EP2172063B1 - User-adaptable hearing aid comprising an initialization module - Google Patents

Description

The invention relates to a hearing aid, in particular a medical hearing aid, with a device for compensating for noise. Preferably, the hearing aid compensates for a hearing loss. The invention further relates to a corresponding method for operating and adjusting a hearing device according to the invention.

The medical need for hearing aids is large and increasing, and the available devices cover a wide spectrum from simple, behind-the-ear broadband amplifiers to sophisticated and highly miniaturized devices that fit into the user's ear canal ,

An essential quality feature of hearing aids of every degree of miniaturization is the adaptability of the amplification and the frequency response of the internal amplifiers to the individual hearing impairment of the user. In practice, no hearing loss equals the other (apart from total deafness, which, however, can not be corrected with the hearing aids described here), so that a corresponding adaptability of the hearing aid is required to correct a hearing loss. If this adjustment is omitted and the sound is uniformly amplified only over the entire processable frequency range, this leads to tones in frequency ranges in which the user still hears well being amplified too much, and in the worst case the hearing is even further damaged. On the other hand, in the affected frequency bands where higher amplification would be required, the broadband amplification is usually too low with respect to the undamaged spectral regions.

The adjustment of the gain of a hearing aid according to the prior art is made by a hearing care professional on the basis of an audiogram which he or an ENT doctor has previously detected by the patient. Among other things, with the help of a calibrated headphones, different sounds with increasing volume played to the patient, where he should indicate from which volume a sound is perceived. In this way, the individual frequency response, in particular the lower hearing threshold of the patient's hearing at different frequencies is determined. The more different frequencies are used, the higher the spectral resolution of the audiogram; and the more often the measurement and the same tone is repeated, the greater the statistical certainty for this reading. The thus determined audiogram provides information about the areas of the auditory spectrum in which the patient requires amplification; and the hearing care professional then adjusts the gain of the hearing aid for different spectral ranges accordingly. As a check, an audiogram with a hearing aid should then be recorded again in order to document its use and to check its setting. Ideally, this new audiogram will be equivalent to that of average normal hearing. However, this ideal is rarely achieved because the acoustician's settings are usually not precise enough, and most hearing aids do not allow a sufficiently high-resolution adjustment of the gain's frequency response. Most of the devices used have only three separately adjustable ranges for high, medium and low frequencies, which forces the hearing care professional to make significant compromises in his work.

In addition to balancing the hearing curve of the patient, the "pain threshold" of the patient must be taken into account when setting the hearing aid. Even if the patient's hearing impairment is perfectly adjusted but linearly amplified, the patient would be able to follow quiet conversations, but loud sound events would be amplified to such an extent that painful or even harmful volumes would result. This is particularly relevant if the present perceived hearing impairment reduces the perceived painful volume. The prior art usually solves this problem by structurally limiting the maximum output volume of a hearing aid. Due to the small size and the limited available electrical energy, the maximum volume is naturally limited. In addition, even the simplest devices usually have a volume control that allows the user to adjust the volume of their hearing aid, e.g. to different environmental situations. High-quality hearing aids make such a situation-dependent adjustment automatically and not only adjust the volume but also optimize your frequency response to the respective situation (for example, conversation, music, street noise). However, such situation-dependent adaptation, whether automatic or manual, goes beyond the medical aspect of restoring normal hearing.

The audiogram and the volume-pain threshold are the decisive data for the analytical characterization of hearing loss. The by ENT doctor or

Hearing aid acousticians often additionally recorded data during the hearing test to understand syllables (for example the Freiburg word test) correspond to the state of the art, but can certainly be regarded as superfluous with regard to the possibilities and limitations of a hearing aid.

