BACKGROUND OF THE INVENTION
1. Field of the invention
[0001] 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. In addition the
invention relates to a system of reducing noise in a hearing aid.
2. Description of the related art
[0002] 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.
[0003] Unpublished PCT application
PCT/EP2005/055305 titled "Method and system for fitting a hearing aid", provides a method for estimating
otherwise unknown functions such as the vent effect and the direct transmission gain
for an in-situ hearing aid. The vent effect estimate is used for correcting the in-situ
audiogram and the hearing aid gain.
US 2005/013456 refers to a hearing aid wherein the penetration of direct sound through a ventilation
channel of the hearing aid device is prevented.
[0004] WO-A-2005/051039 provides a dynamic speech enhancement technique, where speech intelligibility in
noise is improved by optimizing a speech intelligibility index, such as SII (see also
Methods for Calculation of the Speech Intelligibility Index. ANSI S3.5-1997), AI (see
also American National Standard Methods for the Calculation of the Articulation Index.
ANSI S3.5-1996). Noise reduction techniques, where speech intelligibility in noise
is improved by optimizing a speech intelligibility index, increase of decrease the
gain in selected frequency bands, taking into account human auditory masking.
[0005] 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.
[0006] Therefore, acoustic effects of the ventilation canal and possible leakage paths between
the hearing aid and the ear canal are still challenges in today's hearing aid fitting
strategies.
[0007] Thus, there is a need for improved hearing aids as well as improved techniques for
utilizing noise reduction in hearing aids.
SUMMARY OF THE INVENTION
[0008] It is therefore an object of the present invention to provide hearing aids and methods
of processing signals in a hearing aid taking in particular the mentioned requirements
and drawbacks of the prior art into account.
[0009] It is in particular an object of the present invention to provide a hearing aid and
a respective method providing a noise reduction technique that take the relative amount
of directly transmitted sound through the vent into account.
[0010] It is a further object of the present invention to provide a hearing aid and a respective
method providing a SII optimization where speech intelligibility in noise is improved.
[0011] According to a first aspect of the present invention, there is provided 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.
[0012] 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.
[0013] According to other aspects of the present invention, 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.
[0014] According to another aspect of the present invention, there is provided 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.
[0015] According to another aspect of the present invention, there is provided 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.
[0016] According to still another aspect of the present invention, there is provided 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.
[0017] 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.
[0018] It may be seen as a true advantage that 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.
[0019] According to an embodiment of the present invention 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.
[0020] The invention, according to further aspects, provides a system of reducing noise
in a hearing aid, a computer program and a computer program product as recited in
claims 26, 27 and 28.
[0021] Further specific variations of the invention are defined by the further claims. Other
aspects and advantages of the present invention will become more apparent from the
following detailed description taken in conjunction with the accompanying drawings
which illustrate, by way of example, the principles of the invention.
BRIEF DESCRIPTION OF THE DRAWINGS
[0022] The invention will be readily understood by the following detailed description in
conjunction with the accompanying drawings, wherein like reference numerals designate
like structural elements, and in which:
- Figs. 1a
- 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; and
- Fig. 6
- depicts a block diagram of part of a hearing aid according to an embodiment of the
present invention.
DESCRIPTION OF EMBODIMENTS OF THE INVENTION
[0023] Reference is first made to Fig. 1 a 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. Document
PCT/EP2005/055305 (mentioned above) describes this in detail.
[0024] Reference is now made to Fig. 1 b, which shows a hearing aid 200 according to the
first embodiment of the present invention.
[0025] 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.
[0026] 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.
[0027] 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.
[0028] For the sake of computations, 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. As the direct transmission gain is typically less than one, i.e. the
log value expressed in dB, will normally be a negative number. However, as there is
a natural Helmholz resonance by an earpiece placed in an ear canal there will be frequencies
where the DTG is above one, i.e. 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 methods described in the document
PCT/EP2005/055305 to calculate a direct transmission gain for the hearing aid gain used by a certain
user.
[0029] 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.
[0030] Reference is now made to Fig. 2, which 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.
