Cochlear implant hearing aid system

A cochlear implant hearing aid system is disclosed. The system comprises an external unit configured to receive acoustic sound and process the acoustic sound into a coded audio signal, and an implantable unit configured to receive the coded audio signal. Further, the external unit comprises a power supply unit connected via a first path to a switching unit, wherein the switching unit is connected via a second path to ground and is connected at an output thereof to a first coil. Furthermore, the switching unit is configured to operate as a switching element configured to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil. The coded audio signal is supplied to the switching unit as a control signal, and the first coil is inductively linked to a second coil arranged in the implantable unit. The system further comprises a measuring unit connected to at least one of the first and second paths and configured to measure a dissipative current occurring in relation to the switching states of the switching unit. Based on the at least one measured dissipative current, a resonance frequency of the inductive link between the first and second coils is adapted.

FIELD

The present disclosure relates to a cochlear implant hearing aid system. More particularly, to such a system equipped with a dissipative current measuring unit, and to a method for a cochlear implant hearing aid system with a dissipative current measuring unit.

BACKGROUND

Cochlear implants (CIs) are devices containing electrodes inserted in the inner ear (the cochlea) to recover the sensation of audition to people suffering from severe to profound hearing loss. CIs are bypassing most of the functional hearing chain, and generate a series of electrical pulse train inside the cochlea to initiate action potentials from the hair cells. Those devices are thus mostly considered as biocompatible electronic machines. Depending on their implementation, they can either be totally implanted, or composed of two main parts. A first part is the sound processor, often placed near the ear. It contains microphones that capture the environmental sound, which is processed in real time into a series of codes usable by the second part, implanted into the patient. The implant receives both power and sound information though radiofrequency from the sound processor, and generates electrical pulses sent into the cochlea via electrodes inside the cochlea.

Reducing power losses occurring in the sound processor defines a topic of high interest, but is as such often quite difficult to achieve. However, prior to reducing such power losses, identification thereof is often an even bigger challenge.

Therefore, there is a need to provide a solution that addresses at least some of the above-mentioned problems.

SUMMARY

According to an aspect, a cochlear implant hearing aid system is disclosed. The system comprises an external unit configured to receive acoustic sound and process the acoustic sound into a coded audio signal, and an implantable unit configured to receive the coded audio signal. Further, the external unit comprises a power supply unit connected via a first path to a switching unit, wherein the switching unit is connected via a second path to ground and is connected at an output thereof to a first coil. Furthermore, the switching unit is configured to operate as a switching element configured to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil. The coded audio signal is supplied to the switching unit as a control signal, and the first coil is inductively linked to a second coil arranged in the implantable unit. The system further comprises a measuring unit connected to at least one of the first and second paths and configured to measure a dissipative current occurring in relation to the switching states of the switching unit. Based on the at least one measured dissipative current, a resonance frequency of the inductive link between the first and second coils is adapted.

This allows for adapting the resonance frequency (herein below also referred to as tuning of the switching unit) based on measured dissipative currents.

Particularly, such cochlear implant hearing aid system further allows for measuring a tuning/mistuning (no occurrence/occurrence of dissipative current) of the switching unit. As a result thereof, by usage of one measurement, it is directly derivable which type of first and second coils (antennas) to use for establishing the inductive link between the external unit and the implantable unit, thereby for example addressing patients with different skin thicknesses by distinguishing e.g. between skin thickness ranges of 1 mm to 6 mm and 4 mm to 11 mm.

Specifically, the cochlear implant hearing aid system further allows for monitoring an adaptation of the resonance frequency, and even further allows for continuously monitoring the adaptation during use of the cochlear implant hearing aid system, which is unique for a cochlear implant hearing aid system. This leads to even further improve power efficiency of the system, since the inductive link becomes more efficient due to the ability of continuously adapting the resonance frequency (keeping the resonance frequency of the first coil aligned or almost aligned to a target resonance frequency). In particular, the resonance frequency of the first coil is adapted in order to minimize the at least one measured dissipative current.

As result of continuously monitoring a resonance frequency adaptation, adaptation of the resonance frequency (tuning of the switching unit) may be performed in real time, allowing for reducing power losses even further.

Furthermore, the first path may further comprise a shunt resistor, and the measuring unit may further be configured to measure the dissipative current across the shunt resistor. Moreover, the measuring unit may further comprise an amplification element configured to amplify a current to be measured by the measuring unit.

Optionally, the second path may further comprise a third coil inductively coupled to a fourth coil, and the measuring unit comprises the fourth coil and may further be configured to measure a current induced in the fourth coil and to derive the dissipative current therefrom. Optionally, the measuring unit is further configured to comprise an amplification element configured to amplify the current induced in the fourth coil. Furthermore, the measuring unit may further be configured to comprise a resistor connected to an output of the amplification element and a capacitor connected at one out of its terminals to the connection between the output of the amplification element and the resistor and at the other one of its terminals to ground.

