Method of operating a hearing aid system and a hearing aid system

Abstract

A method (100) of operating a hearing aid system while providing that the loudness of the amplified hearing aid input signal as experienced by the hearing aid user is the same as the loudness of the sound providing the unprocessed hearing aid input signal as experienced by a normal hearing person. The invention also refers to a hearing aid system adapted to carry out said method.

Description

[0001] METHOD OF OPERATING A HEARING AID SYSTEM AND A HEARING AID

[0002] SYSTEM

[0003] The present invention relates to a method of operating a hearing aid system. The present invention also relates to a hearing aid system adapted to carry out said method.

[0004] BACKGROUND OF THE INVENTION

[0005] Generally a hearing aid system according to the invention is understood as meaning any device which provides an output signal that can be perceived as an acoustic signal by a user or contributes to providing such an output signal, and which has means which are customized to compensate for an individual hearing loss of the user or contribute to compensating for the hearing loss of the user. They are, in particular, hearing aids which can be worn on the body or by the ear, in particular on or in the ear, and which can be fully or partially implanted. However, some devices whose main aim is not to compensate for a hearing loss, may also be regarded as hearing aid systems, for example consumer electronic devices (televisions, hi-fi systems, mobile phones, MP3 players etc.) provided they have, however, measures for compensating for an individual hearing loss.

[0006] Within the present context a traditional hearing aid can be understood as a small, battery-powered, microelectronic device designed to be worn behind or in the human ear by a hearing-impaired user. Prior to use, the hearing aid is adjusted by a hearing aid fitter according to a prescription. The prescription is based on a hearing test, resulting in a so-called audiogram, of the performance of the hearing-impaired user’s unaided hearing. The prescription is developed to reach a setting where the hearing aid will alleviate a hearing loss by amplifying sound at frequencies in those parts of the audible frequency range where the user suffers a hearing deficit. A hearing aid comprises one or more microphones, a battery, a microelectronic circuit comprising a signal processor, and an acoustic output transducer. The signal processor is preferably a digital signal processor. The hearing aid is enclosed in a casing suitable for fitting behind or in a human ear.

[0007] Within the present context a hearing aid system may comprise a single hearing aid (a so called monaural hearing aid system) or comprise two hearing aids, one for each ear of the hearing aid user (a so called binaural hearing aid system). Furthermore, the hearing aid system may comprise an external device, such as a smart phone having software applications adapted to interact with other devices of the hearing aid system. Thus within the present context the term “hearing aid system device” may denote a hearing aid or an external device.

[0008] However, the audio device system may also include a remote microphone system (which generally can also be considered an external device) comprising additional microphones and / or may even include a remote server providing abundant processing resources and generally these additional devices will also include link means adapted to operationally connect to the various other devices of the hearing aid system.

[0009] The mechanical design has developed into a number of general categories. As the name suggests, Behind-The-Ear (BTE) hearing aids are worn behind the ear. To be more precise, an electronics unit comprising a housing containing the major electronics parts thereof is worn behind the ear. An earpiece for emitting sound to the hearing aid user is worn in the ear, e.g. in the concha or the ear canal. In a traditional BTE hearing aid, a sound tube is used to convey sound from the output transducer, which in hearing aid terminology is normally referred to as the receiver, located in the housing of the electronics unit and to the ear canal. In some modern types of hearing aids, a conducting member comprising electrical conductors conveys an electric signal from the housing and to a receiver placed in the earpiece in the ear. Such hearing aids are commonly referred to as Receiver-In-The-Ear (RITE) hearing aids. In a specific type of RITE hearing aids the receiver is placed inside the ear canal. This category is sometimes referred to as Receiver-In-Canal (RIC) hearing aids.

[0010] In-The-Ear (ITE) hearing aids are designed for arrangement in the ear, normally in the funnel-shaped outer part of the ear canal. In a specific type of ITE hearing aids the hearing aid is placed substantially inside the ear canal. This category is sometimes referred to as Completely-In-Canal (CIC) hearing aids. This type of hearing aid requires an especially compact design in order to allow it to be arranged in the ear canal, while accommodating the components necessary for operation of the hearing aid.

