EP2180726A1 - Sound localization in binaural hearing aids - Google Patents

Abstract

The method has the signal propagation time between an input transducer (2) for an input acoustic signal and an output transducer (4) converting an electrical signal into an output acoustic signal measured for the hearing aid (1) for a first ear, with transmission of a corresponding signal to the hearing aid (1') for the second ear, used for matching the signal propagation time between its corresponding input transducer (2') and output transducer (4'). An independent claim for a binaural hearing aid system is also included.

Description

The invention relates to a hearing aid system and a method for setting a hearing aid system having at least one first and one second hearing aid device, each having at least one input transducer for receiving an acoustic input signal and conversion to an electrical signal, a signal processing unit for processing the electrical signal and an output transducer for conversion comprise the electrical signal in an output signal and between which a signal path for data transmission is provided.

Directional hearing is the ability of a person to distinguish the direction in which a sound source is located. If a sound source is not in front of or behind the person, the finite propagation speed of the sound inevitably results in a difference in transit time between the two ears, and thus a time difference with which the ears perceive a sound wave coming from one direction. If a sound is e.g. from the perspective of the peson comes from the right, this reaches the right ear by a fraction of a second rather than the left ear. This time difference is much shorter than the person concerned can consciously recognize. The effect occurs through an automatic integration process in the acoustic nervous system.

In addition to the time difference, there is still a difference in the volume with which the ears perceive a sound coming from one side. A sound source on one side of the head gives the ear a slightly louder sound on that side. Also, this minimal difference in volume is sufficient so that the sound source from the perspective of the person can be located on the left or right.

In binaural hearing aid care, loss of directional hearing often occurs. This is mainly due to the fact that depending on the hearing situation, which is detected by the respective hearing aid, the signal processing of the two hearing aids may include different steps. Furthermore, in a hearing aid wearer usually the hearing loss of both ears is pronounced to different degrees. Accordingly, the settings of the hearing aids to compensate for the hearing loss of each ear are set differently. However, different settings of the signal processing of both hearing aids usually result in different signal propagation times within the hearing aid devices. There is therefore an unnatural shift in the important for directional listening phase of an acoustic input signal. As already mentioned, the transit time of a sound signal between the two ears in addition to the difference in the volume for the directional hearing of great importance. Even slight changes in this natural transit time shift, as caused for example by different signal propagation times within the hearing aid devices, can therefore lead to a loss of directional hearing.

To solve this problem, it is known to process the recorded on the two ears acoustic signals in a common central signal processing device. That's how it looks US 5,479,522 in addition to two hearing aids worn in each case on one ear, an additional processor unit which can be designed, for example, as a breast appliance or wristwatch. The acoustic signals received at the two ears undergo the same signal processing steps, so that the phase relationship between the two signals is maintained.

From the US 5,434,924 It is known to perform the signal processing in binaural supply substantially only in one of the two hearing aids. These are the one ear received signals transmitted to the hearing of the other ear, processed there together and then fed to both ears (master-slave solution).

The former solution has the disadvantage that a further assembly is necessary and the hearing aid wearer now requires three instead of two devices, which means a significant limitation of wearing comfort, maintenance and handling. The second solution requires that all signal processing must be done by a single signal processing unit on a single page. While in the solution with a third device enough space is available to provide a correspondingly powerful signal processing and ensure their energy needs, the space is limited in a hearing aid located on the ear. Therefore, a master-slave solution with two differently designed hearing aids must necessarily have a lower computational capacity than would be available when using both hearing aids.

Another approach to solving the above problem is to transmit the incoming sound signals on the hearing aids of both sides to the other device and to process both signals on each side. In this way, the acoustic signals received at both ears undergo the same steps of signal processing together and therefore automatically experience the same signal delay. This approach is for example from the WO 97/14268 as well as the WO 99/43185 out. Although the transmission of the microphone signals on both sides of a binaural hearing aid system to the other side and the simultaneous processing of both signals on both sides solves the problem of a transit time difference, but is subject to the same limitations as the master-slave approach.

Another major disadvantage of all the solutions mentioned lies in the fact that they are all transmitting large Require datasets. This requires a considerable amount of time, space and energy consumption. In particular, in wireless data transfer, as required in the current state of the art, this represents a significant disadvantage.

From the EP 0 941 014 A2 a hearing aid system with a first and a second hearing device is known in which by actuation of a control element on the first hearing aid control signals are generated and transmitted to the second hearing aid. As a result, the simultaneous adjustment of both hearing aids is effected by the operation of the control element on one of the hearing aids.

From the DE 100 48 354 A1 For example, a method for operating a hearing aid system is known in which sound field characteristics are transmitted from one hearing aid to another. These can be signal levels.

From the DE 197 04 119 C1 a hearing aid is known in which a signal transmission from one to the other hearing aid via optical fibers is made. In this case, control signals can be transmitted.

The object of the present invention is to support the natural directional hearing in a hearing aid system for binaural care and to keep the additional computational effort required for this purpose low.

This object is achieved by methods with the method steps according to claims 1 or 11.