Such measuring unit configurations as outlined above each allow for simplicity of a design of the measuring unit, by exploiting existing components of the switching unit and/or the cochlear implant hearing aid system. Particularly, by using the existing components, application of new components to the switching unit/cochlear implant hearing aid system is therefore avoidable, which leads to a more complex design of the measuring unit.

In addition, this further allows for avoiding that the measuring unit has an impact on the switching unit adapting the resonance frequency, as well as for avoiding that the measuring unit creates a parasitic resistance to the circuitry constituting the switching unit/cochlear implant hearing aid system.

Moreover, the switching unit may for example be an E-Class amplifier.

This allows for theoretically 100% efficiency of the switching unit, which means no power losses and, therefore, no dissipative current in the transistor constituting the E-Class amplifier.

According to at least an implementation, the resonance frequency of the inductive link may be adapted based on the at least one measured dissipative current by selectively connecting at least one further switched capacitor to a circuitry constituting the switching unit.

Optionally, according to a further implementation, the switching unit may further comprise at least one of a first switched capacitor and a second switched capacitor, wherein the first switched capacitor is connected in parallel to a first circuitry element, and the second switched capacitor is connected in serial to a second circuitry element. Further, the resonance frequency is adapted based on the at least one measured dissipative current by selectively connecting at least one of the first and second switched capacitors to the circuitry constituting the switching unit.

Furthermore, the switching unit may further comprise a transistor element, wherein the transistor element is connected at an output thereof via the first circuitry element to ground, and via the second circuitry element to the first coil. Further, the first circuitry element may be a third capacitor, and the second circuitry element may be a fourth capacitor.

Applying a switched capacitor for adapting the resonance frequency during the above mentioned monitoring process allows for adjusting a tuning of the switching unit in real time.

According to another aspect, a method for a cochlear implant hearing aid system comprising an external unit receiving acoustic sound and processing the acoustic sound into a coded audio signal and an implantable unit receiving the coded audio signal is disclosed. The method comprises the steps of supplying in the external unit power from a power supply unit via a first path to a switching unit, wherein the switching unit is connected via a second path to ground and is connected at an output thereof to a first coil. Further, the switching unit operates as a switching element to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil. The coded audio signal is supplied to the switching unit as a control signal, and the first coil is inductively linked to a second coil arranged in the implantable unit. The method further comprises the steps of measuring across at least one of the first and second paths a dissipative current occurring in relation to the switching states of the switching unit. The method further comprises adapting a resonance frequency of the inductive link between the first and second coils based on the at least one measured dissipative current.

This allows for a method for adapting the resonance frequency based on measuring at least one dissipative current.

Particularly, as derivable from above, such method for a cochlear implant hearing aid system further allows for measuring a tuning/mistuning of the switching unit and further allows for monitoring a resonance frequency adaptation, especially during use of the cochlear implant hearing aid system, which is unique for a cochlear implant hearing aid system. Such method therefore leads to even further improving power efficiency of the system, since the inductive link becomes more efficient due to continuously adapting the resonance frequency, which allows for real time resonance frequency adaptation, allowing for reducing power losses even further.

In addition, the method may further comprise the steps of measuring a first dissipative current across a shunt resistor contained in the first path, and/or measuring a second dissipative current based on a current induced in a fourth coil inductively coupled to a third coil, wherein the third coil is comprised in the second path.

This allows for a method for measuring a dissipative current, by exploiting existing components of the switching unit and/or the cochlear implant hearing aid system. Particularly, by applying the method on the existing components, application of new components to the switching unit/cochlear implant hearing aid system is therefore avoidable, which leads to a more simplified measurement.

This further allows for a method for avoiding the measurement having an impact on a tuning of the switching unit, as well as for avoiding the measurement creating a parasitic resistance to the circuitry constituting the switching unit/cochlear implant hearing aid system.

Furthermore, the method may further comprise the steps of deriving a weighted dissipative current from the measured first and second dissipative currents based on applying a predetermined weighting algorithm. Further, the method comprises adapting the resonance frequency of the inductive link based on the derived weighted dissipative current.

This allows for increasing accuracy in resonance frequency adaptation (in tuning of the switching unit).

Moreover, the method may further comprise the steps of adapting the resonance frequency of the inductive link based on the at least one measured dissipative current by selectively connecting at least one further switched capacitor to a circuitry constituting the switching unit.