[0011] With time new functionality, such as advanced noise reduction algorithms, is implemented in hearing aid systems. One example of a contemporary advanced noise reduction algorithm is the Speech Intelligibility Index (SII) based noise reduction that seeks to select the optimum gain based on the frequency dependent hearing threshold of a specific user and based on a model of speech intelligibility for a given frequency dependent signal to noise ratio. However, the gain hereby selected may not be an optimum selection for all users.

[0012] WO 2001 / 069504 discloses a hearing aid system wherein an additional personalized gain is applied on top of and independent of the various other gains (including e.g. adaptive gains from noise reduction algorithms) applied as part of the signal processing. However, this additional personalized gain is directed at adjustments of the prescribed gain curve and as such does not provide control of the actual hearing aid output level in any given sound environment because this also depends on the various other applied gains.

[0013] Hearing loss of a hearing impaired person is quite often frequency-dependent. This means that the hearing loss of the person varies depending on the frequency. Therefore, when compensating for hearing losses, it can be advantageous to utilize frequencydependent amplification. Hearing aids therefore often provide to split an input sound signal received by an input transducer of the hearing aid, into various frequency intervals, also called frequency bands, which are independently processed. In this way, it is possible to adjust the input sound signal of each frequency band individually to account for the hearing loss in respective frequency bands. The frequency dependent adjustment is normally done by implementing a band split filter and compressors for each of the frequency bands, so-called band split compressors, which may be summarised to a multi-band compressor. In this way, it is possible to adjust the gain individually in each frequency band depending on the hearing loss as well as the input level of the input sound signal in a specific frequency range. For example, a band split compressor may provide a higher gain for a soft sound than for a loud sound in its frequency band.

[0014] The filter banks used in such multi-band compressors are well known within the art of hearing aids, but are nevertheless based on a number of tradeoffs. Most of these tradeoffs deal with the frequency resolution as will be further described below. There are some very clear advantages of having a high resolution filter bank. The higher the frequency resolution, the better individual periodic components can be distinguished from each other. This gives a much finer signal analysis and enables more advanced signal processing. Especially noise reduction and speech enhancement schemes may benefit from a higher frequency resolution.

[0015] However, a filter bank with a high frequency resolution generally introduces a correspondingly long delay, which for most people will have a detrimental effect on the perceived sound quality.

[0016] It has therefore been suggested to reduce the delay incurred by filter banks, such as Discrete Fourier Transform (DFT) and Finite Impulse Response (FIR) filter banks by: applying a time-varying filter with a response that corresponds to the desired frequency dependent gains that were otherwise to be applied to the frequency bands provided by the filter banks.

[0017] Thus, basically, a hearing aid picks up an input signal and provides a processed output signal. Said processing involves amplification of said input signal according to the user’s needs. Amplification is carried out in an amplifier, usually including a compressor having a compressor gain.

[0018] Generally, the hearing loss of a hearing impaired is not linear. That is to say, the hearing ability may be almost normal at some sound pressure levels - typically at louder sound pressure levels, while being quite poor at other sound pressure levels - typically at softer sound pressure levels. The fact that amplification is needed especially for the softer sound pressure levels while not so much for the louder sound pressure levels is quite typical for many of the hearing impaired.

[0019] State of the art hearing aids are adapted to compensate for this common pattern of hearing loss, by means of a compressor. The compressor is adapted for adjusting the gain so as to vary with the current sound pressure level of the input signal. The variation of the level dependent compressor gain is defined in a compression characteristic. A state of the art hearing aid may include a compression characteristic for each frequency band of the input signal. Examples of hearing aids wherein the input signal is amplified in a compressor having a compressor gain that varies with sound pressure level in accordance with a compression characteristic are described in EP -B 1-1059016 and EP-B 1-0824845.