[0031] 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. This general phenomenon is illustrated
in Fig. 2, which illustrates the addition of two signals with differing magnitude
and phase.
[0032] At a certain frequency, the sum of two harmonic signals can be written as

[0033] In our example, A1 = 1, ϕ
1 = 0 and A2 ∝ f. ϕ
2 is either 0, π or ∝ f. With simple calculations, 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. However,
since 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 ϕ
2 ∝ f. It is to be noted that 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.
[0034] The lower dash-dotted gray line shows that in case the two signals are out of phase
by π with the exact same amplitude, the total 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
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. These two cases, however, are rarely met
with respect to the hearing aid sound and the direct sound, since both have a varying
frequency dependent phase. The black line therefore exemplifies how the total sound
pressure might look like, if the relative phase depends linearly on frequency. Note,
that at some frequencies, constructive interference increases the magnitude of the
total signal, whereas for other frequencies, destructive interference diminishes the
total signal. Since the signals do not cancel out as such at frequencies where the
relative phase is almost π and the relative amplitude is not quite 1, this phenomenon
is called
phase disruption.
[0035] 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.
[0036] As is observed in Fig. 2, 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. Thus there is a lower threshold for the gain setting,
equal to the directly transmitted gain +
k, as suggested by the scale in Fig. 4 to the right. The safety margin is the factor
k, 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 SII-optimization.
[0037] 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. A
1 is DTG and A
2 is HA. Note, that 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.
[0038] For example it may be read from the curve that if one signal is 10 dB larger than
the other, the 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. Of course, a value of
about 1 dB would incur a high risk of phase disruption. A value of
k = 7 or
k = 8 gives a phase disruption range of about +-3 dB, which may be considered acceptable.
[0039] If 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. On the contrary, if 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 ≈ A
1), that the strength of the summed signal may vary significantly as indicated in Fig.
3.
[0040] Thus, in the current invention, the factor k, 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.
[0041] 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.
[0042] According to an embodiment, 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.
[0043] 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.
[0044] The Fig. 4 shows the directly transmitted sound in dB. This gain function, called
the direct transmission gain, 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.
[0045] 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 in 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.
[0046] The letter k in these figures 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. In the Fig. 4 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).
[0047] In the diagrams of Fig. 5, the minimum amplification for
k = 8 dB for the two frequency bands are marked on the graphs, containing the hearing
aid gain adjustment necessary to find the optimum gain setting with regards to speech
intelligibility. 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.
[0048] 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. By varying the gain in one or more
frequency bands an optimum setting of the gain in each frequency band is determined
leading to an optimum SII for the hearing aid.
[0049] The diagrams in 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.
[0050] The SII optimization according to the prior art does not presently take the direct
sound arriving through e.g. the vent into account. However, 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.
[0051] A further explanation on how SII is used for noise reduction in a hearing aid is
found in
WO-A-20051051039.
[0052] The diagrams in 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
ø and 3 mm
ø) in combination with two hearing losses (flat 40 dB HL and flat 80 dB HL).
[0053] 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.
[0054] 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 SI-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. However,
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.
[0055] In the graph (upper right pane) for a severe hearing loss (HTL 80dB) combined with
a small vent (1 mm), no changes occur to the optimum parameter setting resulting in
the optimum SII when the minimum amplification is taken into consideration compared
to the conventional optimum parameter setting where the gain can be varied in the
entire area. In contrary, a large vent (3 mm) and a mild hearing loss (HTL=40 dB)
may allow enough direct sound to enter through the vent to influence or even dominate
the total sound pressure at the eardrum (lower left pane), such that the optimum gain
setting of the frequency bands is quite different when the minimum amplification is
used to limit the gain settings of the frequency bands, than if the frequency bands
are varied without taken this into account. In such cases this would lead to a much
better parameter setting of the gain in the various frequency bands.
[0056] Therefore 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.
[0057] A main advantage for the present invention is therefore that the SII is optimized
under consideration of the actual in-situ acoustic surroundings.
[0058] It is evident for the person skilled in the art that the shown iterative path may
vary greatly from a real iterative path, both due to the optimization method and to
the fact that optimization occurs in all bands.