Furthermore, the object is achieved by a hearing aid system having the features according to claims 19 or 21.

In one known from the cited prior art hearing aid system with two hearing aids is an equal signal delay in the signal paths of both hearing aids between generated in each case the microphone and the listener, without explicitly knowing this signal propagation time. Disadvantages are the high computational effort and the high data transfer rates required.

In contrast to this, an embodiment of the invention provides for the signal propagation time of the electrical signal in the signal path between the input transducer and the output transducer of the first hearing aid device to be determined at least in a relevant subarea in which transit time differences are to be expected, and the signal propagation time of the electrical signal between adapt the input transducer and the output transducer of the second hearing aid to the determined signal delay of the first hearing aid.

Delay differences between the two hearing aids of a hearing aid system arise in particular by different settings of the hearing aid during operation. These different settings can be e.g. be caused by a different hearing loss of the two ears of a hearing aid wearer. In addition to the adaptation to the hearing device wearer, the settings of a hearing aid device can also be used to adapt to the respective environmental situation in which the hearing aid device is currently located. Since the last-mentioned settings are adaptive and automatic in modern hearing aid devices, the transit time differences during operation of the hearing aid devices may fluctuate.

In one embodiment of the invention, signal transit times of the first and second hearing aid devices are determined. Preferably, suitable transit time measurements are carried out for this purpose. Advantageously, the determined signal propagation times also include the transit times of the input and output transducers, so that their delays can also be taken into account. However, it can also only a portion of the Signal path between the input and the output transducer are measured.

Hearing aids generally have several hearing programs to adapt the signal processing to different listening situations (e.g., "quiet environment", "background noise environment", "car drive", "telephone conversation", etc.). Advantageously, the hearing aids are first individually adapted to the hearing aid wearer and determines the signal propagation times for the various hearing programs. From a comparison of the signal propagation times thus determined for both hearing aid devices, a signal delay can be ascertained for each possible program pairing (two hearing aids of one hearing aid system can be operated at different times in different hearing programs) by which the signal propagation time in the "faster" hearing aid device, ie the hearing aid device The shorter cycle time of an acoustic input signal from the input transducer to the output transducer, must be extended so that adjusts the same signal propagation time in both hearing aids. If the information about signal propagation times for the individual hearing programs of both hearing aids of a hearing aid system and the current program pairing are known in a hearing aid device and this hearing aid device has means for changing the propagation time, the transit time difference can thereby be compensated. Therefore, only data for labeling the current hearing program of one hearing aid device must be transmitted to the other hearing aid device. Alternatively, it is also possible to transmit data identifying runtime or runtime differences directly.

Another embodiment of the invention provides that a signal propagation time of an electrical signal is automatically determined in a hearing aid device. Elaborate measurements on the hearing aid can be omitted. Furthermore, this embodiment is advantageous in a hearing aid which adapts adaptively to different hearing situations, without having a fixed predetermined division into hearing programs for this purpose. The transit time of an electrical signal between the input transducer and the output transducer is then variable within a certain range and can assume virtually any desired values within this range of variation. Advantageously, the signal propagation time is determined internally in such a hearing aid. For this purpose, for example, the signal propagation time can be known by internal components, such as signal processors or filters, under the current settings, so that only the signal propagation times of the individual components must be added to determine the total signal delay. On the other hand, the signal propagation time can also be determined by means of a test signal, which is advantageously fed to the input transducer and tapped before the output transducer.

In a further embodiment of the invention, a correlation analysis is carried out to determine the signal propagation time of an acoustic input signal through the hearing aid device or the signal propagation time of an electrical signal through a subregion of the hearing aid device. For example, the correlation analysis between the electrical output signal of the microphone and the electrical input signal of the listener can be performed. The result of the correlation analysis shows the phase shift between the two signals. From the phase shift turn can be closed to the signal delay.

In a further variant of the invention, it is provided that in each case the envelope is formed at two points one behind the other in the signal path of the hearing aid and the phase shift or the signal propagation time of the electrical signal between the two points is determined from a comparison of the envelope.

Data regarding the signal propagation time determined in a first hearing aid device are finally provided by the first hearing aid device transferred to the second hearing aid. Likewise, the determination of the current signal propagation time and the transmission of related data from the second hearing aid device to the first hearing aid device also takes place in the case of the second hearing aid device.

Different methods can be used to adapt different signal propagation times of two hearing aid devices of a hearing aid system: On the one hand, it is possible to change the clock frequency of at least one component of the hearing aid device in a digital hearing aid device. Thus, the hearing aid with the larger signal propagation time clocked higher or the hearing aid can be operated with the lower signal delay with reduced clock frequency. On the other hand, it is possible to increase the signal propagation time of the hearing aid with the lower signal propagation time through additional components. For example, in the case of a digital hearing aid device, a shift register may be provided which delays the electrical signal in the signal path of the hearing aid by a specific number of cycles. Furthermore, filter means can be used which cause a certain phase shift and thus delay. Preferably allpasses are used as filter means, by means of which the phase shift can be adjusted and which show no effects on the frequency response.

