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A novel mechanism of cochlear excitation during simultaneous stimulation and
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pressure relief through the round window
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Thomas D. Weddell1, Yury M. Yarin2, Markus Drexl3, Ian J. Russell1, Stephen J.
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Elliott4 and Andrei N. Lukashkin1
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BN2 4GJ, UK
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Dresden, Fetscherstr. 74, D-01307 Dresden, Germany
School of Pharmacy and Biomolecular Sciences, University of Brighton, Brighton,
Clinic of Otorhinolaryngology, Department of Medicine, Universitätsklinikum
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15, 81377 Munich, Germany
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Southampton, SO17 1BJ, UK
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Short title: round window stimulation of the cochlea
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Authors for correspondence:
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Andrei N. Lukashkin
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e-mail: [email protected]
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Stephen J. Elliott
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e-mail: [email protected]
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German Vertigo Center, Ludwig-Maximilians University Munich, Marchioninistr.
Institute of Sound and Vibration Research, University of Southampton,
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Summary
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The round window membrane (RW) provides pressure relief when the cochlea is
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excited by sound. Here we report measurements of cochlear function from guinea pigs
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when the cochlea was stimulated at acoustic frequencies by movements of a miniature
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magnet which partially occluded the RW. Maximum cochlear sensitivity,
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corresponding to subnanometer magnet displacements at neural thresholds, was
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observed for frequencies around 20 kHz, which is similar to that for acoustic
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stimulation. Neural response latencies to acoustic and RW stimulation were similar
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and taken to indicate that both means of stimulation resulted in the generation of
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conventional travelling waves along the cochlear partition. It was concluded that the
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relatively high impedance of the ossicles, as seen from the cochlea, enabled the region
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of the RW not occluded by the magnet, to act as a pressure shunt during RW
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stimulation. We propose that travelling waves, similar to those due to acoustic far-
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field pressure changes, are driven by a jet-like, near-field component of a complex
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fluid-pressure field, which is generated by the magnetically vibrated RW. Outcomes
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of research described here are theoretical and practical design principles for the
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development of new types of hearing aids, which utilise near-field, RW excitation of
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the cochlea.
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Key words: cochlear round window; guinea pig cochlea; cochlear excitation; active
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middle ear prosthesis; implantable hearing aid, near-field excitation.
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1. Introduction
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The round window membrane (RW) acts as a pressure relief valve for the almost
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incompressible fluids of the cochlea, making possible movement of the stapes and,
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hence, movement of the inner ear structures. It has long been known (e.g. see Culler et
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al [1]) that cochlear function is impaired when the RW membrane is immobilised,
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thickened, congenitally malformed or absent [2-4]. These observations signify the
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physiological importance of the RW membrane for audition, and the necessity to
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retain it, which has been revealed during surgery of the middle and inner ear [5].
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Traumatic damage or rupture of the RW membrane can cause hearing loss and
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deafness due to perilymph aspiration or loss of the normal pressure-releasing function
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of a compliant RW [6-9].
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It has been established, however, that while normal function of the RW is important
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for effective stimulation of the cochlea through the conventional, oval window, route,
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the cochlea can be stimulated successfully in non-conventional ways (e.g. through
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bone conduction, through the RW, and through perforations in the cochlea’s apical
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turn). All of these techniques produce similar patterns of cochlear sensitivity and
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excitation [10-12]. In recent years significant recovery of hearing thresholds has been
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achieved in human patients through the use of active middle ear implants (vibrating
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electro-mechanical devices) to stimulate the cochlea through the RW at the
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frequencies of the incoming sound [13-17]. The RW approach could be an effective
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and preferable alternative to conventional hearing aids for hearing rehabilitation in
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patients with, for example, chronic inflammatory middle ear diseases, recurrent
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cholesteatoma, when the anatomy of the middle ear is highly distorted [14], and for
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patients with congenital malformations of the outer and middle ear, i.e. in cases of
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conductive hearing loss. In the latter situation dysplasia and immobilization of the
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ossicles, often in combination with malformations of the oval window, can make
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hearing reconstruction through the oval window impossible. The feasibility of the RW
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approach for the above disorders, and for conditions when the oval window is
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inaccessible, has been demonstrated in successful attempts to combine outer and
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middle ear reconstruction with implantation of active middle ear prostheses on the
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RW [13, 15, 16]. An understanding of the mechanisms of the cochlear excitation
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through the RW is essential to ensure that the technique is to be predictable and
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effective in the clinic.
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One uncertainty relates to the mechanisms of basilar membrane (BM) excitation
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using probes that cover only a part of the RW membrane, whereby the area of the RW
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not covered by the probe provides an effective pressure shunt (Békésy [11], page
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197). This might make excitation of the cochlea problematic through shunting
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pressure changes that would otherwise excite the BM. Mobility of the stapes is
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probably not essential for successful cochlear excitation through the RW if most of
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the pressure relief is provided by the non-occluded area of the RW. The paper
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presented here deals with the mechanisms of cochlear excitation, using a miniature
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magnet that only partially occludes the RW. With the stapes mobility decreased, a
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novel form of cochlear excitation is achieved with a sensitivity that matches that of
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conventional acoustic stimulation.