[0020] Traditionally, fitting of the hearing aid includes adjusting the compressor gain according to a general compression characteristic, in the following referred to as the standard rationale. The standard rationale takes into account the individual hearing loss, but is apart from that intended to accommodate the average hearing aid user.

[0021] However, many different amplification schemes or gain rationales, to compensate for a given hearing loss, exist, such as NAL, DSL, and various proprietary rationales. As a result, there is not a wide consensus among hearing aid experts and audiologists on which strategies are the best for all users. Rather, the preferred gain rationale for hearing aid experts, hearing care professionals and audiologists seem to be a matter of taste and / or experience.

[0022] However, even though the hearing loss of many hearing impaired follow the abovedescribed pattern regarding the need for a larger amplification of softer sound pressure levels but not necessarily an equally large amplification of louder sound pressure levels, individual differences exist. The need for amplification for one hearing impaired may even vary greatly from that of another having a similar hearing loss.

[0023] In an effort to make a conventional hearing aid compensate better for the specific hearing loss and preferences of the individual user, the hearing aid is furthermore fine- fitted to the individual user. The fine-fitting is traditionally carried out as additional adjustments to the standard fitting according to the standard rationale.

[0024] One of the problems with the existing way of fine-fitting a hearing aid to the individual user is that the compressor only provides limited possibility for fine adjusting the compression characteristic so as to fit the hearing loss of the individual user sufficiently accurate. This is due to the fact that the number of adjustment points, in each of which the compression characteristic of the compressor can be adjusted independently of the other adjustment points, is traditionally very limited. In many cases, the compression characteristic only has two adjustment points. Hence, adjustment of the compressor gain for one sound pressure level influences that of many other sound pressure levels, which may not be desirable. Hence, only a crude fitting of the compression characteristic to the hearing loss of the individual user is possible. This means that when fitting a hearing aid to an individual user, a compromise must be made between on one hand providing a sufficient amplification of the input signal for some sound input levels while on the other hand avoiding to amplify the input signal for other sound pressure levels to such an extent that the comfort level of the user is exceeded.

[0025] Another problem concerning the existing way of processing the input signal of a hearing aid is associated with signal optimisation. Implementation of various types of adaptive processing such as for instance speech intelligibility optimisation are becoming more widespread in the signal processing in hearing aids.

[0026] Unfortunately, fine-adjustments carried out on the compression characteristic during the fine-fitting of the hearing aid to the individual hearing aid user may be regarded as deviations from the optimal compression characteristic and may therefore to a great extent be eliminated or reduced by the adaptive processing. Hence, the effect of the fine-fitting of the hearing aid to the individual user is to a great extent never experienced by the user.

[0027] All in all, the existing method of processing and fitting has difficulties in meeting any requirement of individual deviation from the above described typical pattern of hearing loss of many hearing impaired.

[0028] Hence, a need for a more flexibly adjustable gain exists so as to be able to fit the actual hearing loss and individual preferences of the hearing aid user better.

[0029] Also, there is a need for a manner of avoiding that the effect of any fine-fitting is reduced or eliminated by other processing in the hearing aid, such as for instance adaptive processing.

[0030] It is therefore an object of the present invention to provide a method of operating a hearing aid system capable of providing improved sound quality and speech intelligibility for the hearing aid user.

[0031] SUMMARY OF THE INVENTION

[0032] The invention is set out in the appended set of claims. BRIEF DESCRIPTION OF THE DRAWINGS

[0033] By way of example, there is shown and described a preferred embodiment of this invention. As will be realized, the invention is capable of other embodiments, and its several details are capable of modification in various, obvious aspects all without departing from the invention. Accordingly, the drawings and descriptions will be regarded as illustrative in nature and not as restrictive. In the drawings:

[0034] Fig. 1 illustrates highly schematically a method of operating a hearing aid system

[0035] DETAILED DESCRIPTION

[0036] In the present context the term “hearing aid input signal” will generally be construed to mean an electrical (analog or digital) signal representing a sound. A beamformed signal (either monaural or binaural) is one example of such an electrical signal representing a sound. Another example is an electrical signal wirelessly streamed to the audio device system.