Another technical problem independent of the patient's hearing damage arises from the fact that - especially in the case of highly integrated devices - only a small spatial distance between sound recording (microphone) and sound generation (miniature loudspeakers, frequently referred to as "transducers" in the hearing device) is always simply " Speaker ") is present. There is a risk that for individual frequencies, the loop gain between the speaker and microphone is greater than one and this leads to feedback whistles. This problem is often solved by attenuating critical frequencies with additional narrow band filters ("notch filters"). Although this suppresses the feedback or oscillation tendency of the system, these additional filters have an undesirable effect on the frequency response, in particular they possibly counteract the actually required high amplification in the regions of the spectrum in which the patient hears badly.

In the case of hearing aids of the highest price categories, other methods of digital signal processing that go beyond the described methods are known in the prior art. For example, one tries to differentiate between speech and noise components in the sound signal in order to remove or at least reduce the latter. In such methods for noise suppression, which are also known from other fields of application, but there are different side effects to consider: For example, in some methods with the attenuation of the noise also alienation of the useful portion, such as the voice portion, accompanied, and the sound of the muted sounds is noticeably changed. In addition, some methods cause a signal delay that can be accepted in a hearing aid only within very narrow limits, since otherwise the seen and the heard are no longer time-synchronized, which can lead to perception disorders in the hearing aid wearer.

The EP 1 713 302 A1 shows a hearing aid to the arrangement of and / or in an ear, wherein the hearing aid for the device of the hearing aid is connectable to a control unit. Furthermore, the hearing device includes a microphone for converting acoustic signals into electrical signals, a hearing module for processing the electrical signals, a loudspeaker for converting the electrical signals output by the hearing module into acoustic signals. The hearing module has a device for noise suppression. To determine control parameters (in particular the parameters of the noise suppression), the hearing aid uses a classification of the hearing situation by means of a surrounding detector. The user can change the parameters of the hearing aid during the device, in particular the parameters of the noise suppression.

The EP 1 542 500 A shows a hearing aid with noise reduction, which makes a noise estimate.

The object of the invention is to provide an improved hearing aid which overcomes the disadvantages mentioned above. In particular, a hearing aid is to be provided which provides improved noise suppression and preferably in interaction with can be adjusted to the user. Furthermore, a corresponding method is to be provided.

The object is solved by the features of the claims.

According to the invention, parameters of an adjustable filter are changed by means of a noise estimation so that a noise suppression can be performed, which leads to the hearing aid user to a real acoustic perception image. For this purpose, damping factors, for example at certain time intervals or on a continuous basis, can be determined. The parameters of an optional hearing loss compensation and noise suppression can be combined in such a way that the signal to be processed is adjusted in one calculation step per frequency band or discrete frequency.

According to one embodiment of the invention, an audiogram, ie the spectral characteristic of the user's hearing, is automatically determined in interaction with the user during an initialization phase and the data obtained is used for internal signal processing, preferably digital signal processing such as multiband equalizers and limiters. Compressor to adjust so that an ideal compensation of individual hearing damage results. The data obtained, i. Correction factors for compensation of the hearing damage are stored, preferably in a non-volatile storage medium. Preferably, the user can perform the determination of an audiogram again at any time or optimize existing data. The correction factors may also already be fixed or predefined as the starting point for a setting of the audiogram by the user, for example by a physician or hearing aid acoustician.

The hearing test, ie the determination of the audiogram of a patient, can be transferred to the hearing aid itself. This makes it possible for the hearing aid to automatically adjust the frequency response of its amplification in a closed system without an audiogram being interpreted by a hearing aid acoustician. Thereby, the individual hearing damage of a patient can be compensated exactly, because the parameters of the internal signal processing are determined by the hearing aid itself in an initialization mode, which is to be distinguished from the operating mode in which the parameters are applied. in the Initialization mode, the hearing aid outputs test signals; signals picked up by the own microphone are preferably at least partially not supplied to the sound output of the hearing device. There is no need for calibrated gauges as in a traditional listening test, prior calibration of the hearing aid itself is also unnecessary, and the impact of the physical presence of the hearing aid in the auditory canal on hearing is considered intrinsically.