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2. Material and methods
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Pigmented guinea pigs (280-390 g) were anaesthetised with the neurolept anaesthetic
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technique (0.06 mg/kg body weight atropine sulphate s.c., 30 mg/kg pentobarbitone
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i.p., 500 l/kg Hypnorm i.m.). Additional injections of Hypnorm were given every 40
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minutes. Additional doses of pentobarbitone were administered as needed to maintain
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a non-reflexive state. The heart rate was monitored with a pair of skin electrodes
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placed on both sides of the thorax. The animals were tracheotomized and artificially
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respired, and their core temperature was maintained at 38C with a heating blanket
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and a heated head holder. The middle ear cavity of the ear used for the measurements
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was opened to reveal the RW. Compound action potentials (CAPs) of the auditory
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nerve were measured from the cochlear bony ridge in the proximity of the RW
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membrane using Teflon-coated silver wire (S in figure 1). Thresholds of the N1 peak
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of the CAP were estimated visually using 10 ms tone stimuli at a repetition rate of 10
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Hz. Latencies of the N1 peak for acoustic and RW stimulation were estimated off-line
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using recording of the CAP using 50 averages.
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For acoustic stimulation sound was delivered to the tympanic membrane by a closed
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acoustic system comprising two Bruel and Kjaer 4134 ½” microphones for delivering
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tones and a single Bruel and Kjaer 4133 ½” microphone for monitoring sound
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pressure at the tympanum. The microphones were coupled to the ear canal via 1 cm
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long, 4 mm diameter tubes to a conical speculum, the 1 mm diameter opening of
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which was placed about 1 mm from the tympanum. The closed sound system was
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calibrated in situ for frequencies between 1 and 50 kHz. Known sound pressure levels
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were expressed in dB SPL re 210-5 Pa.
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A neodymium iron boron disk magnet (M in figure 1) (diameter 0.6 mm, thickness 0.2
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mm), was placed on the RW to stimulate the cochlea through the RW. The magnet
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covered about one fourth of the RW surface of total area of 1.18 mm2 [18]. A
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miniature coil (C in figure 1) made of two turns of copper wire (0.15 mm in diameter)
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was placed above the magnet. The magnet was driven with a magnetic field created
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by AC current through the coil. Stimulating current through the coil was generated by
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a Data Translation 3010 board, attenuated and fed to the coil through a current buffer.
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Maximum voltage applied to the coil in this study was 10 V which corresponded to 0
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dB attenuation. High-frequency cut-off of the system for electrical stimulation in situ
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was above 100 kHz. No signals associated with the coil current were recorded in the
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absence of the magnet which confirms an absence of electrical interference between
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the coil and the CAP electrode. The efficacy of the floating magnet RW stimulation
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varies between preparations (see below). The floating magnet does not, however,
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impose a variable DC load on the RW, which undergoes large, slow, periodic,
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movements caused by middle ear muscle contraction in anaesthetised animals. This
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would not be the case for probes that employ a rigid lever in contact with the RW,
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which would provide a variable DC load depending on the phase of the large RW
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movement, with the possibility of causing damage to the RW.
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Displacements of the magnet and the stapes were measured using a displacement-
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sensitive laser diode interferometer without the need for reflective beads [19]. The
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beam of the interferometer was focused centrally on the exposed surface of the
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magnet or the head of the stapes. The output signal of the interferometer was
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processed using a signal conditioning amplifier, digitised at 250 kHz with Data
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Translation 3010 board and instantaneous amplitude and phase of the wave were
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recorded and averaged 10 times using a digital phase-locking algorithm.
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Voltage signals to the coil were sinusoidal stimuli of 40 ms in duration (with 110 ms
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between stimuli) shaped with raised cosines of 0.5 ms duration at the beginning and at
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the end of each sinusoidal pip. All acoustic stimuli in this work were shaped with
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raised cosines of 0.5 ms duration at the beginning and at the end of stimulation. White
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noise for acoustical calibration and tone sequences for auditory and mechanical
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stimulation were synthesised by a Data Translation 3010 board at 250 kHz and
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delivered to the microphones or to the coil through low-pass filters (100 kHz cut-off
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frequency). Signals from the acoustic measuring amplifier were digitised at 250 kHz
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using the same board and averaged in the time domain. Experimental control, data
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acquisition and data analysis were performed using a PC with programmes written in
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TestPoint (CEC, MA, USA). All procedures involving animals were performed in
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accordance with UK Home Office regulations with approval from the local ethics
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committee.
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3. Results
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3.1 Pressure distribution in a model of cochlear stimulation through partially
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occluded RW
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The RW is assumed to be driven by a piston (magnet) over a part of its area and the
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remaining part of the RW is flexible and generates an equal and opposite volume
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velocity, since the fluid is assumed to be incompressible and the cochlear wall is rigid.
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The internal pressure in the cochlea is then made up of two components.
