WO2007099115A1 - Procede et systeme permettant de reduire le bruit dans un dispositif d'aide auditive - Google Patents

Procede et systeme permettant de reduire le bruit dans un dispositif d'aide auditive Download PDF

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Publication number
WO2007099115A1
WO2007099115A1 PCT/EP2007/051890 EP2007051890W WO2007099115A1 WO 2007099115 A1 WO2007099115 A1 WO 2007099115A1 EP 2007051890 W EP2007051890 W EP 2007051890W WO 2007099115 A1 WO2007099115 A1 WO 2007099115A1
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WO
WIPO (PCT)
Prior art keywords
hearing aid
gain
adjusting
speech intelligibility
direct transmission
Prior art date
Application number
PCT/EP2007/051890
Other languages
English (en)
Inventor
Morten Agerbak Nordahn
Carsten Paludan-Muller
Original Assignee
Widex A/S
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Widex A/S filed Critical Widex A/S
Priority to EP07712378.4A priority Critical patent/EP1992195B1/fr
Priority to CA2643326A priority patent/CA2643326C/fr
Priority to JP2008545022A priority patent/JP4870780B2/ja
Priority to AU2007220497A priority patent/AU2007220497B2/en
Priority to DK07712378.4T priority patent/DK1992195T3/en
Publication of WO2007099115A1 publication Critical patent/WO2007099115A1/fr
Priority to US12/201,544 priority patent/US8422709B2/en

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Classifications

    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R25/00—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception
    • H04R25/45—Prevention of acoustic reaction, i.e. acoustic oscillatory feedback
    • H04R25/453—Prevention of acoustic reaction, i.e. acoustic oscillatory feedback electronically
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R2225/00—Details of deaf aids covered by H04R25/00, not provided for in any of its subgroups
    • H04R2225/43—Signal processing in hearing aids to enhance the speech intelligibility
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R2460/00—Details of hearing devices, i.e. of ear- or headphones covered by H04R1/10 or H04R5/033 but not provided for in any of their subgroups, or of hearing aids covered by H04R25/00 but not provided for in any of its subgroups
    • H04R2460/11—Aspects relating to vents, e.g. shape, orientation, acoustic properties in ear tips of hearing devices to prevent occlusion
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R25/00—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception
    • H04R25/35—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception using translation techniques
    • H04R25/356—Amplitude, e.g. amplitude shift or compression
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R25/00—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception
    • H04R25/70—Adaptation of deaf aid to hearing loss, e.g. initial electronic fitting

