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Avdelningen för logopedi, foniatri och audiologi
Institutionen för kliniska vetenskaper, Lund
Simulated real-ear measurements
of benefit from digital feedback suppression
Steen Østergaard Olsen
Audionomutbildningen, 2006
Vetenskapligt arbete, 20 poäng
Handledare: Arne Leijon och Ingrid Lennart
Olsen, Measurement of DFS benefit
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Abstract
Digital feedback suppression (DFS) enables users of hearing instrument to benefit from
amplification levels that normally would provoke whistling or poor hearing instrument sound
quality. A standardized test for the measurement of DFS benefit is not available. In this study
an objective method for assessment of extra feedback-free amplification (headroom) provided
by a DFS system is described and evaluated. Simulated real-ear measurements on KEMAR
with the substitution and the modified pressure methods provided data from which the
whistle free loop gain could be derived. The standard deviation of the difference between
repeated measurements on a test hearing instrument was about 2 dB. It is concluded that ways
of bringing the test-retest variance down should be investigated.
Keywords: hearing instrument, acoustical feedback, digital feedback suppression, headroom,
simulated real-ear measurements,
Abbreviations:
AD
AGC-O
BTE
CIC
dB
DFS
DSP
HI
HIT
HL
Hz
ITE
KEMAR
MFG
MSG
NR
adaptive directionality
automatic gain controlled by the output level
behind the ear
completely in the canal
Decibel
digital feedback suppression
digital signal processing
hearing instrument
hearing instrument test
hearing level
Hertz
in the ear
Knowles electronics manikin for acoustic research
maximum feasible gain
maximum stable gain
noise reduction
Olsen, Measurement of DFS benefit
REAR
REIG
REM
SD
SPL
sREM
WDGR
WDRC
real-ear aided response
real-ear insertion gain
real-ear measurement
standard deviation
sound pressure level
simulated real-ear measurement
whistle detection with gain reduction
wide dynamic range compression
iii
Olsen, Measurement of DFS benefit
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Contents
1.
Background ........................................................................................................................1
1.1.
Aim of study ..............................................................................................................1
1.1.1.
Research questions.............................................................................................2
2. Hearing instruments ...........................................................................................................2
2.1. Amplification for the hearing impaired individual ....................................................2
2.2.
Analogue hearing instruments ...................................................................................3
2.2.1.
Linear hearing instruments ................................................................................3
2.2.2.
Nonlinear hearing instruments...........................................................................5
2.3.
Digital hearing instruments........................................................................................5
2.4. Occlusion and hearing instrument earmolds and shells.............................................6
2.4.1.
Effect of occluding earmolds .............................................................................6
2.4.2.
How to avoid occlusion effects..........................................................................7
2.5.
Acoustic feedback......................................................................................................7
2.5.1.
The feedback path ..............................................................................................7
2.5.2.
Loop gain ...........................................................................................................8
2.5.3.
Artifacts..............................................................................................................9
2.5.4.
Algorithm interaction with feedback .................................................................9
2.6.
Feedback management.............................................................................................10
2.6.1.
Acoustic modifications and gain reduction......................................................10
2.6.2.
Digital Feedback Suppression..........................................................................12
2.7.
Evaluation of DFS....................................................................................................18
2.7.1.
Currently used method.....................................................................................18
2.8.
Real-ear Measurements............................................................................................19
2.8.1.
The substitution method...................................................................................20
2.8.2.
The modified pressure method.........................................................................20
2.8.3.
REM of DFS benefit ........................................................................................21
2.8.4.
Simulated real-ear measurements ....................................................................23
3. Material and methods.......................................................................................................24
3.1.
Equipment and test environment .............................................................................24
3.1.1.
The test room ...................................................................................................25
3.2.
Calibration................................................................................................................26
3.3.
Test hearing instrument and fitting software ...........................................................26
3.4.
General procedures ..................................................................................................27
3.5.
Experiments .............................................................................................................28
3.5.1.
Reduction of loudspeaker sound pressure........................................................28
3.5.2.
Definition of maximum gain with acceptable feedback ..................................31
3.5.3.
sREM test-retest variance. ...............................................................................36
3.5.4.
Test-retest variance with the currently used method .......................................39
4. Discussion ........................................................................................................................40
5. Conclusion .......................................................................................................................42
6. References........................................................................................................................43
7. Acknowledgements..........................................................................................................48
8. Appendix..........................................................................................................................50
8.1.
Calibration of the Aurical System............................................................................50
8.2.
ReSoundAIR technical data sheet............................................................................52
8.3. Method for selection of “maximum feasible gain” (Excel sheet)............................54
Olsen, Measurement of DFS benefit
1.