Figur 1

eine schematische Darstellung der Komponenten eines erfindungsgemäßen Hörgeräts;

Figur 2

eine schematische Darstellung eines Hörmoduls eines Hörgeräts gemäß Figur 1;

Figur 3

eine schematische Darstellung eines Initialisierungsmoduls eines Hörgeräts gemäß Figur 1;

Figur 4

eine schematische Darstellung einer Hörkurvenkorrektur in einem Hörmodul gemäß Figur 2;

Figur 5a

eine schematische Darstellung einer ersten Ausführungsform einer Geräuschunterdrückung in einem Hörmodul gemäß Figur 2;

Figur 5b

eine schematische Darstellung einer zweiten Ausführungsform einer Geräuschunterdrückung in einem Hörmodul gemäß Figur 2;

Figur 6

eine schematische Darstellung einer Lautstärkenbegrenzung in einem Hörmodul gemäß Figur 2;

Figur 7

ein Flussdiagramm zur Ermittlung eines Audiogramms gemäß der Erfindung;

Figur 8

ein Flussdiagramm zur Ermittlung einer maximal akzeptablen Lautstärke gemäß der Erfindung; und

Figur 9

eine schematische Darstellung der Bestimmung des Anti-Feedback-Filters gemäß der Erfindung.

The functioning of the hearing device and the corresponding method are described below with reference to preferred embodiments with reference to the figures; show it:

FIG. 1

a schematic representation of the components of a hearing aid according to the invention;

FIG. 2

a schematic representation of a hearing module of a hearing aid according to FIG. 1 ;

FIG. 3

a schematic representation of an initialization module of a hearing aid according to FIG. 1 ;

FIG. 4

a schematic representation of a hearing curve correction in a hearing module according to FIG. 2 ;

FIG. 5a

a schematic representation of a first embodiment of a noise suppression in a hearing module according to FIG. 2 ;

FIG. 5b

a schematic representation of a second embodiment of a noise suppression in a hearing module according to FIG. 2 ;

FIG. 6

a schematic representation of a volume limit in a hearing module according to FIG. 2 ;

FIG. 7

a flowchart for determining an audiogram according to the invention;

FIG. 8

a flow chart for determining a maximum acceptable volume according to the invention; and

FIG. 9

a schematic representation of the determination of the anti-feedback filter according to the invention.

FIG. 1 shows a schematic representation of a hearing aid according to the invention, which is located on or in the human ear, with its components microphone 1, 2 initialization module 2, hearing module 3 and 4 speakers, the initialization module 2 is in communication with a controller 5, via which the user during the initialization interacts with the device. In a preferred embodiment, the Hearing aid further an analog-to-digital converter 6 and a digital-to-analog converter 7, as in FIG. 1 shown on. As a feedback path, the acoustic feedback path is shown, via which sound from the speaker 4 passes back into the microphone 1 and can lead to feedback whistles.

It is noted that the initialization module 2 and the control unit 5 represent optional features of the hearing aid according to the invention.

The hearing module 3 has a device for noise suppression, which performs a noise estimation for determining the parameters of a signal-dependent filter.

For optional adjustment of the hearing module 3 of the hearing aid to the individual hearing damage of a user - ie the deviation from the normal hearing curve - by a correspondingly amplified speaker output of the recorded sound from the microphone 1, an initialization is performed. For this purpose, according to one embodiment, an interaction between the user and the hearing device is provided, which is provided by controls on the hearing aid itself or by a wireless or wired connection to an operating aid, e.g. a PC takes place; This operating aid is generally referred to below as the control unit 5. The control unit has at least one actuating device, which has a switch and / or a push button.

The signal flow in the hearing device is as follows: The microphone signal s _M (t) is preferably discretized and digitized by an analog-to-digital converter 6 and fed to the hearing module 3 and the initialization module 2, where the signal processing, preferably a digital signal processing, takes place. Subsequently, in the case of a digital signal processing, a digital-to-analog converter 7 generates an output signal s _L (t), with which a loudspeaker 4 sonicates the user's ear.