Definitions

  • the present invention relates to the field of hearing aids and more specifically to hearing aids utilizing noise reduction techniques.
  • the invention further relates to methods for adjusting the hearing aid gain for noise reduction.
  • the invention relates to a system of reducing noise in a hearing aid.
  • Hearing aids are adapted for providing at the users eardrum a version of the acoustic environment that has been amplified according to the users prescription. This is normally achieved by providing a device with a microphone, an amplifier and a miniature loudspeaker situated in an earpiece placed in the users ear canal. It is well known that there may be acoustic leaks around the earpiece. There may e.g. be a non-sealed fit or there may be a vent deliberately arranged in the ear piece for considerations about user comfort, e.g. for relieving the sound pressure created by the users own voice. Such leaks may cause a loss in sound pressure and they may allow sound to bypass the hearing aid to reach the ear drum.
  • the sound input to the hearing aid user is a combination of the sound amplified according to the hearing aid gain as well as the direct transmitted sound. As long as the amplified sound dominates the direct transmitted sound in all frequency bands the noise reduction techniques will provide good results. Noise reduction according to the state of the art to enhance SII is based on an assumption that the earplug provides a tight fit between the earplug and the ear canal. However a ventilation canal or a leakage path allows for the sound to be directly transmitted into the ear. Thus, at a certain threshold the sound input to the hearing aid user may be dominated by the direct transmitted sound, so that a decrease of the hearing aid gain will not affect the sound input to the user. If the direct transmitted sound is not taken into account the speech intelligibility may suffer as a consequence.
  • a hearing aid that comprises at least one microphone, a signal processing means and an output transducer, wherein the signal processing means is adapted to receive an input signal from the microphone, wherein the signal processing means is adapted to apply a hearing aid gain to the input signal to produce an output signal to be output by the output transducer, and wherein the signal processing means further comprises means for adjusting the hearing aid gain according to a direct transmission gain calculated for the hearing aid.
  • This hearing aid with means for adjusting the hearing aid gain according to a direct transmission gain gives a knowledge about the amount of directly transmitted sound and provides information about how much a certain frequency band may be attenuated before the direct sound becomes dominant over the amplified sound.
  • the hearing aid and the method are capable of incorporating knowledge of the amount of direct sound into the applied noise reduction algorithm, which thereby is optimized taking the knowledge of vent effect and leakage into account. This provides a more accurate and effective noise reduction than would be otherwise obtainable.
  • a hearing aid that is capable of avoiding phase disruption in the output signal by taking the direct transmitted sound into account when calculating the hearing aid gain to produce the output signal.
  • a method of compensating direct transmitted sound in a hearing aid which comprises the steps of estimating an effective vent parameter for the hearing aid, calculating a direct transmission gain based on the effective vent parameter, and applying a hearing aid gain to produce an output signal from an input signal wherein the direct transmission gain is used as a lower gain limit below which the hearing aid gain is not set.
  • a method of determining direct transmitted sound in a hearing aid which comprises the steps of estimating an effective vent parameter for the hearing aid, and calculating a direct transmission gain based on the effective vent parameter.
  • the provided methods enable a calculation of the direct transmission gain once when fitting the hearing aid which may then be used according to further methods and systems according to the present invention for the dynamic correction of also other hearing aid parameters than gain.
  • the hearing aids, systems and methods according to the present invention provide the ability to dynamically adjust the applicable speech intelligibility index gain and the resulting noise reduced hearing aid gain for the direct transmission gain in real time and, thus, the amount of gain that the hearing aid or system may apply at any given instance.
  • the hearing aid is able to adjust the hearing aid gain in each frequency band based on the instantaneous gain level, the further SII input parameters and the direct transmission gain in order to improve the overall speech intelligibility.
  • This offers a new approach according to which the direct transmission gain is taken into account in the noise reduction technique, giving the user a better speech intelligibility in noise.
  • the invention provides a system of reducing noise in a hearing aid, a computer program and a computer program product as recited in claims 27, 28 and 29.
  • Figs. 1 a depicts a schematic diagram regarding calculation of the direct transmitted sound
  • Fig. 1b depicts a block diagram of a hearing aid according to the present invention.
  • Fig. 2 depicts the level of signal versus frequency that results by adding contributions of two sound signals
  • Fig. 3 depicts the phase disruption range as a function of the difference between the amplitude of the two signals
  • Fig. 4 shows a graph of the directly transmitted sound versus frequency
  • Fig. 5 shows diagrams illustrating the principle of optimizing the SII (Speech Intelligibility Index) taking into account the directly transmitted sound, according to the present invention.
  • Fig. 6 depicts a block diagram of part of a hearing aid according to an embodiment of the present invention.
  • Fig. 1a for an explanation regarding calculating the DTG.