Optionally, the switching unit further comprises at least one of a first switched capacitor and a second switched capacitor, wherein the first switched capacitor is connected in parallel to a first circuitry element, and the second switched capacitor is connected in serial to a second circuitry element. The method may further comprises the steps of adapting the resonance frequency of the inductive link based on the at least one measured dissipative current by selectively connecting at least one of the first and second switched capacitors to the circuitry constituting the switching unit. The switching unit may be an E-Class amplifier and further comprises a transistor element, wherein the transistor element is connected at an output thereof via the first circuitry element to ground, and at an output thereof via the second circuitry element to the first coil. Further first circuitry element may be a third capacitor and the second circuitry element may be a fourth capacitor.

Applying a switched capacitor for adapting the resonance frequency during the above mentioned monitoring process allows for a method for adjusting a tuning of the switching unit in real time.

DETAILED DESCRIPTION

The electronic hardware may include microprocessors, microcontrollers, digital signal processors (DSPs), field programmable gate arrays (FPGAs), programmable logic devices (PLDs), gated logic, discrete hardware circuits, and other suitable hardware configured to perform the various functionality described throughout this disclosure. Computer program shall be construed broadly to mean instructions, instruction sets, code, code segments, program code, programs, subprograms, software modules, applications, software applications, software packages, routines, subroutines, objects, executables, threads of execution, procedures, functions, etc., whether referred to as software, firmware, middleware, microcode, hardware description language, or otherwise.

A hearing device may include a hearing aid that is adapted to improve or augment the hearing capability of a user by receiving an acoustic signal from a user's surroundings, generating a corresponding audio signal, possibly modifying the audio signal and providing the possibly modified audio signal as an audible signal to at least one of the user's ears. The “hearing device” may further refer to a device such as an earphone or a headset adapted to receive an audio signal electronically, possibly modifying the audio signal and providing the possibly modified audio signals as an audible signal to at least one of the user's ears. Such audible signals may be provided in the form of an acoustic signal radiated into the user's outer ear, or an acoustic signal transferred as mechanical vibrations to the user's inner ears through bone structure of the user's head and/or through parts of middle ear of the user or electric signals transferred directly or indirectly to cochlear nerve and/or to auditory cortex of the user.

The hearing device is adapted to be worn in any known way. This may include i) arranging a unit of the hearing device behind the ear with a tube leading air-borne acoustic signals into the ear canal or with a receiver/loudspeaker arranged close to or in the ear canal such as in a Behind-the-Ear type hearing aid, and/or ii) arranging the hearing device entirely or partly in the pinna and/or in the ear canal of the user such as in a In-the-Ear type hearing aid or In-the-Canal/Completely-in-Canal type hearing aid, or iii) arranging a unit of the hearing device attached to a fixture implanted into the skull bone such as in Bone Anchored Hearing Aid or Cochlear Implant, or iv) arranging a unit of the hearing device as an entirely or partly implanted unit such as in Bone Anchored Hearing Aid or Cochlear Implant.

A “hearing system” refers to a system comprising one or two hearing devices, and a “binaural hearing system” refers to a system comprising two hearing devices where the devices are adapted to cooperatively provide audible signals to both of the user's ears. The hearing system or binaural hearing system may further include auxiliary device(s) that communicates with at least one hearing device, the auxiliary device affecting the operation of the hearing devices and/or benefitting from the functioning of the hearing devices. A wired or wireless communication link between the at least one hearing device and the auxiliary device is established that allows for exchanging information (e.g. control and status signals, possibly audio signals) between the at least one hearing device and the auxiliary device. Such auxiliary devices may include at least one of remote controls, remote microphones, audio gateway devices, mobile phones, public-address systems, car audio systems or music players or a combination thereof. The audio gateway is adapted to receive a multitude of audio signals such as from an entertainment device like a TV or a music player, a telephone apparatus like a mobile telephone or a computer, a PC. The audio gateway is further adapted to select and/or combine an appropriate one of the received audio signals (or combination of signals) for transmission to the at least one hearing device. The remote control is adapted to control functionality and operation of the at least one hearing devices. The function of the remote control may be implemented in a SmartPhone or other electronic device, the SmartPhone/electronic device possibly running an application that controls functionality of the at least one hearing device.

In general, a hearing device includes i) an input unit such as a microphone for receiving an acoustic signal from a user's surroundings and providing a corresponding input audio signal, and/or ii) a receiving unit for electronically receiving an input audio signal. The hearing device further includes a signal processing unit for processing the input audio signal and an output unit for providing an audible signal to the user in dependence on the processed audio signal.

