EP2172063A1

EP2172063A1 - User-adaptable hearing aid comprising an initialization module

Info

Publication number: EP2172063A1
Application number: EP08775194A
Authority: EP
Inventors: Dietmar Ruwisch
Original assignee: Individual
Current assignee: Individual
Priority date: 2007-07-18
Filing date: 2008-07-17
Publication date: 2010-04-07
Anticipated expiration: 2028-07-17
Also published as: DK2172063T3; WO2009010572A1; CN101755468A; EP2172063B1; DE102007033484A1; CN101755468B; US20100183177A1; US8406441B2

Abstract

The invention relates to a hearing aid that is to be placed on and/or inside an ear and comprises a microphone (1) for converting acoustic signals into electric signals, a hearing module (3) for supplying the electric signals, and a loudspeaker (4) for converting the electric signals output by the hearing module (3) into acoustic signals. The hearing module (3) is fitted with a noise suppressing device which estimates noise in order to determine a signal-dependent filter and supply a noise-reduced output signal.

Description

HEARING DEVICE WITH INITIALIZING MODULE AND USER ADAPTATION

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 gain and frequency response of the internal amplifiers to the individual hearing loss 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 patient's hearing curve, the patient's "pain threshold" must be taken into account when adjusting the hearing aid, even if the patient's hearing loss is perfectly adjusted, but linear amplification would allow the patient to follow quiet conversations, noisy Sound events, however, are amplified to such an extent that they result in painful or even harmful volumes, which is particularly relevant if the present hearing impairment reduces the perceived painful volume level Due to the small size and the limited amount of available electrical energy, the maximum volume is of course limited, and even the simplest devices usually have a volume control that the user can use can adjust the volume of his 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 a situation-dependent adaptation, whether automatic or manual, goes beyond the medical aspect of normal hearing recovery.

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 exhausters often recorded in the course of a hearing test data for syllabic comprehension (eg, Freiburg Wortter test) Although the state of the art, but with regard to the possibilities and limitations of a hearing aid can be considered quite unnecessary.

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 Rreisverstärkung between speaker and microphone is greater than one and it comes thereby to feedback whistles. This problem is often solved by attenuating critical frequencies with additional narrow-band filters ("notch filters"), which suppress the feedback or oscillation tendency of the system, but these additional filters have an undesirable effect on the frequency response, in particular counteracting it if necessary, the actually required high amplification in the areas 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 an alienation of the useful portion, such as the voice portion, accompanied, and the sound of muffled 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 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 a real acoustic perception image for the hearing device user. 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.

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:

Figure 1 is a schematic representation of the components of an inventive

hearing;

Figure 2 is a schematic representation of a hearing module of a hearing aid according to Figure 1; FIG. 3 shows a schematic representation of an initialization module of a hearing aid according to FIG. 1;

FIG. 4 shows a schematic representation of a hearing curve correction in a hearing module according to FIG. 2; Figure 5 a is a schematic representation of a first embodiment of a

Noise suppression in a hearing module according to Figure 2; Figure 5b is a schematic representation of a second embodiment of a

Noise suppression in a hearing module according to Figure 2; Figure 6 is a schematic representation of a volume limit in a hearing module according to Figure 2;

FIG. 7 shows a flowchart for determining an audiogram according to the invention; FIG. 8 shows a flowchart for determining a maximum acceptable volume according to the invention; and Figure 9 is a schematic representation of the determination of the anti-feedback filter according to the invention.

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, initialization module 2, hearing module 3 and loudspeaker 4, wherein the initialization module 2 is connected to a control unit 5, via which the User interacts with the device during initialization. In a preferred embodiment, the Hearing aid on an analog-to-digital converter 6 and a digital-to-analog converter 7, as shown in Figure 1, 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, a

Interaction between the user and the hearing aid 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 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 O) 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 signal processing, preferably digital signal processing, takes place. Subsequently generated, in the case of a digital

Signal processing, a digital-to-analog converter 7, an output signal s _L (t), with which a speaker 4 sonicates the user's ear.