1
Background
Digital signal processing (DSP) used in modern hearing instruments allows algorithms to
manipulate the input signal in complex ways. Examples of such algorithms are digital
feedback suppression (DFS), noise reduction (NR), adaptive directionality (AD) and data
logging. It was impossible to implement some of these algorithms in analogue hearing
instruments without increasing the physical size of the device. Wide dynamic range
compression (WDRC) - another advanced hearing aid signal processing technique - has
become a common method for compensation of the reduced dynamic range of the impaired
cochlea.
The audiologist often desires to compare the performance of new features across brands,
models, and styles for the individuals with hearing losses. As signal-processing advances
quickly, standardized tests are often not yet developed. DFS is a technique that enables the
hearing-impaired individual to benefit from amplification levels that without DFS would
provoke whistling or poor hearing instrument sound quality. No standardized test for
measurement of DFS benefit is currently available.
1.1. Aim of study
The aim of this study was to describe an objective method for measuring the amount of extra
feedback-free amplification (headroom) provided by DFS systems and to evaluate the
validity and reliability of the obtained measurements.
Olsen, Measurement of DFS benefit
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1.1.1. Research questions
x
Is it possible to define a gain level at which the acoustic feedback does not
interfere with repeatable measurements?
x
Is it possible to develop an objective method for evaluation of DFS benefit in a
reliable and repeatable way?
x
How well does the proposed new method perform compared to the previouslyused method?
2.
Hearing instruments
2.1. Amplification for the hearing impaired individual
Hearing impairment is the loss of sound quality and sensitivity to sounds. Amplification is
used to compensate for the loss of sensitivity by increasing the amplitude of sounds.
According to nonlinear models of the human cochlea (Yates, 1995), inner hair cells in the
cochlea are sensitive only to louder inputs. The research has shown that cochlear outer hair
cells mechanically amplify soft inputs so they can be detected by the inner hair cells. Inner
hair cell damage leads to loss of loudness of both loud and soft sounds. Damage to outer hair
cells will cause lost audibility of soft and intermediate sounds, but the loudness of louder
sounds will not – or only slightly – be affected. A hearing impairment caused by outer hair
cell damage can therefore be regarded as a reduced dynamic hearing range into which as
much as possible of speech and other signals should be amplified in a way that makes it
audible without being uncomfortably loud.
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When using a hearing instrument, considerations must be given to the effect on the ear’s
natural resonance. The resonance frequency can be calculated from the canal properties to be
approximately 3 kHz (the frequency that has a wavelength four times the length of the ear
canal).
When a hearing instrument is connected to the ear canal by an earmold, the resonance peak of
the occluded ear canal is shifted upwards (since the remaining occluded ear canal is a tube
closed at both ends). The resonance peak is at the frequency that has a wavelength that is the
double of the remaining ear canal. In the aided ear usually also a resonance at a lower
frequency, around 1500 Hz, is seen. This resonance is caused by the combined acoustic
effects of earmold tubing and the remaining ear canal volume.
To compensate for the change of the ear canal resonance when blocking the ear canal with an
earmold, up to 20 dB extra gain is applied in the 3-kHz-region, when fitting a hearing
instrument (Dillon, 2001).
2.2. Analogue hearing instruments
2.2.1. Linear hearing instruments
Electro-acoustic amplification has been used to compensate for hearing losses since the first
commercially available electronic hearing instrument was introduced in the USA around the
turn of the 20th century (Dillon, 2001). The first devices obtained amplification by using a
carbon microphone with a large membrane and the collected energy was delivered to a
smaller earphone attached to the hearing impaired subject’s ear. The amount of amplification
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available depended on the relationship between the sizes of the microphone membrane and
the earphone membrane. Only relatively small gains could be achieved using hearing
instruments. With the inventions of vacuum tubes and transistors, increasing levels of
amplification were achieved.
The original amplification technique was described as linear since the system over a certain
range of signal input levels was linear. A system is linear if it is homogeneous and additive
(Rosen & Powell, 1997). For a hearing instrument that would imply the following: Increasing
the hearing instrument input by x dB makes the output increase by x dB independently of the
value of x and of the initial input. Subsequently the gain will be the same for all inputs. This
relationship holds true when the input plus the hearing instrument gain is less than the
maximum output level of the device. When the output reaches the maximum output level, the
system will start clipping the peaks of the output signal resulting in a harmonic distortion.
The gain will decrease with increasing output for such levels of input. To avoid distortion as
a result of peak clipping, automatic gain control steered by the output level (AGC-O) is used
to decrease the gain for high outputs in many hearing instruments.