  • the calculation of the DTG is done by performing a feedback test (FBT) as schematically illustrated in Fig. 1 a. Then, the in-situ vent effect is estimated and the DTG is calculated from the vent effect.
  • FBT feedback test
  • Document PCT/EP2005/055305 (mentioned above) describes this in detail.
  • Fig. 1 b shows a hearing aid 200 according to the first embodiment of the present invention.
  • the hearing aid comprises an input transducer or microphone 210 transforming an acoustic input signal into an electrical input signal 215, and an A/D-converter (not shown) for sampling and digitizing the analogue electrical signal.
  • the processed electrical input signal is then fed into signal processing means 220, which includes an amplifier with a compressor for generating an electrical output signal 225 by applying a compressor gain in order to produce an output signal suitable for compensating a hearing loss according to the users requirements.
  • the compressor gain characteristic is, according to an embodiment, non-linear to provide more gain at low input signal levels and less gain at high signal levels.
  • the signal path further comprises an output transducer 230, i.e. a loudspeaker or receiver, for transforming the electrical output signal into an acoustic output signal.
  • the compressor operates to compress the dynamic range of the input signals. It is useful for treatment of presbyscusis (loss of dynamic range due to haircell-loss). Actually, compressing hearing aids often apply expansion for low level signals, in order to suppress microphone noise while amplifying input signals just above that level.
  • the compressor may also include a soft-limiter in order to limit maximum output level at safe or comfortable levels.
  • the compressor has a non-linear gain characteristic and, thus, is capable of providing less gain at higher input levels and more gain at lower input levels. Hearing aids embodying a compressor in the signal processor are often referred to as non-linear-gain or compressing hearing aid.
  • the signal processing means further comprises memory 240 and adjusting means 250 for adjusting the hearing aid gain further over what the processor basically decides based on the users hearing deficit and the prevailing sound environment. This further adjustment is intended to take into account certain effects of sounds bypassing the hearing aid, e.g. by bypassing the earpiece or by propagating through the vent, as will be explained below.
  • the sound bypassing the hearing aid is expressed in terms of direct transmission gain (DTG).
  • the direct transmission gain (DTG) is defined as the sound pressure at the ear drum that is generated by an acoustic source outside the ear relative to a sound pressure at the exterior vent opening generated by the same source.
  • the direct transmission gain is typically less than one, i.e. the log value expressed in dB, will normally be a negative number.
  • the log value is a positive number.
  • Information about the direct transmitted sound in the single frequency bands can be estimated by e.g.
  • the DTG 245 calculated for the hearing aid as a set of frequency dependent gain values is stored in memory 240 of the hearing aid.
  • the DTG is then used by the adjusting means 250 to adjust the hearing aid gain in order to reduce noise, avoid phase disruption or provide any other useful optimization or improvement of the signal quality in the combined acoustic signal on the ear drum resulting from the amplified output signal and the direct transmitted sound.
  • Fig. 2 depicts the level of signal versus frequency that results by adding contributions of two sound signals, and more specifically shows two frequency dependent signals with a relative phase which are added here, to clarify the principle of adding two sound signals at the eardrum.
  • the black dotted lines are the magnitude of the two signals.
  • the gray dash-dotted line represents the sum of these signals, when the two signals are in phase for all frequencies (upper curve), and when they are out of phase for all frequencies (lower curve), respectively.
  • the full line shows what happens, if the phase difference varies linearly with frequency.
  • the sound level at the eardrum of the user is a superposition of the unaided direct sound and the sound amplified by the hearing aid.
  • the interference of the two sound sources may lead to phase disruptions, i.e. fluctuations in the sound input at frequencies where the unaided direct sound and the amplified sound from the hearing aid has about the same magnitude but has opposite phase.
  • Fig. 2 illustrates the addition of two signals with differing magnitude and phase.
  • the sum of two harmonic signals can be written as
  • both constructive and destructive interference can be made clear, whereas the sum of two signals with frequency dependent phase and amplitude is more complex to describe analytically. In this case, the resulting phase disruption will depend on the amplitudes and phases of the signals.
  • constructive and destructive interference constitutes the upper and lower limit of the phase disruption, respectively, we know, that a phase disrupted signal lies somewhere in between these lines, as shown in Fig. 2 for the case 92 °c f.
  • the ratio of the absolute amplitude corresponds to the difference of the amplitudes in dB, since dB is calculated as 20log10(A). An amplitude of 0 thus corresponds to - ⁇ dB.
  • the lower dash-dotted gray line shows that in case the two signals are out of phase by ⁇ with the exact same amplitude, the total signal cancels out and becomes infinitely small. This is called destructive interference or phase cancellation.
  • the amplitudes simply add up in a constructive interference, and gives 6 dB more sound pressure at the frequency where the two signals have the same amplitude, which can be seen in the upper dash-dotted gray line at 5 kHz.
  • phase disruption the phenomenon in which the signals do not cancel out as such at frequencies where the relative phase is almost ⁇ and the relative amplitude is not quite 1 , this phenomenon is called phase disruption.