FIG. 2 shows the hearing module 3 with a summation unit 31 which adds a negative pseudo-feedback calculated by the anti-feedback filter 32 to the microphone signal, an optional hearing curve correction 33 by frequency-dependent signal amplification, noise suppression 34 and a volume limit 35 of the loudspeaker signal to be output. 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 components and the function of the hearing module 3 shown in FIG. 2 are explained in more detail below with reference to FIGS. 4 to 6.

The hearing module 3 receives the digital microphone signal SM (X) and adds to this the negative pseudo-feedback S _F (X), which by means of the impulse response of the feedback path h (t) as a discrete convolution with the loudspeaker signal S _L (X) ZU sp (t) = h (t) * S _L (t) is calculated to remove the feedback of the loudspeaker signal into the microphone 1 from the microphone signal and thus prevent feedback whistling. Subsequently, the optional auditory curve correction 33, as shown in greater detail in FIG. 4, is performed by using a system of different filters with center frequencies and gains V (fj) are applied to the signal, the quality of the filters being chosen such that the superposition of all filters gives the most constant frequency response if all gains V (f;) have the same value, ie V (fi) = V (f ₂ ) = V (f ₃ ) = ... = V (f _n ). The V (f _; ) values should be matched as closely as possible to the individual hearing damage, so that when using the hearing aid, the hearing curve of the user 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 (fi) values for the correction of the hearing damage is carried out with the aid of the initialization module 2.

After the optional hearing curve correction 33, a first embodiment as shown in FIG. 5a is followed by a noise suppression 34, as is known, for example, from DE 199 48 308 A1. 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. Instead of using IIR filters, the optional hearing curve correction can alternatively also be implemented as a filter in the spectrum, as will be explained below with reference to a second embodiment according to FIG. 5b. 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 multiplication in the signal spectrum, under the boundary condition that the frequencies f; lie in the frequency grid of the Fourier transform. Preferably, the correction factors K (f) correspond to the gain values V (fj). 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). According to FIG. 5b, 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) to compensate for the hearing impairment, a multiplication in the spectrum with attenuation factors G (f) is performed according to FIG. 5b. 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 shown in Figure 5b, 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. To this end, the formula given in Figure 5b can be modified to G (f) = 1- c * R (f) / S (f), where 0 <c <1. Thus, the extent of noise suppression can be automatic or manual by the user be set. 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, a characteristic curve is preferably used, as shown in FIG. 6, which runs linearly for subcritical signal volumes, and approaches the threshold M when the pain threshold is reached, without exceeding the threshold even 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 shown schematically in FIG. 3, the hearing curve of the patient is measured in the initialization module 2 by output of tones or acoustic signals of 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 for measuring the

Listening curve of the user serve. 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, i. a maximum acceptable volume, which is output to the speaker 4, the initialization module 2 outputs electrical signals, preferably according to a predetermined program, which will be explained later with reference to Figure 8. The hearing module 3 limits the speaker output according to the maximum acceptable volume.

FIG. 7 shows a flow chart of an audiogram measurement and determination of the amplifications V (f;) for hearing curve correction according to an embodiment of the invention. In order to determine the hearing curve of the user and the amplification parameters V (f;) for hearing curve correction, in a first step S1 different test tones are played whose frequencies correspond to the center frequencies fi of the filters which are available for hearing curve correction. For a selected frequency f; The volume is initially set to A = A _N , which preferably corresponds to the average audible just audible volume, ie a lower threshold hearing. 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 (fi) is stored in step S4. Subsequently, the procedure is repeated in step S 5 -nein another frequency f _f 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 threshold of hearing at least once, but preferably several times determined for all frequencies f, f z, f _3, ..., f _n order to achieve a certain confidence level 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 (fj), that is Amplification values at the same frequency, an average value is then formed in step S6, preferably the median, since this means "outliers" - ie completely erroneous measured values - are not included in the mean value .From this the gains V (fϊ) are calculated in step S 7 ,

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 fi 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 pseudo-random 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 an inventive determination of the maximum volume M. For this purpose, in step S10 R noise (Rr (t), preferably white noise, with a