[0037] In the following the terms “hearing impaired (person)”, (specific) hearing aid user” and “(specific) hearing aid system user” may be used interchangeably.

[0038] It is also noted that in the following most if not all quantities are time and frequency dependent, e.g. MN= MN(t, f) but for simplicity (t,f) are generally excluded from the notation.

[0039] In the following an individualized masking model (or Masking model for persons with hearing impairment characterization) is considered to include an individual characterization of: auditory filter widening (i.e. a measure of frequency selectivity loss), hearing threshold in quiet (i.e. a measure of audibility loss) and post-masking (i.e. a measure of temporal resolution loss)

[0040] A masking model for normal hearing is obtained from a set of parameters that best matches average psychoacoustic data for normal hearing listeners.

[0041] According to an embodiment both masking models share a common gammatone filterbank and mask calculation. Thus, for improved efficiency the broadening, absolute threshold and forward masking stages only needs to be calculated in by the individualized masking model, whereby processing resources are saved.. Thus, by applying these masking models the masked thresholds for the reference normal hearing person (MN) and the hearing impaired user (M;) are provided, together with the input signal levels Lin), at each time instant and for each auditory band.

[0042] Overall, the present invention suggests a wide dynamic range compression (or hearing loss compensation) scheme to restore loudness in auditory bands. For this, the sound is analyzed with two masking and loudness models, the first masking and loudness model is tuned to an average or reference normal hearing listener, and the other masking and loudness model is tuned to the specific hearing aid users hearing loss in terms of audibility loss (i.e. audiogram), loss of frequency selectivity (or widening of auditory filters) and loss of temporal resolution (or increased post-masking).

[0043] A key aspect of the invention is that the loudness function grows from the masked threshold, and the calculated gain is the one that, applied to the input signal, brings the same loudness for the aided hearing aid user as for an unaided normal hearing listener in each of the auditory bands.

[0044] This requires a gain calculation method that aims at retrieving, for a given input level, the gain G required to produce the same loudness for a hearing impaired person (in the following abbreviated NH) (T() that a normal hearing person (in the following abbreviated NH) would experience without a hearing aid (T’w).

[0045] The loudness function grows from the masked threshold, i.e.

[0046] This loudness is then matched by the loudness for a hearing impaired person after applying the gain G, i.e.:

[0047] After solving for G, we obtain Additionally, the present invention provides a consistent prescription scheme that can restore the loudness experienced by a normal hearing listener - while also compensating for a specific hearing aid system users loss of audibility, frequency selectivity and temporal resolution.

[0048] Thus another advantage of the present invention is that it can provide a natural sound experience that is close to the sound experience of normal hearing persons.

[0049] This can be applied to all hearing aid fittings, but it is of particular interest for selffitting hearing aids and for first time hearing aid users (of which some will be tempted to start their hearing journey with a self-fitting hearing aid) because these users normally will be more sensitive to an un-natural sound experience.

[0050] Additionally, the present invention is advantageous in that the compensation rationale (i.e. the prescription) can be adapted to have significantly fewer parameters to control than most of the contemporary proprietary compensation rationales. This distinguishing feature obviously makes the this invention especially attractive with respect to the ease of delivering a hearing aid system to a user without the need of an audiologist intervening and fine-tuning many different amplification parameters.

[0051] Reference is now made to Fig. 1, that highly schematically illustrates a method 100 of operating a hearing aid system according to an embodiment of the invention.

[0052] In a first step 101 a first frequency dependent masking threshold for a normal hearing person is determined for a given hearing aid input signal;

[0053] In a second step 102 a second frequency dependent masking threshold for a hearing aid user is determined for a given hearing aid input signal;

[0054] In a third step 103 a first frequency dependent hearing aid gain is determined, wherein said first frequency dependent hearing aid gain is determined as the gain that, when applied to said hearing aid input signal, provides that the loudness of the amplified hearing aid input signal as experienced by the hearing aid user is the same as the loudness of the sound, providing the unprocessed hearing aid input signal, as experienced by a normal hearing person.