The adaptation of the signal propagation times of two hearing aids of a hearing aid system takes place according to a variant of the invention at periodic intervals. This ensures that the signal propagation times of the two hearing aids differ at most for a short period of time. The time intervals in which an adjustment takes place can be in the range of minutes or hours.

Another variant of the invention provides that the determination and adaptation of the signal propagation times of two hearing aid devices of a hearing aid system following a parameter and / or functional change in at least one of the hearing aid devices he follows. This change can be made for example by manual operation of at least one of the hearing aids by the user. For example, the user may cause the hearing aids to switch to another listening program. Since the signal processing in the hearing aids thus fundamentally changes, changes in the signal propagation times in the individual hearing aids are to be expected. Thus, a new adjustment of the signal propagation times is performed to compensate for the changed signal propagation times.

As with the parameter and / or function change triggered by manual actuation of a control element, compensation of the signal propagation times between the hearing aid devices also takes place in the case of an automatic parameter and / or function change of at least one hearing aid device. This makes it possible to adapt the signal propagation time to the respective other hearing aid device even in hearing aid devices in which an automatic situation analysis and adaptation to the current hearing situation takes place. This is particularly important because in such hearing aid systems due to the frequently occurring in both hearing aids independently of each other signal analysis different settings and functions are selected. Thus, for example, in a hearing aid in which feedback-induced oscillations are detected, a feedback suppression algorithm may be active or a filter for feedback cancellation may be set without affecting the other hearing aid. This may result in runtime differences for the electrical signals in the individual hearing aids.

Preferably, whenever a parameter and / or functional change results in at least one of the hearing aid devices, an adaptation of the signal propagation times of the two hearing aid devices is carried out. This is particularly advantageous when in both hearing aid devices automatically a continuous and continuous adjustment to the particular environmental situation takes place without there being a fixed predefined division into specific listening situations and an associated assignment to predefined hearing programs.

In a prior art hearing aid, signal processing often takes place in parallel in several parallel channels of a signal processing unit. As a rule, each channel usually comprises a specific frequency band of the signal to be processed. Since a different signal processing usually takes place for the individual frequency bands, the signal propagation times between the individual bands can vary. In order to compensate for this effect, a variant of the invention provides to determine the signal propagation times or the amplitude transmission behavior for the individual frequency bands and to adjust them between the hearing aid devices.

In a hearing aid system according to the invention, the directional hearing in the case of a binaural hearing aid supply is improved by matching the signal propagation times of the hearing aid devices attached to the two ears. However, the signal propagation times are only one factor affecting directional hearing. In an advantageous development of the hearing aid system according to the invention, an adaptation of the amplitude response of the two hearing aids is also carried out. Differences in the amplitudes of signals that occur from different directions are mainly caused by the shading effect of the head. The differences in the amplitudes are very small and can not be consciously perceived. Only by a very fine adjustment of the hearing aids of a hearing aid system, these minimum differences in amplitude, which are caused by different directions of incidence remain maintained. The exact amount of these differences is rather secondary. It is especially important that an amplitude difference in a signal from a certain direction is largely retained, even if in one or both hearing aids Change settings. If, for example, the volume is increased in a hearing aid, the volume should also be adjusted in the case of the other hearing aid. However, since often both ears of a hearing aid wearer are equally affected by hearing loss, the volume adjustment usually can not be done equally in both hearing aids. Rather, the adjustment has to be done taking into account the individual hearing curves that were measured on the two ears of a hearing aid wearer. So it is crucial that a hearing aid wearer is always given a slightly higher volume at the ear with the shorter distance to the signal source at a signal that comes from a certain direction.

In one embodiment of the invention, in a hearing aid system having two hearing aid devices that can be worn on the head, a gain or gain change of an electrical signal in at least one of the hearing aid devices is determined. The gain change may, for example, have been caused by the change of a parameter of the signal processing of the hearing aid. Then, data for identifying the current amplification or for characterizing the gain change is transmitted from the hearing aid device to the other hearing aid device of the hearing aid system. Also in this hearing aid then the gain is adjusted accordingly. This may mean that the gain is changed by the same amount. Preferably, however, the amplification in the second hearing aid device is changed such that the same loudness impression is produced again in the case of a sound signal arriving from the 0 degree direction (directly from the front) on both ears through the supply of the hearing aid devices. Deviating from the 0-degree direction sound signals are then perceived again with different loudness impression, so that the hearing aid wearer can perceive the direction from which the sound signal arrives.

The value of a gain change in a hearing aid according to the invention may be permanently assigned to certain settings or functions of the hearing aid. For example, in a feedback suppression algorithm, there may always be a 10 dB reduction in gain. Data for identifying this gain change can then, as soon as the algorithm is actively switched, be transferred to the other hearing aid device of the hearing aid system, so that a corresponding gain reduction is also performed in this case. In many applications, however, there is no fixed association between certain functions of the hearing aid and related gain changes. The amplification or gain change is then first determined automatically in the hearing aid. For this purpose, signal amplitudes or signal levels of an electrical signal can be detected and evaluated at points located one behind the other in the signal path of the hearing aid device. For this, too, a test signal is preferably fed into the signal path, which at least partially passes through the signal processing unit of the hearing aid. Preferably, the amplification in both hearing aids is determined and related data transferred to the other hearing aid in the gain adjustment. To adapt the gain in a hearing aid to a gain change in a second hearing aid of a hearing aid system filter media are preferably set. Preferably, an adjustment of the amplification of the two hearing aids of a hearing aid system is always carried out even when the gain setting when at least one of the hearing aids a parameter and / or functional change. However, the gain adjustment can also be done at periodic intervals. As well as the determination and adaptation of the signal propagation time, the determination and adaptation of the amplification or of the amplitude transmission behavior in a hearing device system with multi-channel hearing aid devices can in each case only relate to specific frequency bands.

