Patent Publication Number: US-9426582-B2

Title: Automatic real-time hearing aid fitting based on auditory evoked potentials evoked by natural sound signals

Description:
TECHNICAL FIELD 
     The present application relates to hearing aids, and to the monitoring of auditory evoked potentials (AEP). The disclosure relates specifically to a hearing aid comprising means for picking up and analysing auditory evoked potentials, e.g. an auditory brainstem response (ABR). The application furthermore relates to a method of operating a hearing aid and to the use of a hearing aid. The application further relates to a data processing system comprising a processor and program code means for causing the processor to perform at least some of the steps of the method. 
     The disclosure may e.g. be useful in hearing aids or hearing aid systems where a continuous evaluation of a user&#39;s hearing thresholds is needed. 
     BACKGROUND 
     Fitting of a hearing aid to a particular person&#39;s hearing impairment generally requires knowledge of clinically measured hearing thresholds for the person in question. The auditory brainstem response (ABR) can be used as an objective estimate of audiometric hearing thresholds (e.g. [Stürzebecher et al., 2006]). ABR signals are traditionally measured by surface electrodes mounted on the head with one electrode at the vertex or in the middle of the forehead, one behind the ear on the mastoid or on the earlobe, and one ground electrode on the opposite side of the head. Future hearing aids may, however, include electrodes on the surface of the hearing aid shell facing the ear canal to record electric brain wave signals such as an electroencephalogram (EEG) (cf. e.g. [Lunner, 2010]). 
     A portable EEG monitoring apparatus is described in [Kidmose and Westermann, 2010]. A hearing aid comprising electrodes for detecting electrical signals such as brain waves is described in [Kidmose and Mandic, 2011]. The design of stimuli for a system for the recording of an auditory brainstem response (ABR) of a person is e.g. described in WO 2006/003172 A1. 
     EP1073314A1 deals with a method of fitting a hearing aid. The method involves detecting auditory, involuntary body signals from the wearer of the hearing aid using a measuring device connected to the wearer. An evaluation device generates starting parameters for the hearing aid based on the measured values. These parameters are transferred to the signal processing unit of the hearing aid for individual adjustments. A signal may be fed to at least one ear of the wearer of the hearing aid. The measuring device detects the involuntary body signals caused by this signal. The signal may be acoustic, electrical, electromagnetic or mechanical. 
     The present disclosure builds further on the idea of using auditory models to continuously be able to present an auditory ‘test signal’ that does not disturb the user of the hearing aid, as described in EP2581038A1. A large part of the description and drawings of the present disclosure is taken from EP2581038A1 and presented here in identical or slightly modified form. 
     SUMMARY 
     Auditory evoked potentials (AEPs) are a subclass of event-related potentials (ERP)s, such as auditory brainstem responses (ABR). ERPs are brain responses that are time-locked to some “event”, such as a sensory stimulus, a mental event (such as recognition of a target stimulus), or the omission of a stimulus. For AEPs, the “event” is a sound. AEPs (and ERPs) are very small electrical voltage potentials originating from the brain, e.g. recorded from the scalp in response to an auditory stimulus, such as different tones, speech sounds, etc. 
     The analysis of measured AEPs for a person can be used to estimate audiometric hearing thresholds (HTL) of that person. A fitting algorithm can be executed in the hearing aid using the estimated hearing thresholds as inputs to determine an appropriate frequency dependent gain for the user wearing the hearing aid. The analysis of the measured AEPs for a person can also be used to estimate other characteristics (herein especially supra-threshold characteristics) of the person&#39;s hearing, e.g. sensitivity to a) pitch changes, b) interaural b) time and b2) level differences, c) spectro-temporal modulations (modulations in frequency (f) AND time (t)), etc. A user&#39;s hearing ability may e.g. (to a certain extent) be defined by a number of (typically frequency (f) dependent) ‘threshold’ values, e.g. hearing thresholds (HT(f), minimum level for perception of a sound), frequency discrimination threshold (Δf TH (f), minimum difference in frequency for perception of difference (spectral masking)), time discrimination threshold (Δt TH (f), minimum difference in frequency for perception of different sounds (temporal masking)). 
     According to the present disclosure, signal processing parameters of the forward path (and thus the time and frequency dependent gain applied to the electric input signal before being provided to user via an output transducer) may be adaptively changed (e.g. over hours or days or months) to optimize the user&#39;s hearing ability (e.g. by optimizing (minimizing) one or more of the mentioned threshold values) as determined by the ongoing analysis of the user&#39;s brain response to (selected) sound (segments). 
     In the present disclosure, the hearing aid creates and maintains a database of observed sound stimuli and recorded AEPs, which enables the hearing aid to average AEPs over a number of naturally occurring sounds. Thus, the hearing aid is not (necessarily) generating additional test sounds. Instead (or primarily), the hearing aid selects a number of similar instances of naturally occurring sounds and averages over those. 
     The hearing aid is configured to decide how similar two naturally sound instances are, e.g. by comparing one or more of the frequency content, temporal envelope, the instantaneous frequency, localization cues, spectro-temporal modulation cues. This is similar to the theory of Just Noticeable Differences [Fastl &amp; Zwicker, 2007] that describes listener capabilities by the smallest difference (from the above list) that the listener can detect. 
     The describing models can be purely mathematically defined, as the short time Fourier Transform (STFT), which describes/encodes the sounds observed by the hearing aid according to the energy and phase in different frequency bands at specific times (c.f. Goodwin, Chapter 12, Benesty, Handbook of Speech Processing). In a further analysis single STFT channels can be transformed once more into the frequency domain to derive the corresponding temporal modulation spectrum [McWalter, 2012, pp 6-7]. 
     In a further refined analysis the full STFT, or if more biologically inspired the Auditory Spectogram, output undergoes a 2-dimensional Fourier Transformation which provides the spectrotemporal representation [Elihilai, 2004, pp 18]. 
     An alternative model is described by [McWalter 2012, pp 19], where sound is analysed in a biologically inspired manner that reveals the frequency content, temporal envelopes, modulation spectra, and cross-channel statistics. (c.f. e.g. [Blauert; 1983], [Elhilali; 2004], [McWalter; 2012], [Fastl &amp; Zwicker, 2007]). 
     Herein the hearing aid considers two sound instances to be equal, if psychoacoustic tests of just noticeable differences or AEP measures have shown that normal hearing listeners cannot discriminate between them. 
     The term ‘naturally occurring sounds’ is in the present context taken to mean sounds that are picked up by an input unit of the hearing aid, be they vocalized (speech, non-speech), music, naturally or artificially generated sounds, etc. 
     The low voltage electric signal from the user&#39;s brain picked up during a particular time segment or the corresponding amplified (or otherwise processed) brain signal(s) are in the following termed the ‘brain signal data’. 
     The use of (time segments of) naturally occurring sounds as ‘test signals’ for exciting an auditory response may be combined with the use of artificially generated test signals, e.g. during times of lack of appropriate input sounds, or during times of similar sound character (e.g. only low frequency content, with some hearing ability parameters NOT being challenged). This may e.g. be achieved either by ‘re-using’ characteristics of already identified input signal (segments) to generate similar test signals, or otherwise by generating and playing appropriate test signals, e.g. from a predefined database of signals or parameterized signals, e.g. to supplement brain data provided by the ‘naturally occurring sounds’. Such artificially generated test signals may then be played alone or mixed with the normally processed electric input signal. 
     In an embodiment, the hearing aid is configured—in a predefined mode of operation—to play auditory test signals (e.g. chirps, clicks, or narrowband signals such as tones making auditory steady state response, ASSR) from a hearing aid receiver under daily use and with the hearing aid equipped with electrodes to electrically measure brain signals. 
     An object of the present application is to provide a hearing aid capable of monitoring a user&#39;s hearing ability over time. 
     Objects of the application are achieved by the invention described in the accompanying claims and as described in the following. 
     A Hearing Aid: 
     In an aspect of the present application, an object of the application is achieved by a hearing aid comprising
         an ear part adapted for being mounted fully or partially at an ear or in an ear canal of a user (or partially implanted in the head of a user), the ear part comprising
           a housing,   at least one electrode located at a surface of said housing to allow said electrode(s) to contact the skin of a user when said ear part is operationally mounted on the user, the at least one electrode being adapted to pick up a low voltage electric signal from the user&#39;s brain,   
           an amplifier unit operationally connected to said electrode(s) and adapted for amplifying said low voltage electric signal(s) to provide amplified brain signal(s),   an input unit, e.g. an input transducer, for providing an electric audio input signal,   a signal processing unit, an output transducer for converting an electric output signal to an acoustic output sound,   said signal processing unit being operationally connected
           to said amplifier unit, and adapted to process said amplified brain signal(s) to provide a processed brain signal,   to said input unit, and adapted to apply a time and frequency dependent gain to said electric audio input signal or a signal originating therefrom and to provide a processed audio output signal, and   to said output transducer, allowing said processed audio output signal to be presented to the user as a processed acoustic signal,   
           said signal processing unit being configured to
           identify characteristic signal parameters describing a current time segment of the electric audio input signal, and   to decide whether said characteristic signal parameters fulfill a predefined criterion, and, if said predefined criterion is fulfilled,   to store (qualified) brain signal data representative of the low voltage electric signal from the user&#39;s brain picked up during said current time segment, or corresponding (qualified) amplified brain signal(s), in a memory, and   to use the stored (qualified) brain signal data to determine a (qualified) processed brain signal related to said predefined criterion, and   to modify said time and frequency dependent gain depending on said (qualified) processed brain signal.   
               

     This has the advantage of providing a hearing aid wherein at least a part of the fitting process of a hearing aid to a particular user can be automated and/or continuously updated. A further advantage is that the fitting of the hearing aid is not disturbing in any way as it solely (or mainly) uses the surrounding naturally occurring sounds. 
     In an embodiment, the hearing aid comprises at least two electrodes. In an embodiment, the hearing aid comprises a reference electrode. 
     In an embodiment, the signal parameters that describe the natural sound (‘the characteristic signal parameters’) comprise one or more of a) spectro-temporal modulations, b) pitch changes, and c) spatial cues (such as level and/or timing). If the present sound (segment) and the describing signal parameters corresponds to a usable test signal, the natural sound is considered a test signal with those characteristics. 