FIG. 2 shows the hearing module 3 with a summer 31, which adds a calculated from the anti-feedback filter 32 negative pseudo feedback to the microphone signal, an optional hearing curve correction 33 by frequency-dependent signal amplification, a noise suppression 34 and a volume limit 35 of the output speaker signal. The calculation of the negative pseudo-feedback is done by discrete Convolution of the impulse response of the feedback path with the loudspeaker signal s _L (t) to be output.

The following are the components and function of the in FIG. 2 shown listening module 3 with respect to the FIGS. 4 to 6 explained in more detail.

The hearing module 3 receives the digital microphone signal s _M (t) and adds to this the negative pseudo-feedback s _F (t), which with the aid of the impulse response of the feedback path h (t) as a discrete convolution with the loudspeaker signal s _L (t) to s _F ( t) = h (t) * s _L (t) is calculated in order to remove the feedback of the loudspeaker signal into the microphone 1 from the microphone signal and thus prevent feedback whistling. Subsequently, the optional hearing curve correction 33, as more detailed in FIG. 4 shown by a system of different filters with center frequencies f _i = f ₁ ... f _n and gains V (f _i ) are applied to the signal, the quality of the filter is chosen so that the superposition of all filters one gives a constant frequency response if all gains V (f _i ) have the same value, ie V (f ₁ ) = V (f ₂ ) = V (f ₃ ) = ... = V (f _n ). The V (f _i ) values are to be adapted as exactly as possible to the individual hearing damage, so that when using the hearing device, the user's hearing curve comes as close as possible to that of an average normal hearing aid.

The optional hearing curve correction in the hearing module 3 is performed by a series of independent filters, preferably IIR filters. The individual adjustment of the V (f _i ) values for the correction of the hearing damage takes place with the aid of the initialization module 2.

After the optional hearing curve correction 33 follows a as in FIG. 5a illustrated first embodiment of a noise suppression 34, as for example DE 199 48 308 A1 is known. The signal is subjected to a Fourier transformation in order, for example, to obtain an estimate of the noise spectrum by minimadetection in the spectrum. This noise estimate is used to determine a noise and signal dependent filter or the filter coefficients of a filter applied to the signal spectrum. The latter is then converted back to a noise-reduced time signal by inverse Fourier transform, which is provided at the output of the noise suppression 34.

Alternatively, with the aid of IIR filters, the optional hearing curve correction can alternatively also be implemented as a filter in the spectrum, as described below with reference to a second embodiment according to FIG FIG. 5b is set out. For this purpose, the signal is first subjected to a Fourier transformation, so that the correction factors K (f) can be used to compensate for a hearing impairment, directly as a multiplication in the signal spectrum, under the boundary condition that the frequencies f _i lie in the frequency raster of the Fourier transformation. Preferably, the correction factors K (f) correspond to the gain values V (f _i ). This embodiment can be advantageously combined with the application of noise cancellation. For this purpose, the signal spectrum is additionally multiplied by signal- and noise-dependent damping factors (gain factors) G (f). The attenuation factors are preferably determined from a noise estimate R (f) and the current signal spectrum S (f), eg as G (f) = 1-R (f) / S (f). The noise estimate is formed from the signal spectrum by averaging it over those time intervals in which the signal consists essentially only of noise, and no or only a negligible useful signal component (voice) is present. For example, a good noise estimate in a speech break, in which no useful signal component is present, are performed. FIG. 5b shows the combined application of hearing curve correction by means of correction factors K (f) and noise suppression by means of damping factors G (f). To FIG. 5b For example, the signal spectrum S (f) is supplied both to a determination of a noise estimate R (f) and to a multiplication in the spectrum with correction factors K (f). After the determination of the noise estimate R (f), a determination is made of attenuation factors G (f) based on the noise estimate R (f). After the multiplication of the microphone signal in the spectrum with the correction factors K (f) for compensation of the hearing impairment is in accordance with FIG. 5b a multiplication in the spectrum with damping factors G (f) performed. As a result, the signal, for example, the external conditions, such as subway, apartment, concert hall, etc., to be adjusted. After the corresponding arithmetic operations in the spectrum, the thus modified signal spectrum is converted back into a hearing curve-corrected, noise-reduced time signal which is provided at the output of the filter module 34 by means of inverse Fourier transformation.