In an advantageous embodiment of the invention, in addition to the determination of signal propagation times in the hearing aid devices of a hearing aid system, the transmission behavior of signal amplitudes is also measured. Here, too, a test signal can be fed into the signal path at one point and read out again at the following point. Preferably, this measurement is also carried out for different signal frequencies. If a parameter or function change then takes place in at least one of the hearing aid devices, the transmission behavior with respect to the signal amplitudes can be measured again and differences in the transmission behavior can be ascertained. Data that are characteristic of the signal amplitudes are then transferred to the respective other hearing aid device of the hearing device system for adaptation to the changed transmission behavior.

The invention applies equally to behind the ear portable (BTE), in the ear portable (ITE) or implantable hearing aid systems application.

• Figur 1 ein Hörgerätesystem mit zwei Hörhilfegeräten, zwischen denen ein Signalpfad vorgesehen ist und bei denen unterschiedliche Hörprogramme einstellbar sind,
• Figur 2 ein Hörhilfegerät mit einer Signallaufzeit-Messeinrichtung und einem einstellbaren Verzögerungselement,
• Figur 3 ein Hörhilfegerät mit einem Signallaufzeit- und Amplituden- Messelement und einstellbarer Taktfrequenz und
• Figur 4 ein Hörhilfegerät, bei dem die Signalverarbeitung parallel in mehreren Frequenzkanälen erfolgt, mit einer Signalanalyse- und Steuereinheit.

Further details of the invention will be explained in more detail with reference to embodiments. Show it:

• FIG. 1 a hearing aid system with two hearing aids, between which a signal path is provided and in which different hearing programs can be set,
• FIG. 2 a hearing aid device with a signal transit time measuring device and an adjustable delay element,
• FIG. 3 a hearing aid with a signal delay and amplitude measuring element and adjustable clock frequency and
• FIG. 4 a hearing aid, in which the signal processing takes place in parallel in several frequency channels, with a signal analysis and control unit.

FIG. 1 shows a schematic representation of a hearing aid system with two hearing aids 1 and 1 '. The hearing aids 1 and 1 'each comprise an acousto-electrical input transducer (microphone) 2 or 2' for receiving an acoustic input signal and conversion into an electrical signal. The processing of the electrical signal to compensate for the hearing loss of a hearing aid wearer takes place in the signal processing units 3 or 3 '. The processed signal is finally converted back into a sound signal by an electric-acoustic output transducer (earphone) 4 or 4 'and fed to the ears of a hearing aid wearer.

To adapt to different listening situations, such as "speech in a quiet environment", "speech with noise", "driving in the car", etc., the hearing aids 1 and 1 'each include a control unit 5 and 5'. The control units 5 and 5 'are connected to memory units 6 and 6' in which different parameter sets for adapting the signal processing units 3 and 3 'to different listening situations are stored.

The adjustment of the hearing aid devices 1 and 1 'to the respective hearing situation takes place by actuation of a control element 7 or 7' on at least one of the hearing aid devices 1 or 1 '.

According to the invention, in the hearing aid devices 1 and 1 'signal propagation times of the signal processing units 3 and 3' for the respective hearing programs and taking into account the respective settings of the hearing aid devices 1 and 1 'to compensate for the individual hearing loss of a hearing aid wearer determined. This can be done, for example, by transit time measurements during the adaptation of the hearing aid devices 1 and 1 '. If the signal propagation times for both hearing aid devices 1 and 1 'are known under the selected settings for the respective hearing programs, the hearing programs are assigned data for characterizing the signal propagation times and also stored in the storage units 6 and 6 '. These data can be both the signal propagation times as such and also the respective transit time differences between the individual hearing programs or the hearing aid devices 1 and 1 '. If, for example, the hearing aid device 1 is switched between two hearing programs, then not only the parameters of the new hearing program are read from the memory unit 6, but also the data assigned to the newly set hearing program for characterizing the signal delay time. The latter are then transmitted via a transmitting and receiving unit 8 to the hearing aid 1 '. The hearing aid device 1 ', in turn, receives the data transmitted by the hearing aid device 1 by means of the transmitting and receiving unit 8' and supplies it to the control unit 5 '. This in turn compares the transmitted data with the information stored in the memory unit 6 'regarding the running time of the currently set hearing program. For example, by controlling a delay means, which is designed in the exemplary embodiment as an all-pass filter 9 or 9 ', then any differences in transit time can be compensated. Advantageously, both hearing aid devices 1 and 1 'thus have the same signal propagation delay between the input transducer 2 and the output transducer 4 or the input transducer 2' and the output transducer 4 '. Thus, with the hearing aid system 1, 1 'always directional hearing allows, regardless of the currently active program pairing of the hearing aids both hearing aids 1 and 1'.