     In an embodiment, the signal processing unit is configured to decide whether the characteristic signal parameters describing the current time segment of the electric audio input signal fulfill one of a multitude of predefined criteria, and, if a particular predefined criterion is fulfilled, to store the low voltage electric signal from the user&#39;s brain picked up during the current time segment or the corresponding amplified brain signal(s) in the memory together with the characteristic signal parameters describing that current time segment of the electric audio input signal. 
     In an embodiment, the hearing aid comprises a first memory for temporarily storing (or recording) a first time segment of brain signal data, e.g. the low voltage electric signal(s) from the user&#39;s brain picked up by the at least one electrode (or processed, e.g. amplified versions thereof). Thereby it is possible to identify characteristic signal parameters describing a current time segment of the electric audio input signal, to decide whether said characteristic signal parameters fulfill the predefined criterion, and, if the predefined criterion is fulfilled, to store (or relocate) the brain signal data picked up during said current time segment in a (more permanent, e.g. a non-volatile) memory. Preferably, the first memory is adapted to allow that the first time segment is larger than or equal to the current time segment. 
     In an embodiment, a (current) time segment considered by the hearing aid and used as test signal (i.e. for which the recorded brain signal data are stored and used in the determination of parameters defining the user&#39;s hearing ability) is determined by the time range for which a given predefined criterion is fulfilled. In an embodiment, a (current) time segment in the range between 10 ms and 5 s, such as between 10 ms and 1 s, such as between 10 ms and 500 ms, such as between 10 ms and 100 ms, such as larger than 10 s. 
     In an embodiment, the hearing aid comprises a memory for storing a multitude of different sets of signal processing parameters, each set being configured to determine (or to contribute to the determination of) a different time and frequency dependent gain of the signal processing unit. In an embodiment, the different sets of signal processing parameters correspond to different hearing aid programs, e.g. optimized for different acoustic situations (in dependence of audiology parameters defining the user&#39;s hearing ability). At a given point in time, a specific set of signal processing parameters (program) is loaded (selected) to provide a given time and frequency dependent gain in the signal processing unit. In an embodiment, the hearing aid is configured to allow a user to select a currently used set of signal processing parameters. In an embodiment, the hearing aid is configured to automatically select a currently used set of signal processing parameters (e.g. depending on inputs from a number of sensors/detectors, or according to a predefined scheme). 
     In an embodiment, the hearing aid (e.g. the signal processing unit) is configured to classify the stored brain signal data according to a multitude of criteria. Further, the memory is preferably configured to contain a number of different classes of brain signal data. Each class corresponds to specific values (or ranges of values) of the characteristic signal parameters, such values (or ranges of values) fulfilling one of the predefined criteria. The signal processing unit is configured to store a multitude of brain signal data for a given class. Each individual stored brain signal data of a given class correspond to a time segment of the electric audio input signal (or an artificial test signal) that fulfill the criterion corresponding to the class in question. Although belonging to the same class, the different brain signal data of a class may have been recorded for different sets of signal processing parameters (either selected by the user or automatically at the given point in time). Preferably, each class of brain signal data is hence associated with the set of signal processing parameters used in the signal processing unit during the recording of the brain signal data in question. 
     In an embodiment, at least some of the different predefined criteria for ‘using a given time segment of the electric input signal as test signal’ are designed with a view to a given criterion being specifically suitable for a testing a specific parameter defining the user&#39;s hearing ability (e.g. hearing threshold or frequency or time differentiation, etc.). In that way a number of different sets of brain signal data will be created over time, each set of brain signal data being particularly appropriate for determining a specific parameter defining the user&#39;s hearing ability. 
     In an embodiment, at least some of the different predefined criteria for ‘using a given time segment of the electric input signal as test signal’ are designed with a view to a given criterion being characteristic of a ‘type of acoustic environment’. 
     In an embodiment, the hearing aid is configured to statistically average brain signal data associated with a specific class (fulfilling a predefined criterion). 
     Preferably, each brain signal data stored in the memory is associated with 
     a) the characteristic signal parameters describing the time segment of the electric audio input signal, which have resulted in the brain signal data, and 
     b) the class corresponding to the predefined criterion that the characteristic signal parameters have fulfilled 
     In other words, the brain signal data are associated with data representative of the acoustic environment (criterion, signal parameters). 
     In an embodiment, the hearing aid, e.g. the signal processing unit, is configured to, e.g. adaptively (over time, e.g. over days or weeks), determine a set of signal processing parameters from the different sets of brain signal data (e.g. related to a given acoustic environment) that optimize the user&#39;s perception of the electric input signal, e.g. as determined by minimized threshold parameters related to the user&#39;s hearing ability. 
     In an embodiment, the signal processing unit is adapted to estimate the user&#39;s hearing thresholds based on said (qualified) processed brain signals, to provide current estimated hearing thresholds. This has the advantage that the hearing aid can be fully or partially self fitting. In an embodiment, the estimate of the user&#39;s hearing threshold is based on (qualified) processed brain signals from low voltage electric signals picked up by the at least one electrode over a period of time, termed the measurement time. In an embodiment, the measurement time is longer than 8 hours, such as longer than one day, such as longer than one week, such as longer than one month. 
     In an embodiment, the auditory evoked potential is an auditory brainstem response. The auditory brainstem response is an auditory evoked potential extracted from ongoing electrical activity in the brain and recorded via electrodes placed on the scalp. AEP and ABR are e.g. described in Wikipedia [Wiki-AEP] and [Wiki-ABR], respectively. 
     In an embodiment, the auditory evoked potential is an auditory steady state response. Auditory Steady State Response (ASSR) is an auditory evoked potential, elicited with modulated tones that can be used to predict hearing sensitivity in patients of all ages. It is an electrophysiologic response to rapid auditory stimuli and creates a statistically valid estimated audiogram (evoked potential used to predict hearing thresholds for normal hearing individuals and those with hearing loss). The ASSR uses statistical measures to determine if and when a threshold is present and is a “cross-check” for verification purposes prior to arriving at a differential diagnosis [Wiki-ABR] (see e.g. U.S. Pat. No. 7,035,745 or [Stürzebecher et al., 2006]). 
     In the same way, a frequency specific hearing threshold level (HTL) estimate can be provided through ASSR. 
     With such estimate, where the ASSR (or ABR) signals have been presented through the hearing aid output and recorded using the ear electrodes, the frequency specific ASSR response provides an estimate of the hearing sensitivity as a function of frequency. Furthermore, these HTLs can then be used to apply conventional hearing threshold based prescription rules. In an embodiment, the signal processing unit is adapted to run a fitting algorithm, such as NAL-RP, NAL-NL2 (National Acoustic Laboratories, Australia), DSL (National Centre for Audiology, Ontario, Canada), ASA (American Seniors Association), VAC (Veterans Affairs Canada), etc. using said estimated hearing thresholds. The fitting algorithm uses the estimated hearing thresholds to determine the appropriate frequency dependent gain for the user. In an embodiment, the hearing aid is adapted to execute the fitting algorithm in real-time. In an embodiment, the hearing aid is adapted to execute the fitting algorithm automatically. 
     In certain cases where behavioral thresholds cannot be attained, ABR thresholds can be used for hearing aid fittings. New fitting formulas such as DSL v5.0 allow the user to base the settings in the hearing aid on the ABR thresholds. Correction factors do exist for converting ABR thresholds to behavioral thresholds. The Desired Sensation Level multistage input/output algorithm (DSLm[i/o]) is an electroacoustic fitting algorithm particularly aimed at children (National Centre for Audiology, Ontario, Canada). 
     So when the hearing aid records the AEP, e.g. ASSR, response during daily use for some time (e.g. a few days, e.g. at least 3-5 days) (without amplification or just minor amplification) the estimate of the ASSR response is accurate and the hearing aid can use the given prescription rule (NAL, DSL, etc.) to provide an individually prescribed amplification scheme (without having to measure and enter a clinically obtained audiogram). 
     AEP measurements may further be used to measure supra-threshold effects. Supra-threshold AEPs can help to determine whether the signal processing applied in a hearing aid is appropriate to make certain acoustic information not just audible but usable to the user. In other words, online supra-threshold AEP measurements may be used to steer the signal processing, e.g. it could be made more aggressive if the hearing loss worsens (as verified by means of objective hearing threshold measurements, for example), so that the important acoustic information still gets through. 
     One example of such supra-threshold measure is the measure of the ABR activity as is it is (without dedicated periodic acoustic stimuli). However, since it will not be based on repeated sound, it is difficult to use in this application. However, Frequency Following Response (FFR) may advantageously be used. It has been shown that trained musicians have a more distinct and pronounced FFR compared to untrained subjects. Hearing impaired subjects have a poorer FFR. 
     Frequency following response (FFR), also referred to as Frequency Following Potential (FFP), is an evoked response generated by continuous presentation of low-frequency tone stimuli. Unlike the Acoustic Brain Reflex (ABR), the FFR reflects sustained neural activity; integrated over a population of neural elements. It is phase locked to the individual cycles of the stimulus waveform and/or the envelope of the periodic stimuli. 
     The collection and analysis of auditory brainstem responses to complex sounds (cABR) may be used to track the systemic changes due to intervention (e.g. by the use of ordinary hearing aids) (cf. e.g. [Skoe &amp; Kraus, 2010]). According to the present disclosure, the cABR generation and brainstem recording can be made through the electrode equipped hearing aids, where the hearing aid settings are changed in order to maximize the measured FFR response. This means that the compression settings and gain as a function of frequency are altered in the direction of increased FFR response. 
     In general, it has been assumed that the AEP, e.g. ABR, signals are objective in the sense that they are automatically (reflexively) generated by the person&#39;s perceptive system and not influenced by the person&#39;s will. This assumption is the basis of an independent determination of hearing thresholds from such measurements. It has, however, been indicated (c.f. e.g. [Sörqvist et al.; 2012]) that also AEP, e.g. ABR, signals—under certain circumstances—may be sensitive to a person&#39;s will, and thus that such assumingly ‘objective’ measurements may be distorted. According to the present disclosure, however, due the relatively long measurement times (e.g. continuous measurement), such ‘incidents’ of non-reflexive action may be eliminated from influencing the results due to long term averaging of the AEP-signals or specific identification of such ‘distorted’ time segments and elimination from the calculation. 