It is noted that the hearing curve correction is optional and the corresponding device feature or step may be omitted.

The signal processing as in FIG. 5b shown, can be changed. For example, the order of multiplication in the spectrum with correction factors K (f) and the multiplication in the spectrum with damping factors G (f) can be reversed. According to a further alternative, the multiplication in the spectrum with correction factors K (f) and the multiplication in the spectrum can be combined with attenuation factors G (f) and preferably take place in one step per frequency band or discrete frequency. For this purpose, the attenuation factors G (f) are preferably multiplied by the correction factors K (f), and only then is the signal spectrum S (f) multiplied by the result of this multiplication of the two factors. This has the advantage that the real-time signal (microphone signal) only has to undergo multiplication, so that overall the signal processing time can be shortened.

The damping factors G (f) are determined based on the noise estimate R (f), which is preferably renewed at certain intervals and / or adaptively in order to be able to take into account a change in the noise environment. Under adaptive, a continuous automatic, i. automatic, noise estimation understood. In addition to fixed time intervals, dynamic factors can be used, which trigger a new noise estimate. A dynamic trigger factor may be a manual user input. A user preferably chooses a moment in which as little useful signal as possible is present. Also, a pre-selection of the environment by the user with a subsequent optimization of the noise estimate can be performed. Fixed time intervals for determining a new noise estimate can be combined with dynamic triggering factors.

The damping factors can also be applied only partially or not at all, ie they can be changed. This can be like in FIG. 5b given formula can be modified to G (f) = 1 - c * R (f) / S (f), where 0 <c <1. Thus, the extent of noise suppression can be set automatically or manually by the user. At c = 0 the noise suppression is deactivated and at c = 1 the noise suppression is fully active. With the adjustability of the noise suppression, the signal emitted by the loudspeaker (4) can be matched as accurately as possible to a real acoustic environment. For example, noise suppression could detect sea noise or the noise of leaves in the forest as a jamming signal and consequently suppress it, although it would not be desirable by the user in this case.

The last step of the signal processing in the hearing module 3 before the output of the signal to the digital-to-analog converter 7 and the loudspeaker 4 is to limit the maximum output volume to a maximum value M so as not to exceed the user's individual pain threshold. For this purpose, preferably a characteristic curve, as in FIG. 6 is used, which is linear for subcritical signal volumes, approaching the threshold M when the threshold of pain is reached, without exceeding even the threshold for even greater input levels. The threshold M is preferably determined in the initialization module in interaction with the user.

According to the invention, with the aid of the optional initialization module 2, the individual setting of the parameters of the hearing module 3 for the ideal compensation of the user's personal hearing deficit is undertaken. For this purpose, the optional control unit 5 is used, with which the user interacts with the initialization module 2. As in FIG. 3 shown schematically, the hearing curve of the patient is measured in the initialization 2 by outputting tones or acoustic signals increasing volume. In particular, the initialization module outputs electrical signals which are converted into tones or acoustic signals. Then the hearing curve is determined relative to the hearing curve of an average normal listener and determines the appropriate filter to compensate for the individual Hördefekts. In addition, the pain threshold of the user is measured by outputting noise of increasing volume. So the maximum tolerable output volume is determined, which is also individual for each user. Preferably, at the same time - by resuming the test signals output by the loudspeaker 4 from the microphone 1 - the impulse response of the feedback path is determined, which is used to eliminate feedback in the anti-feedback filter 32 of the hearing module 3.