Another embodiment of the invention shows FIG. 2 , Since in this case both hearing aids of a hearing aid system have the same equivalent circuit, is in FIG. 2 only one of the two, in the embodiment, the hearing aid 11, shown. Also this includes as the hearing aids 1 and 1 'in the embodiment according to FIG. 1 a microphone 12 for receiving an acoustic signal and conversion into an electrical signal, a signal processing unit 13 for frequency-dependent processing of the electrical signal and a receiver 14 for converting the electrical signal into a acoustic output signal. The hearing aid device 11 further comprises an A / D converter 15 for converting the output signal of the microphone into a digital signal, and a D / A converter 16 for reconverting the digital signal into an analog signal before the signal output via the receiver 14.

In contrast to the embodiment according to FIG. 1 takes place according to the hearing aid device 11 FIG. 2 a signal analysis of the digital electrical input signal in an analysis and control unit 17. Also, this is connected to a memory unit 18 in which different, the signal processing processing memory sets are stored. In addition to the possibility of controlling the signal processing in the hearing aid 11 by a complete set of parameters, which is stored in the memory unit 18, it is provided in the hearing aid 11, even adaptively change only individual settings and parameters for adjusting the signal processing to the respective hearing situation. Also, if necessary, certain functions or algorithms can be switched on or off. Thus, in the case of a recognized speech, an algorithm for speech enhancement can be set in the hearing aid device, or an algorithm for noise suppression can be activated when detected noises. Thus, a multiplicity of different settings and functions are possible which mostly have effects on the signal propagation time of a signal by the hearing aid device 11. Therefore, in the hearing aid device 11, the signal propagation time is automatically determined taking into account the current settings and functions. For this purpose, the hearing aid device 11 has a transit time determination unit 19. This includes a signal generator for generating and injecting a synthetic signal into the signal path. The fed-in signal passes through the signal processing unit 13 and is tapped before the output via the handset 14 and fed to the transit time determination unit 19. Preferably, the generated signal lies in a frequency range which is not audible acoustically by the hearing aid wearer. By the runtime determination unit 19 can now the Signal transit time measured by the signal processing unit 13 and transmitted to the analysis and control unit 17. The transit time measurement is advantageously carried out whenever a parameter or functional change has occurred in the hearing aid device 11. The determined data relating to the signal transit time are finally transmitted via a transmitting and receiving unit 20 to the second hearing aid device (not shown) of the hearing aid system. Likewise, the hearing aid device 11 also receives the instantaneous signal transit time by means of the transmitting and receiving unit 20 through the signal processing unit of the second hearing aid device. In the analysis and control unit 17 is thus the information regarding the signal propagation times of both hearing aids of the hearing aid system. In the hearing aid with the shorter signal propagation time determined, in the exemplary embodiment the hearing aid device 11, a signal delay is subsequently performed by the difference of the signal propagation times determined in both hearing aid devices. For this purpose, the hearing aid device 11 comprises a delay unit designed as a shift register 21. In this case, the number of delay clocks by the analysis and control unit 17 is adjustable. Advantageously, it is also achieved in this embodiment that the passage of an acoustic input signal in parallel through two hearing aid devices of a hearing aid system, the same signal delay is needed.

Another embodiment of the invention is in FIG. 3 shown. In this case, a hearing aid 22 shows an embodiment according to FIG. 2 very similar construction. In contrast to the hearing aid 11 according to FIG. 2 However, the hearing aid 22 has a clock generator 23 with adjustable clock frequency. By means of the adjustable clock generator 23, the system clock of the hearing aid 22 is adjustable. Depending on the system clock so that the signal delay of a signal by the hearing aid 22 is changeable. Is used in an analogous manner to the in FIG. 2 described hearing aid that the signal propagation time compared to a second Hearing aid of the hearing aid system is longer, so the clock frequency is increased to compensate for the difference in transit time until the delay difference is balanced. Accordingly, the clock frequency of the hearing aid device 22 is reduced so far that the signal propagation times are equalized at a shorter signal propagation time determined for the hearing aid device 22.

In a preferred embodiment of the invention, in addition to the compensation of the signal propagation times with changed settings and functions of at least one hearing aid device, an amplitude compensation also takes place. For this purpose, for example, analogously to the compensation of the signal propagation times in the hearing aid devices 1 and 1 'according to FIG. 1 Gain values are determined and related data stored in the memory units 6 and 6 '. In a gain change in one of the two hearing aids as a result of a parameter and / or functional change (eg changing the hearing program) then the gain in the other hearing aid is adjusted accordingly.