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The first component is an average alternating pressure, 𝑃𝑀 , which is the same
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throughout the cochlea. The magnitude of this will depend on the stiffness, 𝑆, of the
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freely moving area of the RW. 𝑃𝑀 is required to be large enough to force this area to
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have a volume velocity equal and opposite to that of the driving piston. The force on
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the freely moving part is thus 𝑆𝑥, where 𝑥 is its mean displacement. Its volume
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velocity, 𝑞, is thus 𝑖𝜔𝐴𝑥, where 𝐴 is the freely moving area and 𝜔 is the stimulation
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frequency. Because the force on the freely moving part is equal to 𝑃𝑀 𝐴, we can define
𝑃𝑀 =
𝑆𝑞
2
𝑖𝜔𝐴
.
(1)
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Since the average pressure is uniform throughout the volume, it does not cause any
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excitation of an incompressible cochlear partition. It is worth noting that if the
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stiffness, 𝑆, of the freely moving area of the RW is small then the mean pressure
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generated within the fluid volume is also small.
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Even if the volume velocity of the freely moving part of the RW is equal and opposite
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to that of the piston, the second pressure component, a near-field pressure, is
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generated close to the RW. This is due to the pressure required to accelerate the fluid
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from below the piston into the freely moving part of the RW and is thus dependent on
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the fluid inertia. The spatial distribution of this near-field pressure will depend on the
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details of the geometry, but its magnitude will be proportional to the fluid density, 𝜌,
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and to the acceleration of the piston, 𝑖𝜔𝑞. An indicative overall magnitude of 𝑃𝑁 can
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then be defined as
𝑃𝑁 ∝ 𝑖𝜔𝜌𝑞.
(2)
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For a given piston volume velocity, the average pressure (Eq. 1) decreases with
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increasing the stimulation frequency, whereas the average near-field pressure (Eq. 2)
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increases with the frequency.
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It is worth noting that if the cochlear wall is not perfectly rigid, for example if the
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stapes is not completely fixed, the internal pressure also generates a volume velocity
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on the oval window, as illustrated in figure 2A. This would generate a pressure
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component due to the flow of fluid across the cochlea that would depend on the fluid
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inertia rather than the stiffness of the window. A near-field pressure would then be
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generated in the vicinity of both the oval window and the RW.
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The on-axis near-field pressure, 𝑃𝑁 , due to RW stimulation can be estimated using a
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simple model of cancelling piston sources. The assumed geometry is shown in figure
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2B in which an inner piston with a radius of 𝑎1 has a velocity of 𝑣1 and an annular
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piston with an inner radius of 𝑎1 and an outer radius of 𝑎2 has a velocity of 𝑣2 .
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The volume velocity of the inner piston thus is
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𝑞1 = 𝜋𝑎21 𝑣1 ,
(3)
𝑞2 = 𝜋(𝑎22 − 𝑎21 )𝑣2 .
(4)
and that of the outer piston is
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If the cochlea is sealed, the volume velocities are equal and opposite so that 𝑞2 =
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−𝑞1, and the ratio of the two linear velocities must be
𝑣2
𝑎21
=
,
𝑣1 𝑎22 − 𝑎21
(5)
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which is plotted in figure 3.
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Assuming free field conditions, the complex on-axis pressure a distance of 𝑟 from a
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single piston of radius 𝑎 vibrating with velocity 𝑣 is given (e.g. see Kinsler et al [20])
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as
𝑝(𝑟) = −𝜌𝑐𝑣 (𝑒−𝑖𝑘√𝑟
2 +𝑎2
− 𝑒−𝑖𝑘𝑟 ),
(6)
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where 𝑘 = 2𝜋/𝜆 is the wavenumber, 𝜆 is the acoustic wavelength and 𝑐 is the speed
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of sound in the fluid. Distance between the RW and the BM, which is <1 mm in both
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guinea pigs and humans [21], is much smaller than the wavelength, which is equal to
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about 1.5m at 1 kHz assuming a speed of sound of 1500 ms-1 in the cochlear fluid.
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Hence, we can assume 𝜆 ≫ 𝑟, and
𝑘𝑟 =
2𝜋𝑟
≪ 1.
𝜆
(7)
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In this case we can take the first order series approximation to the exponentials in
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equation (6) to give the pressure due to the inner piston as
𝑝1 (𝑟) ≈ 𝑖𝜌𝜔𝑣1 (√𝑟2 + 𝑎21 − 𝑟).
(8)
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The pressure due to the annular piston is equal to that due to a piston of radius 𝑎2 and
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velocity 𝑣2 minus that due to a piston of radius 𝑎1 and velocity 𝑣1
𝑝2 (𝑟) ≈ 𝑖𝜌𝜔𝑣2 (√𝑟2 + 𝑎22 − √𝑟2 + 𝑎21 ).
(9)
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The total near-field pressure 𝑃𝑁 (𝑟) is the superposition of those due to the inner and
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outer pistons in equations (8) and (9) with 𝑣2 /𝑣1 given by equation (5)
𝑃𝑁 (𝑟) = 𝑖𝜌𝜔𝑣1 [(√𝑟2 + 𝑎21 − 𝑟) −
𝑎21
(√𝑟2 + 𝑎22 − √𝑟2 + 𝑎21 )].