In the initialization mode according to an embodiment of the invention, the initialization module 2 outputs a sequence of electrical signals to the loudspeaker 4, which are converted into acoustic signals, the acoustic signals serving to measure the auditory curve of the user. The acoustic signals have a certain frequency or a specific frequency spectrum with a certain center frequency in order to determine a lower hearing threshold of the user as a function of the respective frequency. In a preferred embodiment, the transmission from the microphone 1 to the loudspeaker 4 interrupted while the initialization module 2 is operated to measure the hearing curve of the user.

Preferably, the hearing aid according to the invention further comprises a comparator for comparing a user's lower hearing threshold at a certain center frequency with a stored lower hearing threshold of a normal hearing and setting means for setting a gain at the respective frequency, so as to compensate for a hearing defect of the user.

To determine a pain threshold of the user, ie a maximum acceptable volume, which is output to the speaker 4, the initialization module 2 outputs electrical signals, preferably according to a predetermined program, with reference to FIG FIG. 8 will be explained in more detail later. The hearing module 3 limits the speaker output according to the maximum acceptable volume.

FIG. 7 FIG. 10 shows a flowchart of an audiogram measurement and determination of the amplifications V (f _i ) for hearing curve correction according to an embodiment of the invention. In order to determine the hearing curve of the user and the gain parameter V (f _i ) for hearing curve correction, in a first step S1 different test tones are played whose frequencies correspond to the center frequencies f _{i of} the filters which are available for hearing curve correction. For a selected frequency f _i , the volume is initially set to A = A _N , which preferably corresponds to the volume which is just audible by an average normal listener, ie a lower hearing threshold. In step S2, the volume A is now successively increased with the rate of increase to be specified, until the user in step S3-yes signals by pressing a button on the control unit 5 that he has perceived the sound. The corresponding individual lower hearing threshold A (f _i ) is stored in step S4. Subsequently, the procedure is repeated in step S5-no with a different frequency f _i until the hearing curve measurement is terminated by a corresponding user interaction at the control unit 5 and / or a termination condition in step S5-yes. The individual hearing threshold is determined for all frequencies f ₁ , f ₂ , f ₃ ,..., F _n at least once, but preferably several times in order to achieve a certain statistical certainty for the measured values. A possible termination condition can therefore be, for example, a sufficient amount of data collected, ie all lower hearing thresholds of the user at the respective frequency are detected at least once. About the different values of A (f _i ), that is Amplification values at the same frequency, an average value is then formed in step S6, preferably the median, since in this means "outliers" - ie completely erroneous measured values - are not included in the mean value. From this, the gains V (f _i ) are calculated in step S7.

The number of test tones or acoustic signals of the sequence of electrical signals for measuring the hearing curve of the user is preferably between 4 and 128, or between 8 and 64, or between 16 and 48 and particularly preferably 32 different tones, ie that in the particularly preferred Number of tones 32 different frequencies f ₁ to f _{32 are} measured. The amplitude of a sound becoming louder in the measurement of the user's hearing curve is stepped from a minimum volume to a maximum volume preferably in 10 to 200, or in 50 to 150, and more preferably in 100 amplitude values, ie the amplitude is louder At the most preferred number of steps, the incoming tones are changed 100 times from the minimum to the maximum volume.

In a preferred embodiment, the frequencies of the successive test tones or acoustic signals are changed in the measurement in a random order or defined pseudorandom order.

In addition to the optional correction of the personal hearing curve, another element of the digital signal processing of the hearing aid is the limitation of the maximum output volume, which is also individually adapted to the user's hearing. FIG. 8 shows a flowchart of a determination according to the invention of the maximum volume M. For this purpose, in step S 10 noise Rr (t), preferably white noise, generated with an initial volume R = R _N , which corresponds to a volume which is approximately in the middle between the hearing threshold and the pain threshold of an average normal listener. Before the noise signal reaches the user's ear, it is frequency-dependent amplified in step S11 by the previously determined hearing curve correction with the aid of the appropriately set filter V (f _i ). This step is preferred so that the measurement of pain threshold is already tuned to the user's personal hearing. The volume R of the noise signal is now successively increased in step S 12 until the user in step S13-yes via a keystroke on the control unit signals that a volume is reached, which is perceived as painful. If so, the current value of R is stored as the maximum volume M in step 14. This measurement is also preferably repeated several times (step S15-yes) in order to be able to form an average over the different measurements in step S16, so that a certain statistical certainty arises. Preferably, the median for averaging is determined.