  • the above example is general, and can be extrapolated to the situation in a users ear, where the amplified sound and the direct sound superpose. This in turn means that the amplified sound has to exceed a certain level before the total sound pressure at the eardrum remains unperturbed by the direct sound with respect to phase disruption. Maintaining the hearing aid gain at a similar magnitude to the direct sound would result in an increased risk of phase disruption, which is avoided with the current invention.
  • the difference in amplitude between the amplified sound and the unaided direct sound must be higher than a certain amount (a safety margin) to minimize phase disruption.
  • a safety margin is the factor /c, which in principle could be set to anything. If k is negative and numerically large, the interaction between direct and amplified sound is neglected and nothing extraordinary is ever done to take the interaction into account. If k is large and positive, measures are taken all the time, which is also not optimal. Choosing the factor k is therefore a trade-off between minimizing the risk of phase disruption and limiting the Sll-optimization.
  • Fig. 3 shows the phase disruption range versus signal amplitude ratio.
  • Fig. 3 more specifically shows the difference in dB between the amplitude of the in-phase summed signal and the out-of-phase summed signal as a function of the difference between the amplitudes of the two signals shown in Fig. 2.
  • the curve thus shows the uncertainty or possible spread of the total sound pressure due to phase disruption.
  • the signal amplitude ratio in dB is the difference between the hearing aid sound (expressed in terms of gain) and the directly transmitted sound (expressed in terms of gain) in each band, i.e. HA - DTG (Direct Transmitted Gain) in dB, i.e. Ai is DTG and A 2 is HA.
  • the DTG is fixed once the earplug is made, whereas the hearing aid gain may change with the sound input.
  • the hearing aid sound is thus the only variable, once the vent has been chosen.
  • phase disruption may in a worst case scenario cause the amplitude of the summed signal to vary up to -5 dB from the in-phase summed signal.
  • Values from 1 and upward are applicable, preferably between 5 and 15 dB.
  • a value of about 1 dB would incur a high risk of phase disruption.
  • the hearing aid was turned off, the sound from the hearing aid would be - ⁇ (completely silent), obviously meaning that the DTG would dominate totally. This would correspond to - ⁇ on the x-axis in Fig. 3, which gives no phase disruption problems, as we would expect.
  • the hearing aid gain is e.g. 60 dB and the direct transmitted sound -10 dB, the direct sound is negligible in comparison, and no phase disruption is risked. It is only when the sound level of the direct sound and the hearing aid sound are comparable (A 2 «Ai), that the strength of the summed signal may vary significantly as indicated in Fig. 3.
  • the factor /c which is indicated as an example in Fig. 3, constitutes a lower limit, below which the hearing aid gain should not be set during the optimization process, without risking a large amount of phase disruption.
  • Information about the direct transmitted sound in the single frequency bands can be estimated by e.g. the methods described in the document PCT/EP2005/055305 to calculate a direct transmission gain for the hearing aid gain used by a certain user. This knowledge will then be used to optimize SII. If the direct sound e.g. dominates the lowest band, it is possible to find a new optimum for SII by changing the gain in some of the bands where the amplified sound dominates.
  • the adjusting means is a means for optimizing a speech intelligibility index (SII) by applying a respective noise reduction technique taking the DTG into account to give the user a better speech intelligibility in noise, as will now be described in detail.
  • SII speech intelligibility index
  • the Figs. 4 and 5 show the principle in the combination of SII (Speech Intelligibility Index) - based noise reduction technique and the directly transmitted sound through the vent.
  • the Fig. 4 shows the directly transmitted sound in dB.
  • This gain function represents the sound pressure at the eardrum relative to the sound pressure at the entrance of the vent by a sound source external to the ear.
  • the direct transmission gain may be determined during the feedback test, as in the above-mentioned PCT/EP2005/055305.
  • the values in this example are calculated for 15 frequency bands between 100 Hz and 10 kHz.
  • the figure has two y-scales, where the left represents the direct transmission gain, and the right represents a minimal amplification, which the hearing aid gain must exceed order to dominate the total sound at the eardrum.
  • the minimum amplification is determined as the hearing aid gain necessary to avoid the risk of phase disruption problems caused by adding two sound pressures of same magnitude but opposite phase. Such phase disruption results in bad sound quality, which may be described as metallic or raspy at the frequencies in which phase disruption occurs.
  • k refers to a limit in dB where the amplified sound is large enough to dominate the total sound pressure at the eardrum relative to the direct sound
  • k is a limit that divides the action of the algorithm into two states: one, where actions need to be taken to avoid phase disruption, and one where no action is needed. If the amplified sound - k is less than the direct sound, there is a risk of phase disruption, and something must be done. See Fig. 3 for clarification on the k- factor.
  • the direct transmission gain and the minimum amplification is emphasized for frequency band 4 and frequency band 5 for an estimated vent diameter of 1 mm (dark color) respectively 3 mm (light color).
  • These graphs show how the direct transmission gain interacts and interferes with the hearing aid gain in the search for the optimum gain setting with regards to the SII.
  • the graphs illustrate how the SII varies as a function of the hearing aid gain for two frequency bands, with a given vent diameter and hearing loss.