The input unit may include multiple input microphones, e.g. for providing direction-dependent audio signal processing. Such directional microphone system is adapted to enhance a target acoustic source among a multitude of acoustic sources in the user's environment. In one aspect, the directional system is adapted to detect (such as adaptively detect) from which direction a particular part of the microphone signal originates. This may be achieved by using conventionally known methods. The signal processing unit may include amplifier that is adapted to apply a frequency dependent gain to the input audio signal. The signal processing unit may further be adapted to provide other relevant functionality such as compression, noise reduction, etc. The output unit may include an output transducer such as a loudspeaker/receiver for providing an air-borne acoustic signal transcutaneously or percutaneously to the skull bone or a vibrator for providing a structure-borne or liquid-borne acoustic signal. In some hearing devices, the output unit may include one or more output electrodes for providing the electric signals such as in a Cochlear Implant.

A “cochlear implant hearing aid system” represents a particular type of “hearing system” comprising an external unit, which receives acoustic sound and processes the acoustic sound into a coded audio signal, and an implantable unit, receives the coded audio signal.

Now referring toFIG.1, which illustrates a schematic diagram of a cochlear implant hearing aid system according to a first embodiment of the disclosure, i.e. a cochlear implant hearing aid system with a measuring unit configured to measure a dissipative current.

Particularly, according toFIG.1, a cochlear implant hearing aid system100is disclosed. The system100comprises an external unit110(diagram according toFIG.1exclusive implantable unit120) configured to receive acoustic sound and process the acoustic sound into a coded audio signal, and an implantable unit120configured to receive the coded audio signal. Further, the external unit110comprises a power supply unit111connected via a first path to a switching unit112, wherein the switching unit112is connected via a second path to ground and is connected at an output thereof to a first coil113. Furthermore, the switching unit112is configured to operate as a switching element configured to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil113. The coded audio signal is supplied to the switching unit112as a control signal, and the first coil113is inductively linked to a second coil121arranged in the implantable unit120. The system100further comprises a measuring unit130connected to the first path and configured to measure a dissipative current occurring in relation to the switching states of the switching unit112, wherein the first path further comprises a shunt resistor114, and the measuring unit130is configured to measure the dissipative current across the shunt resistor114. Based on the measured dissipative current, a resonance frequency of the inductive link between the first and second coils113,121is adapted.

A measuring element135is arranged in the measuring unit130for schematically illustrating a measurement performed by the measuring unit130. In addition, a feedback control mechanism (dashed arrow according toFIG.1) is indicated, running from the measuring unit135to the switching unit112, enabling to continuously monitor the resonance frequency adaptation. Further, the second path comprises at least one of a coil and a resistor.

This allows for adapting the resonance frequency based on a measured dissipative current.

Particularly, such cochlear implant hearing aid system100further allows for measuring a tuning/mistuning of the switching unit112. As a result thereof, by usage of one measurement, it is directly derivable which type of first and second coils113,121to use for establishing the inductive link between the external unit110and the implantable unit120, thereby for example addressing patients with different skin thicknesses by distinguishing e.g. between skin thickness ranges of 1 mm to 6 mm and 4 mm to 11 mm.

Specifically, the cochlear implant hearing aid system100further allows for monitoring an adaptation of the resonance frequency, and even further allows for continuously monitoring the adaptation during use of the cochlear implant hearing aid system100, which is unique for a cochlear implant hearing aid system. This leads to even further improve power efficiency of the system100, since the inductive link becomes more efficient due to the ability of continuously adapting the resonance frequency (keeping the resonance frequency of the first coil121aligned or almost aligned to a target resonance frequency). In particular, the resonance frequency of the first coil121is adapted in order to minimize the measured dissipative current.

As result of continuously monitoring a resonance frequency adaptation, adaptation of the resonance frequency (tuning of the switching unit112) may be performed in real time, allowing for reducing power losses even further.

The system100according toFIG.1further allows for simplicity of a design of the measuring unit130, by exploiting existing components of the switching unit112and/or the cochlear implant hearing aid system100. Particularly, by using the existing components, application of new components to the switching unit112/cochlear implant hearing aid system100is therefore avoidable, which leads to a more complex design of the measuring unit130.

Additionally, the system100according toFIG.1further allows for avoiding that measuring unit130has an impact on the switching unit112in adapting the resonance frequency (an impact on tuning of the switching unit). Moreover, such cochlear implant hearing aid system100allows for avoiding that measuring unit130creates a parasitic resistance to the circuitry constituting the switching unit112/cochlear implant hearing aid system100.

FIG.2illustrates a schematic diagram of a cochlear implant hearing aid system200according to the first embodiment of the disclosure. Particularly,FIG.2differs fromFIG.1in thatFIG.2illustrates a measuring unit230with an alternative structure in comparison to the measuring unit130according toFIG.1. Specifically, the measuring unit230according toFIG.2may further comprise an amplification element231configured to amplify a current to be measured by the measuring unit230.