𝑎22 − 𝑎21
(10)
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Although this expression is not strictly valid in the present case unless the near-field
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pressure has died away over the width of the cochlea, the modulus of equation (10) is
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plotted in figure 4 as a function of distance from the centre of the source distribution,
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𝑟, for various values of 𝑎1 /𝑎2 and with a constant value of 𝑣1 . The pressure drops
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quickly with the distance from the piston, especially for small 𝑎1 /𝑎2, but the shortest
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distance between the RW and the BM is also only 0.2 mm in guinea pigs [21]. Hence,
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excitation of the cochlea with a probe which covers just a part of the RW is possible
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by the near-field pressure in vicinity of the RW. A natural consequence of the simple
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model of cancelling piston sources is higher near-field pressure for larger 𝑎1 /𝑎2 ratios
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for constant piston velocity. In the real cochlea, however, the RW piston velocity is
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likely to drop for larger 𝑎1 /𝑎2 ratios as a consequence of an increase in cochlear
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impedance as seen by the piston. A model is needed, which takes into account this
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effect and measurements of the RW impedance for different proportions of the RW
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covered by the magnet, in order to find an optimum 𝑎1 /𝑎2 ratio which provides the
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most efficient stimulation of the cochlea through the RW. It also should be noted that
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within the validity of equation (10), the total near-field pressure 𝑃𝑁 (𝑟) is proportional
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to the acceleration of the inner piston. The BM is a pressure detector [22], it should,
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therefore, be expected that BM stimulation is proportional to piston acceleration.
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3.2 Placement of magnet on the RW does not alter the sensitivity of the cochlea to
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acoustic stimulation
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The cochlear neural threshold to acoustic stimulation, measured as a threshold for the
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N1 peak of the CAP, did not change after placement of the magnet on the RW and
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during the entire experiment (up to 3 hours after the placement) (figure 5). It is well
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established that increases in the impedance of the RW leads to elevation of the
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hearing threshold (e.g. see Culler et al [1]), thus the observed stability of the CAP
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thresholds after placement of the magnet revealed that the RW impedance did not
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increase and the area of the RW not covered by the magnet could provide effective
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pressure relief at acoustic frequencies when the cochlea is stimulated conventionally.
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The RW impedance is much lower than any other leakage impedance in the cochlea
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(e.g. see Stieger et al [23]). Thus, in view of the stability of the N1 thresholds
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following placement of the magnet, we can conclude that the RW impedance
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remained the smallest impedance in the system, even after this manipulation.
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3.3 Frequency dependence of neural thresholds in response to subnanometer RW
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and acoustic stimulation are similar
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The coil voltage - magnet displacement relationship was linear in all six preparations
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where this characteristic was studied (an example is given in figure 6). In all
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preparations, the frequency dependence of the magnet displacement resembled that of
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a low-pass filter with the high-frequency slope being between 12 - 18 dB/octave
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(figure 7A). A slope of 12 dB/octave is to be expected if the magnet is behaving as a
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mass subject to a constant force caused by the magnetic field due to the voltage. In
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this case, the magnet acceleration should not depend on frequency, which is indeed
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observed for most of the frequency range used in this study (figure 7B). Because the
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high-frequency cut-off of the system for electrical stimulation was above 100 kHz, the
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relatively steep high-frequency slope of the coil voltage - magnet displacement
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relationship for frequencies below 10 kHz (figure 7A) most likely reflects the multiple
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degrees of freedom of the mechanical system formed by the inductively coupled coil
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and the magnet inertia, together with the RW and the cochlear fluid.
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There were no major differences between CAP threshold curves for acoustic
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stimulation between all preparations used in this study (figure 7C). However, in the
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same preparations, CAP threshold curves expressed as a function of the coil voltage
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required to generate threshold CAP for magnetically driven RW stimulation varied by
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~10 dB over the 8 – 20 kHz range (figure 7D). Likely bases for these CAP threshold
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variations could include differences, between preparations, in the relative position of
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the magnet and the coil and to slightly different locations of the magnet on the RW.
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The smallest coil voltages required to generate the threshold CAP was observed for
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frequencies between 10-11 kHz. Taking into account the linearity of the magnet
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displacement (figure 6), the corresponding magnet displacements for the threshold
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coil voltage could be readily derived from the iso-voltage response curves at 10 V
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(figure 7A) and the CAP threshold curves (figure 7D) using the following equation:
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displacement at threshold = displacement at 10 V / 10^(attenuation at threshold/20).
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The derived magnet displacement threshold curves (figure 7E) reveal that the
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sensitivity of the guinea pig cochlea to RW stimulation is in the sub nanometre range.
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Maximum sensitivity, which corresponds to the smallest magnet displacements at
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threshold, was observed for frequencies around 20 kHz, which is similar to the
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frequency of maximum sensitivity for acoustic stimulation (figures 5, 7C, note
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different frequency ranges in these figures). The absolute sensitivity of the magnet
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displacement threshold curves varied between preparations (figure 7E), probably, as
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pointed out above, because of variations between preparations in the relative locations
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of the coil, the magnet and the RW membrane. Magnet displacement threshold curves
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within the range of those shown in figure 7E were obtained for three further
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preparations but, for clarity, are not shown.