The white noise preferably used to determine the maximum volume is preferably output in a frequency band of 0-8 kHz from the initialization module 2 via the loudspeaker 4. The sampling rate used for detecting the feedback signal via the microphone 1 is therefore greater than 16 kHz according to the sampling theorem of Nyquist-Shannon.

The sampling rate when using the hearing aid after initialization is preferably 16 kHz, i. a hearing deficit of a user is corrected in a frequency band of preferably 0 kHz to approximately 8 kHz.

Since white noise is one of the most unpleasant sounds for the human ear, it can be assumed that all other sound events, which are output with the determined maximum volume M, are less critical. Another advantage is the use of (white) noise: the signal is very well suited for determining the impulse response of the feedback path h (t) used in the antifeedback filter 32. For this purpose, the microphone signal s _M (t) is evaluated, preferably while the output speaker signal s _L (t) as described consists of noise signals of different volume for determining the maximum volume M. For example, as can be inferred from the simultaneous analysis of microphone and loudspeaker signal to the impulse response h (t) of the acoustic path between loudspeaker 4 and microphone 1-that is, the feedback path DE 101 40 523 or DE 100 43 064 described.

FIG. 9 shows the determination of the anti-feedback filter 32 and the filter coefficients. From both signals s _M (t) and s _L (t), spectra S _M (f) and S _L (f) are formed on frames of the length to be given by means of Fourier transforms; from S _L (f) also the complex conjugate, S * _L (t) is determined. The product S _M (f) S * _L (f) and the absolute square S _L (f) S * _L (f) are each time-averaged and divided by each other. Thus one obtains the transfer function H (f) of the feedback path, from which by inverse Fourier transform, the impulse response h (t) arises.

After all the desired individual parameters have been determined in the manner described, the initialization module 2 switches to the hearing module 3, and the circle closes: the last-determined impulse response h (t) is required first in the digital signal processing of the hearing module 3. The control unit 5 is not required by the hearing module 3 according to the invention after the initialization, nevertheless it can be used for trivial interactions not described here in more detail, e.g. for user-controlled volume change or a situation-dependent equalizer selection.

This invention has been described by way of examples. It should be emphasized that individual features, examples and embodiments can be combined with each other as desired and thereby further preferred features, examples and embodiments can be formed.

Claims (15)

1. Hearing aid for arrangement at and/or in an ear and connectable to a control device (5), comprising:

   a microphone (1) for converting acoustic signals into electrical signals,

   a hearing module (3) for processing the electrical signals,

   a loudspeaker (4) for converting the electrical signals outputted by the hearing module (3) into acoustic signals, wherein the hearing module comprises means for noise suppression conducting a noise estimation to determine parameters of a signal-dependent adjustable filter and to provide a noise-reduced output signal, wherein

   the means of the hearing module (3) is suitable to ascertain damping factors for the noise suppression, based on the noise estimation, in order to account for a change in the noise environment and wherein