[0055] In a fourth and final step 104 said first frequency dependent hearing aid gain is applied to said hearing aid input signal or a signal derived therefrom. It is generally assumed that the input signal is a representation of the eardrum signal without hearing aid. This is a fair assumption for in-the-ear formfactors or for a so called Mic-Ric i.e. a hearing aid system with a microphones both behind and in the ear.

[0056] For a RiC or BTE microphone placement, this fulfilling this assumption is direction dependent and the assumption weaker.

[0057] Thus a masking model for normal hearing persons determines said first frequency dependent masking threshold (for each time instant and in each auditory band) and an individualized masking model determines said second frequency dependent masking for said (specific) hearing aid user. Finally said first frequency dependent hearing aid gain is determined by solving the problem of: “what is the gain that, applied to the input signal, brings the loudness of the processed signal as experienced by a given hearing aid user to be the same loudness as experienced by a normal hearing person when exposed to the unprocessed signal input signal ? ”

[0058] According to an embodiment a post-processing step ensures the following:

[0059] “The gain applied to sounds that are inaudible by normal hearing persons should not make the sound audible for a hearing aid used"

[0060] This is obtained by the following (additional) steps of operating a hearing aid system:

[0061] - determining whether a hearing aid input signal level, for a given frequency range, is below said first frequency dependent masking threshold (see step 101 above) for a normal hearing person; and if this is the case:

[0062] - adapting the magnitude of said first frequency dependent hearing aid gain, within said given frequency range, such that an amplified hearing aid input signal level is below said second frequency dependent masking threshold for said hearing aid user (see step 102 above), whereby it is ensured that sound inaudible for a normal hearing person is also inaudible for the hearing aid user.

[0063] In an embodiment the applied gains are smoothed in time and frequency in order to reduce audible artifacts.

[0064] More specifically, and according to an embodiment, said smoothing in time is carried out independently for each auditory band, with a first-order recursive symmetric smoother with a time constant of 20 milliseconds or in the range between 10 and 30 milliseconds or between 5 and 50 milliseconds.

[0065] Said first-order recursive symmetric smoother is especially advantageous with respect to:

[0066] - its low computational cost (which is ideal for real-time, low-power devices like hearing aids);

[0067] - its smooth and continuous output (that helps to avoids abrupt changes in the provided gain);

[0068] - its generally adjustable time constant (that allows tuning responsiveness vs. smoothness), and

[0069] - the filter is casual and consequently well suited for real-time implementations.

[0070] However, in alternative embodiments said first-order recursive symmetric smoother, could be replaced by one of: a moving average filter, a Gaussian smoothing filter or a Kalman filter.

[0071] In an embodiment a combined gain is determined and applied to the hearing aid input signal, wherein said combined gain comprises said first frequency dependent hearing aid gain with other gains comprising at least one of speech enhancement gain and noise reduction gain.

[0072] In an embodiment, an acclimatization period is used to scale the applied gains by a factor A (where A preferably is a number between zero and one), and where A is configured to be progressively increased during a period, possibly extending during weeks. Acclimatization is especially beneficial for this invention because the applied gains can become quite large.

[0073] In one embodiment, said hearing aid system is based on an analysis-synthesis system and in another (alternative) embodiment the hearing aid system is a low delay hearing aid system wherein the gains are synthesized as filter coefficients to be applied to the input signal via a FIR filter, that may or may not be warped.

[0074] In an embodiment only a sensorineural hearing loss is considered when determining for said hearing aid input signal, a second frequency dependent masking threshold for a hearing aid user. This means that if there is a mixed sensorineural / conductive hearing loss (given by a gap between bone-conducted and airborne thresholds of more than ~20 dB), only the sensorineural hearing loss should be considered for compensation according to the present invention (including the various embodiments disclosed above).

[0075] Thus only a fixed linear amplification shall be applied for the conductive hearing loss.