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Near-field pressure, which is likely to excite the BM responses in our experiments, is
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proportional to the acceleration of the magnet (figure 7F) which, at threshold, was
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calculated from the threshold displacement (figure 7E). The threshold acceleration
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changes only by about 20 dB within the frequency range studied (figure 7F). This
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change corresponds better to the changes in the threshold SPL for acoustic stimulation
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within the same frequency range (figure 7C) while threshold displacement changes by
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more than 40 dB for the same frequencies (figure 7E). This last observation provides
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an additional confirmation that in our experiments the BM is excited by the near-field
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pressure which is proportional to the magnet acceleration.
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3.4 Latencies of neural responses to acoustic and RW stimulation are similar
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To gain insight into the mechanisms by which the cochlea is excited through RW
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stimulation, we compared the CAP latencies for suprathreshold stimulation within 10
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dB above the thresholds for acoustic and RW stimulation of the cochlea (figure 8).
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The latencies were essentially the same. The CAP latency for low-level acoustic
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stimulation depends on the traveling wave delay and is inversely related to the
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characteristic frequency of the fibre (e.g. see Goldstein et al [24]). Hence, we can
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conclude that the mechanisms of energy propagation to the characteristic frequency
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place were similar for both acoustic and RW stimulation.
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3.5 Increase in the stapes impedance does not affect efficiency of cochlear
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stimulation through the partially occluded RW
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As suggested previously, the relatively high acoustic impedance of the middle ear
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ossicles (as seen from the cochlea), combined with a magnet that only partially
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occluded the RW, may allow the region of the RW not covered by the magnet to act
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as a pressure shunt during RW stimulation. It may thus be suggested that stapes
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mobility plays little or no role in the stimulation of the cochlea through the RW.
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Indeed, when the cochlea was excited through the RW, movements of the stapes at the
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frequency of stimulation were observed above the noise floor of the interferometer
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only at low frequencies <7 kHz and high stimulation amplitudes (figures 9A-B).
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Preparations with the largest and smallest stapes responses are shown in figure 9A-B.
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Similar results were observed in two further preparations. When the stapes impedance
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was increased by filling the ear canal with superglue, and stapes displacements could
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not be detected above the noise floor of the interferometer regardless of the frequency
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and magnitude of the control voltage to the stimulating coil (figures 9A-B), thresholds
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of RW elicited CAPs before and after increase in the stapes impedance were largely
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similar (figure 9C). According to this observation it is suggested that stapes mobility
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is not required for effective excitation of the cochlea in our experimental
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configuration, i.e. when relatively large part of the RW was not occluded.
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4. Discussion
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On the basis of CAP threshold measurements at stimulus frequencies that span almost
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its entire auditory range, the guinea pig cochlea is sensitive to sub-nanometre
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displacements and minute accelerations of the RW (figure 7). The CAP data, set in the
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context of a model of cochlear stimulation through the partially occluded RW, support
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the viability of RW stimulation as an effective route for exciting the mammalian
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cochlea, further building on previous studies of the mechanical properties of the
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cochlea and recent work concerning the implementation of RW implantable hearing
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devices [13, 25, 26]. Quantification of the RW transducer displacement/acceleration
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presented here provides parameters essential for the design of future devices, although
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these parameters vary within individuals (figures 7 and 9). These factors may include
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the shape of the transducer and the area of the RW it covers. Most significant is
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whether the transducer completely covers the RW, as demonstrated by other groups
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[27-29], or leaves a RW area partially free [26, 30-32]. The outcome of our
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experiments reveals that partial coverage of the RW may not affect RW impedance
341
sufficiently to affect cochlear stimulation with the exposed portion of the RW
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membrane acting as a pressure shunt (figure 5). Complete coverage would probably
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remove this shunt, possibly resulting in a different mechanism of BM stimulation.
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Our experiments demonstrate that middle ear prostheses, which partially cover the
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RW, can be used effectively for simultaneous cochlear stimulation and pressure relief
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through the RW. The existence of a “third window” (an additional pressure shunt,
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𝒁𝟑,𝑺𝑽 , 𝒁𝟑,𝑺𝑻 , figure 10) in the cochlea has been postulated to account for the efficiency
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of RW stimulation because of the relatively high impedance, as seen from the cochlea,
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presented by the ossicles [30]. The proposed identity of the third window ranges from
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the vasculature of the cochlea to the cochlea and vestibular aqueducts [17, 30].
