode Contact Length

Computer simulations were conducted to explore the effect of side electrode contact lengths on both

SFAP and noise. In the first part of the study, the distance between middle electrode contact and side

electrode contacts were kept constant while increasing the side electrode contact lengths. While a very

slight (around 1%) increase was found for the Vpp of the SFAP, an approximately 4% increase was found

for the RI value of the SFAP when the length of the side electrode contacts increased from 0.25 mm

to 2 mm. First, it could be argued that the effect of middle electrode contact length on the Vpp of

the SFAP is significantly more dominant than the effect of side electrode contact length. One possible

reason behind this is that the maximum and minimum values of the SFAP are realized when the nodes

of Ranvier closer to the middle electrode contact are active. Therefore, the average potential recorded

by the middle electrode is so larger than the average potential recorded by side electrodes (because of

the distances between the current source and each of the electrode contacts) that changes in the length

of the latter would have negligible effects on the differential value (i.e., Vmiddlecontact - Vsidecontacts).

However, side electrode contact length had a more notable effect on the RI value of the SFAP. This

effect was slightly more than the effect of middle electrode contact length. The potential of the side

electrode contacts were averaged over the two electrodes and therefore changing each electrode about

1.75 mm yielded a total of 3.5 mm change in length. However, since the effect of side electrode contacts

in the recorded SFAP is less dominant than the effect of the middle electrode contact, only a slight

increase in the percent difference was found compared to changing the middle electrode contact length.

Despite the minor effect on SFAP, increasing the length of the side electrode contacts while keeping the

inter-electrode distance constant resulted in a huge effect on noise. As the side electrode contact length

was increased, the recorded noise significantly increased. However, as the increase in electrode contact

length resulted in the contacts becoming closer to the ends of the cuff, and since the interfering fields are

more nonlinear closer to the ends of the cuff [50], it could be argued that the increase in recorded noise

was a consequence of the end-of-the-cuff effect and not the length of the electrode contact. Moreover,

as supported by our monopolar simulations, changes in electrode size result in modest (4-7%) changes

for both signal and noise. Therefore, such huge percent differences (around 40% while changing from

0.25 mm to 2 mm) can not be the result of changes in electrode contact length. One possible solution

to this problem would be to increase the overall length of the cuff ([50, 59]). However, long nerve cuff

electrodes were not useful for the current study and manipulating the length of the cuff was also out of

the scope of this work. An alternate solution would be to keep the distance between the side electrode

contacts and the end of the cuff constant and repeat the simulations.

In the next part of the study, the distance between the ends of the side electrode contacts and the

ends of the cuff were kept constant while changing the length of the side electrodes. Meaning that

increasing the length of the side electrodes caused a significant decrease in the tripole length and the

significant changes in SFAPs’ Vpp and duration showed the interfering effect of this variable (for δd mm

change in the length of the side electrode contact, the inter-electrode distance was changed δd mm and

the tripole length was changed 2δd mm ). Moreover, the reduction in noise that resulted by increasing

the length of the side electrode contact could be the effect of both electrode contact length, and moving

the side electrodes toward the centre of the cuff where the noise field is more linear. Previous studies have

also showed that moving the side electrode contacts toward the centre of the cuff (by adding an extra

pair of short circuiting electrodes [3]), has the same effect as increasing the cuff length (i.e. increasing

the distance from the edges while keeping the tripole length constant ) [51]. Due to the very minor


improvements of the SFAPs, and since the end-of-the-cuff effect overrides the effect of the side electrode

contact length for noise recording, this idea was not further investigated in our rat experiments.

4.3 Use of an External Shielding Layer in Conventional Nerve

Cuff Electrodes

The design of an enhanced nerve cuff electrode that involved an external shielding layer was investigated.

The use of external shielding for noise reduction was proposed by [40] for noise reduction in intrafascicular

recordings where use of a carbon shielding layer around the nerve significantly reduced the noise. The

same approach was also used in other studies involving intrafascicular recordings, in which significant

noise reduction was achieved ([76]). The results of this study show that the use of an external shield

can also be applied to nerve cuff electrodes, where the distance between the nerve fiber and electrode

contacts are significantly greater than that of intrafascicular electrodes.

Simulation studies showed that using an external shielding for a conventional nerve cuff electrode

will significantly decrease the noise while not affecting the SFAP. The most effective configuration for

the external shielding was to cover the whole cuff which gave an approximately consistent 70% noise

reduction throughout a range of up to a 100-fold impedance mismatch.

Figures depicting the electric potential fields showed how the external shielding rearranged the electric

field generated by an outside cuff source. It could be argued that the external shielding shunts away

the current and prevents electrical field gradients from being created inside the cuff. However, when an

electrical source is inside the cuff (modelled as an active node of Ranvier ), the shielding does not affect

the corresponding extracellular potential field.