Also with the example in the FIGS. 2 and 3 illustrated hearing aids, an amplitude compensation can be done. Here, a test signal is advantageously fed via the measuring device 19 into the signal path and picked up again at a later point in the signal path, preferably after the signal processing unit 13. In addition to the signal propagation time, the signal transmission behavior with regard to the signal amplitudes is also measured so advantageously. Preferably, the measurement is carried out at different frequencies. Thus, a different gain value can be determined for different frequencies. Data relating to the amplification values thus determined are then transmitted to the respective other hearing aid device of the hearing aid system. Subsequently, an adjustment of the signal amplitudes takes place, with at least one of the hearing aid devices changing the gain or setting filter means. Advantageously, the adjustment follows the signal amplitudes taking into account the audiograms measured in both ears. Data regarding these audiograms may also be stored in the storage units 18. The loudness compensation then takes place in relation to the audiograms, which ensures that, for example, a slight change in loudness caused by a parameter change on a hearing aid causes a loudness change that is subjectively the same to the hearing aid wearer on the other hearing aid. As a result, slight loudness differences in the two ears of a hearing aid wearer are always perceived the same regardless of the current hearing aid settings.

Another embodiment of the invention is in FIG. 4 shown. Also FIG. 4 shows only a hearing aid 24 of a hearing aid system with two identically constructed hearing aids. The hearing aid device 24 comprises two microphones 25 and 26 whose output signals are fed to a signal preprocessing unit 27. In the signal pre-processing unit 27, an A / D conversion and an electrical connection of the microphone signals to produce a directional microphone characteristic. A filter bank 28 serves to split the electrical signal into frequency bands. In signal processing units 29A, 29B, 29C and 29D, a frequency band-specific signal processing of the electrical signals in the individual frequency bands then takes place. Finally, the output signals of the signal processing units 29A to 29D are added and post-processed in a signal post-processing unit 30. The signal post-processing may include, for example, an end gain and D / A conversion. Finally, the analog electrical output signal is converted back into an acoustic output signal by a receiver 31. The individual signal processing blocks of the hearing aid device, that is to say the signal preprocessing unit 27, the filter bank 28, the signal processing units 29A to 29D in the individual channels and the signal postprocessing unit 30, are shown in FIG Embodiment summarized referred to as signal processing unit 29.

Also in the hearing aid device 24 in this embodiment, different hearing programs for adapting the signal processing in the hearing aid to different listening situations are provided. Corresponding parameter sets are stored in a memory unit 32. For recognizing the instantaneous hearing situation, the hearing aid device 24 has a signal analysis and control unit 33, into which the electrical input signal before dividing into different frequency bands and the electrical output signal after passing through the signal processing units 29A to 29D. By means of the signal analysis and control unit 33, for example, feedback-related oscillations in the electrical input signal can be detected. As a countermeasure against detected feedback oscillations, for example, in a frequency band in which the oscillation frequency is located, the gain may be reduced. Data relating to this gain change in the relevant channel are then acquired by the signal analysis and control unit 33 and transmitted by means of a transmitting and receiving unit 34 to the second hearing aid (not shown). This receives the transmitted data and in turn lowers the gain in the corresponding channel by means of a signal analysis and control unit of the hearing aid 24 corresponding signal analysis and control unit. Likewise, data concerning a change in gain in the second hearing aid device of the hearing aid system can also be transmitted to the hearing aid device 24, which acts on the components (for example the signal processing units 29A to 29D in the individual channels) by means of the signal analysis and control unit 33 and the amplification in the hearing aid device 24 adapts accordingly.

The gain change can be made in both hearing aids by the same amount. Preferably, however, it is done taking into account the individual hearing loss of the hearing aid wearer and the signal transmission characteristics of the hearing aid devices. The hearing aid wearer then subjectively perceives the same gain reduction on both hearing aids. Natural loudness differences in the acoustic input signals remain largely preserved for the hearing aid wearer.

Frequently, parameter or functional changes in hearing aids do not result in predetermined gain changes due to the current listening situation. This is the case, for example, with hearing aid devices in which complete sets of parameters for adaptation to different listening situations are not predetermined, but in which an adaptive and continuous adaptation of individual parameters takes place. A gain change is then advantageously determined by a hearing aid internal measurement. Thus, in the hearing aid according to FIG. 4 the gain change can be determined from measurements of the gain before and after a parameter change. For this purpose, the electrical input signal and the electrical output signal in the signal analysis and control unit 33 are evaluated. In the embodiment according to FIG. 4 Both an evaluation of the total input or output signal and the electrical input and output signals of the signal processing units 29A to 29D of the individual channels is possible, depending on whether a change in the parameter affects the entire frequency range or only signal frequencies within a frequency band.

Analogous to the adjustment of the amplification can in a hearing aid system with two hearing aids with a schematic block diagram according to the exemplary hearing aid device 24, as in FIG. 4 represented, the signal amplitudes or the signal propagation times of the two hearing aids are adapted to each other, so that the natural directional hearing is maintained even with worn hearing aids. Here are for the amplitude or delay compensation over the gain compensation merely provide other signal analysis methods in the signal analysis and control unit 33. For example, the amplitude compensation is preceded by amplitude or level measurements or the delay compensation phase or signal propagation time measurements on the overall signal or in the individual channels of the hearing aid device 24. The compensation is then preferably carried out by adjustable filter means within the signal processing unit 29, which are set by the signal analysis and control unit 33.

