JP4870780B2 - Hearing aid noise reduction method and system - Google Patents

Description

  The present invention relates to the field of hearing aids and, more particularly, to hearing aids that utilize noise reduction techniques. The present invention further relates to a hearing aid gain adjustment method for reducing noise. The present invention also relates to a noise reduction system in a hearing aid.

  The hearing aid is configured to provide the user's eardrum with an acoustic environment that is amplified according to the user's prescription. This is typically accomplished by providing a device with a microphone, amplifier and small speaker located in an earpiece placed in the user's ear canal. It is well known that acoustic leaks occur around the earpiece. Sound leakage occurs, for example, in a non-sealed fit or when the earpiece is intentionally vented to reduce the sound pressure caused by the user's own voice, taking into account user comfort, for example. Occur. Such leaks can cause a loss in sound pressure and can bypass the hearing aid to reach the eardrum.

  PCT / EP2005 / 05305305, an unpublished PCT application entitled “Method and system for fitting a hearing aid”, is incorporated herein by reference. This application provides a method for estimating unknown functions such as the vent effect and direct transmission gain of an in-situ hearing aid. The vent effect estimate is used to correct the worn audiogram and hearing aid gain.

  WO-A-2005 / 051039 provides a dynamic speech enhancement technique. SII (Methods for Calculation of the Speech Intelligibility Index) See also ANSI S3.5-1997), AI (American National Standard Methods for the Calculation of the Speech Intelligibility Index) By optimizing speech intelligibility index such as the Articulation Index (see also ANSI S3.5-1996), speech intelligibility in noise is improved. Noise reduction technology that improves speech intelligibility in noise by optimizing the speech intelligibility index increases gain reduction in selected frequency bands taking into account human auditory masking.

  The sound input to the hearing aid user is a combination of the sound amplified according to the hearing aid gain and the directly transmitted sound. As long as the amplified sound dominates over the direct transmission sound in all frequency bands, noise reduction techniques will give good results. Noise reduction with current technology to enhance SII is based on the premise that the earplug provides a tight fit between the earplug and the ear canal. However, sound can be transmitted directly into the ear by a ventilation canal or leakage path. Therefore, at a certain threshold, the sound input to the hearing aid user is dominated by the direct transmission sound, and the reduction of the hearing aid gain adversely affects the sound input to the user. If direct transmission sound is not taken into account, speech intelligibility will be poor.

  Therefore, the acoustic effects of leakage paths that may exist between vent tubes and hearing aids and the ear canal remain a challenge in today's hearing aid fitting methods.

  Therefore, there is a need for improved hearing aids and improved techniques that utilize noise reduction in hearing aids.

  Accordingly, an object of the present invention is to provide a hearing aid and a signal processing method for the hearing aid, particularly considering the above-mentioned requirements and the drawbacks of the prior art.

  In particular, an object of the present invention is to provide a hearing aid that performs a noise reduction technique that takes into account the relative amount of directly transmitted sound through the vent, and a method corresponding thereto. To do.

  It is another object of the present invention to provide a hearing aid that performs SII optimization that improves speech intelligibility in noise, and a method corresponding thereto.

  According to a first aspect of the present invention, there is provided a hearing aid including at least one microphone, signal processing means, and an output transducer, wherein the signal processing means is configured to receive an input signal from the microphone. The signal processing means is configured to apply a hearing aid gain to the input signal to generate an output signal to be output at the output transducer, and the signal processing means further includes a direct transmission gain calculated for the hearing aid. A hearing aid is provided having means for adjusting the hearing aid gain according to gain).

  This hearing aid, with means to adjust the hearing aid gain according to the direct transmission gain, gives information (knowledge) about the amount of direct transmitted sound and before the direct sound becomes dominant over the amplified sound (before the direct sound provides information on how much a certain frequency band may be attenuated).

  In another aspect of the invention, the hearing aid and the method incorporate knowledge of the amount of direct sound into an applied noise reduction algorithm and vent it. It can be optimized taking into account effects and leakage knowledge. It provides more accurate and effective noise reduction than can be obtained with conventional methods.