  • the SII is illustrated as contour curves.
  • the SII varies between 0 and 1. It is approximately monotonous though it may have some local minima or maxima.
  • FIG. 5 illustrate the gain for a frequency band 4, having a center frequency of 500 Hz, and for a frequency band 5, having a center frequency of 634 Hz.
  • the contour curves show how the SII is a function of the setting of the gain in each frequency band.
  • the SII optimization according to the prior art does not presently take the direct sound arriving through e.g. the vent into account.
  • the direct sound adds to the hearing aid amplified sound and thus in practice it will not be possible to obtain a gain lower than the gain originating from the direct sound.
  • the presence of a large vent in the ear mould in combination with a relatively mild hearing loss may thus imply that only the direct sound is heard, since it might overwhelm the amplified sound.
  • FIG. 5 also illustrate and exemplify the actual interval of the gain when k has been chosen to 8 for each of the frequency bands 4 and 5, for two vent diameters (1 mm 0 and 3 mm 0 ) in combination with two hearing losses (flat 40 dB HL and flat 80 dB HL).
  • the optimization of the SII in the hearing aid is performed in all bands, i.e. 15 dimensions in this example. However, illustrating an optimization procedure in 15 dimensions rather impedes than facilitates an easily understandable visualization of the principle.
  • the diagrams in Fig. 5 are therefore limited to illustrate a way of optimizing the SII in two selected bands (bands 4 and 5).
  • An example of a linear optimization method where the gain for frequency band 4 is kept constant and where the gain of frequency band 5 is varied in steps until an optimum SII for that setting has been detected, then the gain of frequency band 4 is varied and the previously detected optimum setting of frequency band 5 is kept constant until an optimum setting of frequency band 4 has been detected.
  • the diagrams in Fig. 5 illustrate an optimization procedure where the optimization is continued until it is not possible to obtain a better SII.
  • Naturally other optimization methods can be implemented, as long as the method takes the direct sound into account.
  • the contour plot shows the Sl-index as a function of the absolute gain in each band.
  • the theoretical optimum i.e., when it is assumed that the sound at the eardrum is provided only by the hearing aid, is easily detected as an 'island' in the plot.
  • the direct sound (plus k) which is illustrated on the axes by use of the same symbols as in the top plot, influences not only whether that optimum is attainable or not, but also the path leading to the optimum.
  • the gray area illustrates the region, which would be counterproductive to enter.
  • the iterative optimization process which could be performed in many ways, is here illustrated as a sequential adjustment of each band.
  • a star indicates the result of the optimization method.
  • the iterative optimization path may be different from what would otherwise be carried out and the optimum parameter setting may also be different from what would else be determined as optimum according to other embodiments.
  • a main advantage for the present invention is therefore that the SII is optimized under consideration of the actual in-situ acoustic surroundings.
  • Fig. 6 shows a part of a hearing aid 300 according to another embodiment of the present invention.
  • SII optimization block 610 as means for optimizing a speech intelligibility index produces the SII gain 615, which is fed to the combiner or summation block 620, where the signal 615 is subtracted from the amplified sound signal 605 produced by the signal processor or compressor by applying the hearing aid gain.
  • the output of the combiner may be considered as the noise reduced output signal 625 fed to the output transducer and also fed to the comparator 630.
  • the comparator 630 compares the noise reduced output signal 625 plus the safety margin k in block 640 with the direct transmitted sound according to the DTG in block 245, both also supplied to the comparator.
  • the hearing aid comprises a band-split filter for converting the input signal into band-split input signals of a plurality of frequency bands and the hearing aid is adapted to process the band-split input signals in each of the frequency bands independently.
  • systems and hearing aids described herein may be implemented on signal processing devices suitable for the same, such as, e.g., digital signal processors, analogue/digital signal processing systems including field programmable gate arrays (FPGA), standard processors, or application specific signal processors (ASSP or ASIC).
  • FPGA field programmable gate arrays
  • ASSP application specific signal processors
  • Hearing aids, methods, systems and other devices according to embodiments of the present invention may be implemented in any suitable digital signal processing system.
  • the hearing aids, methods and devices may also be used by, e.g., the audiologist in a fitting session.
  • Methods according to the present invention may also be implemented in a computer program containing executable program code executing methods according to embodiments described herein. If a client-server- environment is used, an embodiment of the present invention comprises a remote server computer, which embodies a system according to the present invention and hosts the computer program executing methods according to the present invention.
  • a computer program product like a computer readable storage medium, for example, a floppy disk, a memory stick, a CD-ROM, a DVD, a flash memory, or another suitable storage medium, is provided for storing the computer program according to the present invention.
  • the program code may be stored in a memory of a digital hearing device or a computer memory and executed by the hearing aid device itself or a processing unit like a CPU thereof or by any other suitable processor or a computer executing a method according to the described embodiments.

