At least partially implantable sound pick-up device with ultrasound emitter

There is provided an at least partially implantable device for picking up sound impinging onto a skin area of a person, comprising means for generating an audio signal corresponding to the change in time of the distance between a position of the device and the outer surface of the skin area, wherein the device position is adjacent to the skin area.

The invention relates to an at least partially implantable microphone, in particular of a hearing aid.

Fully implantable hearing aids require bio-compatibility of all components due to the need of implanting all components of the device. This applies, in particular, also to the sound input transducer, which usually is a microphone.

Conventional hearing aid microphones are designed to have an acoustic impedance similar to air in order to reduce reflection losses and to obtain high sound sensitivity, low noise and low vibration sensitivity. To provide the possibility for implantation of a microphone into the body two different approaches are known: The first approach is to provide a implanted sensor, such as displacement sensor, a velocity sensor (U.S. Pat. No. 6,636,768), an acceleration sensor (US 2005/0137447 A1), an electric sensor or a hydrostatic sensor (U.S. Pat. No. 6,473,651B1) as a sound pick-up means at the ossicular chain, the tympanic membrane (U.S. Pat. No. 6,554,761 B1) or inside the cochlea (U.S. Pat. No. 5,782,744). The second approach is to build a hermetic microphone suitable for locations under the skin (U.S. Pat. No. 5,859,916, U.S. Pat. No. 6,516,228, U.S. Pat. No. 5,814,095) or the mucosa of the middle ear (U.S. Pat. No. 6,216,040, U.S. Pat. No. 6,636,768). Both approaches are currently used in practice, but involve design-specific problems.

Sound sensors for implantable hearing aids at the ossicular chain or tympanic membrane suffer from the significant drawback that they are acoustically connected to the output (for example, a transducer at the ossicular chain or at the cochlear). In order to avoid output being coupled back into the pick-up means (i.e. to avoid feedback), the ossicular chain has to be interrupted, which may result in a permanent damages for the patient. It is to be noted that the feedback issue is not relevant to hearing aids having a non-mechanical actuator, such as cochlear implants.

Subdermal microphones are mostly derived from conventional designs and have a bio-compatible housing with an inert microphone membrane (U.S. Pat. No. 6,422,991 B1, U.S. Pat. No. 6,093,144, U.S. Pat. No. 6,626,822 B1). Piezo- and electrodynamic mechanical transduction principles and, more rarely, mechano-optical conversion (US 2007/0161848 A1), have been suggested. In any case, the acoustic reflective losses of about 55 dB at the air/tissue interface and the mass loading on the microphone membrane by the overlying skin have to be compensated for. Even in applications with minimal skin thickness, for example, when placing the microphone in the outer ear canal wall (U.S. Pat. No. 6,516,228, U.S. Pat. No. 6,381,336, U.S. Pat. No. 5,814,095), the mass loading by skin is by several magnitudes higher than for conventional microphone membranes when used for sound pick-up in air. Due to the lower sensitivity caused by significant reflections, implanted microphones must have larger integration surfaces (WO 2005/046513 A2) and larger size in order to lower the noise level (WO 02/49394 A1, WO 2007/008259 A2). In some applications, corresponding closed volumes are used to increase the amplitude at the implanted microphone (U.S. Pat. No. 6,736,771). In addition, the mass loading of the overlying skin will be subject to normal biological changes like temperature-induced thickness changes, blood flow and muscular activity.

On the other hand, the skin is suspended by the microphone membrane from below, and accelerations of the body will lead to significant artificial amplitudes in the audio signals produced by the implanted microphone; this effect is taken into account by some designs (US 2006/0155346 A1, US 2005/0197524 A1). It has been proposed to use soft tissue placement with a movable microphone position in order to provide for less acceleration-induced relative movements between the skin surface and the microphone membrane, compared to fixation at a bone (WO 2007/001989 A2).