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Evidence for the existence of the third window and its ability to shunt pressure is
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claimed from experiments in which the RW membrane is completely blocked during
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normal acoustic stimulation or the stapes is immobilized and the cochlea is stimulated
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through the RW route. Despite blockage or fixation, the air conduction thresholds are
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raised by between 20-50 dB depending on the study cited [3, 30, 33-35]. However,
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even the comparatively low experimental value of a 20 dB rise in threshold level is
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significant when compared with the low thresholds achieved in the current study
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(figure 9). Additionally, it is unlikely that a hypothetical third window is responsible
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and, indeed, required for pressure relief in our experiments. The RW impedance
360
(𝒁𝑹𝑾 , figure 10A) is much lower than any other leakage impedance in the cochlea
361
during normal acoustic stimulation (e.g. see Stieger et al [23]), i.e. 𝒁𝑹𝑾 ≪ 𝒁𝟑,𝑺𝑻 and
362
pressure at point 𝑁 is determined mainly by 𝒁𝑹𝑾 . Moreover, the neural thresholds are
363
highly sensitive to increases in RW impedance (e.g. see Culler et al [1]). In our
364
experiments, the neural thresholds during acoustic stimulation did not change after the
365
magnet placement (figure 5). We took this finding to indicate that the RW impedance
366
remained the smallest impedance in the system even after this manipulation (i.e.
367
𝒁𝑹𝑾 ≪ 𝒁𝟑,𝑺𝑻 after the magnet placement) and provided an effective pressure shunt
368
minimising the pressure at point 𝑁 (figure 10B) and the far-field pressure drop across
369
the impedance of the cochlear partition (𝒁𝑩𝑴 , figure 10B). This outcome might also
370
be expected from the direct finding that intracochlear pressure differences between the
371
scala vestibuli and scala tympani remained low during RW stimulation [23]. On the
372
basis of our findings reported here, we propose that the pressure shunting occurred
373
through the area of the RW which remained exposed. This proposal is also supported
374
by the findings of Schraven et al [32] who tested different sized actuator tips and their
375
coupling to the RW. Schraven et al [32] found that stimulation by actuators with tip
376
sizes well below the dimensions of the RW resulted in reduced stapedial movements
377
compared with those tips that covered most of the RW. This result could be accounted
378
for if the free exposed area of the RW acted as a pressure shunt. This does not
379
discount other “windows” contributing to pressure relief but it is clear from results
380
presented here that their contribution appears to be negligible in our experiments.
381
We found that CAP thresholds did not change after stapes impedance (𝒁𝑴𝑬 , figure
382
10B) was increased (figure 9). CAP threshold elevation was, however, observed in
383
earlier studies using RW stimulation through a partially blocked RW [30].
384
Furthermore, where we found RW stimulation through the partially occluded RW
385
elicited stapes displacements only at high intensities below 7 kHz, under similar
386
stimulus conditions, Lupo et al [30] observed stapes movement in chinchillas across a
387
wider frequency range of 0.25 -16 kHz. We suggest that tighter hydromechanical
388
coupling between the RW transducer and the stapes, and consequent changes in the
389
hydromechanical properties of the cochlea after stapes immobilisation, may account
390
for elevation of the CAP thresholds after stapes fixation observed by Lupo et al [30].
391
Correspondingly, differences in transducer shape and the proportion of the RW
392
covered, may potentially account for differences in the transducer-stapes
393
hydromechanical coupling reported in this study and by Lupo et al [30].
394
Pre-tension of the RW, which can potentially increase stiffness of the exposed area of
395
the RW, and hence the average pressure within the cochlear (Eq. 1), can enhance
396
stapedial movement during RW stimulation [31, 36]. It is apparent that differences in
397
the outcomes of different studies, including ours presented here, depend on a number
398
of factors including the choice of both the size and shape of the transducer and pre-
399
tension of the RW. The finding by Stieger et al [23] that the stapes velocity is not a
400
good measure of the effectiveness of reverse stimulation of the cochlea, might reflect
401
sensitivity of the stapedial responses to variations in the parameters of RW
402
stimulation mentioned above. Control of these parameters is important to ensure that
403
mobility of the ossicular chain is reduced to levels that negate the need for additional
404
intrusive procedures that might cause sensorineural hearing loss.
405
In our experiments the RW provided an effective common pathway for both cochlear
406
stimulation and pressure relief. Therefore, we conclude that the mode of cochlear
407
excitation through the RW in this case is different from that observed during
408
conventional, acoustical cochlear stimulation or RW stimulation, which does not
409
allow pressure relief through the RW [28, 29]. Our results together with direct
410
measurements of the pressure in the cochlea during RW stimulation [23] using a
411
transducer that leaves the RW partially exposed, indicate that under this condition the
412
transducer displacement probably is not able to cause a far-field pressure difference
413
between the cochlear scalae, which is the normal stimulus for the cochlea [12]. Far-
414
field pressure differences between the scalae are unlikely to be caused by the type of
415
RW stimulation we employed because of the high impedance of the ossicular chain,
416
as seen from the cochlea [27, 28], and pressure backflow around the transducer [11].