Due to the amplification, this allows for easier measurement of the dissipative current, as well as for reducing measurement errors based on background noise. Regarding a quality of the inductive link between the first coil113and the second coil121, the higher the measured current, the worse the inductive link, because more current is drawn from the power supply unit111for being able to transmit the coded audio signal.

Referring toFIG.3,FIG.3illustrates a schematic diagram of a cochlear implant hearing aid system according to a second embodiment of the disclosure, i.e. a cochlear implant hearing aid system with a measuring unit configured to measure a dissipative current.

Particularly, according toFIG.3, a cochlear implant hearing aid system300is disclosed. The system300comprises an external unit310(diagram according toFIG.3exclusive implantable unit320) configured to receive acoustic sound and process the acoustic sound into a coded audio signal, and an implantable unit320configured to receive the coded audio signal. Further, the external unit310comprises a power supply unit311connected via a first path to a switching unit312, wherein the switching unit312is connected via a second path to ground and is connected at an output thereof to a first coil313. Furthermore, the switching unit312is configured to operate as a switching element configured to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil313. The coded audio signal is supplied to the switching unit312as a control signal, and the first coil313is inductively linked to a second coil321arranged in the implantable unit320. The system300further comprises a measuring unit330connected to the second path and configured to measure a dissipative current occurring in relation to the switching states of the switching unit312, wherein the second path comprises a third coil315inductively coupled to a fourth coil332, and the measuring unit330comprises the fourth coil332. The measuring unit330is further configured to measure a current induced in the fourth coil332and to derive the dissipative current therefrom. Based on the measured dissipative current, a resonance frequency of the inductive link between the first and second coils313,321is adapted.

A measuring element335is arranged in the measuring unit330for schematically illustrating a measurement performed by the measuring unit330. In addition, a feedback control mechanism (dashed arrow according toFIG.3) is indicated, running from the measuring unit335to the switching unit312, enabling to continuously monitor the resonance frequency adaptation.

Similar to the first embodiment, the system300according toFIG.3allows for monitoring an adaptation of the resonance frequency in real time, therefore allowing to reduce power losses even further. Additionally, similar to the first embodiment, the system300according toFIG.3allows for simplicity of a design of the measuring unit330, by exploiting existing components of the switching unit312and/or the cochlear implant hearing aid system300.

Moreover, further similar to the first embodiment, this system300further allows for avoiding the measuring unit330to have an impact on the switching unit312adapting the resonance frequency (tuning of the switching unit), as well as for avoiding the measuring unit330to create a parasitic resistance to the circuitry constituting the switching unit312/cochlear implant hearing aid system300.

FIG.4illustrates a schematic diagram of a cochlear implant hearing aid system400according to the second embodiment of the disclosure. Particularly,FIG.4differs fromFIG.3in thatFIG.4illustrates a measuring unit430with an alternative structure in comparison to the measuring unit330according toFIG.3. Specifically, the measuring unit430according toFIG.4may further comprise an amplification element431configured to amplify the current induced in the fourth coil432.

Furthermore,FIG.5illustrates a schematic diagram of a cochlear implant hearing aid system500according to the second embodiment of the disclosure. Particularly,FIG.5differs fromFIGS.3and4in thatFIG.5illustrates a measuring unit530with an alternative structure in comparison to the measuring units330,430according toFIGS.3and4. Specifically, in comparison to the measuring unit430according toFIG.4, the measuring unit530according toFIG.5may further comprise a resistor533connected to an output of an amplification element531and a capacitor534connected at one out of its terminals to the connection between the output of the amplification element531and the resistor533and at the other one of its terminals to ground.

Similar to the first embodiment, due to the amplification, this allows for easier measurement of the dissipative current, as well as for reducing measurement errors based on background noise.

Specifically, regarding cochlear implant hearing aid system500according toFIG.5, at maximum efficiency (perfect tuning), there is theoretically no current in a (not shown) drain of a transistor of the switching unit512, thus there is no dissipative current. However, if the switching unit is mistuned, dissipative current occurs in the transistor, leading to power losses. The dissipative current will go through the third coil515and image of this dissipative current will be created in the fourth coil523because of the inductive coupling between the third and fourth coils515,523. The amplification element531provides an impedance adaptation between the very low impedance of the fourth coil532and the measurement circuit comprising the resistor533and the capacitor534. The amplification element531amplifies the current (created image of the dissipative current) in the fourth coil532. The current in the fourth coil532is an AC current with high frequency content. Thus, such current is complex to measure. Therefore, after the amplification by the amplification element531, the resistor533and the capacitor532allow to create from this AC current a DC voltage (average value of the signal). As a result, a voltage used for determining the dissipative current may be effectively measured across the capacitor534.