Initial volume R = RN generated, 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 Sil by the previously determined hearing curve correction with the aid of the appropriately set filter V (fi). This step is preferred so that the measurement of

Pain threshold is already tuned to the personal hearing of the user. The volume R of the noise signal is now successively increased in step S 12 until the user in step S13-jaüber a key press 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 various 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 © is evaluated, preferably while the output loudspeaker signal SiXt) as described consists of noise signals of different volume for determining the maximum volume M. As can be concluded from the simultaneous analysis of microphone and loudspeaker signal on the impulse response h (t) of the acoustic path between loudspeaker 4 and microphone 1 - ie the feedback path - is, for example in detail in DE 101 40 523 or DE 100 43 064 described.

FIG. 9 shows the determination of the anti-feedback filter 32 or the filter coefficients. Of the two signals S _M (O and S _L (O), spectra S _M (Q and S ₁ Xf) are formed on frames of the length to be given by means of Fourier transforms, and S _L (Q is also the complex conjugate, S * _L (I) The product S _M (f) S * _L (f) and the absolute square SL (f) S * L (f) are each time averaged and divided in order to obtain the transfer function H (f) of the feedback path from which by inverse Fourier transformation results in the impulse response h (t).

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 used in the digital signal processing of the

Hearing module 3 is needed first. The control unit 5 is not required by the inventive hearing module 3 after the initialization, nevertheless it can be used for trivial interactions not described here, 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

claims

1. Hearing aid for arrangement on and / or in an ear, which is connectable to a control unit (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 output by the hearing module (3) into acoustic signals, wherein the hearing module has a Has a noise suppression device that makes a noise estimate for determining a signal-dependent filter and provides a noise-reduced output signal.

2. Hearing aid according to claim 1, wherein the device of the hearing module (3) is adapted to determine based on the noise estimate attenuation factors for noise suppression to account for a change in the noise environment.

3. Hearing aid according to claim 2, wherein the device of the hearing module (3) is adapted to multiply the signal from the microphone in the spectrum with correction factors for the compensation of a hearing impairment and / or attenuation factors of the noise suppression in at least one step per frequency.

4. Hearing aid according to claim 2 or 3, wherein the attenuation factors for noise suppression in addition, preferably with a user-adjustable factor, can be modified to vary the strength of the noise suppression.

5. Hearing aid according to one of the preceding claims, wherein a noise estimate at fixed time intervals and / or continuously performed automatically.

6. Hearing aid according to one of the preceding claims, wherein the hearing module (3) has a Einrichö «-G to compensate for a hearing impairment by means of correction factors.

7. Hearing aid according to one of claims 1 to 6, further comprising an initialization module (2) for outputting initialization signals to the speaker (4) and a control unit (5) via which a user can interact with the hearing aid to parameters of the hearing module (3 ) individually.

8. Hearing aid according to claim 7, wherein the initialization module (2) outputs a sequence of electrical signals to the loudspeaker (4), which are converted into acoustic signals, which serve to measure the hearing curve of a user.

9. Hearing aid according to claim 8, wherein the transmission from the microphone (1) to the loudspeaker (4) is interrupted, while the initialization module (2) is operated to measure the hearing curve of the user.

10. Hearing aid according to claim 8 or 9, wherein the hearing module (3) compensates for the loudspeaker output in accordance with the deviation of the measured hearing curve from a normal hearing curve.

11. Hearing aid according to one of the preceding claims 7 to 10, wherein the initialization module (2) outputs electrical signals to the speaker (4), which are converted into acoustic signals, which serve to determine a pain threshold of the user for a maximum acceptable volume.

12. Hearing aid according to claim 11, wherein the hearing module (3) according to the maximum acceptable volume limits the speaker output.

13. Hearing aid according to one of the preceding claims, wherein the hearing module (3) has different filters with different center frequencies and respective adjustable gain.

14. Hearing aid according to one of claims 8 to 13, wherein the sequence of electrical signals corresponding to acoustic signals of a particular frequency or a frequency spectrum with a certain center frequency for the interactive determination of a lower threshold of the user depending on the respective frequency.