It is an object of the invention to provide for an at least partially implantable sound pick-up device which has only little sensitivity to body acceleration and which is small and easy to implant. It is a further object to provide for an at least partially implantable hearing aid comprising such implantable sound pick-up device. It is a further object to provide for a corresponding sound pick-up method.

According to the invention, these objects are achieved by a sound pick-up device as defined in claim1, a hearing aid as defined in claim16and a sound pick-up method as defined in claim18, respectively.

The present invention is beneficial in that, by generating an audio signal corresponding to the change in time of the distance between the position of the device and the outer surface of a skin area adjacent to the device position the need of a subcutaneous microphone membrane is eliminated, whereby the impact of body acceleration on the audio signal and the size of the device can be reduced; also, the lower size makes implantation easier.

Preferred embodiments of the invention are defined in the dependent claims.

According to the invention, sound impinging onto a skin area of a patient is picked-up by generating an audio signal corresponding to the change in time of the distance between a position of the device and the outer surface of the skin area, wherein the device position is adjacent to the skin area.

According to one embodiment, an ultrasound signal is emitted towards the outer surface of the skin area from an ultrasound emitter fixed to a bone or in soft tissue, and an ultrasound signal reflected at the outer surface of the skin area is received by an ultrasound sensor fixed to a bone or in soft tissue. Preferably, the audio signal is generated as an output signal which is proportional to the velocity of the outer surface of the skin area, as detected by analyzing the reflected ultrasound signal. This principle is schematically shown inFIGS. 1 and 2.

According toFIG. 1, an ultrasound emitter10which is fixed on an underlying bone12or in soft tissue emits a frequency modulated or constant frequency sine wave14towards the skin surface16which is impressed by outer audible sound waves17in the air and thus acts as a low-compliant microphone membrane to modulate and reflect the incident ultrasound wave14. The reflected and thereby modulated ultrasound wave18is received by ultrasound sensors20which are likewise fixed on the underlying bone12or in soft tissue. The velocity of the reflecting skin surface16can be extracted by using an interferometer or a heterodyne interferometer method, as will be explained in more detail by reference toFIGS. 2 and 3, respectively. The device22ofFIG. 2serves to measure the sound-evoked velocity of the skin surface16by ultrasound reflection using a interferometer principle, thereby avoiding the need for a subcutaneous microphone membrane. When sound impinges on the skin, most of the intensity is reflected at the surface due to the pronounced impedance difference between air and tissue. Due to the specific impedance for tissue and the specific impedance of air, the loss in sound transmission is approximately 55 dB, and the sound reflection for perpendicular incidence is almost 100%. The amount of reflection only depends on the impedance difference and is independent of the direction. Hence, ultrasound from the tissue side is reflected with high efficiency to the interior at the skin surface.

The external sound impinging on the skin surface causes an indention of the skin, which is a relatively small effect requiring an adequate measurement technique. The skin velocity resulting from hearing aid relevant sound pressure levels can be estimated, for example, to be about 1 μm/s for a sound pressure level of 100 dB and to be 0.1 nm/s for a sound pressure level of 20 dB. The detector uses the fact that the Doppler frequency shift for a stationary emitter and a moving reflector having a velocity ν is proportional to the surface velocity. More precisely, Doppler frequency shift is given by Δf=2ν/2λ0, where λ0is the wavelength. Consequently, the resolution increases with increasing sound frequencies (corresponding to decreasing wavelengths).