Referring toFIG.6,FIG.6illustrates a schematic diagram of a cochlear implant hearing aid system600according to a third embodiment of the disclosure. In general, the third embodiment refers to combinations of the first and second embodiments.FIG.6illustrates a particular combination of the first and second embodiments.

Namely, according toFIG.6, a cochlear implant hearing aid system600is disclosed. The system600comprises an external unit610(diagram according toFIG.6exclusive implantable unit620) configured to receive acoustic sound and process the acoustic sound into a coded audio signal, and an implantable unit620configured to receive the coded audio signal. Further, the external unit610comprises a power supply unit611connected via a first path to a switching unit612, wherein the switching unit612is connected via a second path to ground and is connected at an output thereof to a first coil613. Furthermore, the switching unit612is configured to operate as a switching element configured to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil613. The coded audio signal is supplied to the switching unit612as a control signal, and the first coil613is inductively linked to a second coil621arranged in the implantable unit620.

The system600further comprises a first measuring unit630aconnected to the first path and configured to measure a first dissipative current occurring in relation to the switching states of the switching unit612. The first measuring unit630afurther comprises a first amplification element631aconfigured to amplify a current to be measured by the first measuring unit630a. The first path further comprises a shunt resistor614, and the first measuring unit630ais configured to measure the first dissipative current across the shunt resistor614.

In addition, the system600further comprises a second measuring unit630bconnected to the second path and configured to measure a second dissipative current occurring in relation to the switching states of the switching unit612. The second path comprises a third coil615inductively coupled to a fourth coil632b, and the second measuring unit630bcomprises the fourth coil632b. In addition, the second measuring unit630bfurther comprises a second amplification element631bconfigured to amplify the current induced in the fourth coil632b. Moreover, the second measuring unit630bfurther comprises a resistor633bconnected to an output of a the second amplification element631band a capacitor634bconnected at one out of its terminals to the connection between the output of the second amplification element631band the resistor633band at the other one of its terminals to ground. Further, the second measuring unit630bis configured to measure a current induced in the fourth coil632band to derive the second dissipative current therefrom.

A first measuring element635ais arranged in the first measuring unit630afor schematically illustrating a measurement performed by the first measuring unit630aand a second measuring element635bis arranged in the second measuring unit630bfor schematically illustrating a measurement performed by the second measuring unit630b. In addition, a feedback control mechanisms (dashed arrows according toFIG.6) are indicated, running from the measuring units635a,635bto the switching unit612, enabling to continuously monitor the resonance frequency adaptation and/or allowing further assessing/evaluation of the measured dissipative currents.

According to the cochlear implant hearing aid system600as illustrated inFIG.6, based on at least one of the first and second measured dissipative current, a resonance frequency of the inductive link between the first and second coils613,621is adapted.

In addition to the first and second embodiments, the cochlear implant hearing aid system600according toFIG.6further allows for, basically, comparing the first and second measured dissipative currents in order to even further improve a tuning of the switching unit612based on a result of the comparison. For example, the system600according toFIG.6allows for deriving a weighted dissipative current from the measured first and second dissipative currents based on applying a predetermined weighting algorithm. As a result for example, an effect of background noise may be reduced. Therefore, the resonance frequency of the inductive link may be adapted based on the derived weighted dissipative current.

Such predetermined weighting algorithms are for example the following, but not limited thereto, averaging the first and second measured dissipative currents (weighting of 50% per measured current), selecting only the measured dissipative current of highest/lowest intensity (weighting of 100% for one measured current), or weighting the first and second measured dissipative currents relative to respective measured intensities.

Moreover, according to various exemplary embodiments, the switching unit112,212,312,412,512,612may be for example an E-Class amplifier.

An E-Class amplifier is an amplifier that uses resonating load in order to minimize loss and allows for a theoretical maximum efficiency of 100%. This is achieved by having no losses during transistor switches performed by a transistor contained by the E-Class amplifier. Thus, if a current dissipated by the transistor is relatively small, the E-Class amplifier is well-tuned (the resonance frequency is well-adapted). However, if the current dissipated by the transistor comprises for example larger spike currents than in the well-tuned case, the E-Class amplifier is mistuned (the resonance frequency adapted incorrectly).

Therefore, such E-Class amplifier allows for theoretically 100% efficiency of the switching unit112,212,312,412,512,612, which means no power losses and, therefore, no dissipative current.

Referring toFIG.7,FIG.7illustrates a cochlear implant hearing aid system700where the switching unit712is an E-Class amplifier, and tuning of the E-Class amplifier712is performed via switched capacitors743,744.

In addition, as an example with reference toFIG.7, according to various exemplary embodiments, the resonance frequency of the inductive link may be adapted based on the at least one measured dissipative current by selectively connecting at least one further switched capacitor743,744to a circuitry constituting the switching unit112,212,312,412,512,612,712.