417
We propose, instead, that during RW stimulation through the partially occluded RW,
418
the BM is stimulated via the near-field complex pressure generated in the vicinity of
419
the RW (figure 4), i.e. a fluid-jet flow, which, however, results in the generation of
420
conventional travelling waves along the cochlear partition. In this sense, the cochlear
421
excitation due to the RW stimulation in our experiments is similar to excitation due to
422
rocking movement of the stapes, which creates only local pressure gradients/fluid
423
motion in the cochlea [37, 38] resulting, nevertheless, in generation of the travelling
424
waves and neural excitation [39, 40]. In our experiments, the near-field pressure is
425
proportional to the acceleration of the transducer which allows effective cochlear
426
stimulation even at extremely small RW displacements at high frequencies. This
427
conclusion should be taken into account during the design of hearing-aid devices
428
which stimulate the cochlea through the partially occluded RW. The near-field
429
pressure variations will be limited to the vicinity of the RW membrane (figure 4), but
430
the close spatial relationship between the RW and the BM [21] make excitation of the
431
BM by these localised near-field pressure variations possible. Once the BM is
432
stimulated in the vicinity of the RW, a conventional travelling wave is generated
433
along the cochlear partition, which is supported by our CAP recordings. This
434
conclusion is based on our finding that the 12-25 kHz region of the cochlea, which is
435
most sensitive to RW magnet stimulation (figure 7E,F), is located in the middle of the
436
basal turn of the cochlea and this region is not adjacent to, or cannot be observed
437
through the RW. Indeed, the CAP latency for low-level acoustic stimulation, which
438
depends on the travelling wave delay (e.g. see Goldstein et al [24]), is similar for both
439
acoustic and RW stimulation (figure 8). Furthermore, CAP threshold tuning curves
440
derived through acoustic (figures 5 and 7C) and RW magnet (figure 7E,F)
441
stimulations have similar maxima in sensitivity between 12-25 kHz.
442
It should be noted that it is not essential for the excitation that the motion of the
443
magnet is strictly piston-like. It is likely that the magnet underwent some rocking
444
motion in our experiments. This possibility, however, does not affect our conclusion
445
about the mode of cochlear excitation and, in terms of the model (figure 2B), simply
446
corresponds to a different ratio 𝑎1 /𝑎2 which should still provide effective stimulation.
447
Our investigation of the mechanisms underlying cochlea stimulation through the RW
448
using a transducer that partially covers the RW membrane has important significance
449
for future clinical use of such devices. RW stimulation does not even require mobility
450
of the ossicular chain, as demonstrated in this study, to provide input to the cochlea
451
that is similar to that obtained acoustically in a ‘normal’ ear. RW stimulation does,
452
however, involve a novel mechanism of generating travelling waves along the
453
cochlear partition, which we believe has not previously been considered. Travelling
454
waves are initiated as a consequence of near-field pressure in the immediate vicinity
455
of the RW, rather than far-field pressure differences between the scalae vestibule and
456
tympani.
457
ACKNOWLEDGMENTS
458
This work was supported by the Medical Research Council. We thank J. Hartley for
459
technical assistance.
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Maier H, Salcher R, Schwab B, Lenarz T. 2013 The effect of static force on
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Edom E, Obrist D, Henniger R, Kleiser L, Sim JH, Huber AM. 2013 The
effect of rocking stapes motions on the cochlear fluid flow and on the basilar
membrane motion. J. Acoust. Soc. Am. 134, 3749-3758.
39.
Huber AM, Sequeira D, Breuninger C, Eiber A. 2008 The effects of complex
stapes motion on the response of the cochlea. Otol. Neurotol. 29, 1187-1192.
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Eiber A, Huber AM, Lauxmann M, Chatzimichalis M, Sequeira D, Sim JH.
2012 Contribution of complex stapes motion to cochlea activation. Hear. Res. 284,
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571
FIGURE LEGENDS
572
Figure 1. Schematic view of the cochlea (CH) through a lateral opening in the
573
temporal bone. M indicates a neodymium iron boron disk magnet (diameter 0.6mm,
574
thickness 0.2mm) placed on the surface of the round window (RW). C labels a
575
miniature coil situated above the magnet. S denotes a Teflon-coated silver wire used
576
for recording of activity of the auditory nerve. IS indicates incudostapedial joint.
577
Figure 2. Pressure distribution during cochlear stimulation through partially occluded
578
RW. A Pressure distribution in presence of a mobile stapes. The effect of the mean
579
internal pressure on an imperfectly fixed stapes is to generate a small volume velocity,
580
−𝛼𝑞 , where 𝛼 < 1. In this case, the freely moving area of the RW generates a
581
volume velocity of −(1 − 𝛼)𝑞 and these two components cancel the volume velocity,
582
𝑞, of the driving piston on the RW. Then the internal pressure is the sum of that due to
583
a fixed oval window and that due to two pistons with volume velocities +𝛼𝑞 and −𝛼𝑞
584
on either side of the volume. B Sketch of cancelling piston sources.
585
Figure 3. Dependence of the ratio of the two linear velocities, 𝑣1 /𝑣2 , on the ratio of
586
inner piston to outer piston radii, 𝑎1 /𝑎2.
587
Figure 4. The modulus of the total near-field pressure, |𝑃𝑁 |, as a function of the
588
distance from the piston for different ratio of inner piston to outer piston radii, 𝑎1 /𝑎2.