[0059] Reference is now made to Fig. 6, which shows a part of a hearing aid 300 according
to another embodiment of the present invention.
[0060] 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. If
the level of the noise reduced output signal plus the safety margin k is at or below
the DTG, the comparator produces an error signal 635 which is fed to the SII optimizer
610 as a further input parameter which is taken into account during optimization of
the SII so that the noise reduced output signal will not be attenuated below the threshold
anymore in order to avoid phase disruption.
[0061] In a modified embodiment 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.
[0062] According to embodiments of the present invention, 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). Obviously, it is preferred that the whole system
is implemented in a single digital component even though some parts could be implemented
in other ways - all known to the skilled person.
[0063] 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. According to another embodiment, 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.
[0064] According to a further embodiment, 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.
[0065] Having described and illustrated the principles of the present invention in embodiments
thereof, it should be apparent to those skilled in the art that the present invention
may be modified in arrangement and detail without departing from such principles.
1. A hearing aid (200) comprising at least one microphone (210), a signal processing
means (220) and an output transducer (230), wherein said signal processing means is
adapted to receive an input signal from the microphone, wherein said signal processing
means is adapted to apply a hearing aid gain to said input signal to produce an output
signal to be output by said output transducer,
characterised in that
said signal processing means further comprises means for adjusting said hearing aid
gain according to direct transmission gain calculated for the hearing aid;
wherein said means for adjusting said hearing aid gain is adapted to adjust said hearing
aid gain to a value not below said direct transmission gain.
2. The hearing aid according to claim 1, wherein said means for adjusting said hearing
aid gain comprises means for applying dynamic noise reduction techniques.
3. The hearing aid according to claim 2, wherein said means for adjusting said hearing
aid gain are adapted to improve speech intelligibility in noise of said output signal.
4. The hearing aid according to claim 3, wherein said means for adjusting said hearing
aid gain further comprises means for optimizing a speech intelligibility index.
5. The hearing aid according to claim 4, wherein said means for adjusting said hearing
aid gain is adapted to optimize said speech intelligibility index to produce a set
of frequency dependent speech intelligibility index gain values for each time sample
of said input signal.
6. The hearing aid according to one of the preceding claims, wherein said means for adjusting
said hearing aid gain provides a safety margin k and is adapted to adjust said hearing
aid gain to a value not below said direct transmission gain plus said safety margin.
7. The hearing aid according to one of claims 4 to 5, wherein said means for optimizing
a speech intelligibility index is adapted to calculate a speech intelligibility index
gain as a function of a plurality of input parameters.
8. The hearing aid according to claim 7, wherein said input parameters comprises at least
one of a frequency dependent hearing threshold level, an estimated noise level, and
an estimated speech level.
9. The hearing aid according to claim 5 or 7 to 8, wherein said means for adjusting said
hearing aid gain is adapted to calculate a noise reducing hearing aid gain from an
initial hearing aid gain and said optimized speech intelligibility index gain, and
to adjust said noise reducing hearing aid gain to a value not below a threshold level.
10. The hearing aid according to claim 9, wherein said threshold level is the level of
said direct transmission gain.
11. The hearing aid according to claim 9, wherein said threshold level is the level of
said direct transmission gain plus a safety margin.
12. The hearing aid according to one of claims 9 to 11, wherein said means for adjusting
said hearing aid gain is adapted to detect the level of said noise reducing hearing
aid gain before adjustment and, if said noise reducing hearing aid gain would be below
said threshold level, to input said noise reducing hearing aid gain before adjustment
as a further input parameter to said means for calculating a speech intelligibility
index.
13. The hearing aid according to claim 6, wherein said safety margin is a gain value in
the range of 0 to 15 dB, preferably in the range of 5 to 15 dB.
14. The hearing aid according to claim 6, wherein said safety margin is a gain value of
5 to 8 dB, preferably 7 to 8 dB.
15. The hearing aid according to one of the preceding claims, further comprising a band-split
filter for converting said input signal into band-split input signals of a plurality
of frequency bands and wherein said hearing aid is further adapted to process said
band-split input signals in each of said frequency bands independently.