  In another aspect of the present invention, the phase disruption in the output signal is avoided by considering the direct transmission sound when calculating the hearing aid gain for generating the output signal. A hearing aid is provided.

  According to another aspect of the present invention, a method of compensating direct transmitted sound in a hearing aid, wherein the effective vent parameter of the hearing aid is estimated. And calculating a direct transmission gain based on the effective vent parameter and applying a hearing aid gain for generating an output signal from the input signal, wherein the direct transmission gain is less than or equal to the hearing aid gain. A method is provided that is used as a lower gain value for which is not set.

  In another aspect of the present invention, there is provided a method for determining a direct transmission sound in a hearing aid, wherein a method for estimating an effective vent parameter of the hearing aid and calculating a direct transmission gain based on the effective vent parameter is provided. Is done.

  The method provided above allows the calculation of the direct transmission gain when fitting a hearing aid, and by using the calculated gain in accordance with further methods and systems according to the present invention, other methods of hearing aid parameters other than gain. Dynamic correction can also be performed.

  Hearing aids, systems and methods according to the present invention provide the ability to dynamically adjust the speech intelligibility exponent gain, resulting in noise reduction hearing aid gain for direct transmission gain in real time, and in any case Even so, providing the amount of gain that a hearing aid or system can apply can be considered a real advantage.

  In an embodiment of the invention, the hearing aid adjusts the hearing aid gain in each frequency band based on the instantaneous gain level, additional SII input parameters and direct transmission gain. The overall speech intelligibility can be improved. A direct approach is taken into account in noise reduction techniques, and a new approach is provided that gives users better speech intelligibility in noise.

  In a further aspect, the present invention provides a hearing aid noise reduction system, a computer program and a computer program product according to claims 27, 28 and 29.

  Further particular variations of the invention are defined in the other claims.

  Other aspects and advantages of the invention will become more apparent from the following detailed description, taken in conjunction with the accompanying drawings, illustrating by way of example the principles of the invention.

  The present invention will be readily understood from the following detailed description when taken in conjunction with the accompanying drawings, in which like reference numerals designate like structural elements.

  First, FIG. 1a will be referred to for the explanation on the calculation (calculation) of the DTG. The calculation of DTG is performed by performing a feedback test (FBT) as schematically shown in FIG. 1a. Here, an in-situ vent effect is estimated, and the DTG is calculated from the vent effect. The PCT / EP2005 / 0555305 document (described above) describes this in detail.

  Reference is now made to FIG. 1b, which shows a hearing aid 200 according to a first embodiment of the present invention.

  The hearing aid includes an input transducer or microphone 210 that converts an acoustic input signal into an electrical input signal 215, and an A / D converter (not shown) that samples and digitizes the analog electrical signal. The processed electrical input signal is then provided to a signal processing means 220 that includes an amplifier with a compressor that produces an electrical output signal 225 by applying a compressor gain (compression gain) and is subject to user requirements. Accordingly, an output signal suitable for compensating for hearing loss is generated. In one embodiment, the compressor gain characteristic is non-linear and provides more gain for low input signal levels and less gain for high signal levels. The signal path further includes an output transducer 230, for example a speaker or receiver, that converts the electrical output signal into an acoustic output signal.

  The compressor operates to compress the dynamic range of the input signal. This is useful for the treatment of presbyscusis (loss of dynamic range due to auditory cell loss). In practice, in most cases, the hearing aid compression operation expands the low level signal and suppresses microphone noise while amplifying the input signal just above it. The compressor can also be equipped with a soft-limiter that limits the maximum output level to a safe or comfortable level. The compressor has a non-linear gain characteristic and can therefore provide less gain at higher input levels and more gain at lower input levels. Hearing aids that incorporate a compressor in the signal processor are often referred to as non-linear gain hearing aids or compression hearing aids.

  The signal processing means further comprises a memory 240 and adjusting means 250 for further adjusting the hearing aid gain in addition to the processor's basic determination based on the user's hearing impairment and primary acoustic environment. This additional adjustment takes into account certain effects of sounds that bypass the hearing aid, eg by bypassing the earpiece or propagating the vent, as described below. The purpose is that.