     In an embodiment, the measuring time during which brain wave signals, e.g. AEP data (e.g. ABR data) are recorded and (possibly continuously) processed is longer (such as much longer) than a normal clinical AEP recording session. In an embodiment, the measuring time is longer than 8 hours, such as longer than one day, such as longer than one week, such as longer than one month. 
     In an embodiment, the measuring time is an accumulated measuring time, in case the measurements have been interrupted (and/or that time segments of the data have been eliminated). 
     In an embodiment, a measurement time comprises a large number of recorded AEP-responses, e.g. more than one hundred or more than one thousand or more than ten thousand responses. In an embodiment, the complex values of the large number of recorded AEP-responses are added in magnitude and phase (and possibly averaged). 
     In general, due to long measurement times, the ambient real world sounds will be cancelled/averaged out of the averaging process since the only repeated response is the evoked potential/event related potential. In an embodiment, however, the hearing aid comprises one or more filters, such as one or more variable filters, adapted to filter the low voltage electric signal(s) (as picked up by the electrode(s)) and/or the amplified brain signal(s) before being further processed to estimate the user&#39;s hearing thresholds. In an embodiment, the hearing aid is adapted to use or NOT use the voltages or data from the electrodes depending on an indication of the user&#39;s current environment, e.g. acoustic environment, and/or cognitive load, or e.g. depending on an input from the user. 
     In an embodiment, the hearing aid comprises a number of hearing aid programs adapted for providing a signal processing of the input audio signal in various specific acoustic environments or situations (e.g. speech in noise, speech in silence, live music, streamed music or sound, telephone conversation, silence, ‘cocktail party’, etc.). In an embodiment, the hearing aid comprises different transfer functions for the variable filter(s) corresponding to the different hearing aid programs, so that a transfer function corresponding to a particular acoustic situation is applied to the variable filter, when the program for that acoustic situation is used in the hearing aid. Alternatively or additionally, the hearing aid may comprise one or more detectors for identifying the acoustic environment. In an embodiment, the hearing aid is adapted to apply a transfer function corresponding to a particular acoustic situation to the variable filter depending on the acoustic situation indicated by said detector(s). 
     In an embodiment, the hearing aid is adapted to provide that the filtering of the low voltage electric signal(s) and/or the amplified brain signal(s) is dependent on an estimate of the current cognitive load of the user. A hearing aid wherein the processing of an audio input is adapted in dependence of an estimate the present cognitive load of the user is e.g. discussed in [Lunner, 2010], which is hereby incorporated by reference. 
     In an embodiment, the hearing aid comprises a user interface adapted for allowing a user to activate or deactivate a specific mode (e.g. termed an AEP- or ABR-mode) where the voltages or data from the electrodes are recorded for further processing to determine an estimate of the user&#39;s hearing thresholds. In an embodiment, the user interface is adapted to allow a user to start an estimation of new hearing thresholds (ignoring previously recorded values). 
     Furthermore, such online estimated ABR response can be used to monitor (also in a relatively long-term perspective, i.e. over days, or months) whether the hearing thresholds deteriorate (i.e. increase) over time (and if so, to possibly inform the user thereof). 
     In an embodiment, the hearing aid comprises a memory for logging values of said estimated hearing thresholds of the user over time. In an embodiment, values of the estimated hearing thresholds are stored with a predefined log frequency, e.g. at least once every hour, such as at least once every day. 
     In an embodiment, the signal processing unit is adapted to determine whether said estimated hearing thresholds or a hearing threshold measure derived therefrom change over time, e.g. by determining corresponding rates of change (e.g. a rate of increase or decrease). 
     ERPs (including AEPs) can be reliably measured using electroencephalography (EEG), a procedure that measures electrical activity of the brain through the skull and scalp. As the EEG reflects thousands of simultaneously ongoing brain processes, the brain response to a single stimulus or event of interest is not usually visible in the EEG recording of a single trial. To see the brain response to the stimulus, the experimenter must conduct many trials (100 or more) and average the results together, causing random brain activity to be averaged out and the relevant ERP to remain. While evoked potentials reflect the processing of the physical stimulus, event-related potentials are caused by the “higher” processes that might involve memory, expectation, attention, or changes in the mental state, among others (cf. [Wiki-ERP]). 
     Such (automatic) real time AEP (e.g. ABR) may be used for temporal fitting, meaning that the hearing aid initially provided to the user may comprise no or little amplification. Over time, when the AEP response grows through averaging, the hearing threshold estimates become more and more valid, and reliable values for such threshold estimates emerge (possibly replacing previous clinically measured hearing thresholds). Thereby automatic hearing threshold based prescription of a hearing aid through ‘online AEP’ may be implemented. 
     In an embodiment, the signal processing unit is adapted to modify the presently used (time and) frequency dependent gains of the hearing aid, based on said estimated hearing thresholds. In an embodiment, such modification of the intended frequency dependent gain values is performed according to a predefined scheme, e.g. with a predefined update frequency, and/or if said currently estimated hearing thresholds deviate with a predefined amount from the presently used hearing thresholds. In an embodiment, a hearing threshold difference measure is defined and used to determine said predefined amount. In an embodiment, the hearing threshold difference measure comprises a sum (ΔHT cur ) of the differences between the currently estimated hearing thresholds (CEHT(f)) and the presently used hearing thresholds (PUHT(f)), where f is frequency. In an embodiment, the hearing thresholds are estimated at a number NHT of predefined frequencies, f 1 , f 2 , . . . , f NHT . In an embodiment, NHT is smaller than or equal to 12, e.g. in the range from 2 to 10. In an embodiment, the predefined frequencies comprise one or more of (such as a majority or all of) 250 Hz, 500 Hz, 1 kHz, 1.5 kHz, 2 kHz, 3 kHz, 4 kHz and 6 kHz. In an embodiment, the gain adaptation is performed with a predefined update frequency in the range from once every 6 months to once every month, or even up to once every day, or more often. In an embodiment, the update frequency is defined in relation to (e.g. determined by) the measurement time. In an embodiment, the measurement time is defined in relation to (e.g. determined by) the update frequency. In an embodiment, the gain adaptation is performed, if the relative hearing threshold difference measure (ΔHT cur /SUM(PUHT(f)) is larger than 10%, such as larger than 25%. In an embodiment, the gain adaptation is performed, if the rate of change (increase) of the hearing threshold difference measure (or the individual estimated hearing thresholds) is above a predefined rate, e.g. if ΔHT(t 2 ,t 1 )]/(t 2 −t 1 ) is larger than a predefined rate, ΔHT(t 2 ,t 1 )=SUM(EHT(f i ,t 2 )−EHT(f i ,t 1 )), where EHT(f i ,t n ) is the estimated hearing threshold at frequency f i  (i=1, 2, . . . , NHT) and time t n  (n=1, 2) and where the summation (SUM) is over frequencies f i . In an embodiment, the gain adaptation is performed at the request of a user via a user interface of the hearing aid (e.g. a remote control). 
     The ABR estimates may be used to monitor (possibly relatively short term, e.g. within hours or days) temporal threshold shifts (TTS) as a consequence of being subject to excessively loud sounds. Also here a warning to the user can be appropriate. 
     In an embodiment, the hearing aid is configured to issue an alarm, when a threshold value of an acoustic dose is exceeded. US 2010/141439 A1 deals with determining an accumulated sound dose and issuing an alarm to a user of a hearing aid. 
     In an embodiment, the hearing aid comprises an alarm indication unit adapted for issuing an alarm signal to the user in case said estimated hearing thresholds deteriorate over time. 
     In an embodiment, the deterioration is identified in that said estimated hearing thresholds (e.g. in dB sound pressure level (SPL)) increase above predetermined relative or absolute levels or that said rates of change of the hearing thresholds are above predefined values. Alternatively, the deterioration is identified in that said hearing threshold difference measure exceeds a predetermined threshold value. 
     In an embodiment, an absolute hearing threshold difference measure (ΔHT abs ) is defined as a sum of the differences between the originally stored (or estimated) hearing thresholds (OSHT(f)) and the currently estimated hearing thresholds (CEHT(f)), where f is frequency (e.g. ΔHT abs =SUM(OSHT(f i )−CEHT(f i )), i=1, 2, . . . , NHT). The term ‘originally stored (or estimated) hearing thresholds’ is taken to mean hearing thresholds that were used when the hearing aid was initially taken into operation by the user (or at a later point in time, where the thresholds have been updated in a normal fitting procedure); such original hearing thresholds e.g. being clinically determined and stored in the hearing aid or estimated and stored by the hearing aid itself (first time estimation′). In an embodiment, the absolute hearing threshold difference measure is used as an indicator of the (long term) hearing threshold deterioration. 
     In an embodiment, the hearing aid is adapted to determine at least an estimate of the real or absolute time elapsed between two time instances where estimates of hearing thresholds are determined and possibly stored. In an embodiment, the hearing aid is adapted to receive a signal representative of the present time from another device, e.g. from a cell phone or from a transmitter of a radio time signal (e.g. DCF77 or MSF). In an embodiment, the hearing aid comprises a real time clock circuit and a battery ensuring a constant functioning of the clock. In an embodiment the hearing aid comprises an uptime clock for measuring an uptime in which the hearing aid is in operation, and/or a power-up counter for counting a number of power-ups of the hearing aid, and the hearing aid is adapted to estimate a real time range elapsed from the uptime and/or the number of power-ups of the hearing aid. 
     In an embodiment, the alarm indication unit is adapted to issue a first alarm signal, if said deterioration rate or if said current hearing threshold difference measure is above a predefined threshold value (indicating that the user may have been exposed to an excessive acoustic dose, possibly over a relatively short period of time, and that the user should take measures to minimize such exposure). 
     In an embodiment, the alarm indication unit is adapted to issue a second alarm signal, if said absolute hearing threshold difference measure exceeds a predefined threshold value (indicating that the user&#39;s hearing ability has deteriorated, possibly over a relatively long period of time, and that the user should act to verify the cause of such deterioration and identify a proper remedy). 