The figures depicting the electric potential also showed how an impedance mismatch will interfere with

linearizing the electric field of the noise inside the cuff and introduce noise into the system. In the ideal

case, the interfering field is completely linearized inside the cuff and therefore the averaged potential

of the side electrodes would be equal to the average potential of the middle electrode, and therefore

differential recording will eliminate noise. When an impedance mismatch is applied into the system,

the interfering field is no longer (completely) linearized and differential recording can not completely

eliminate the noise([50, 2]).

Rat experiments also showed promising results for the use of external shielding in conventional nerve

cuff electrodes. It was found that external shielding reduced both the stimulus artifact and EMG artifacts

in the recorded signal. The two stimulus amplitudes for which the percent decrease values were reported

all belong to the saturation interval of the recruitment curve; meaning that the RI value of the ENG

signal does not increase with increased stimulus amplitude (from 40µA to 120µA).

4.4 Bipolar Nerve Cuff Electrode

The preliminary results of computationally generated SFAPs and noise using shielded bipolar nerve cuff

electrodes were reported in the study. While bipolar electrodes have the advantage of having smaller

lengths (where they could be used for recording from smaller nerve branches e.g., thoracic branch of the

vagus), they lack the improved signal to noise ratios achieved by pseudo-tripolar nerve cuff electrodes

and therefore are not currently used for peripheral nerve recordings. The results of this study showed


that electrically shielded bipolar nerve cuff electrodes could yield a comparable noise level to that of a

conventional tripolar nerve cuff electrode. A bipolar electrode of a length half of a tripolar electrode was

expected to give a comparable SFAP. It was shown that both Vpp and RI values of a SFAP generated

with a bipolar nerve cuff with length l are significantly smaller than those of a SFAP generated with

a tripolar nerve cuff with a length of 2l. However, for impedance mismatches greater than 3-fold, the

signal to noise ratios were larger for the shielded bipolar nerve cuff electrode, compared to that of a

conventional tripolar nerve cuff electrode. Further work is still required to explore the use of shielded

bipolar nerve cuff electrodes in peripheral nerve recording, both computationally and experimentally

4.5 Conclusion and Future Approach

The results of this study suggest that modifications to conventional nerve cuff electrodes can have

potentially significant effects on the fidelity (i.e. SNR) of the recorded ENG. Further work is required to

test these ideas under more realistic conditions, such as using spontaneous (non-evoked) neural activity

in long-term implant studies.

Furthermore, according to the results of this study, it is suggested that the use of the combination

of the two methods, i.e., modifying the length of the middle electrode contact while using an electric

shielding around the cuff for noise reduction, presents an optimal configuration for ENG signal recording.

While the electric shielding would significantly reduce the noise, a smaller middle electrode contact could

be used to increase the Vpp and RI values of the CNAP and subsequently increase the ”signal” component

of the signal to noise ratio.

Moreover, the results of this study showed how the edge effect can override the effect of the side

electrode contact length on noise. For future studies, it is suggested to use a very long cuff in which the

side electrodes could be placed as far from the edges as possible, such that changes in noise levels can

be attributed to either the effect of side electrode contact length or the edge effect.

For the future studies, an alternative rat experimental setup is recommended, in which two indepen-

dent stimulation sources are used to generate both neural (signal) and muscular (noise) activity. For

instance, using one constant current source for stimulating the recurrent laryngeal nerve and another

one for stimulation of the Sternohyoid muscle. Using this method, a more exact evaluation of the nerve

cuff electrode could be achieved since the effect on noise and ENG could be separately examined as they

could be separately elicited by independent stimulators. It is also recommended to stimulate muscles

the EMG responses of which have time delays different from the time delay of the ENG response to be

able to better discriminate between the pure ENG response and the EMG contamination of it.

In experimental studies which involve recording from small nerve branches, the space limitation poses

a significant challenge for using tripolar nerve cuff electrodes. Preliminary computational studies using

shielded bipolar nerve cuff electrodes showed that it is possible to significantly reduce the noise down

to levels comparable to that of conventional tripolar nerve cuff electrodes. Further work is required to

fully develop this idea and test it in experimental studies.

Improving the SNR of conventional nerve cuff electrodes is an important step towards using ENG

signals as a feedback signal in closed loop neuroprosthetic systems. The main objective of this study

was to improve the design of nerve cuff electrodes to enhance signal quality. Moreover, we envision using

these novel electrode designs for recording neural activity from the various different branches of the

vagus nerve in studies aimed at better understanding the mechanism(s) behind vagus nerve stimulation


for the treatment of epilepsy. With smaller (bipolar) nerve cuff electrodes, recording from the smaller

branches would become feasible.

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