The device22ofFIG. 2comprises a signal generator24which drives an ultrasound emitter10in such a manner that it emits ultrasound waves at a constant carrier frequency f0. The ultrasound wave14is reflected at the skin surface16which moves at a velocity v. The reflected ultrasound wave18has a frequency which is modulated by the vibration velocity of the skin surface16by 2v(t)/λ0. The modulated ultrasound wave18is detected by an ultrasound sensor20. The output signal of the sensor20undergoes band pass filtering in a band pass26and thereafter is demodulated in a mixer/demodulator28which is fed by the signal generator24with the demodulator reference. The output signal of the mixer/demodulator28undergoes low pass filtering in a low pass30. The elements24,26,28and30form an audio signal unit36which creates an output signal which is proportional to the actual skin velocity v and hence can be used as a microphone signal for audio signal processing in a hearing aid. The required modulation band width can be estimated as 4fskinwhere fskinis the vibration frequency of the skin surface16. Various demodulation techniques can be used, such as analogue demodulation, phase locked loop (PLL) demodulation and digital demodulation utilizing digital signal processing (DSP) techniques. Velocity resolutions and noise may be optimized sufficiently in order to obtain relative resolutions far below the ultrasound wavelength by integration.

For example, an ultrasound frequency f0of 40 MHz and a travel distance of 2 cm from the emitter10to the skin surface16back to the receiver20can be assumed. Assuming a sound velocity in tissue as 1,600 m/s leads to a wavelength λ0=40 μm. Assuming the damping coefficient as αskin=0.5 leads to an attenuation of 40 dB at the receiver site. This damping also restricts the vibration sensitive skin area to a reasonable size and reduces reflection effects from other sites in the head (preferably, the sound pick-up device22will be located in the patient's head). Damping of the reflected ultrasound wave18by transmission to the air side is negligible.

Increasing the carrier frequency will result in better resolution and advantages for filtering, while the amplitude of the reflected wave18will decrease. It depends on the specific geometry and skin thickness whether such trade-off in reflective amplitude is tolerable. Increasing the carrier frequency also will allow reducing the size of the transducers10,20, whereby implantation is facilitated. Probably the size could be reduced to such an extent that minimal invasive implantation, for example, by syringe needle application, is enabled. Such reduced size devices may allow realizing arrays for directed emission and taped delay lines for directional hearing.

Preferably, the carrier frequency f0is between 10 MHz and 100 MHz to increase resolution and reduce crosstalk between multiple implanted microphones of the mentioned type.

InFIG. 3an alternative embodiment is shown which uses a heterodyne interferometer method for extracting the skin velocity. While in the interferometer method inFIG. 2ultrasound waves of constant frequency are emitted, in the embodiment ofFIG. 3the constant frequency carrier signal generated by the signal generator24is frequency modulated by a modulator32at a modulation frequency fM, which modulated signal is supplied to the ultrasound emitted10in order to emit frequency modulated ultrasound waves rather than constant frequency ultrasound waves. Accordingly, the demodulator28is supplied with the signal of the modulator32(rather than with the signal of the signal generator24) as the demodulation reference. In the embodiment ofFIG. 3the required modulation band width can be estimated as 2(fM+2fskin) where fskinand fMare the vibration frequency of the skin surface16and the modulation frequency.

Preferably, the band pass filter26blocks frequencies differing from the carrier frequency f0by more than 4fskin(for the interferometer principle ofFIG. 2) or 2(fM+2fskin) (for the heterodyne principle ofFIG. 3).

FIG. 4is a schematic block diagram of an example of a fully implantable hearing aid using an implantable sound pick-up device100according to the invention. The hearing aid comprises an implantable sound pick-up device100, an implantable audio signal processing unit50, an implantable power receiving coil52, an implantable power management unit54including a rechargeable battery, and an implantable actuator56. The audio signals picked up by the implantable sound pick-up device100are supplied to the audio signal processing unit50which converts the audio signals into a signal for driving the actuator56which stimulates the patient's hearing according to the sound picked up by the implantable sound pick-up device100. The actuator56may be, for example, a cochlear electrode or an electromechanical transducer acting on the ossicular chain or directly on the cochlea.

The power receiving coil52receives power from an external charging device58comprising a power transmission coil60via an inductive transcutaneous power link (typically, the external charging device58may be worn at night to recharge the implantable battery of the power management unit54).