     In an embodiment, the hearing aid comprises at least two separate physical bodies, each comprising a housing. In an embodiment, one part is adapted for being mounted fully or partly in an ear canal of a user (a so-called ITE-part). In an embodiment, one part is adapted for being mounted behind an ear of a user (a so-called BTE-part). In an embodiment, the ITE-part as well as the BTE-part comprises at least one electrode located at a surface of the housing of the part in question to allow the electrode or electrodes to contact the skin of a user&#39;s head when the part is operationally mounted on the user. 
     In an embodiment, time and frequency dependent gain of the signal processing unit is adapted to compensate for a hearing loss of a user. Various aspects of digital hearing aids are described in [Schaub; 2008]. 
     In an embodiment, the output transducer comprises a receiver (speaker) for providing the stimulus as an acoustic signal to the user. 
     In an embodiment, the input unit comprises an input transducer. In an embodiment, the input transducer comprises a microphone. In an embodiment, the input transducer comprises a directional microphone system adapted to separate two or more acoustic sources in the local environment of the user wearing the hearing aid. 
     In an embodiment, the input unit comprises a receiver of a direct electric signal. In an embodiment, the hearing aid, e.g. the input unit, comprises a (possibly standardized) electric interface (e.g. in the form of a connector to implement a wired interface or a wireless interface and/or an antenna and transceiver circuitry to implement a wireless interface) for receiving a direct electric input signal from another device, e.g. a communication device or another hearing aid. In an embodiment, the direct electric input signal represents or comprises an audio signal and/or a control signal and/or an information signal. In an embodiment, the hearing aid comprises demodulation circuitry for demodulating the received direct electric input to provide the direct electric input signal representing an audio signal and/or a control signal e.g. for setting an operational parameter (e.g. volume) and/or a processing parameter of the hearing aid. In general, the wireless link established by a transmitter and antenna and transceiver circuitry of the hearing aid can be of any type. In an embodiment, the wireless link is used under power constraints, e.g. in that the hearing aid comprises a portable (typically battery driven) device. In an embodiment, the wireless link is a link based on near-field communication, e.g. an inductive link based on an inductive coupling between antenna coils of transmitter and receiver parts. In another embodiment, the wireless link is based on far-field, electromagnetic radiation. In an embodiment, the communication via the wireless link is arranged according to a specific modulation scheme, e.g. an analogue modulation scheme, such as FM (frequency modulation) or AM (amplitude modulation) or PM (phase modulation), or a digital modulation scheme, such as ASK (amplitude shift keying), e.g. On-Off keying, FSK (frequency shift keying), PSK (phase shift keying) or QAM (quadrature amplitude modulation). 
     In an embodiment, the electric audio input signal is provided by an input transducer. In an embodiment, the electric audio input signal is provided by wired connector. In an embodiment, the electric audio input signal is provided by wireless receiver. Hence, streamed sound may replace (or supplement) sounds picked up with one or more microphone(s). 
     In an embodiment, the ear part of the hearing aid is a device whose maximum physical dimension (and thus of a possible antenna for providing a wireless interface to the device) is smaller than 10 cm, such as smaller than 5 cm, such as smaller than 2 cm. 
     In an embodiment, the hearing aid comprises a forward or signal path between the input transducer (microphone system and/or direct electric input (e.g. a wireless receiver)) and the output transducer. In an embodiment, the signal processing unit is located (at least partially) in the forward path. In an embodiment, the signal processing unit is adapted to provide a frequency dependent gain according to a user&#39;s particular needs. In an embodiment, the hearing aid comprises an analysis path comprising functional components for analyzing the input signal (e.g. determining a level, a modulation, a type of signal, an acoustic feedback estimate, etc.). The analysis path may further comprise functionality (e.g. implemented in the signal processing unit) that is not directly related to the current signal of the forward path, e.g. the processing of the brain signals picked up by the one or more electrodes. In an embodiment, some or all signal processing of the analysis path and/or the signal path is conducted in the frequency domain. In an embodiment, some or all signal processing of the analysis path and/or the signal path is conducted in the time domain. In an embodiment, some or all signal processing of the forward path is conducted in the time domain, whereas some or all signal processing of the analysis path in the frequency domain. 
     In an embodiment, an analogue electric signal representing an acoustic signal is converted to a digital audio signal in an analogue-to-digital (AD) conversion process, where the analogue signal is sampled with a predefined sampling frequency or rate f s , f s  being e.g. in the range from 8 kHz to 40 kHz (adapted to the particular needs of the application) to provide digital samples x n  (or x[n]) at discrete points in time t n  (or n), each audio sample representing the value of the acoustic signal at t n  by a predefined number N s  of bits, N s  being e.g. in the range from 1 to 16 bits. A digital sample x has a length in time of 1/f s , e.g. 50 μs, for f s =20 kHz. In an embodiment, a number of audio samples are arranged in a time frame. In an embodiment, a time frame comprises 64 audio data samples. Other frame lengths may be used depending on the practical application. 
     In an embodiment, the hearing aids comprise an analogue-to-digital (AD) converter to digitize an analogue input with a predefined sampling rate, e.g. 20 kHz. In an embodiment, the hearing aids comprise a digital-to-analogue (DA) converter to convert a digital signal to an analogue output signal, e.g. for being presented to a user via an output transducer. 
     In an embodiment, the hearing aid, e.g. the input transducer, comprises a TF-conversion unit for providing a time-frequency representation of an input signal. In an embodiment, the time-frequency representation comprises an array or map of corresponding complex or real values of the signal in question in a particular time and frequency range. In an embodiment, the TF conversion unit comprises a filter bank for filtering a (time varying) input signal and providing a number of (time varying) output signals each comprising a distinct frequency range of the input signal. In an embodiment, the TF conversion unit comprises a Fourier transformation unit for converting a time variant input signal to a (time variant) signal in the frequency domain. In an embodiment, the frequency range considered by the hearing aid from a minimum frequency f min  to a maximum frequency f max  comprises a part of the typical human audible frequency range from 20 Hz to 20 kHz, e.g. a part of the range from 20 Hz to 12 kHz. In an embodiment, a signal of the forward and/or analysis path of the hearing aid is split into a number NI of frequency bands, where NI is e.g. larger than 5, such as larger than 10, such as larger than 50, such as larger than 100, such as larger than 500, at least some of which are processed individually. In an embodiment, the hearing aid is/are adapted to process a signal of the forward and/or analysis path in a number NP of different frequency channels (NP≦NI). The frequency channels may be uniform or non-uniform in width (e.g. increasing in width with frequency), overlapping or non-overlapping. 
     In an embodiment, the hearing aid comprises one or more detectors for classifying an acoustic environment around the hearing aid and/or for characterizing the signal of the forward path of the hearing aid. Examples of such detectors are a level detector, a speech detector, a feedback detector (e.g. a tone detector, an autocorrelation detector, etc.), a directionality detector, etc. 
     In an embodiment, the hearing aid comprises an acoustic (and/or mechanical) feedback suppression system. 
     In an embodiment, the hearing aid further comprises other relevant functionality for the application in question, e.g. compression, noise reduction, etc. 
     In an embodiment, the hearing aid comprises a hearing instrument, e.g. a hearing instrument adapted for being located at the ear or fully or partially in the ear canal of a user, e.g. a headset, an earphone, an ear protection device or a combination thereof. 
     Use: 
     In an aspect, use of a hearing aid as described above, in the ‘detailed description of embodiments’ and in the claims, is moreover provided. In an embodiment, use is provided in a system comprising one or more hearing instruments, headsets, ear phones, active ear protection systems, etc. 
     A Method: 
     In an aspect, a method of operating a hearing aid is furthermore provided by the present application, the hearing aid comprising
         an ear part adapted for being mounted fully or partially at an ear or in an ear canal of a user, the ear part comprising
           a housing,   at least one electrode located at a surface of said housing to allow said electrodes to contact the skin of a user when said ear part is operationally mounted on the user, and adapted to pick up a low voltage electric signal from the user&#39;s brain,   
           an amplifier unit operationally connected to said electrode(s) and adapted for amplifying said low voltage electric signal(s) to provide amplified brain signal(s),   an input unit, e.g. an input transducer, for providing an electric audio input signal,   an output transducer for converting an electric output signal to an acoustic output sound to a user,   a signal processing unit, said signal processing unit being operationally connected   to said amplifier unit,
           to said input transducer, and   to said output transducer.   
               

     The method comprises mounting said hearing aid on said user;
         applying a time and frequency dependent gain to said electric audio input signal or a signal originating therefrom and providing a processed audio output signal;   recording said low voltage electric signal from the user&#39;s brain   identifying characteristic signal parameters describing a current time segment of the electric audio input signal, and   deciding whether said characteristic signal parameters fulfill a predefined criterion, and, if said predefined criterion is fulfilled,   storing (qualified) brain signal data representative of the low voltage electric signal from the user&#39;s brain picked up during said current time segment or corresponding (qualified) amplified brain signal(s) in a memory, and   using the stored (qualified) brain signal data to determine a (qualified) processed brain signal related to said predefined criterion, and   modifying said time and frequency dependent gain depending on said (qualified) processed brain signal.       

     It is intended that the structural features of the hearing aid described above, in the ‘detailed description of embodiments’ and in the claims can be combined with the method, when appropriately substituted by a corresponding process and vice versa. Embodiments of the method have the same advantages as the corresponding devices. 
     In an embodiment, the method comprises that the user&#39;s hearing thresholds (and/or other parameters characteristic of a user&#39;s hearing ability) are estimated based on the (qualified) processed brain signals. 
     In an embodiment, the method comprises running a fitting algorithm using said estimated hearing thresholds to determine the appropriate frequency dependent gain for the user. In an embodiment, the method comprises executing the fitting algorithm in real-time. 
     In an embodiment, the method comprises that the (time and) frequency dependent gain is modified (e.g. optimized) based on said estimated hearing thresholds. 
     In an embodiment, the method of measuring auditory evoked potentials is selected among Auditory Brainstem Response (ABR), including Auditory Brainstem Responses to complex sounds (cABR), Auditory Steady State Response (ASSR), and Frequency Following Response (FFR). 
     A Computer Readable Medium: 
     In an aspect, a tangible computer-readable medium storing a computer program comprising program code means for causing a data processing system to perform at least some (such as a majority or all) of the steps of the method described above, in the ‘detailed description of embodiments’ and in the claims, when said computer program is executed on the data processing system is furthermore provided by the present application. In addition to being stored on a tangible medium such as diskettes, CD-ROM-, DVD-, or hard disk media, or any other machine readable medium, the computer program can also be transmitted via a transmission medium such as a wired or wireless link or a network, e.g. the Internet, and loaded into a data processing system for being executed at a location different from that of the tangible medium. 