In a preferred variant, a correlation analysis is performed for the transit time measurement. For this purpose, the signal analysis and control unit 33 electrical signals from successive points in the signal path between the microphones 25 and 26 and the handset 31 are supplied. By means of the correlation analysis, the phase shift and thus the signal propagation time can then be determined in a simple manner.

In a further preferred variant, the envelopes of the supplied signals are first determined in the signal analysis and control unit. The comparison of the envelopes in the signal analysis and evaluation unit 33 also makes it easy to deduce the phase shift of the relevant signals and thus the signal propagation time between the points of interest in the signal path of the hearing aid device 24.

The measurements are carried out in each case shortly before and shortly after parameter or functional changes in the hearing aid device 24 in order to detect the resulting amplification and / or amplitude and / or signal propagation time changes in the hearing aid device 24, related data to the second hearing aid device of the hearing aid system transmit, receive there, evaluate and finally compensate for the changes.

In summary, it is stated:

In the binaural care of a hearing aid wearer with two hearing aids that can be worn on the ear, directional hearing should be improved. For this purpose, the invention proposes to measure in each case signal propagation times and / or signal amplitudes and / or amplifications of an electrical signal in a signal path between an input transducer and an output transducer of a hearing aid and data relating to the measured signal propagation times and / or signal amplitudes and / or amplifications to the other Hearing aid to transfer. As a result, the signal propagation times and the signal amplitudes of the electrical signals can be matched to one another by the two hearing aid devices. Thus, no phase or amplitude distortion caused by the hearing aids and the natural phase shift and the natural amplitude difference of an incident from a particular direction sound signal are maintained. Thus, the direction information for the hearing aid wearer is retained.

Claims (24)