16. A method of reducing noise in a hearing aid (200) comprising at least one microphone
(210) producing an input signal, a signal processing means (220) producing an output
signal from said input signal, and an output transducer (230) outputting said output
signal, wherein said method comprises:
- storing a direct transmission gain calculated for said hearing aid and its user
in a memory of said hearing aid; and
- applying a hearing aid gain to said input signal to produce said output signal,
characterised in that
said hearing aid gain is adjusted by said direct transmission gain so that said hearing
aid gain is not set to a value below said direct transmission gain.
17. The method according to claim 16, wherein said step of adjusting said hearing aid
gain comprises the step of applying dynamic noise reduction techniques.
18. The method according to claim 17, wherein said step of adjusting said hearing aid
gain comprises improving speech intelligibility in noise of said output signal.
19. The method according to claim 18, wherein said step of adjusting said hearing aid
gain further comprises a step of optimizing a speech intelligibility index.
20. The method according to one of claims 16 to 19, wherein said step of adjusting said
hearing aid gain comprises calculating a speech intelligibility index gain reducing
the noise in said output signal and adjusting said hearing aid gain by said speech
intelligibility index gain.
21. The method according to claim 19, wherein said step of adjusting said hearing aid
gain comprises optimizing said speech intelligibility index to produce a set of frequency
dependent speech intelligibility index gain values for each time sample of said input
signal.
22. The method according to claim 20, wherein said speech intelligibility index gain is
calculated with said direct transmission gain as a constraint to ensure that said
hearing aid gain is not set to a value below said direct transmission gain.
23. The method according to one of claims 16 to 22, wherein said hearing aid gain is not
set to a value below said direct transmission gain plus a safety margin k.
24. The method according to one of claims 20 to 23, further comprising the step of converting
said input signal into band-split input signals of a plurality of frequency bands
and wherein said method is further carried out for each of said frequency bands.
25. The method according to one of claim 20 or 22, wherein said speech intelligibility
index gain comprises a set of frequency dependent gain values which are calculated
simultaneously for an actual time sample of said frequency dependent input signal.
26. A system of reducing noise in a hearing aid comprising means adapted to carry out
a method according to one of claims 16 to 25.
27. A computer program comprising executable program code which, when executed on a computer,
executes a method according to one of claims 16 to 25.
28. A computer program product containing a computer readable medium with executable program
code which, when executed on a computer, executes a method according to one of claims
16 to 25.
1. Hörgerät (200) umfassend zumindest ein Mikrofon (210), eine Signalverarbeitungseinrichtung
(220) und einen Ausgangsmesswandler (230), wobei die Signalverarbeitungseinrichtung
geeignet ist zum Empfangen eines Eingangssignals von dem Mikrofon, wobei die Signalverarbeitungseinrichtung
geeignet ist, zum Anwenden einer Hörgeräteverstärkung auf das Eingangssignal, um ein
Ausgangssignal zu erzeugen, das von dem Ausgangsmesswandler ausgegeben werden soll,
dadurch gekennzeichnet, dass
die Signalverarbeitungseinrichtung weiterhin Mittel zum Justieren der Hörgeräteverstärkung
entsprechend einer für das Hörgerät berechneten direkten Übertragungsverstärkung umfasst;
wobei das Mittel zum Justieren der Hörgeräteverstärkung geeignet ist zum Justieren
der Hörgeräteverstärkung auf einen Wert, der nicht unterhalb der direkten Übertragungsverstärkung
liegt.
2. Hörgerät nach Anspruch 1, wobei das Mittel zum Justieren der Hörgeräteverstärkung
Mittel zum Anwenden von dynamischen Rauschverminderungstechniken umfasst.
3. Hörgerät nach Anspruch 2, wobei die Mittel zum Justieren der Hörgeräteverstärkung
geeignet sind zum Verbessern der Sprachverständlichkeit im Rauschen des Ausgangssignals.
4. Hörgerät nach Anspruch 3, wobei das Mittel zum Justieren der Hörgeräteverstärkung
weiterhin Mittel zum Optimieren eines Sprachverständlichkeitsindexes umfasst.