     A Data Processing System: 
     In an aspect, a data processing system comprising a processor and program code means for causing the processor to perform at least some (such as a majority or all) of the steps of the method described above, in the ‘detailed description of embodiments’ and in the claims is furthermore provided by the present application. 
     A Hearing Aid System: 
     In a further aspect, a hearing aid system comprising a hearing aid as described above, in the ‘detailed description of embodiments’, and in the claims, AND an auxiliary device is moreover provided. 
     In an embodiment, the system is adapted to establish a communication link between the hearing aid and the auxiliary device to provide that information (e.g. control and status signals, possibly audio signals) can be exchanged or forwarded from one to the other. 
     In an embodiment, the auxiliary device comprises an audio gateway device adapted for receiving a multitude of audio signals (e.g. from an entertainment device, e.g. a TV or a music player, a telephone apparatus, e.g. a mobile telephone or a computer, e.g. a PC) and adapted for selecting and/or combining an appropriate one of the received audio signals (or combination of signals) for transmission to the hearing aid. In an embodiment, the auxiliary device comprises a remote control for controlling operation of the hearing aid. In an embodiment, the auxiliary device comprises a communication device, e.g. a cellular telephone, e.g. a SmartPhone. 
     In an embodiment, the auxiliary device is or comprises another hearing aid. In an embodiment, the hearing aid system comprises two hearing aids adapted to implement a binaural hearing aid system. 
     In an embodiment, the hearing aid system comprises another hearing aid as described above, in the ‘detailed description of embodiments’, and in the claims and an auxiliary device, e.g. an audio gateway and/or a remote control for the hearing aids. In an embodiment, the two hearing aids implement or form part of a binaural hearing aid system. 
     In an embodiment, the hearing aid system is adapted to transmit values of the amplified or processed brain signals from at least one of the hearing aids to the other. Thereby the electrodes of both hearing aids may be used together in the estimation of the hearing thresholds. In an embodiment, at least one of the electrodes is a reference electrode. 
     In an embodiment, the hearing aid system is adapted to transmit values of the amplified or processed brain signals from the hearing aids to the auxiliary device. The processing of the low voltage EEG-signals from the electrodes (e.g. including the estimation of hearing thresholds and resulting gains) may be fully or partially performed in the auxiliary device. This has the advantage of removing power consuming operations from the listening devices to the auxiliary device for which the size limitations and thus the power consumption constraints are less strict. 
     When picking up brain-wave signals (potentials) using in-the-ear encephalogram (EEG) electrodes located on the housing of a single ear piece, the distances between the electrodes are necessarily small (e.g. ≦10 mm). An electric potential created by accumulation of charge on a given electrode (provided by neurons in connection with brain activity) at a given location can be determined from a sum of the contributions of each point charge except for a constant. To be able to compare different brain-wave signals, here termed ‘EEG-potentials’ V EEGi , i=1, 2, . . . , N EEG , where N EEG  is the number of EEG-potentials (electrodes) at play, each ‘EEG-potential’ V EEGi  is preferably referred to a common reference potential V REF , e.g. defined by a reference electrode of the system. To achieve a larger potential difference, ΔV EEGi =V EEGi −V REF , between the active EEG-electrodes (EEGe i ) and the reference electrode (REFe, and possibly to reduce or avoid ‘cross-talk’ between them), the reference electrode is preferably located at a distance from the EEG-electrodes, e.g. larger away than the typical distance between the active EEG-electrodes (e.g. larger than 10 mm, such as larger than 50 mm). In an embodiment, the reference electrode is located on a housing of an opposite ear piece with respect to the ear piece where the active EEG electrodes (to be associated with the reference electrode) are located. Thereby, a sufficiently large distance between the active EEG electrodes and the reference electrode is provided. One way or the other, there is, however, a need to communicate one or more electric potentials between two separate parts (e.g. two ear pieces or one ear pieces and another part) to allow the mentioned voltage difference ΔV EEGi  to be determined, so that they can be further processed (and possibly transmitted to another device, e.g. in a conventional way). 
     The ‘brain-wave signals (here also termed ‘EEG-signals’ or ‘EEG-potentials’) at a given electrode reflect the variation (over time) in electric potentials resulting from varying ionic currents associated with brain activity. The amplitude variation with time is typically in the 10s of microvolt range e.g. from 10 μV to 100 μV, when the electrodes are attached to the skin (scalp) of the user&#39;s head. The variation in time of EEG-signals may comprise oscillations at frequencies in the Hz range, e.g. in the range from 1 Hz to 20 Hz, such as between 5 Hz and 10 Hz. 
     Further objects of the application are achieved by the embodiments defined in the dependent claims and in the detailed description of the invention. 
     As used herein, the singular forms “a,” “an,” and “the” are intended to include the plural forms as well (i.e. to have the meaning “at least one”), unless expressly stated otherwise. It will be further understood that the terms “includes,” “comprises,” “including,” and/or “comprising,” when used in this specification, specify the presence of stated features, integers, steps, operations, elements, and/or components, but do not preclude the presence or addition of one or more other features, integers, steps, operations, elements, components, and/or groups thereof. It will also be understood that when an element is referred to as being “connected” or “coupled” to another element, it can be directly connected or coupled to the other element or intervening elements may be present, unless expressly stated otherwise. Furthermore, “connected” or “coupled” as used herein may include wirelessly connected or coupled. As used herein, the term “and/or” includes any and all combinations of one or more of the associated listed items. The steps of any method disclosed herein do not have to be performed in the exact order disclosed, unless expressly stated otherwise. 
    
    
     
       BRIEF DESCRIPTION OF DRAWINGS 
       The disclosure will be explained more fully below in connection with a preferred embodiment and with reference to the drawings in which: 
         FIGS. 1A-1B  show a first embodiment of a hearing aid according to the present disclosure, 
         FIG. 2  shows a second embodiment of a hearing aid according to the present disclosure, 
         FIG. 3  shows an embodiment of a binaural hearing aid system comprising first and second hearing instruments according to the present disclosure, 
         FIGS. 4A-4D  show various elements of embodiments of a binaural hearing aid system according to the present disclosure, 
         FIGS. 5A-5B  show an application scenario comprising an embodiment of a binaural hearing aid system comprising first and second hearing instruments and an auxiliary device according to the present disclosure, 
         FIG. 6  shows a third embodiment of a hearing aid according to the present disclosure, 
         FIG. 7  shows three sound examples /A, B, C) that give rise to different spectro-temporal representations (=FIG. 3.1 in [Elhilali 2004]), 
         FIG. 8  shows a biologically inspired analysis model of sound textures (=FIG. 3.7 in [McWalter; 2012]), and 
         FIG. 9  illustrates recorded brain signal data stored together with characteristic signal parameters describing a current time segment of the electric audio input signal, and the signal processing parameters that were used to process the input signal while the brain signal data were recorded. 
     
    
    
     The figures are schematic and simplified for clarity, and they just show details which are essential to the understanding of the disclosure, while other details are left out. Throughout, the same reference signs are used for identical or corresponding parts. 
     Further scope of applicability of the present disclosure will become apparent from the detailed description given hereinafter. However, it should be understood that the detailed description and specific examples, while indicating preferred embodiments of the disclosure, are given by way of illustration only. Other embodiments may become apparent to those skilled in the art from the following detailed description. 
     DETAILED DESCRIPTION OF EMBODIMENTS 
     Event related potentials. ERPs can be reliably measured using electroencephalography (EEG). The (weak) EEG-potentials are buried deep into the electrical activity of the brain. Actually, the interesting signals have magnitudes below the ‘brain activity noise’. This is where the benefit of systematically evoked potentials comes in. If a pre-defined sound stimulus is sent out, and if we know exactly when the wearer heard the sound, we can expect the weak interesting response signal after some delay, and can thus associate audio signal and brain wave response. If we now experience the same (or a sufficiently similar) signal again, the weak response should be (roughly) the same again. but the brain activity noise was different. If we then add (or average) the two (or more) responses and assume that the weak response is independent of the brain activity noise then the weak response will be added in magnitude and phase while the noise most probably will cancel out. If this procedure is repeated hundreds or thousands or more times the estimate of the weak response will be more and more certain since several thousand or more responses are added in magnitude and phase, while the uncorrelated noise parts will cancel each other in the adding/averaging process since they are uncorrelated samples. The very nice property with the electrode equipped hearing aid doing this procedure is that this procedure/averaging can be sustained for very long time (days, weeks, months) and thereby a much more certain estimate of the response can be obtained compared to a time limited clinical measure (in the clinic you will need a silent room resting on a couch to be able to get a stable ERP in a few minutes due to limited clinical time). 
       FIG. 1  shows a first embodiment of a hearing aid according to the present disclosure.  FIG. 1A  shows a hearing aid (HA) comprising a forward or signal path (FP) from an input transducer (IT) to an output transducer (OT) the forward path being defined there between and comprising a signal processing unit (SPU) for (among other things) applying a (time and) frequency dependent gain to the audio signal picked up by the input transducer (IT, e.g. as in  FIG. 1B  a microphone unit) and providing an enhanced signal (OUT) to the output transducer (OT, e.g. as in  FIG. 1B  a loudspeaker). The hearing aid comprises an EEG unit (DEEG) for picking up and amplifying low voltage electric signals from the user&#39;s brain and providing corresponding (here digital) amplified brain signals (DAEI 1 −DAEI N , where N is the number of electrodes picking up the low voltage electric signal, cf. FIG.  1 B). The signal processing unit is adapted to determine a user&#39;s hearing thresholds (e.g. thresholds related to minimum perception of level, differences in level, frequency, time, etc.) based on the amplified brain signal(s) from the EEG unit (DEEG). The hearing aid, e.g. the signal processing unit (SPU), comprises an analysis unit, e.g. a spectro-temporal analysis unit. The analysis unit is adapted for decomposing the electric input signal (IN) representing the incoming sound into spectro-temporal modulations, and to use the output of this analysis for 1) deciding whether the brain wave signal evoked by the current input signal is suitable for determining hearing threshold(s) of the user, and 2) for identifying the sounds that are sufficiently similar so that the system can treat them as being the same (i.e. that they fulfil the same predefined criterion and/or are representative for a type of acoustic environment). The analysis unit may e.g. comprise a frequency analyser. 