A method for setting a hearing aid system having at least one first (1, 11, 22, 24) and a second (1 ') hearing aid, each having at least one input transducer (2, 2', 12, 25, 26) for receiving an acoustic input signal and Conversion into an electrical signal, a signal processing unit (3, 3 ', 13, 29) for processing the electrical signal and an output transducer (4, 4', 14, 31) for converting the electrical signal into an output signal and between which a signal path (10) is provided for data transmission, characterized
characterized in that a gain or gain change of the electrical signal in the signal path between the input transducer (2, 12, 25, 26) and the output transducer (4, 14, 31) of the first hearing aid (1, 11, 22, 24) is determined and a signal via the signal path (10) to the second hearing aid (1 ') is transmitted to adjust the gain of the electrical signal in the signal path between the input transducer (2') and the output transducer (4 ') of the second hearing aid (1') the determined amplification of the electrical signal in the first hearing aid (1, 11, 22, 24). Method according to claim 1, characterized
characterized in that the gain or gain change of the electrical signal for a portion of the signal path between the input transducer (2, 12, 25, 26) and the output transducer (4, 14, 31) of the first hearing aid (1, 11, 22, 24) determined becomes. A method according to claim 1 or 2, characterized
in that the gain or gain change of the electrical signal in the first hearing aid device (1, 11, 22, 24) is automatically determined and a signal is transmitted to the second hearing aid device (1 '). Method according to one of claims 1 to 3,
characterized in that for determining the amplification or gain change, a test signal is generated which determines the signal path between the input transducer (2, 12, 25, 26) and the output transducer (4, 14, 31) of the first hearing aid device (1, 11, 22, 24) at least partially passes. Method according to one of claims 1 to 4,
characterized in that signal amplitudes and / or signal levels of the electrical signal are determined to determine the gain or gain change. Method according to one of claims 1 to 5,
characterized in that the gain or gain change in the first (1, 11, 22, 24) and the second (1 ') hearing aid device is determined and in each case a signal is transmitted to the other hearing aid device. Method according to one of claims 1 to 6,
characterized in that for adjusting the gain filter means (9, 9 ') can be adjusted. Method according to one of claims 1 to 7,
characterized in that the determination and adjustment of the gain or gain change take place at periodic intervals. Method according to one of claims 1 to 8,
characterized in that the determination and adjustment of the gain following a parameter and / or functional change in at least one of the hearing aid devices (1, 1 ', 11, 22, 24) take place. Method according to one of claims 1 to 9,
characterized in that the signal processing in the first (1, 11, 22, 24) and the second (1 ') hearing aid in several parallel frequency channels of the respective signal processing unit (3, 13, 29) and the determination and adjustment of the gain in each case at least one frequency channel. A method for setting a hearing aid system having at least one first (1, 11, 22, 24) and a second (1 ') hearing aid, each having at least one input transducer (2, 2', 12, 25, 26) for receiving an acoustic input signal and Conversion into an electrical signal, a signal processing unit (3, 3 ', 13, 29) for processing the electrical signal and an output transducer (4, 4', 14, 31) for converting the electrical signal into an output signal and between which a signal path (10) is provided for data transmission, thereby
characterized in that a signal amplitude of the electrical signal in the signal path between the input transducer (2, 12, 25, 26) and the output transducer (4, 14, 31) of the first hearing aid (1, 11, 22, 24) is determined and a signal is transmitted to the second hearing aid (1 ') via the signal path (10) to adapt the signal amplitude of the electrical signal in the signal path between the input transducer (2') and the output transducer (4 ') of the second hearing aid (1') to the determined Signal amplitude of the electrical signal in the first hearing aid (1, 11, 22, 24). A method according to claim 11, characterized in that the signal amplitude of the electrical signal in the first hearing aid (1, 11, 22, 24) is automatically determined and a signal to the second hearing aid (1 ') is transmitted. A method according to claim 11 or 12, characterized
characterized in that for determining the signal amplitude, a test signal is generated, the signal path between the input transducer (2, 12, 25, 26) and the output transducer (4, 14, 31) of the first hearing aid (1, 11, 22, 24) at least partially passes through. Method according to one of claims 11 to 13,
characterized in that the signal amplitude in the first (1, 11, 22, 24) and the second (1 ') hearing aid device is determined and in each case a signal is transmitted to the other hearing aid device. Method according to one of claims 11 to 14,
characterized in that for adjusting the signal amplitude filter means (9, 9 ') can be adjusted. Method according to one of claims 11 to 15,
characterized in that the determination and adaptation of the signal amplitude take place at periodic intervals. Method according to one of claims 11 to 16,
characterized in that the determination and adaptation of the signal amplitude following a parameter and / or function change in at least one of the hearing aid devices (1, 1 ', 11, 22, 24) take place. Method according to one of claims 11 to 17,
characterized in that the signal processing in the first (1, 11, 22, 24) and the second (1 ') hearing aid in several parallel frequency channels of the respective signal processing unit (3, 13, 29) and the detection and adaptation of the signal amplitudes in each case at least one frequency channel. Hearing aid system comprising at least a first (1, 11, 22, 24) and a second (1 ') hearing aid, each having at least one input transducer (2, 2', 12, 25, 26) for receiving an acoustic input signal and conversion into an electrical Signal, a signal processing unit (3, 3 ', 13, 29) for processing the electrical signal and an output transducer (4, 4', 14, 31) for converting the electrical signal into an output signal and between which a signal path (10) for Data transmission is provided,
characterized in that the first hearing aid (1, 11, 22, 24) comprises means for measuring and means for transmitting data relating to a gain or gain change of an electrical signal in the signal path between the input transducer (2, 12, 25, 26) and the Output transducer (4, 14, 31) of the first hearing aid (1, 11, 22, 24) and the second hearing aid (1 ') means for receiving the transmitted data and means for adjusting a gain in the signal path between the input transducer (2') and the output transducer (4 ') of the second hearing aid (1') to the gain or gain change of the electrical signal in the first hearing aid (1, 11, 22, 24). Hearing aid system according to claim 19, characterized
characterized in that the signal processing in the first (1, 11, 22, 24) and the second (1 ') hearing aid device takes place in several parallel frequency channels of the respective signal processing unit (3, 3', 13, 29) and at least the first hearing aid device (1 , 11, 22, 24) means for determining the gain or gain change and at least the second hearing aid (1 ') comprises means for adjusting the gain in at least one frequency channel. Hearing aid system with at least one first (1, 11, 22, 24) and a second (1 ') hearing aid, each having at least one input transducer (2, 2', 12, 25, 26) for receiving an acoustic input signal and conversion into an electrical signal, a signal processing unit (3, 3 ', 13, 29) for processing the electrical signal and an output transducer (4, 4', 14, 31) for converting the electrical signal into an output signal, and between which a signal path (10) is provided for data transmission,
characterized in that the first hearing aid device (1, 11, 22, 24) means for measuring and means for sending data regarding a signal amplitude of an electrical signal in the signal path between the input converter (2, 12, 25, 26) (and the output transducer 4, 14, 31) of the first hearing aid (1, 11, 22, 24) and the second hearing aid (1 ') means for receiving the transmitted data and means for adjusting a signal amplitude in the signal path between the input transducer (2') and the Output transducer (4 ') of the second hearing aid device (1') to the signal amplitude of the electrical signal in the first hearing aid device (1, 11, 22, 24). Hearing aid system according to claim 21, characterized
in that the signal processing in the first (1, 11, 22, 24) and the second (1 ') is carried hearing aid device in several parallel frequency channels of the respective signal processing unit (3, 3', 13, 29) and at least the first hearing aid device (1 , 11, 22, 24) means for determining the signal amplitude and at least the second hearing aid device (1 ') comprises means for adjusting the signal amplitude in at least one frequency channel. Hearing aid system according to claim 21 or 22,
characterized in that the first hearing aid device (1, 11, 22, 24) at least one transmitting unit (8, 8 ', 20, 34) and the second hearing aid device (1') at least one receiving unit (8 ') for wireless signal transmission between the first (1, 11, 22, 24) hearing aid device and the second (1 ') hearing aid device. Hearing aid system according to one of claims 21 to 23,
characterized in that at least the first (1, 11, 22, 24) hearing aid device comprises means for generating a test signal.

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