15. Hearing aid according to claim 14, comprising a comparator for comparing a lower hearing threshold of a user at a certain center frequency with a stored lower hearing threshold of a normal hearing and a setting device for setting a gain at the respective frequency.

16. Hearing aid according to one of the preceding claims 8 to 15, wherein in an initialization electrical signals from the initialization module (2) are output to the loudspeaker (4) according to a predetermined program.

17. Hearing aid according to claim 16, wherein the control unit (5) has a button which a user actuates as soon as the lower hearing threshold is reached at a center frequency.

18. Hearing aid according to one of the preceding claims, comprising an A / D converter (6) for digitizing electrical signals from the microphone (1) and a D / A converter (7) for

Converting digital signals into analog signals for the loudspeaker (4).

19. Hearing aid according to one of claims 11 to 18, wherein the electrical signals for determining a pain threshold of the user have white noise, wherein preferably the noise signal is frequency-dependent amplified according to

Deviation of the determined hearing curve from a normal hearing curve.

20. A method for noise suppression in a hearing aid according to any one of claims 1 to 19, comprising the steps of: converting acoustic signals into electrical signals with a microphone (1);

Processing the electrical 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 the passage of a noise estimate for determining a signal-dependent filter and the provision of a noise-reduced output signal.

21. The method of claim 20, comprising determining attenuation factors for noise suppression by the listening module based on the noise estimate to account for a change in the noise environment.

22. The method of claim 21, comprising multiplying the signal from the microphone in the spectrum with correction factors to compensate for a hearing impairment and / or attenuation factors of the noise suppression in at least one step per frequency by the listening module (3).

23. A method according to claim 21 or 22, in addition to modifying the

Noise reduction attenuation factors, preferably with a user adjustable factor, to vary the amount of noise cancellation.

24. The method according to any one of claims 20 to 23, with the performing of a noise estimate at fixed time intervals and / or continuously continuously.

25. The method of claim 20, wherein processing the electrical signals further comprises compensating for a hearing impairment with correction factors.

26. The method according to any one of claims 20 to 25, with the outputting of initialization signals to the loudspeaker (4) by means of an initialization module (2) to individually adjust parameters of the hearing module (3), preferably by interacting with a user with the hearing aid by means of a control unit (5).

27. The method of claim 26, comprising outputting a sequence of electrical signals to the loudspeaker (4) by the initialization module (2), wherein the electrical signals are converted into acoustic signals and used to measure the hearing curve of a user.

28. The method of claim 27, comprising interrupting the transmission from the microphone (1) to the loudspeaker (4) while the initialization module (2) is operated to measure the user's auditory curve.

29. Method according to claim 27 or 28, with compensating the

Speaker output by the listening module (3) corresponding to the deviation of the measured hearing curve from a normal hearing curve.

30. The method according to any one of claims 26 to 29, with the determining a

Pain threshold of the user for a maximum acceptable volume through the initialization module (2).

31. The method of claim 30, including limiting the maximum acceptable volume of the speaker output by the listening module (3).

32. The method according to any one of claims 20 to 31, wherein the hearing module (3) comprises different filters with different center frequencies and respective adjustable gain.

33. The method according to any one of claims 27 to 32, with the interactive determination of a lower hearing threshold of the user as a function of the respective frequency, wherein the sequence of electrical signals corresponding to acoustic signals of a particular frequency or a frequency spectrum with a certain center frequency.

34. The method of claim 33, comprising comparing a user's lower threshold of hearing at a particular center frequency to a stored lower hearing threshold of a normal listener and adjusting a gain at the particular frequency.

35. The method according to any one of claims 27 to 34, with the outputting of electrical signals to the loudspeaker (4) according to a predetermined program by the initialization module (2) in an initialization mode.

36. The method according to any one of claims 20 to 35, comprising digitizing electrical signals from the microphone (1) by an A / D converter (6) and converting digital signals into analog signals for the loudspeaker (4) by a D / A converter (7).

37. The method of claim 30, wherein the electrical signals for determining a pain threshold of the user comprise white noise, preferably with the frequency-dependent amplification of the noise signal corresponding to the deviation of the determined hearing curve from a normal hearing curve.