     The hearing aid further comprises a memory (MEM) for storing the brain wave signals recorded by the EEG unit, and the sets of hearing thresholds as determined in the signal processing unit from the brainwave signals at different points in time (t 1 , t 2 , . . . , t n ). The hearing aid may further comprise a stimulus signal generator for generating an electric test signal specifically adapted to be used in an auditory evoked potential (AEP, e.g. a brain stem response (ABR)) measurement (e.g. at times when the electric input signal IN is NOT suitable for determine a user&#39;s hearing thresholds). The optional stimulus signal generator is operationally connected to the output transducer (OT) via the signal processing unit (SPU) allowing the electric test signal to be mixed with the processed audio signal and converted to an auditory test stimulus for being presented to a user together with the acoustic version of the processed audio signal (cf. e.g.  FIG. 2 ), and as described in EP2581038A1. In an embodiment, the EEG unit (DEEG) comprising the electrodes is enclosed in an ear part comprising a housing adapted for being mounted fully or partially at an ear or in an ear canal of a user. The electrodes are located at a surface of the housing to allow the electrodes to contact the skin of the user when the ear part is operationally mounted at or in an ear of the user (and/or partially implanted in the head at an ear of the user, e.g. in connection with a bone-anchored hearing aid). In an embodiment, the ear part (ED) further comprises the output transducer (OT) as indicated by the curved enclosure in the embodiment of  FIG. 1A . In an embodiment, all the mentioned components of the hearing aid (HA) and enclosed in the solid rectangle are enclosed in the same common housing adapted for being mounted fully or partially at an ear or in an ear canal of a user. Other partitions of the components in two or more separate bodies may be implemented depending on the application in question. 
     In the more detailed embodiment of the hearing aid of  FIG. 1A  shown in  FIG. 1B , the main part of the signal processing of the hearing aid (HA) is digital, so the forward path further comprises an analogue-to-digital (AD) converter to digitize an analogue audio input from the microphone with a predefined sampling rate, e.g. 20 kHz, and a digital-to-analogue (DA) converter to convert a digital signal from the signal processing unit (SPU) to an analogue output signal, which is fed to the loudspeaker. The forward path of the hearing aid thus converts an input sound (Sound-in) to an analogue electric input signal (by a microphone unit), which is digitized (unit AD), providing digitized input signal IN, and processed in a signal processing unit (unit SPU), and the processed output signal OUT of the signal processing unit is converted to an analogue signal (unit DA), which is converted (by a speaker unit) to an output sound (Sound-out). The EEG unit (cf. dashed rectangular outline DEEG) comprises N electrodes E 1 , E 2 , . . . , E N , (N≧2) each being adapted to pick up a low voltage electric signal from the user&#39;s brain when located in contact with the user&#39;s skin at different locations of the head of the user (e.g. at or in an ear). In the embodiment of  FIG. 1B , the EEG unit further comprises an amplifier unit (AMP) operationally connected to the N electrodes and adapted for amplifying the low voltage electric signals (Electrode inputs) from the electrodes and to provide amplified brain signals AEI 1 , AEI 2 , . . . , AEI N . The EEG unit further comprises an analogue-to-digital (AD) converter to digitize the analogue inputs from the amplifier and to provide digital amplified brain signals DAEI 1 , DAEI 2 , . . . , DAEI N , which are fed to the signal processing unit (SPU) for further processing. The digital amplified brain signals are used in the signal processing unit (SPU) to determine a user&#39;s hearing thresholds (at different frequencies and) at different points in time (t 1 , t 2 , . . . , t n ). These are stored in memory (MEM), cf. e.g.  FIG. 6 . The memory (MEM) is operationally connected to the signal processing unit (SPU), via signal SHT, to allow storage and retrieval of data in/from the memory, including the mentioned sets of hearing thresholds, controlled by the signal processing unit. The resulting output signal OUT is (via DA-converter DA) fed to the output transducer (here a speaker) to be converted to an auditory signal for being presented to a user together 
       FIG. 2  shows a second embodiment of a hearing aid according to the present disclosure. The hearing aid HA of  FIG. 2  comprises the same functional units as the hearing aid discussed in connection with  FIG. 1B . However, the embodiment of a hearing aid of  FIG. 2  further comprises a test stimulus signal generator (ABR-SG) for generating a test stimulus, and an alarm indication unit (ALU) adapted for issuing a warning or information signal to a user of the hearing aid (or to another person in the user&#39;s environment). The stimulus signal generator (ABR-SG) is configured to generate an electric test signal (ABR) specifically adapted to be used in an auditory evoked potential (AEP, e.g. a brain stem response (ABR)) measurement (e.g. at times when the electric input signal IN is NOT suitable for determining a user&#39;s hearing thresholds). The optional stimulus signal generator is operationally connected to the output transducer (here a loudspeaker) via the signal processing unit (SPU) allowing the electric test stimulus signal—alone or mixed with the processed audio signal—to be presented to a user. The signal processing unit (SPU) is adapted to determine a hearing threshold difference measure comprising a sum (ΔHT cur ) of the differences between currently estimated hearing thresholds (CEHT(f)) and presently used hearing thresholds (PUHT(f)), where f is frequency. In an embodiment, the hearing thresholds are estimated at a number NHT of predefined frequencies, f 1 , f 2 , . . . , f NHT . The signal processing unit (SPU) is e.g. adapted to determine when the hearing threshold difference measure exceeds a predetermined threshold value and (in such case) to generate an alarm signal AL which is fed to the alarm indication unit (ALU). The alarm indication unit is adapted to issue a corresponding alarm (e.g. a visual and/or mechanical and/or acoustic alarm) in response to the alarm signal AL. Alternatively or additionally, the alarm signal may be transmitted to another device (e.g. via a network), e.g. for presentation to a caring person or an audiologist. Alternatively or additionally, the alarm signal may trigger a modification of signal processing parameters (resulting in modified time and frequency dependent gains) of the signal processing unit to reflecting the current hearing thresholds of the user. The analogue to digital conversion unit (AD) of the embodiment of  FIG. 1B  is integrated with the amplifier unit (AMP) in  FIG. 2  (cf. unit AMP-AD in  FIG. 2 ). 
       FIG. 3  shows an embodiment of a binaural hearing aid system comprising first and second hearing instruments according to the present disclosure. The binaural hearing aid system comprises first and second (possibly, as in  FIG. 3 , essentially identical) hearing instruments (HI- 1 , HI- 2 ) adapted for being located at or in left and right ears of a user. The hearing instruments (HI- 1 , HI- 2 ) of  FIG. 3  are substantially similar to the embodiment of a hearing aid (HA) shown in  FIG. 2  (without the test signal generator (ABR-SG) of  FIG. 2 ). The hearing instruments (HI- 1 , HI- 2 ) are additionally adapted for exchanging information between them via a wireless communication link, e.g. a specific inter-aural (IA) wireless link (IA-WL). The two hearing instruments HI- 1 , HI- 2  are adapted to (at least) allow the exchange of status signals, e.g. including the transmission of characteristics of the input signal received by a device at a particular ear to the device at the other ear, and/or (e.g. amplified) EEG-data (e.g. signals DAEI(1:N) or signals derived therefrom) picked up by one or more electrodes (E 1 , E 2 , . . . , E N ) of the contra-lateral hearing instrument, cf. signal IAS. To establish the inter-aural link, each hearing instrument comprises antenna and transceiver circuitry (here indicated by block IA-Rx/Tx). Each of the hearing instruments of the binaural hearing aid system of  FIG. 3  comprises two different input transducers, 1) a microphone unit (MIC) for converting an acoustic input sound to a first electric audio signal INm, and 2) a wireless transceiver (at least a receiver) (ANT and Rx/Tx-unit) for receiving (and possibly transmitting) a signal from (to) another device, e.g. an audio signal, resulting in electric input signal INw. The hearing instruments HI- 1 , HI- 2  are (in this embodiment) assumed to process the audio signal of the forward path in the frequency domain, and therefore each comprise analysis (A-FB) and synthesis (S-FB) filter banks after and before the input (MIC and ANT, Rx/Tx) and output (SP) transducers, respectively. The analysis filter bank (A-FB) is adapted for splitting the (time varying) input signals (INm, INw) into a number NI of (time varying) signals IFB 1 , IFB 2 , . . . , IFB NI , each comprising a distinct (possibly overlapping) frequency range of the input signal. The input transducer (MIC and ANT, Rx/Tx) or the analysis filter bank (A-FB) is assumed to comprise an analogue to digital converter (AD). Correspondingly, the synthesis filter bank (S-FB) is adapted for merging the a number NO of (time varying) signals OFB 1 , OFB 2 , . . . , OFB NO , each comprising a distinct (possibly overlapping) frequency range of the output signal into a (time varying) output signal (OUT), which is fed to the output transducer (SP) for conversion to an output sound for presentation to the user. The output transducer (SP) or the synthesis filter bank (S-FB) may comprise a digital to analogue converter (DA). Each of the hearing instruments HI- 1 , HI- 2  comprises the same functional units as discussed for the hearing aid of 1 and  FIG. 2 , including the EEG-data generating units (electrodes E n  and amplifier and analog to digital converting unit AMP-AD). In the embodiments of hearing instruments HI- 1 , HI- 2  of  FIG. 3 , the amplifier blocks further comprise a time to time-frequency conversion functionality (as indicated by the name of the amplifier unit AMP-AD-T→F) to provide the digital amplified brain signals (DAI 1 , DAI 2 , . . . , DAI N ) in the frequency domain to adapt to further signal processing of the brain signals (determining the frequency dependent hearing thresholds) which may be performed by the signal processing unit (SPU) in the frequency domain. The same may be the case for the signal IAS from the opposite hearing instrument, in which case a T→F conversion unit is included in the IA-Rx/Tx transceiver unit. The number of frequency units provided by the time to time-frequency conversion functionality may or may not be equal to the number NI of frequency bands of the forward path (e.g. smaller than NI). Alternatively, the digital amplified brain signals may be further processed in the signal processing unit (SPU) in the time domain, in which case the time to time-frequency conversion functionality of the amplifier can be omitted. 