5. Hörgerät nach Anspruch 4, wobei das Mittel zum Justieren der Hörgeräteverstärkung
geeignet ist zum Optimieren des Sprachverständlichkeitsindexes, um einen Satz von
frequenzabhängigen Verstärkungswerten des Sprachverständlichkeitsindexes für jeden
Zeitabtastwert des Eingangssignals zu erzeugen.
6. Hörgerät nach einem der vorhergehenden Ansprüche, wobei das Mittel zum Justieren der
Hörgeräteverstärkung einen Sicherheitsspielraum k vorsieht und geeignet ist zum Justieren
der Hörgeräteverstärkung auf einen Wert, der nicht unterhalb der direkten Übertragungsverstärkung
plus dem Sicherheitsspielraum liegt.
7. Hörgerät nach einem der Ansprüche 4 bis 5, wobei das Mittel zum Optimieren des Sprachverständlichkeitsindexes
geeignet ist zum Berechnen einer Sprachverständlichkeitsindexverstärkung als eine
Funktion einer Mehrzahl von Eingangsparametern.
8. Hörgerät nach Anspruch 7, wobei die Eingangsparameter zumindest einen von einem frequenzabhängigen
Hörschwellenpegel, einem geschätzten Rauschpegel und einem geschätzten Sprachpegel
umfassen.
9. Hörgerät nach einem der Ansprüche 5 oder 7 bis 8, wobei das Mittel zum Justieren der
Hörgeräteverstärkung geeignet ist zum Berechnen einer rauschvermindernden Hörgeräteverstärkung
aus einer anfänglichen Hörgeräteverstärkung und der optimierten Sprachverständlichkeitsindexverstärkung
und zum Justieren der rauschvermindernden Hörgeräteverstärkung auf einen Wert, der
nicht unterhalb eines Schwellenpegels liegt.
10. Hörgerät nach Anspruch 9, wobei der Schwellenpegel der Pegel der direkten Übertragungsverstärkung
ist.
11. Hörgerät nach Anspruch 9, wobei der Schwellenpegel der Pegel der direkten Übertragungsverstärkung
plus einem Sicherheitsspielraum ist.
12. Hörgerät nach einem der Ansprüche 9 bis 11, wobei das Mittel zum Justieren der Hörgeräteverstärkung
geeignet ist zum Erkennen des Pegels der rauschvermindernden Hörgeräteverstärkung
vor dem Justieren und, wenn die rauschvermindernden Hörgeräteverstärkung unter dem
Schwellenpegel sein sollte, zum Eingeben der rauschvermindernden Hörgeräteverstärkung
vor dem Justieren als einen weiteren Parameter für das Mittel zum Berechnen eines
Sprachverständlichkeitsindexes.
13. Hörgerät nach Anspruch 6, wobei der Sicherheitsspielraum ein Verstärkungswert in dem
Bereich von 0 bis 15 dB, vorzugsweise in dem Bereich 5 bis 15 dB ist.
14. Hörgerät nach Anspruch 6, wobei der Sicherheitsspielraum ein Verstärkungswert in dem
Bereich von 5 bis 8 dB, vorzugsweise in dem Bereich 7 bis 8 dB ist.
15. Hörgerät nach einem der vorhergehenden Ansprüche, weiterhin umfassend einen Bandteilungsfilter
zum Umwandeln der Eingangssignale in bandgeteilte Eingangssignale einer Mehrzahl von
Frequenzbändern und wobei das Hörgerät weiterhin geeignet ist zum unabhängigen Verarbeiten
der bandgeteilten Eingangssignale in jedem der Frequenzbänder.
16. Verfahren zur Rauschverminderung in einem Hörgerät (200) umfassend zumindest ein Mikrofon
(210), das ein Eingangssignal erzeugt, eine Signalverarbeitungseinrichtung (220),
die aus dem Eingangssignal ein Ausgangssignal erzeugt, und einen Ausgangsmesswandler
(230), der das Ausgangssignal ausgibt, wobei das Verfahren umfasst:
- Speichern einer direkten Übertragungsverstärkung, berechnet für das Hörgerät und
seinen Benutzer, in einem Speicher des Hörgerätes; und
- Anwenden einer Hörgeräteverstärkung auf das Eingangssignal, um das Ausgangssignal
zu erzeugen,
dadurch gekennzeichnet, dass
das Hörgerät mittels der direkten Übertragungsverstärkung justiert wird, sodass die
Hörgeräteverstärkung nicht auf einen Wert gesetzt wird, der unter der direkten Übertragungsverstärkung
liegt.