  For calculation purposes, the sound that bypasses the hearing aid is represented by a direct transmission gain (DTG). Direct transmission gain (DTG) is defined as the sound pressure in the eardrum caused by a sound source outside the ear and corresponds to the sound pressure at the exterior vent opening caused by the same sound source. The direct transmission gain is generally less than 1, in other words, the logarithmic value expressed in dB is usually a negative number. However, because there is a natural Helmholz resonance due to the earpiece placed in the ear canal, there may be frequencies (s) with a DTG greater than 1, that is, a positive logarithmic value. . Information about the direct transmitted sound in the single frequency band (Information about the direct transmitted sound in the single frequency band) is, for example, PCT / EP2005 / 0555305 for calculating the direct transmission gain for the hearing aid gain used by a specific user. It can be estimated by the method described.

  The DTG 245 calculated for the hearing aid is stored in the hearing aid memory 240 as a set of frequency dependent gain values. The DTG is then used by the adjusting means 250 to adjust the hearing aid gain to reduce noise, avoid phase disruption, or result from an amplified output signal and direct transmission sound. Other useful optimizations or improvements are made to the signal quality of the combined acoustic signal on the tympanic membrane.

  Referring now to FIG. 2, FIG. 2 represents the level of signal versus frequency, and shows the results by adding contributions of two sound signals. Show. In particular, Figure 2 shows two frequency dependent signals with a relative phase which are added here to clarify the principle of the addition of two sound signals in the eardrum. Is shown. The black broken lines (plurality) are the magnitudes of the two signals. The gray dot-and-dash line represents the sum of these signals, where the two signals are in phase at all frequencies (upper curve) and out of phase at all frequencies. Each case (lower curve) is represented. The solid line indicates what happens if the phase difference varies linearly with frequency.

  The sound level in the user's eardrum is the superposition of the undirected direct sound and the sound amplified by the hearing aid. Interference of two sound sources causes phase destruction, ie sound input at a frequency where the unprocessed direct sound and the amplified sound from the hearing aid have approximately the same magnitude but are in opposite phase May cause fluctuations. FIG. 2 shows a general phenomenon, showing the addition of two signals of different magnitude and phase.

At a certain frequency, the sum of two harmonic signals can be expressed as:
A ₁ cos (2πft + φ ₁ ) + A ₂ cos (2πft + φ ₂ ) (1)

In this embodiment, A ₁ = 1, φ ₁ = 0, and A ₂ ∝f. φ ₂ is either 0 or π, or ∝f. A simple calculation can reveal constructive and destructive interference, but on the other hand an analytical explanation of the sum of two signals with frequency-dependent phase and frequency-dependent amplitude Becomes more complex. In this case, the phase breakdown that occurs will depend on the amplitude and phase of the signal. However, since constructive interference and destructive interference constitute the upper and lower limits of phase destruction, respectively, the phase destruction signal is between these lines, as shown in FIG. 2 for the φ ₂ ∝f case. You can see that it is located somewhere. Note that dB is calculated as 20log10 (A), so the ratio of the absolute amplitude corresponds to the amplitude difference in dB display. Therefore, amplitude 0 corresponds to −∞ dB.

  The lower gray dash-dot line shows that when the two signals have exactly the same amplitude but are out of phase by π, the entire signal cancels out and becomes infinitely small. This is called destructive interference or phase cancellation. On the other hand, if the two signals are in phase at all frequencies, the amplitudes are simply added in a constructive interference, resulting in an additional 6 dB sound pressure at the frequency where the two signals have the same amplitude. . This can be seen in the upper gray dashed line at 5 kHz. However, in these two cases, the hearing aid sound and the direct sound are rarely seen because both the hearing aid sound and the direct sound have a varying frequency dependent phase. Therefore, the black line shows how the overall sound pressure appears when the relative phase is linearly dependent on frequency. Note that constructive interference increases the overall signal magnitude at some frequencies, whereas destructive interference reduces the overall signal at other frequencies. This phenomenon is called phase disruption because the signal is not canceled out at a frequency where the relative phase is approximately π and the relative amplitude is not quite 1 (the relative amplitude is not quite 1).