[0076] Let AT= Ac+ Asbe the airborne threshold (in dB HL) (i.e. the sum of the conductive Acand sensorineural Asthresholds and

[0077] = As(the bone-conducted threshold), with a conductive loss Ac= AT— ABand a sensorineural loss As.

[0078] The gain calculation will only compensate for the sensorineural loss As, so this will be the audibility threshold for the hearing impaired masking model. In addition, the linear gain Acwill be applied.

[0079] According to an embodiment an extra step will transform the calculated gains G into the actual gains G' to be applied dependent on whether open or semi-open fittings are used. Assuming that the hearing aid system comprises a so called closed fittings (such as an earpiece without a vent) then the assumption is that the aided sound at the ear is brought uniquely by the hearing aid, i.e. aid Lin+ G

[0080] Or in linear quantities

[0081] ^aid 9 ' ^in

[0082] Where G = 20 log10g.

[0083] However, when using open or semi-open fittings, the direct sound will come in:

[0084] ^aid (-9 L L)X ii

[0085] Here a models the transformation of the unaided input signal at the ear with and without the presence of the hearing aid.

[0086] This means that, for open and semi -open fittings, an extra step will transform the calculated gains G into the actual gains G' to be applied.

[0087] G = 201og10(^' + a) Isolating, we find:

[0088] Where A = 20 log10a.

Claims

CLAIMS1. A method of operating a hearing aid system comprising the steps of:- determining, for a hearing aid input signal, a first frequency dependent masking threshold for a normal hearing person;- determining, for said hearing aid input signal, a second frequency dependent masking threshold for a hearing aid user;- determining, as a first frequency dependent hearing aid gain, the gain that, when applied to said hearing aid input signal, provides that the loudness of the amplified hearing aid input signal as experienced by the hearing aid user is the same as the loudness of the sound providing the unprocessed hearing aid input signal as experienced by a normal hearing person; and- applying said first frequency dependent hearing aid gain to said hearing aid input signal or a signal derived therefrom.

2. .The method according to claim 1, comprising the further steps of:- determining whether a hearing aid input signal level, for a given frequency range, is below said first frequency dependent masking threshold for a normal hearing person; and if this is the case:- adapting the magnitude of said first frequency dependent hearing aid gain, within said given frequency range, such that an amplified hearing aid input signal level is below said second frequency dependent masking threshold for the hearing aid user, whereby it is ensured that sound inaudible for a normal hearing person is also inaudible for the hearing aid user.

3. The method according to any of the preceding claims, wherein said adaptation of the magnitude of said first frequency dependent hearing aid gain is carried out in accordance with the formula:G(t) = max [Gmin, min(Gmax , G(t - 1) - A)],- wherein the maximum magnitude of said first frequency dependent hearing aid gain Gmax is determined by subtracting said frequency dependent hearing aid input signal level from the second frequency dependent masking threshold for the hearing aid user;- wherein the minimum magnitude of said frequency dependent hearing aid gainGmin may be selected from the range between +5 dB and - 30 dB or between 0 dB and - 20 dB; and- wherein A controls the fixed rate of the gain decay from Gmax and to G™.

4. The method according to any of the preceding claims, wherein said step of determining, for said hearing aid input signal, a second frequency dependent masking threshold for the hearing aid user, comprises the further steps of:- providing an individualized masking model adapted to incorporate at least one of:- a measure of the hearing aid users frequency selectivity loss;- the hearing aid users hearing threshold; and- the hearing aid users temporal resolution loss.

5. The method according to any of the preceding claims, comprising the further steps of:- combining said first frequency dependent hearing aid gain with other gains comprising at least one of speech enhancement gain and noise reduction gain to provide a combined gain; and- applying said combined gain to the hearing aid input signal.

6. The method according to any of the preceding claims, comprising the further step of: smoothing in time and frequency said determined gains and hereby reducing sound artifacts.

7. A hearing aid system, configured to carry out a method according to any one of the claims 1 - 6.

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