17. Verfahren nach Anspruch 16, wobei der Schritt des Justierens der Hörgeräteverstärkung
den Schritt des Anwendens von dynamischen Rauschverminderungstechniken umfasst.
18. Verfahren nach Anspruch 17, wobei der Schritt des Justierens der Hörgeräteverstärkung
umfasst Verbessern der Sprachverständlichkeit im Rauschen des Ausgangssignals.
19. Verfahren nach Anspruch 18, wobei der Schritt des Justierens der Hörgeräteverstärkung
weiterhin einen Schritt des Optimierens eines Sprachverständlichkeitsindexes umfasst.
20. Verfahren nach einem der Ansprüche 16 bis 19, wobei der Schritt des Justierens der
Hörgeräteverstärkung umfasst Berechnen einer Sprachverständlichkeitsindexverstärkung,
die das Rauschen in dem Ausgangssignal vermindert, und Justieren der Hörgeräteverstärkung
mittels der Sprachverständlichkeitsindexverstärkung.
21. Verfahren nach Anspruch 19, wobei der Schritt des Justierens der Hörgeräteverstärkung
umfasst Optimieren des Sprachverständlichkeitsindexes zum Erzeugen eines Satzes von
frequenzabhängigen Verstärkungswerten des Sprachverständlichkeitsindexes für jeden
Zeitabtastwert des Eingangssignals.
22. Verfahren nach Anspruch 20, wobei die Sprachverständlichkeitsindexverstärkung berechnet
wird, mit der direkten Übertragungsverstärkung als eine Bedingung, um sicher zu machen,
dass die Hörgeräteverstärkung nicht auf einen Wert gesetzt wird, der unter der direkten
Übertragungsverstärkung liegt.
23. Verfahren nach einem der Ansprüche16 bis 22, wobei die Hörgeräteverstärkung nicht
auf einen Wert gesetzt wird, der unter der direkten Übertragungsverstärkung plus einem
einen Sicherheitsspielraum k liegt.
24. Verfahren nach einem der Ansprüche 20 bis 23, weiterhin umfassend den Schritt des
Umwandelns der Eingangssignale in bandgeteilte Eingangssignale einer Mehrzahl von
Frequenzbändern und wobei das Verfahren weiterhin für jedes der Frequenzbänder ausgeführt
wird.
25. Verfahren nach einem der Ansprüche 20 oder 22, wobei die Sprachverständlichkeitsindexverstärkung
einen Satz von frequenzabhängigen Verstärkungswerten umfasst, die gleichzeitig für
einen tatsächlichen Zeitabtastwert der frequenzabhängigen Eingangssignale berechnet
werden.
26. System zur Rauschverminderung in einem Hörgerät umfassend Mittel geeignet zum Ausführen
eines Verfahrens nach einem der Ansprüche 16 bis 25.
27. Computerprogramm umfassend ausführbaren Programmcode, welcher, wenn auf einem Computer
ausgeführt, ein Verfahren nach einem der Ansprüche 16 bis 25 ausführt.
28. Computerprogramm-Produkt umfassend ein computerlesbares Medium mit ausführbarem Programmcode,
welcher, wenn auf einem Computer ausgeführt, ein Verfahren nach einem der Ansprüche
16 bis 25 ausführt.
1. Prothèse auditive (200) comprenant au moins un microphone (210), un moyen de traitement
de signal (220) et un transducteur de sortie (230), dans laquelle ledit moyen de traitement
de signal est à même de recevoir un signal d'entrée du microphone, dans laquelle ledit
moyen de traitement de signal est à même d'appliquer un gain de prothèse auditive
audit signal d'entrée pour produire un signal de sortie à délivrer par ledit transducteur
de sortie,
caractérisée en ce que :
ledit moyen de traitement de signal comprend en outre un moyen de réglage dudit gain
de prothèse auditive selon un gain de transmission directe calculé pour la prothèse
auditive ;
dans laquelle ledit moyen de réglage dudit gain de prothèse auditive est à même d'ajuster
ledit gain de prothèse auditive à une valeur qui n'est pas inférieure audit gain de
transmission directe.