Additionally, according to various exemplary embodiments, the switching unit112,212,312,412,512,612,712may further comprise at least one of a first switched capacitor743and a second switched capacitor744, wherein the first switched capacitor743may be connected in parallel to a first circuitry element741, and the second switched capacitor744may be connected in serial to a second circuitry element742. Further, the resonance frequency may be adapted based on the at least one measured dissipative current by selectively connecting at least one of the first and second switched capacitors743,744to the circuitry constituting the switching unit112,212,312,412,512,612,712.

Furthermore, according to various exemplary embodiments, the switching unit112,212,312,412,512,612,712may further comprise a transistor element740, wherein the transistor element740may be connected at an output thereof via the first circuitry element741to ground, and via the second circuitry element742to the first coil713. Further, the first circuitry element741may be a third capacitor, and the second circuitry element742may be a fourth capacitor.

Based on the continuous monitoring process as outlined above, this allows for adjusting a tuning of the switching unit112,212,312,412,512,612,712in real time.

Regarding the switched capacitors743,744is it to be understood that these two switched capacitors serve as an example, and that the switching unit112,212,312,412,512,612,712may comprise more than two such switched capacitors743,744at different positions in the circuitry constituting the switching unit112,212,312,412,512,612,712. Further, a switched capacitor, such as any of the first and second switched capacitor743,744according toFIG.7, may be an arrangement of a plurality of capacitors connected together in parallel or in serial or in any combination thereof. Therefore, allowing for fine-tuning of the switching unit112,212,312,412,512,612,712.

Furthermore, multiple switch-capacitor pairs may be connected in parallel and in parallel to capacitor C2, e.g. a single switch-capacitor pair includes a switch and a capacitor. If a first switch-capacitor pair is turned on the remaining are turned off, the resonant frequency increases, and when turning a second switch-capacitor pair on and the remaining off, the resonant frequency increases even more. For example, when turning the first switch-capacitor pair on is ideal for a first skin thickness, and when turning the second switch capacitor pair on is ideal for a second skin thickness, wherein the second skin thickness is thicker than the first skin thickness.

Referring toFIG.8,FIG.8illustrates a method for a cochlear implant hearing aid system according to embodiments of the disclosure. The method according toFIG.8may be executed by the cochlear implant hearing aid systems100,200,300,400,500,600, and700according toFIGS.1,2,3,4,5,6, and7respectively. However, the method is not limited thereto. Further, the cochlear implant hearing aid systems100,200,300,400,500,600, and700according toFIGS.1,2,3,4,5,6, and7, respectively, may execute the method according toFIG.8. However, the cochlear implant hearing aid systems100,200,300,400,500,600, and700are not limited thereto.

Particularly, according toFIG.8, a method for a cochlear implant hearing aid system100,200,300,400,500,600,700comprising an external unit110,210,310,410,510,610,710receiving acoustic sound and processing the acoustic sound into a coded audio signal and an implantable unit120,220,320,420,520,620,720receiving the coded audio signal is disclosed. The method comprises the steps of supplying (Step S810) in the external unit110,210,310,410,510,610,710power from a power supply unit111,211,311,411,511,611,711via a first path to a switching unit112,212,312,412,512,612,712wherein the switching unit112,212,312,412,512,612,712is connected via a second path to ground and is connected at an output thereof to a first coil113,213,313,413,513,613,713. Further, the switching unit112,212,312,412,512,612,713operates as a switching element to switch between switching states, wherein the switching states include a first state for applying and a second state for not applying a current to the first coil113,213,313,413,513,613,713. The coded audio signal is supplied to the switching unit112,212,312,412,512,612,712as a control signal, and the first coil113,213,313,413,513,613,713is inductively linked to a second coil121,221,321,421,521,621,721arranged in the implantable unit120,220,320,420,520,620,720. The method further comprises the steps of measuring (Step S820) across at least one of the first and second paths a dissipative current occurring in relation to the switching states of the switching unit112,212,312,412,512,612,712. The method further comprises adapting (Step S830) a resonance frequency of the inductive link between the first and second coils113,213,313,413,513,613;121,221,321,421,521,621,721based on the at least one measured dissipative current.

This allows for a method for measuring a dissipative current, by exploiting existing components of the switching unit112,212,312,412,512,612,712and/or the cochlear implant hearing aid system100,200,300,400,500,600,700. Particularly, by applying the method on the existing components, application of new components to the switching unit112,212,312,412,512,612,712/cochlear implant hearing aid system100,200,300,400,500,600,700is therefore avoided, which leads to a more simplified measurement.