     In an embodiment, the hearing aid system further comprises an auxiliary device, e.g. an audio gateway device for receiving a number of audio signals and for transmitting at least one of the received audio signals to the hearing instruments (cf. transceivers (ANT, Rx/Tx in  FIG. 3 ) or a cell phone, e.g. a SmartPhone, as e.g. illustrated in  FIG. 5 . In an embodiment, the listening system is adapted to provide that a telephone input signal (or a dedicated audio signal appropriate for challenging the user&#39;s hearing ability) can be received in the hearing instrument(s) via the auxiliary device. In an embodiment, the hearing aid system comprises a remote control acting as a user interface to the hearing instruments, e.g. to allow a user to change program (e.g. to activate or deactivate the ABR-recording) or otherwise modify operational parameters of the hearing instruments, e.g. output volume of the loudspeaker. In an embodiment, the remote control and the audio gateway are integrated into the same communications device, e.g. a SmartPhone (as e.g. illustrated in  FIG. 5 ). The processing of (possibly amplified) EEG-data (e.g. signals DAEI(1:N) of the hearing instruments, or signals derived therefrom) picked up by the one or more electrodes (E 1 , E 2 , . . . , E N ) may be fully or partially performed in the auxiliary device (e.g. in the audio gateway/remote control/SmartPhone device). 
       FIG. 4  shows various elements of embodiments of a binaural hearing aid system according to the present disclosure.  FIG. 4A  shows an ‘in the ear’ part (ITE) of a hearing aid adapted for being located at or in an ear of a user. In an embodiment, the ITE part constitutes the hearing aid. The ITE part is e.g. adapted for being located fully or partially in the ear canal of the user U (cf.  FIG. 4C, 4D ). The ITE part comprises two electrodes E 1 , E 2  located on (or extending from) the surface of the housing of the ITE part. The ITE part. comprises a mould, e.g. adapted to a particular user&#39;s ear or ear canal. The mould is typically made of a form stable plastic material by an injection moulding process or formed by a rapid prototyping process, e.g. a numerically controlled laser cutting process (see e.g. EP 1 295 509 and references therein). A major issue of an ITE part is that it makes a tight fit to the ear canal. Thus, electrical contacts on the surface (or extending from the surface) of the mould contacting the walls of the ear canal are inherently well suited for forming an electrical contact to the body.  FIG. 4B  shows another embodiment of a (part of a) hearing aid according to the present disclosure.  FIG. 4B  shows a part (BTE) of a ‘behind the ear’ hearing aid, where the BTE part is adapted for being located behind the ear (pinna, EAR in  FIGS. 4C and 4D ) of a user U. The BTE part comprises four electric terminals E 3 , E 4 , E 5 , E 6 , two of which are located on the face of the BTE part, which is adapted for being supported by the ridge where the ear (Pinna) is attached to the skull and two of which are located on the face of the BTE part adapted for being supported by the skull. The electric terminals (electrodes) are specifically adapted for picking up electric (e.g. brain wave) signals from the user&#39;s body, in particular from the brain of the user. The electrical terminals may all serve the same purpose (e.g. measuring EEG) or different purposes, cf. e.g. EP 2 200 347 A2. Electrical terminals (electrodes) for forming good electrical contact to the human body are e.g. described in literature concerning EEG-measurements (cf. e.g. US 2002/028991 or U.S. Pat. No. 6,574,513). 
       FIG. 4C  shows an embodiment of a binaural hearing aid system according to the present disclosure comprising first and second hearing aids comprising (or being constituted by) left and right ear parts ITEl and ITEr, respectively, adapted for being located in left and right ear canals of a user, respectively (each ITE ear part being an ear part as shown in  FIG. 4A ). Alternatively or additionally, the left and right hearing aids may comprise left and right ear parts BTEl and BTEr, respectively, adapted for being located behind left and right ears of a user, respectively (each BTE ear part being an ear part as shown in  FIG. 4B ). The electric terminals (E 1   l , E 2   l  and E 1   r , E 2   r  of the left and right parts, respectively) are adapted to pick up a relatively low voltages (from the body, e.g. the brain) and are operationally connected to an amplifier for amplifying the low voltage signals and to transmit a value representative of the amplified voltage to a signal processing unit of the hearing aid (e.g. located in the ITE-part, in a BTE-part or in an auxiliary device, e.g. unit Aux in  FIG. 4D  or audio gateway/remote control/SmartPhone (Aux) in  FIG. 5 ). Preferably, the hearing aid system comprises a reference terminal. At least one of the left and right hearing aids or hearing aid parts is adapted to allow transmission of signals from the (amplified, EEG) voltages picked up by the electrodes of the hearing aid in question to the other hearing aid (or to an auxiliary device performing the further processing of the voltages from (all) the electrodes) to allow the estimate of hearing thresholds of the user to be based on all available electrodes. The transmission may be wired. Preferably, however, each of the hearing aids (ITEl, ITEr) comprises antenna and transceiver circuitry to establish an interaural wireless link (IA-WL) between the two hearing aids as illustrated in  FIG. 3 . 
       FIG. 4D  shows an embodiment of a binaural hearing aid system according to the present disclosure, which additionally comprises a number of electric terminals or sensors contributing to an estimate of the present cognitive load and/or a classification of the present environment of the user. The embodiment of  FIG. 4D  is identical to that of  FIG. 4C  apart from additionally comprising a (e.g. body-mounted) auxiliary device (Aux), optionally having 2 extra electric terminals, e.g. EEG electrodes, (En) mounted in good electrical contact with body tissue (but NOT on the head). In an embodiment, the auxiliary device (Aux) comprises amplification and processing circuitry to allow a processing of the signals picked up by the electric terminals En. The auxiliary device and at least one of the hearing aids (ITEl, ITEr) each comprise a wireless interface (comprising corresponding transceivers and antennas) for establishing a wireless link (ID-WL) between the devices for use in the exchange of data between the body-mounted auxiliary device (Aux) and the hearing aid(s) (ITEl, ITEr). In an embodiment, the hearing aids (ITEl, ITEr) transmit the amplified voltages picked up by their respective electrodes to the auxiliary device, where the estimate of the hearing thresholds of the (left and/or right ears of the) user is performed. This has the advantage that the (power consuming) ABR-processing can be performed in the (larger) auxiliary device, which typically can be equipped with an energy source of larger capacity than that of a hearing aid (due to the different size constraints). In this case, the interaural link (IA-WL) may be dispensed with for the sake of the calculation of hearing thresholds (and corresponding required frequency dependent gains). In such case, the wireless links ID-WL between the auxiliary device and each of the hearing aids is preferably bidirectional, allowing the auxiliary device to forward revised hearing thresholds or gains to the hearing aids, when the hearing thresholds determined in the auxiliary device have changed more than predefined amounts. The wireless link may be based on near-field (capacitive or inductive coupling) or far-field (radiated fields) electromagnetic fields, and e.g. comply with the Bluetooth standard, e.g. Bluetooth Lowe Energy. The voltages from the electrodes of the auxiliary device may e.g. be used to classify (‘filter’) the voltages from the head mounted electrodes of the hearing aids, e.g. based on the correlation between the signals picked up by the head worn and body worn electrodes, respectively. This may e.g. be used to filter out time segments of the recorded brain wave signals comprising distortions, e.g. ‘incidents’ of non-reflexive (e.g. willful) influence of the user on the brainwave signals. In an embodiment, the voltages picked up by the head worn electrodes (which are used for the estimate of hearing thresholds of the user) are NOT particularly related to hearing, if the correlation with the voltages picked up by the body worn electrodes is large (e.g. above a predefined value, depending on the specific correlation measure used). Such comparison may improve the reliability of the brain signal data used for estimating current hearing thresholds of the user. 
       FIG. 5  shows an application scenario comprising an embodiment of a binaural hearing aid system comprising first and second hearing instruments (HI r , HI l ) and an auxiliary device (Aux) according to the present disclosure. The auxiliary device (Aux) comprises a cellular telephone, e.g. a SmartPhone. In the embodiment of  FIG. 5 , the hearing instruments and the auxiliary device are configured to establish wireless links (WL-RF) between them, e.g. in the form of digital transmission links according to the Bluetooth standard. The links may alternatively be implemented in any other convenient wireless and/or wired manner, and according to any appropriate modulation type or transmission standard, possibly different for different audio sources. The SmartPhone device of  FIG. 5  has the function of a remote control of the hearing instruments, e.g. for changing program or operating parameters (e.g. volume, cf. Vol-button) in the hearing aids, cf. user interface UI. In the context of the present disclosure, the remote control functions of the auxiliary device (Aux) further comprises activation or deactivation of the ABR part of the hearing aid (including disabling the generation of (acoustic) ABR-stimuli and the processing of the voltages picked up by the electrodes of the hearing aids). This can e.g. be defined by one or more special modes that are selectable via mode buttons (Mode 1 , Mode 2 ) of the user interface (UI) on the auxiliary device (or via a touch sensitive display or any other appropriate activation element). Other ‘normal’ modes of operation of the binaural hearing aid system may likewise be selected by the user via the user interface (UI), cf.  FIG. 5B . 
     The hearing instruments (HI- 1 , HI- 2 ) are shown in  FIG. 5A  as devices mounted at the ear (behind the ear) of a user U. Each of the hearing instruments comprise a wireless transceiver to establish an interaural wireless link (IA-WL) between the hearing instruments, here e.g. based on inductive communication. Each of the hearing instruments further comprises a transceiver for establishing a wireless link (CNT r , CNT l ) to the auxiliary device (Aux), at least for receiving signals, e.g. including audio signals from the auxiliary device (e.g. based on radiated fields (RF)). The transceivers are indicated by RF-IA-Rx/Tx-r and RF-IA-Rx/Tx-l in the right and left hearing instruments, respectively. 
     The user interface of the auxiliary device illustrated in  FIG. 5B  is an example of an APP implementing a remote control function and display of information regarding the binaural hearing aid system comprising hearing instruments HI r , HI l . In the illustrated mode of the APP ‘Hearing instrument Remote control’, it is possible to
         Select a mode of operation of the system (program)   Select a test sound (e.g. in case of no other relevant ‘natural sound’)   Present current hearing ability (illustrate a current hearing ability of the user)   Present development in hearing ability (illustrate a development in the hearing ability of the user)       

     The latter option is selected, which provides the option of choosing a time range for the development of the user&#39;s hearing ability. 