   the damping factors for the noise suppression can be modified additionally, preferably with a factor adjustable by the user, in order to vary the extent of the noise suppression.
2. The hearing aid according to claim 1, wherein the means of the hearing module (3) is suitable to multiply the signal from the microphone in the spectrum with correction factors to compensate for a hearing defect and/or with damping factors of the noise suppression in at least one step per frequency.
3. The hearing aid according to any one of the preceding claims, wherein a noise estimation is carried out at fixed time intervals and/or continuously automatically.
4. The hearing aid according to any one of the preceding claims, wherein the hearing module (3) comprises a means for compensating for a hearing defect by means of correction factors.
5. The hearing aid according to any one of claims 1 to 4, further comprising an initialising module (2) for outputting initialising signals to the loudspeaker (4) and a control device (5) via which a user can interact with the hearing aid in order to individually adjust parameters of the hearing module (3).
6. The hearing aid according to claim 5, wherein the initialising module (2) outputs a series of electrical signals to the loudspeaker (4) which are converted into acoustic signals used for the measurement of the auditory curve of a user, wherein the series of electrical signals preferably corresponds to acoustic signals of a certain frequency or a frequency spectrum with a certain centre frequency for the interactive determination of a lower auditory threshold of the user depending on the respective frequency.
7. The hearing aid according to claim 6, wherein the hearing module (3) compensates for the loudspeaker output in accordance with the deviation of the measured auditory curve from a normal auditory curve and/or wherein the hearing module (3) comprises various filters with different centre frequencies and respectively adjustable amplification.
8. The hearing aid according to claim 7 comprising a comparator for comparing a lower auditory threshold of a user at a certain centre frequency with a stored lower auditory threshold of a person with normal hearing ability and an adjustment means for adjusting an amplification at the respective frequency.
9. The hearing aid according to any one of the preceding claims 5 to 8, wherein the initialising module (2) outputs electrical signals to the loudspeaker (4) which are converted into acoustic signals which are used for determining a pain threshold of the user for a maximally acceptable volume, wherein the hearing module (3) preferably limits the loudspeaker output according to the maximally acceptable volume.
10. The hearing aid according to any one of the preceding claims 6 to 9, wherein in an initialising mode electrical signals are outputted from the initialising module (2) to the loudspeaker (4) according to a predetermined program and wherein the control device (5) preferably comprises a push-button operated by a user as soon as the lower auditory threshold at a centre frequency is reached.
11. The hearing aid according to any one of claims 7 to 10, wherein the electrical signals for determining a pain threshold of the user comprise white noise, wherein preferably the noise signal is amplified depending on the frequency according to the deviation of the ascertained auditory curve from a normal auditory curve.
12. The hearing aid according to any one of the preceding claims comprising an anti-feedback filter (32) for calculating a negative pseudo feedback and a summing unit (31) for adding the negative pseudo feedback to the microphone signal.
13. The hearing aid according to claim 12, wherein the negative pseudo feedback is calculated by a discrete convolution of the impulse response of the feedback path with the loudspeaker signal to be outputted, wherein preferably the impulse response of the feedback path is determined upon initialisation of the hearing aid by evaluation of the microphone signal during the output of loudspeaker signals, preferably of noise signals and particularaly preferably of white noise.
14. Method for noise suppression in a hearing aid according to any one of claims 1 to 13, comprising the following steps:

    converting acoustic signals into electrical signals with a microphone (1);

    processing the electric signals in a hearing module (3);

    converting the electrical signals outputted by the hearing module (3) into acoustic signals with a loudspeaker (4),

    wherein the processing of the electrical signals in the hearing module (3) comprises at least carrying out a noise estimation to determine parameters of a signal-dependent adjustable filter and providing a noise-reduced output signal,

    ascertaining damping factors for the noise suppression based on the noise estimation in order to account for a change in the noise environment and

    modification of the damping factors for the noise suppression, with a factor to be adjusted by the user in order to vary the extent of the noise suppression.
15. The method according to claim 14, comprising the steps of calculating a negative pseudo feedback using an anti-feedback filter (32) and adding the negative pseudo feedback to the microphone signal using a summing unit (31) and preferably comprising the step of calculating the negative pseudo feedback by a discrete convolution of the impulse response of the feedback path with the loudspeaker signal to be outputted, wherein preferably the impulse response of the feedback path is determined upon initialisation of the hearing aid by evaluation of the microphone signal during the output of loudspeaker signals, preferably of noise signals and particularly preferably of white noise.

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