In addition, according to various exemplary embodiments, the method may further comprise the steps of measuring a first dissipative current across a shunt resistor114,214,614contained in the first path, and/or measuring a second dissipative current based on a current induced in a fourth coil332,432,532,632binductively coupled to a third coil315,415,515,615, wherein the third coil315,415,515,615is comprised in the second path.

This further allows for a method for avoiding the measurement to have an impact on a tuning of the switching unit112,212,312,412,512,612, as well as for avoiding the measurement to create a parasitic resistance to the circuitry constituting the switching unit112,212,312,412,512,612,712/cochlear implant hearing aid system100,200,300,400,500,600,700.

Furthermore, according to various exemplary embodiments, the method may further comprise the steps of deriving a weighted dissipative current from the measured first and second dissipative currents based on applying a predetermined weighting algorithm. Further, the method may comprise adapting the resonance frequency of the inductive link based on the derived weighted dissipative current.

Since for example an effect of background noise occurring in the measurements may be reduced, evaluating the first and second measured dissipative currents in combination allows for increasing accuracy in tuning of the switching unit112,212,312,412,512,612,712.

Moreover, according to various exemplary embodiments, the method may further comprise the steps of adapting the resonance frequency of the inductive link based on the at least one measured dissipative current by selectively connecting at least one further switched capacitor to a circuitry constituting the switching unit112,212,312,412,512,612,712.

Additionally, according to various exemplary embodiments, the switching unit112,212,312,412,512,612,712may further comprise at least one of a first switched capacitor and a second switched capacitor. The first switched capacitor is connected in parallel to a first circuitry element, and the second switched capacitor is connected in serial to a second circuitry element. The method may further comprises the steps of adapting the resonance frequency of the inductive link based on the at least one measured dissipative current by selectively connecting at least one of the first and second switched capacitors to the circuitry constituting the switching unit112,212,312,412,512,612,712. The switching unit112,212,312,412,512,612,712may be an E-Class amplifier and may further comprise a transistor element, wherein the transistor element is connected at an output thereof via the first circuitry element to ground, and at an output thereof via the second circuitry element to the first coil. The first circuitry element may be a third capacitor and the second circuitry element may be a fourth capacitor.

This allows for a method for adjusting a tuning of the switching unit112,212,312,412,512,612,712in real time.

A Cochlear Implant typically includes i) an external part for picking up and processing sound from the environment, and for determining sequences of pulses for stimulation of the electrodes in dependence on the current input sound, ii) a (typically wireless, e.g. inductive) communication link for simultaneously transmitting information about the stimulation sequences and for transferring energy to iii) an implanted part allowing the stimulation to be generated and applied to a number of electrodes, which are implantable in different locations of the cochlea allowing a stimulation of different frequencies of the audible range. Such systems are e.g. described in U.S. Pat. Nos. 4,207,441 and in 4,532,930.

In an aspect, the hearing device comprises multi-electrode array e.g. in the form of a carrier comprising a multitude of electrodes adapted for being located in the cochlea in proximity of an auditory nerve of the user. The carrier is preferably made of a flexible material to allow proper positioning of the electrodes in the cochlea such that the electrodes may be inserted in cochlea of a recipient. Preferably, the individual electrodes are spatially distributed along the length of the carrier to provide a corresponding spatial distribution along the cochlear nerve in cochlea when the carrier is inserted in cochlea.

In an aspect, the functions may be stored on or encoded as one or more instructions or code on a tangible computer-readable medium. The computer readable medium includes computer storage media adapted to store a computer program comprising program codes, which when run on a processing system causes the data processing system to perform at least some (such as a majority or all) of the steps of the method described above, in the and in the claims.

By way of example, and not limitation, such computer-readable media can comprise RAM, ROM, EEPROM, CD-ROM or other optical disk storage, magnetic disk storage or other magnetic storage devices, or any other medium that can be used to carry or store desired program code in the form of instructions or data structures and that can be accessed by a computer. Disk and disc, as used herein, includes compact disc (CD), laser disc, optical disc, digital versatile disc (DVD), floppy disk and Blu-ray disc where disks usually reproduce data magnetically, while discs reproduce data optically with lasers. Combinations of the above should also be included within the scope of computer-readable media. In addition to being stored on a tangible medium, the computer program can also be transmitted via a transmission medium such as a wired or wireless link or a network, e.g. the Internet, and loaded into a data processing system for being executed at a location different from that of the tangible medium.

Regarding the above described method, adjustment of the resonance frequency may be implemented in software.

In an aspect, a data processing system comprising a processor adapted to execute the computer program for causing the processor to perform at least some (such as a majority or all) of the steps of the method described above and in the claims.

Regarding the above described method, adjustment of the resonance frequency may be implemented in software.