2. Prothèse auditive selon la revendication 1, dans laquelle ledit moyen de réglage dudit
gain de prothèse auditive comprend un moyen pour appliquer des techniques de réduction
du bruit dynamiques.
3. Prothèse auditive selon la revendication 2, dans laquelle ledit moyen de réglage dudit
gain de prothèse auditive est à même d'améliorer l'intelligibilité vocale dans le
bruit dudit signal de sortie.
4. Prothèse auditive selon la revendication 3, dans laquelle ledit moyen de réglage dudit
gain de prothèse auditive comprend en outre un moyen d'optimisation d'un indice d'intelligibilité
vocale.
5. Prothèse auditive selon la revendication 4, dans laquelle ledit moyen de réglage dudit
gain de prothèse auditive est à même d'optimiser ledit indice d'intelligibilité vocale
pour produire un jeu de valeurs de gain d'indice d'intelligibilité vocale fonction
de la fréquence pour chaque échantillon de temps dudit signal d'entrée.
6. Prothèse auditive selon l'une quelconque des revendications précédentes, dans laquelle
ledit moyen de réglage dudit gain de prothèse auditive fournit une marge de sécurité
k et est à même de régler ledit gain de prothèse auditive à une valeur qui n'est pas
inférieure audit gain de transmission directe plus ladite marge de sécurité.
7. Prothèse auditive selon l'une quelconque des revendications 4 à 5, dans laquelle ledit
moyen d'optimisation d'un indice d'intelligibilité vocale est à même de calculer un
gain d'indice d'intelligibilité en fonction d'une pluralité de paramètres d'entrée.
8. Prothèse auditive selon la revendication 7, dans laquelle lesdits paramètres d'entrée
comprennent au moins l'un ou l'autre d'un niveau de seuil auditif fonction de la fréquence,
d'un niveau de bruit estimé et d'un niveau vocal estimé.
9. Prothèse auditive selon les revendications 5 ou 7 à 8, dans laquelle ledit moyen de
réglage dudit gain de prothèse auditive est à même de calculer un gain réducteur de
bruit de prothèse auditive à partir d'un gain initial de prothèse auditive et dudit
gain d'index d'intelligibilité vocale optimisé et de régler ledit gain de prothèse
auditive réducteur de bruit à une valeur non inférieure à un niveau de seuil.
10. Prothèse auditive selon la revendication 9, dans laquelle ledit niveau de seuil est
le niveau dudit gain de transmission directe.
11. Prothèse auditive selon la revendication 9, dans laquelle ledit niveau de seuil est
le niveau dudit gain de transmission directe plus une marge de sécurité.
12. Prothèse auditive selon l'une quelconque des revendications 9 à 11, dans laquelle
ledit moyen de réglage dudit gain de prothèse auditive est adapté pour détecter le
niveau dudit gain de prothèse auditive réducteur de bruit avant réglage et, si ledit
gain de prothèse auditive réducteur de bruit se trouvait en dessous dudit niveau de
seuil, saisir ledit gain de prothèse auditive réducteur de bruit avant réglage comme
autre paramètre d'entrée dans ledit moyen de calcul d'un indice d'intelligibilité
vocale.
13. Prothèse auditive selon la revendication 6, dans laquelle ladite marge de sécurité
est une valeur de gain dans la plage de 0 à 15 dB, de préférence dans la plage de
5 à 15 dB.
14. Prothèse auditive selon la revendication 6, dans laquelle ladite marge de sécurité
est une valeur de gain de 5 à 8 dB, de préférence de 7 à 8 dB.