589
The inner piston velocity, 𝑣1 , is 1 mm/s, the fluid density is 103 kg/m3, the frequency
590
of stimulation is 15 kHz. Cartoons along the 𝑎1 /𝑎2 axis show position of the magnet
591
(grey area) on the RW for different relative probe sizes.
592
Figure 5. Thresholds of acoustic stimulation (dB SPL) vs frequency (kHz) for N1
593
peak of the CAP of the auditory nerve measured before (open symbols) and 2.5 hours
594
after (solid symbols) placement of the magnet on the RW.
595
Figure 6. Displacement of the RW magnet as a function of voltage applied to the coil
596
for three different stimulus frequencies (5, 10 and 20 kHz) measured in situ on the
597
RW of a single preparation. The data at each frequency are pooled and normalized
598
from measurements at different gains of the signal conditioning amplifier.
599
Figure 7. Magnet displacement and neural thresholds as a function of the stimulation
600
frequency in three representative preparations. A Frequency dependence of the
601
magnet displacement at constant voltage (10 V) applied to the coil. B Frequency
602
dependence of the magnet acceleration at constant voltage (10 V) applied to the coil
603
calculated from data at panel A. C CAP thresholds during acoustic stimulation. Note a
604
shorter frequency range than in figure 2. D Dependence of the coil voltage to generate
605
threshold CAP during the RW stimulation. E Corresponding magnet displacement at
606
the CAP threshold. F Corresponding magnet acceleration at the CAP threshold
607
calculated from data at panel E.
608
Figure 8. CAP latency for suprathreshold stimulation within 10 dB of the threshold
609
for the acoustic and RW stimulation of the cochlea (mean  SD, N=3 for each
610
condition of stimulation).
611
Figure 9. Effect of stapes fixation on the CAP thresholds during the RW stimulation.
612
A, B Stapes displacements for different attenuation of voltage (re 10V) applied to the
613
coil. Average of three measurements. Error bars are omitted for clarity. Noise floor of
614
the interferometer is indicated by a dashed line. C Corresponding CAP thresholds
615
measured before (solid symbols) and after (open symbols) stapes fixation.
616
Figure 10. Schematics of impedances of the cochlea during acoustic (A) and RW (B)
617
stimulation. 𝒁𝑩𝑴 and 𝒁𝑹𝑾 are impedances of the cochlear partition and the open area
618
of the RW respectively. 𝒁𝑴𝑬 is the stapes impedance as seen from the cochlea. 𝒁𝟑,𝑺𝑽
619
and 𝒁𝟑,𝑺𝑻 are impedances of “third windows” in the scala vestibuli and scala tympani,
620
respectively. 𝒒 indicates a volume velocity generated by stimulation.
21
Figure 1
IS
C
M
RW
S
CH
22
Figure 2
A
B
23
Figure 3
1
-v2/v1
10
-1
10
-3
10
0.0
0.2
0.4
0.6
a1/a2
0.8
1.0
24
Figure 4
25
CAP threshold (dB SPL)
Figure 5
30
before magnet placement
2.5 hours after
20
10
0
-10
rw2
-20
1
10
Tone frequency (kHz)
26
Displacement amplitude (AU)
Figure 6
10
1dB/dB
1
5kHz
0.1
10kHz
0.01
20kHz
rw3
50
40
30
20
10
Attenuation (dB re 10V)
0
27
Magnet displacement (nm)
100
A
10V input
12dB/oct
10
18dB/oct
1
rw2
rw6
rw17
10
Sound pressure (dB SPL)
20
30
Frequency (kHz)
C
10
0
-10
rw2
rw6
rw17
-20
5
1
Magnet displacement (nm)
20
10
15
20
25 30
Frequency (kHz)
E
0.1
0.01
rw2
rw6
rw17
1E-3
5
10
15
Frequency (kHz)
20
25 30
Magnet acceleration (m/sec^2) Threshold voltage (-dB re 10 V) Magnet acceleration (m/sec^2)
Figure 7
120
B
80
40
rw2
rw6
rw17
0
10
30
D
20
30
20
25 30
20
25 30
Frequency (kHz)
40
50
60
rw2
rw6
rw17
70
5
10
15
Frequency (kHz)
F
1
0.1
rw2
rw6
rw17
0.01
5
10
15
Frequency (kHz)
28
Figure 8
CAP latency (ms)
2.5
2.0
1.5
magnet stimulation
air stimulation
1.0
1
10
Frequency (kHz)
29
Figure 9
Displacement (nm)
10
A
noise
attenuation
30att
20att
10att
0att
1
0.1
1
10
Frequency (kHz)
10
Displacement (nm)
B
noise
attenuation
30att
20att
10att
0att
1
0.1
Threshold voltage (-dB re 10 V)
1
30
10
Frequency (kHz)
C
stapes mobile
stapes fixed
stapes mobile
stapes fixed
40
50
60
70
5
10
15
Frequency (kHz)
20
25 30
30
Figure 10
A
q
Z3,SV
ZBM
ZRW
N
Z3,ST
B
Z3,SV
ZME
ZBM
ZRW
N
Z3,ST
q