  The above example is general, and it is possible to estimate the situation in the user's ear where the amplified sound and the direct sound are superimposed. That is, before the total sound pressure at the eardrum remains unperturbed by the direct sound with respect to phase disruption, Means that must be exceeded. Maintaining the hearing aid gain as large as the direct sound leads to an increased risk of phase destruction, which is avoided by the present invention.

  Referring to FIG. 2, the amplitude difference between the amplified sound and the unprocessed direct sound must exceed a certain amount (safety margin) to minimize phase destruction. Therefore, there is a lower threshold for gain setting equal to the direct transmission gain + k, which is shown on the right scale of FIG. The safety margin is a coefficient k and can be set freely in principle. When k is a large negative number, the interaction between the direct sound and the amplified sound is ignored and no special measures are taken into account. If k is a large positive number, measures are taken all the time, but this is also not optimal. Therefore, the choice of the coefficient k is a trade-off between minimizing the risk of phase destruction and limiting the SII optimization.

FIG. 3 shows the phase disruption range versus signal amplitude ratio. More specifically, FIG. 3 shows the difference in dB between the amplitudes as a function of the amplitude difference between the two signals shown in FIG. of the in-phase summed signal and the out-of-phase summed signal). That is, this curve shows the uncertainly or possible spread of the total sound pressure due to phase destruction. The signal amplitude ratio in dB is the difference between the hearing aid sound (represented by gain) and the directly transmitted sound (represented by gain) in each band, that is, HA-DTG (Direct Transmitted Gain) (Direct Transmitted) in dB representation. Gain), A ₁ is DTG, and A ₂ is HA. The DTG is fixed at the time of manufacturing the earplug, whereas the hearing aid gain changes with sound input. Therefore, only the hearing aid sound is a variable when vent is selected.

  For example, when one signal is 10 dB larger than the other signal, it is read from the above curve that the phase destruction is a worst case scenario and the amplitude of the sum signal can change from the in-phase sum signal to −5 dB. be able to. A value of 1 or more is applicable, preferably between 5 and 15 dB. Of course, a value of about 1 dB will increase the risk of phase breakdown. A value of k = 7 or k = 8 results in a phase disruption range of about ± 3 dB, which can be considered an acceptable range.

When the hearing aid is stopped, the sound from the hearing aid becomes -∞ (complete silence), which clearly means that DTG becomes completely dominant. This corresponds to −∞ on the x-axis of FIG. 3 and does not lead to the expected phase destruction problem. On the other hand, when the hearing aid gain is 60 dB, for example, and the direct transmission sound is -10 dB, the direct sound can be relatively ignored and there is no risk of phase destruction. Only when the sound level of the direct sound and the hearing aid sound is equivalent (A ₂ ≈A ₁ ), the strength of the summed signal may vary significantly as shown in FIG. There is.

  Therefore, in the present invention, the coefficient k represented as an example in FIG. 3 constitutes the lower limit, and during the optimization process (during the optimization process), the hearing aid gain is not set lower than that and the risk of large phase destruction Not accompanied.

  In order to calculate the direct transmission gain for the hearing aid gain used by a specific user, information on the direct transmission sound in a single frequency band is estimated by, for example, the method described in PCT / EP2005 / 0555305. This is later used for SII optimization. If the direct sound is dominant in the lowest band, for example, a new optimum of SII is obtained by changing the gain in some of the bands where the amplified sound is dominant The value can be found.

  In one embodiment, the adjusting means is means for optimizing the speech intelligibility index (SII) by applying each noise reduction technique considering the DTG, as will be described in detail below. Yes, it provides users with better speech intelligibility in noise.

  4 and 5 show the principle of a combination of noise reduction technology based on SII (sound intelligibility index) and direct transmission sound through a vent.

  FIG. 4 shows the directly transmitted sound in dB. This gain function is called the direct transmission gain and represents the sound pressure in the eardrum relative to the sound pressure at the vent entrance by the sound source outside the ear. The direct transmission gain can be determined during a feedback test as described in PCT / EP2005 / 0555305 above.