     A time range of 1 month has been selected and the result No significant change to hearing thresholds (in the selected time period of 1 month) has been experienced by the system (based on brain signal data extracted according to the present disclosure). 
     Other parts of the functionality of the binaural hearing aid system may be controlled and other information displayed (e.g. graphically) via the user interface (UI) of the auxiliary device (Aux) 
       FIG. 6  shows a third embodiment of a hearing aid (HA) according to the present disclosure. The hearing aid of  FIG. 6  comprises the same functional elements as the embodiment of  FIG. 1B . The memory MEM is shown to have different sets of estimated hearing thresholds HT(f,t) of the user of the hearing aid as determined from the on-board auditory evoked potential (e.g. an auditory brainstem response) system. The on-board auditory brainstem response system primarily uses (selected (‘qualified’) time segments) of the natural sound input represented by electric input signal IN received or picked up by the input transducer IT. The input signals IN are processed by the signal processing unit SPU, specifically gain unit G applying a time and frequency dependent gain to the input signal IN (controlled by signal PUG) and providing a processed output signal PAS. In a normal mode of operation of the hearing aid (HA), the output signal OUT of selector SEL (controlled by control signal SC) is equal to the processed output signal PAS from the gain unit G. The processed output signal are converted to acoustic stimuli via output transducer (OT)) and (at specific selected points in time, controlled by control unit CNT based on analysis of the current input signal IN) the resulting evoked potentials are picked up by the EEG-unit comprising electrodes (E 1 -E N ) and corresponding amplifier and AD-converter (AMP-AD) for providing digital amplified brain signals (DAEI 1 -DAEI N ). 
     The signal processing unit (SPU) calculates (cf. sub-unit V 2 HT) a set of estimates of the users hearing threshold HT(f i ,t n ) at frequencies f i  (i=1, 2, . . . , NHT) at time instance n (t n ) based on the digital amplified brain signals (DAEI 1 -DAEI N ) at time t n , such signals being preferably averaged over a measurement time, e.g. over a number of hours or days). These are stored in the memory (MEM), cf. signal SHT. Preferably, each of the stored brain signals is associated with the signal processing parameters (resulting in gain values applied by gain unit G) used during recording of the brain signal in question. Likewise, a classifier of the acoustic environment, e.g. represented by characteristic parameters of the electric input signal IN during recording of the brain signal in question is preferably stored together with each of the stored the brain signals. The signal processing unit is adapted to store such sets of estimates of the users hearing threshold HT(f i ,t n ) in the memory according to a predefined scheme, e.g. with a predefined frequency. Each set of hearing threshold estimates may correspond to a particular measurement time (or accumulated measurement time). The memory (MEM) is shown to include n+1 sets of hearing thresholds corresponding to times t 0 , t 1 , t 2 , . . . , t n . The first set of hearing thresholds corresponding to time t 0  may e.g. be a set of hearing thresholds that are stored during a fitting procedure, e.g. based on clinical measurements. Otherwise, they may represent the first set of hearing thresholds determined by the hearing aid system (e.g. in case NO fitting has been performed). The hearing thresholds may represent minimum levels at a given frequency allowing a user to perceive a sound, but other minimum threshold values may be extracted from the brain wave signals and used in the fitting algorithm to determine a (currently) appropriate frequency dependent gain. The hearing thresholds are used to calculate an appropriate gain to be used in the gain unit (G) and (possibly in amended form depending on the input signal in question) applied to the input audio signal IN from the input transducer (IT) (or a signal derived therefrom, e.g. a feedback corrected input signal) to provide a processed audio signal PAS, which is fed to selector unit SEL. The frequency dependent gains (G(f i ), i=1, 2, . . . , NI) to be used in the gain unit (G) of the forward path are determined in the sub-unit OL-FIT from hearing thresholds PUHT(f i ) defined by the control unit (CNT) to be presently used at a given time. This involves the use of a fitting algorithm, such as e.g. NAL-NL2. The ‘presently used hearing thresholds’ (PUHT) may e.g. be equal to a particular one of the sets of stored hearing thresholds HT(f i ,t n ), e.g. the one last determined by the sub-unit V 2 HT, or to an average of a number of the stored sets of hearing thresholds, etc. The presently used gains G(f i ) determined by the sub-unit OL-FIT are forwarded to the gain unit G via signal PUG. In a particular ‘artificial test signal mode’ (controlled by control unit CNT or by a user via a user interface, e.g. a remote control, cf. e.g.  FIG. 5B ), a test signal for the ABR-system is generated in the ABR-SG unit, e.g. as a sum of a series of three or more pure tones, each having a specified frequency, amplitude and phase, and wherein a frequency difference between the successive pure tones in the series is constant, f s , cf. e.g. WO 2006/003172 A1. The test signal ABR-S is fed to a psycho acoustic model (here a loudness model), cf. unit LM, together with the processed audio signal PAS, to generate a masked test signal MTS, which is preferably inaudible to the user when combined with the processed audio signal PAS. The masked test signal MTS forming the output of the LM unit and comprising the processed audio signal in combination with the test stimuli is fed to the selector unit SEL. The pure test signal ABR-S is also fed to the selector unit SEL. The resulting output OUT from the selector unit (SEL) (for presentation to a user via output transducer OT) can be either of a) the masked test signal MTS, b) the pure (artificial) test signal ABR-S or c) the processed audio signal PAS. The output signal from the selector unit is controlled by control signal SC from control unit (CNT). The signal processing unit further comprises sub-unit DIFM for calculating a hearing threshold difference measure ΔHT=SUM[HT(f i ,t p )−HT(f i ,t q )], i=1, 2, . . . , NHT, where NHT is the number of frequencies where the hearing threshold estimates are determined, and t p  and t q  are different points in time, for which a set of hearing threshold estimates has been stored. From the stored sets of hearing threshold estimates and the corresponding times, various difference measures can be determined, including indications of rates of change of the hearing thresholds. In an embodiment, the hearing aid (HA) comprises an alarm indication unit (cf. e.g.  FIG. 2, 3 ) adapted for issuing an alarm or warning, when a difference measure (as determined in sub-unit DIFM) is above a predefined value. Such alarm indication unit can e.g. be implemented, if operationally coupled to the control unit (CNT). 
     In an embodiment, the signal processing unit (SPU), is configured to adaptively determine a set of signal processing parameters from the different sets of brain signal data (e.g. related to a given acoustic environment) that optimize the user&#39;s perception of the electric input signal, e.g. as determined by minimized threshold parameters related to the user&#39;s hearing ability. 
       FIG. 7  shows three sound examples that give rise to different representations. The input characteristics triggers the recording if the energy of the sound corresponding to a single point in the time/frequency modulation space relative to the overall sound level is sufficiently high that the AEP is useful. 
       FIG. 7  (=FIG. 3.1 in [Elhilali 2004]) panel A) shows a ripple profile (signal) with a single temporal modulation frequency and single spectral modulation frequency and the single point representation, hence the single element in the 2D modulation frequency domain to the right. Panel B) shows a TORC profile, which includes the same spectral modulation frequency but a range of temporal modulation frequencies giving rise to the horizontal line representation in the 2D modulation domain to the right. Panel C) shows a speech spectrogram and the corresponding 2D modulation frequency domain to the right. This example shows that a speech signal is composed of many different ripple profiles with different intensities. The input characteristics triggers the recording if the energy of the sound corresponding to a single point in the time/frequency modulation space relative to the overall sound level is sufficiently high (&gt; T _{RIPPLE}) that the AEP is useful.  T _{RIPPLE} depends on the signal/contrast of interest and is furthermore related to the normal-hearing JND (Just Noticeable Difference) for that signal/contrast. 
       FIG. 8  (=FIG. 3.7 in [McWalter; 2012]) shows the biologically inspired analysis model of sound textures. The signal in the top is processed through the sequential chain of cochlear filtering, non-linear envelope transformation, and modulation filtering. Each processing output undergoes statistical analysis to compute marginal statistics and (M) and (many) pairwise correlations statistics (C). McWalter&#39;s work suggests which statistics contribute to discrimination of types of sound, and consequently suggests which AEP&#39;s and AEP contrasts which are important to measure the listeners ability to discriminate such sound textures. 
       FIG. 9 . illustrates an application scenario of the present disclosure, where recorded brain signal data (Recorded AEP signals in  FIG. 9 ) are stored together with (associated or linked with) characteristic signal parameters describing a current time segment of the electric audio input signal (Characterisation of input sounds according to analysis model), and with the signal processing parameters (Hearing Aid Settings) that was used to process the input signal, while the brain signal data were recorded.  FIG. 9  shows brain wave data AEP recording  1  and AEP recording  2  arise from two examples of the same time segment of the electric input signal (Sound  1 ) and is furthermore processed with the same settings (Settings  1 ). Thus Averaging across AEP recording  1  and AEP recording  2  reduces noise in AEP response to Sound  1  with Settings  1 . If Sound  1  already contains a relevant threshold or contrast (like a pitch shift) this averaging is sufficient to determine the threshold or just noticeable difference (JND). In practice the averaging requires 10, 100, or more recordings. 
     However, if the targeted contrast is not embedded in Sound  1  but occurs as the differences to Sound  2 , then the procedure described just above is repeated for AEP recordings corresponding to Sound  1  and AEP recordings corresponding to Sound  2 , and the just noticeable difference is determined from comparing AEP recording  1  vs AEP recording  3 . 
     Similarly, if the impact of the processing on a given sound is determined from comparing AEP recording  1  with AEP recording  4 , as they correspond to the same sound with different processing (e.g., comparing the hearing threshold with and without amplification). 
     Likewise, four AEP recordings are required to rank the listener&#39;s ability to discriminate between Sound  1  and Sound  2  with Settings  1  and Settings  2 . 
     The invention is defined by the features of the independent claim(s). Preferred embodiments are defined in the dependent claims. Any reference numerals in the claims are intended to be non-limiting for their scope. 
     Some preferred embodiments have been shown in the foregoing, but it should be stressed that the invention is not limited to these, but may be embodied in other ways within the subject-matter defined in the following claims. 
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