15. Prothèse auditive selon l'une quelconque des revendications précédentes, comprenant
en outre un filtre à bande scindée pour convertir ledit signal d'entrée en signaux
d'entrée à bande scindée d'une pluralité de bandes de fréquences et dans laquelle
ladite prothèse auditive est en outre à même de traiter lesdits signaux d'entrée à
bande scindée dans chacune desdites bandes de fréquences indépendamment.
16. Procédé de réduction de bruit dans une prothèse auditive (200) comprenant au moins
un microphone (210) produisant un signal d'entrée, un moyen de traitement de signal
(220) produisant un signal de sortie à partir dudit signal d'entrée et un transducteur
de sortie (230) délivrant ledit signal de sortie, dans lequel ledit procédé comprend
les étapes consistant à :
- stocker un gain de transmission directe calculé pour ladite prothèse auditive et
son utilisateur dans une mémoire de ladite prothèse auditive ; et
- appliquer un gain de prothèse auditive audit signal d'entrée pour produire ledit
signal de sortie,
caractérisé en ce que :
ledit gain de prothèse auditive est réglé par ledit gain de transmission directe de
sorte que ledit gain de prothèse auditive ne soit pas réglé à une valeur inférieure
audit gain de transmission directe.
17. Procédé selon la revendication 16, dans lequel ladite étape de réglage dudit gain
de prothèse auditive comprend l'étape d'application de techniques de réduction de
bruit dynamiques.
18. Procédé selon la revendication 17, dans lequel ladite étape de réglage dudit gain
de prothèse auditive comprend l'amélioration de l'intelligibilité vocale dans le bruit
dudit signal de sortie.
19. Procédé selon la revendication 18, dans lequel ladite étape de réglage dudit gain
de prothèse auditive comprend en outre une étape d'optimisation d'un indice d'intelligibilité
vocale.
20. Procédé selon l'une quelconque des revendications 16 à 19, dans lequel ladite étape
de réglage dudit gain de prothèse auditive comprend le calcul d'un gain d'indice d'intelligibilité
vocale, réduisant le bruit dans ledit signal de sortie et réglant ledit gain de prothèse
auditive par ledit gain d'indice d'intelligibilité vocale.
21. Procédé selon la revendication 19, dans lequel ladite étape de réglage dudit gain
de prothèse auditive comprend l'optimisation dudit indice d'intelligibilité vocale
pour produire un jeu de valeurs de gain d'indice d'intelligibilité vocale fonction
de la fréquence pour chaque échantillon de temps dudit signal d'entrée.
22. Procédé selon la revendication 20, dans lequel ledit gain d'indice d'intelligibilité
vocale est calculé avec ledit gain de transmission directe comme une contrainte pour
s'assurer que ledit gain de prothèse auditive ne soit pas réglé à une valeur inférieure
audit gain de transmission directe.
23. Procédé selon l'une quelconque des revendications 16 à 22, dans lequel ledit gain
de prothèse auditive n'est pas réglé à une valeur située en dessous dudit gain de
transmission directe plus une marge de sécurité k.
24. Procédé selon l'une quelconque des revendications 20 à 23, comprenant en outre l'étape
de conversion dudit signal d'entrée en signaux d'entrée à bande scindée d'une pluralité
de bandes de fréquences et dans lequel ledit procédé est en outre réalisé pour chacune
desdites bandes de fréquences.
25. Procédé selon l'une quelconque des revendications 20 ou 22, dans lequel ledit gain
d'indice d'intelligibilité vocale comprend un jeu de valeurs de gain fonction de la
fréquence, qui sont calculées simultanément pour un échantillon de temps courant dudit
signal d'entrée fonction de la fréquence.
26. Système de réduction de bruit dans une prothèse auditive comprenant des moyens qui
sont à même de réaliser un procédé selon l'une quelconque des revendications 16 à
25.
27. Programme informatique comprenant un code de programmation exécutable qui, lorsqu'il
est exécuté sur un ordinateur, exécute un procédé selon l'une quelconque des revendications
16 à 25.
28. Produit de programmation informatique contenant un support lisible sur ordinateur
avec un code de programmation exécutable qui, lorsqu'il est exécuté sur un ordinateur,
exécute un procédé selon l'une quelconque des revendications 16 à 25.