  In this example, values for 15 frequency bands between 100 Hz and 10 kHz are calculated. The two y-scales in this figure are the direct transmission gain on the left and the minimal amplification on the right that the hearing aid gain must exceed to dominate the total sound at the eardrum. ). The minimum amplification is determined as the hearing aid gain required to avoid the risk of phase disruption problems caused by adding two sound pressures of the same magnitude but in antiphase. Such phase destruction results in bad sound quality that can be expressed as a metallic sound or an annoying sound at the frequency at which the phase destruction occurs.

  The letter k in these figures refers to a dB limit that is large enough for the amplified sound to dominate the overall sound pressure in the eardrum compared to the direct sound. k is the limit to divide the operation of this algorithm into two states, one of which requires action to avoid phase destruction, and one of which does not require any action It is. If the amplified sound -k is smaller than the direct sound, there is a risk of phase destruction and some measures need to be taken. Refer to FIG. 3 to clarify the k-factor. In FIG. 4, the direct transmission gain and minimum amplification for frequency band 4 and frequency band 5 are emphasized for estimated vent diameters of 1 mm (dark color) and 3 mm (light color), respectively.

  In FIG. 5, the minimum amplification of k = 8 dB for the two frequency bands is marked on the graph and includes the hearing aid gain adjustment necessary to find the optimal gain setting for speech intelligibility. Yes. These graphs show how the direct transmission gain interacts with and interferes with the hearing aid gain when determining the optimum gain setting for SII.

  These graphs show how SII varies as a function of hearing aid gain in the two frequency bands at a given vent diameter and hearing loss. SII is represented as a contour curve. SII varies between 0 and 1. SII may have local minima or maxima, but is almost monotonic. By varying the gain in multiple frequency bands, the optimal gain setting that leads to the optimal SII for the hearing aid is determined in each frequency band.

  FIG. 5 shows the gain of frequency band 4 whose center frequency is 500 Hz and the gain of frequency band 5 whose center frequency is 634 Hz. The contour curves (plurality) indicate what kind of gain setting function SII is in each frequency band (how the SII is a function of the setting of the gain).

  Prior art SII optimization currently does not take into account, for example, direct sound arriving through a vent. However, when the direct sound is added to the hearing aid amplified sound, it is actually impossible to obtain a gain smaller than the gain generated from the direct sound. If a large vent is formed in the ear mold in combination with a relatively low hearing loss, the direct sound will overwhelm the amplified sound and only the direct sound will be heard.

  Further explanation on how to use SII for hearing aid noise reduction is described in WO-A-2005 / 051039, which is incorporated herein by reference.

  FIG. 5 also shows that k was selected to be 8 for frequency bands 4 and 5, respectively, for two vent diameters (1 mmφ and 3 mmφ) combined with two types of hearing loss (just 40 dB HL and 80 dB HL). The actual section (interval) (interval) of the gain in the case is exemplarily shown.

  The SII optimization of the hearing aid is performed in all bands, i.e. 15 ranges in this example. However, the illustration of the 15 ranges of optimization procedures hinders the principle rather than making it easier to understand and visualize. Thus, FIG. 5 only shows a method for optimizing the SII of the two selected bands (bands 4 and 5). In the example of the linear optimization method, the gain for the frequency band 4 is kept constant, the gain for the frequency band 5 is gradually changed until the optimum SII relating to the setting is detected, and then the gain of the frequency band 4 is changed. Until the optimum setting of the frequency band 4 is detected, the optimum setting of the frequency band 5 detected in the previous stage is kept constant.

  FIG. 5 shows an optimization procedure in which the optimization is continued until a better SII cannot be obtained. Of course, other optimization methods can be implemented as long as direct sound is considered. The contour plot shows the SI index as a function of the absolute gain in each band. The theoretical optimum, i.e. the optimal value when the sound in the eardrum is provided only by a hearing aid, is easily detected as 'island' in the plot. However, the direct sound (plus k) represented on the axis using the same symbols as the top plot (plus k) can only be obtained if the above optimum value is obtained. It also affects the path leading to the optimum. The gray area is the region, which would be counterproductive to enter. The iterative optimization process can be performed in many ways, but here it is shown as a sequential adjustment of each band. An asterisk (a star) indicates the result of the optimization method.