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  • Health & Medical Sciences (AREA)
  • General Health & Medical Sciences (AREA)
  • Neurosurgery (AREA)
  • Otolaryngology (AREA)
  • Physics & Mathematics (AREA)
  • Engineering & Computer Science (AREA)
  • Acoustics & Sound (AREA)
  • Signal Processing (AREA)
  • Circuit For Audible Band Transducer (AREA)
  • Measurement Of The Respiration, Hearing Ability, Form, And Blood Characteristics Of Living Organisms (AREA)
  • Soundproofing, Sound Blocking, And Sound Damping (AREA)
  • Control Of Amplification And Gain Control (AREA)

Abstract

Un dispositif d'aide auditive (200) comprend au moins un microphone (210), un moyen de traitement du signal (220) et un transducteur de sortie (230). Le moyen de traitement du signal est conçu pour recevoir un signal d'entrée du microphone. Le moyen de traitement du signal est conçu pour appliquer un gain du dispositif d'aide auditive au signal d'entrée pour produire un signal de sortie produit par le transducteur de sortie, et le moyen de traitement du signal comprend un moyen de réglage du gain du dispositif d'aide auditive par un gain de transmission directe calculé du dispositif d'aide auditive.
PCT/EP2007/051890 2006-03-03 2007-02-28 Procede et systeme permettant de reduire le bruit dans un dispositif d'aide auditive WO2007099115A1 (fr)

Priority Applications (6)

Application Number Priority Date Filing Date Title
EP07712378.4A EP1992195B1 (fr) 2006-03-03 2007-02-28 Procede et systeme permettant de reduire le bruit dans un dispositif d'aide auditive
CA2643326A CA2643326C (fr) 2006-03-03 2007-02-28 Procede et systeme permettant de reduire le bruit dans un dispositif d'aide auditive
JP2008545022A JP4870780B2 (ja) 2006-03-03 2007-02-28 補聴器のノイズ低減方法およびシステム
AU2007220497A AU2007220497B2 (en) 2006-03-03 2007-02-28 Method and system of noise reduction in a hearing aid
DK07712378.4T DK1992195T3 (en) 2006-03-03 2007-02-28 A method and system for noise reduction in a hearing aid
US12/201,544 US8422709B2 (en) 2006-03-03 2008-08-29 Method and system of noise reduction in a hearing aid

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
US77837606P 2006-03-03 2006-03-03
US60/778,376 2006-03-03

Related Child Applications (1)

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US12/201,544 Continuation-In-Part US8422709B2 (en) 2006-03-03 2008-08-29 Method and system of noise reduction in a hearing aid

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WO2007099115A1 true WO2007099115A1 (fr) 2007-09-07

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EP (1) EP1992195B1 (fr)
JP (1) JP4870780B2 (fr)
CN (1) CN101356854A (fr)
AU (1) AU2007220497B2 (fr)
CA (1) CA2643326C (fr)
DK (1) DK1992195T3 (fr)
WO (1) WO2007099115A1 (fr)

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CN102148033B (zh) * 2011-04-01 2013-11-27 华南理工大学 一种语言传输系统清晰度测试方法
CN106101969A (zh) * 2016-08-18 2016-11-09 孟玲 增进语音即时输出的助听器
CN110351644A (zh) * 2018-04-08 2019-10-18 苏州至听听力科技有限公司 一种自适应声音处理方法及装置
CN110493695A (zh) * 2018-05-15 2019-11-22 群腾整合科技股份有限公司 一种音频补偿系统
DE102019213809B3 (de) 2019-09-11 2020-11-26 Sivantos Pte. Ltd. Verfahren zum Betrieb eines Hörgeräts sowie Hörgerät

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CN101356854A (zh) 2009-01-28
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CA2643326A1 (fr) 2007-09-07
AU2007220497A1 (en) 2007-09-07
AU2007220497B2 (en) 2009-11-26
EP1992195A1 (fr) 2008-11-19
DK1992195T3 (en) 2016-09-12
CA2643326C (fr) 2013-10-01
JP4870780B2 (ja) 2012-02-08

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