  In the graph (upper right frame) for the combination of a small vent (1 mm) and severe hearing loss (HTL 80 dB), the optimal parameter is considered even when the minimum amplification is taken into account when compared with the conventional optimal parameter setting where the gain varies in the entire region. Optimal SII is obtained without any change in the settings. Conversely, in the case of a large vent (3 mm) and mild hearing loss (HTL = 40 dB), the direct sound can enter through the vent sufficiently, so the total sound pressure in the eardrum (lower left frame) May even be affected, and may even become dominant. Therefore, when the minimum amplification is used to limit the gain setting of the frequency band, the optimum gain setting of the frequency band is considerably different from the case where the frequency band fluctuates without considering the minimum amplification. In such a case, as a result, it is possible to set parameters with better gain in various frequency bands.

  Therefore, the iterative optimization path may differ from that performed by other methods, and the optimal parameter settings may also be determined by other methods as optimal values by other embodiments. May be different.

  Therefore, the main advantage of this invention is that the SII is optimized taking into account the actual in-situ acoustic surroundings.

  It is clear to those skilled in the art that both the optimization method and the fact that optimization occurs in all bands can greatly change the illustrated iterative path from the actual iterative path. .

  Reference is now made to FIG. 6, which shows a portion of a hearing aid 300 according to another embodiment of the present invention.

  The SII optimization block 610 as a means for optimizing the speech intelligibility index generates an SII gain 615 that is supplied to the combiner or summing block 620. In the combiner 620, the signal 615 is subtracted from the amplified sound signal 605 generated by the signal processor or compressor by applying the hearing aid gain. The output of the combiner can be viewed as a noise reduction output signal 625 that is provided to the output transducer and comparator 630. The comparator 630 compares the noise reduction output signal 625 with the safety margin k added in the block 640 and the direct transmission sound according to the DTG in the block 245, both supplied to the comparator. If the level of the noise-reduced output signal with the safety margin k added is below DTG, the comparator generates an error signal, which is provided to the SII optimizer 610 as a further input parameter and considered in the SII optimization. Thus, the noise reduced output signal is no longer attenuated below the threshold and phase destruction is avoided.

  In a variation, the hearing aid includes a band division filter that converts the input signal into a band division input signal of a plurality of frequency bands, and the hearing aid is configured to process the band division input signal independently in each frequency band.

  In embodiments of the present invention, the systems and hearing aids described in this document may include, for example, field programmable gate arrays (FPGAs), standard processors, or application specific signal processors (ASSP or It can be implemented in a signal processing apparatus suitable for the above systems and hearing aids, such as digital signal processors including ASICs, analog / digital signal processing systems. As is well known to those skilled in the art, it is clear that the entire system is preferably implemented with a single digital element even though some parts could be implemented in other ways.

  Hearing aids, methods, systems and other devices according to embodiments of the invention can be implemented in any suitable digital signal processing system. The hearing aid, method and apparatus can also be used, for example, when fitting by an audiologist. The method according to the invention may also be performed by a computer program comprising executable program code that performs the method according to the embodiments described herein. When using a client-server environment, embodiments of the present invention comprise a remote server computer that embodies a system according to the present invention and hosts a computer program that executes the method according to the present invention. In another embodiment, a computer program product such as a computer readable storage medium such as a floppy disk, memory stick, CD-ROM, DVD, flash memory, or other suitable storage medium is in accordance with the invention. Provided for storing a computer program.

  According to a further embodiment, the program code is stored in a digital hearing aid memory or computer memory and is processed by a processing device such as the hearing aid device body or its CPU, or according to any other suitable processor or described embodiment. It can be executed by a computer that executes

  While the principles of the invention have been described and illustrated in these embodiments, it will be apparent to those skilled in the art that the invention can be modified in arrangement and detail without departing from such principles. Changes and modifications within the scope of the invention may be made without departing from the spirit thereof, and the invention includes all such changes and modifications.

The schematic regarding calculation of a direct transmission sound is shown. 1 shows a block diagram of a hearing aid according to the invention. The signal versus frequency level as a result of adding the contribution of two sound signals is shown. The phase breakdown range is a function of the amplitude difference between the two signals. A graph of directly transmitted sound versus frequency is shown. The principle of SII (speech intelligibility index) optimization considering direct transmission sound according to the present invention will be described. 1 shows a block diagram of a hearing aid portion according to an embodiment of the present invention. FIG.

Claims (16)

A hearing aid comprising an earpiece provided with at least one microphone, signal processing means , an output transducer and a vent ,
Receiving an input signal said signal processing means from the microphone, hearing aids gain to generate an amplified sound signal is applied to the input signal, and is amplified sound signal generated is configured to make the output from the output transducer And
Said signal processing means further Ru calculated for said hearing aid comprises means for adjusting the hearing aid gain according to the direct transmission gain represents the sound pressure at the eardrum relative to the sound pressure at the vent inlet according outside of the sound source of the ear,
The hearing aid gain adjustment means is configured to adjust the hearing aid gain to a value not less than the direct transmission gain ;
hearing aid. The hearing aid according to claim 1, further comprising a memory for storing the direct transmission gain. It said hearing aid gain adjustment means is configured safety margin k provides, and to adjust the hearing aid gain to a value not less than the direct transmission gain plus said safety margin k, according to 請 Motomeko 1 hearing aid. Comprising speech intelligibility index optimization means for determining a speech intelligibility index gain that leads to an optimal speech intelligibility index for the hearing aid by varying the hearing aid gain in one or more frequency bands;
The hearing aid according to claim 1. A combiner for generating a noise reduction output signal by subtracting the speech intelligibility index gain determined by the speech intelligibility index optimization means from the amplified sound signal generated by the signal processing means;
The noise reduction output signal is output from the output transducer;
The hearing aid according to claim 4 . Hearing aid according to claim 3 , wherein the safety margin k is a gain value in the range of 0 to 15 dB, preferably in the range of 5 to 15 dB. Hearing aid according to claim 3 , wherein the safety margin k is a gain value of 5 to 8 dB, preferably 7 to 8 dB. A band division filter for converting the input signal into a plurality of band division input signals for a plurality of frequency bands, so that the hearing aid processes the band division input signal independently in each of the plurality of frequency bands; The hearing aid according to any one of claims 1 to 7 , further configured. At least one microphone for generating an input signal , signal processing means for generating an amplified sound signal by applying a hearing aid gain to the input signal , an output transducer for outputting the generated amplified sound signal, and an earpiece provided with a vent A method of controlling a hearing aid,
Storing the direct transmission gain calculated for the hearing aid and its user , representing the sound pressure at the eardrum relative to the sound pressure at the vent inlet by the sound source outside the ear, in the memory of the hearing aid;
It said hearing aid gain as above Symbol hearing aid gain is not set to a value below the direct transmission gain adjusted using the direct transmission gain,
Method. 10. The method of claim 9 , wherein the hearing aid gain is not set to a value less than the direct transmission gain plus a safety margin k. Determining the speech intelligibility index gain that leads to the optimal speech intelligibility index for the hearing aid by varying the hearing aid gain in one or more frequency bands;
The method of claim 9. A noise reduction output signal is generated by subtracting the speech intelligibility index gain determined from the generated amplified sound signal,
The noise reduction output signal is output from the output transducer;
The method of claim 11 . 13. The method according to any one of claims 9 to 12 , further comprising the step of converting the input signal into a band-division input signal of a plurality of frequency bands, wherein the method is performed for each of the plurality of frequency bands. the method of. Comprising means for performing the method according to any one of claims 9 to 13, the hearing aid control system. A computer program comprising executable program code for performing the method of any one of claims 9 to 13 when executed on a computer. A computer-readable recording medium on which the computer program according to claim 15 is recorded.

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