When in-the-ear type earphones were developed for body-worn hearing instruments,
earmolds came in use to seal the ear canal in order to obtain a sufficient sound pressure level
without creating feedback whistling (see chapter 2.5).
Over the years, hearing instruments decreased in size as components became smaller. The
original large non-portable hearing instruments were replaced by body-worn instruments and
again replaced by behind-the-ear (BTE), in-the-ear (ITE) and completely-in-the-canal (CIC)
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hearing instruments. In the first BTEs the microphone inlet was placed at the bottom of the
hearing instrument, later the microphone inlet was placed at the top very close to the hearing
instrument loudspeaker ( in hearing instruments often called the receiver). Each step in this
development meant more challenges regarding acoustic feedback problems because the
distance between the microphone and receiver (hearing instrument loudspeaker) was reduced.
2.2.2. Nonlinear hearing instruments
Nonlinear models describing the human cochlea (Yates, 1995) led to new philosophies for
hearing instrument fittings. Nonlinear hearing instruments were introduced in order to
compensate for the loss of nonlinearity in the cochlea by amplifying soft inputs more than
loud inputs. This would increase the possibility of audibility of low intensity speech and thus,
maximize speech intelligibility. The reduction of gain for moderate and high intensity inputs
would increase user comfort. Early nonlinear instruments (e.g. Killion, 1993) had a
compressor that worked in one broad frequency channel; later developments included hearing
instruments with several compression channels.
2.3. Digital hearing instruments
In 1987, the University of Wisconsin and Nicolet Instrument Cooperation produced the first
hearing instrument with digital sound processing. The instrument was the Nicolet Phoenix
(Heide, 1994), a BTE hearing instrument combined with a body-worn device to house the
digital signal processor (DSP). The instrument was not commercially successful, but around
1987-1988 several hearing instrument manufacturing companies (Diaphon, 3M, Bernafone,
and Widex) introduced digitally programmable hearing instruments, however still with
Olsen, Measurement of DFS benefit
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analogue signal processing (Herbst, 1989; Mangold et al, 1990; Stypulkowski, 1994;
Westermann, 1994). In 1996, Widex and Oticon were the first manufacturers to introduce
digital-processing, ear-level hearing instruments and since then, digital hearing instruments
have taken over the market at almost all price points. With digital technology new features
such as digital feedback suppression (DFS), noise reduction (NR), and adaptive directionality
(AD) were introduced in hearing instruments. Data logging was available already in some of
the first digitally programmable analogue instruments (Ringdahl, 1988) and is now common
in digital instruments.
Most digital hearing instruments use nonlinear amplification schemes. Often digital hearing
instruments have lower compression knee points just below 50 dB SPL and relative fast
attack and release times (e.g. attack time: 5 ms and release time: 70 ms).
2.4. Occlusion and hearing instrument earmolds and shells
2.4.1. Effect of occluding earmolds
Hearing instruments are attached to the ear canal by a physical coupling that depends on the
hearing instrument style and on the hearing loss. For BTE hearing instruments, earmolds
made from ear impressions are used to allow for a sufficient sound pressure in the ear canal
without the risk of feedback whistling (see chapter 2.5). The achievable sound pressure at low
frequencies and the risk of feedback whistling is also dependent on the venting in the
earmold. If a tightly sealing earmold is used for a hearing loss of less than around 40 dB HL
at low frequencies, the hearing instrument user will often complain about changes in the
perception of his own voice (Dillon, 2001). Low-frequency energy from the individuals’ oral
Olsen, Measurement of DFS benefit
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cavities resonates via bone conduction and thus, their own voice sounds boomy. Chewing
(especially crunching sounds) can be very loud in the occluded ear. This phenomenon is
called the occlusion effect (Dillon, 2001), and the related annoyance will typically be
increased when hearing instruments are used bilaterally. In addition to occlusion, an earmold
can also cause earwax build up and create a humid environment in the ear canal that is
annoying for the hearing instrument user.
2.4.2. How to avoid occlusion effects
Earmolds are often made with a ventilation bore that will alleviate occlusion effects. If the
ventilation diameter is sufficiently large the occlusion effect may be reduced to an acceptable
level. Great success has been achieved using “open fittings” where almost no earmold is
used (Courtois, 1988; Kiessling et al., 2003). Open fittings do not allow for amplification at
low frequencies due to leakage of sound. Without algorithm intervention the achievable
feedback free gain at high frequencies is limited because of too little attenuation in the
feedback path (See chapter 2.5).
2.5. Acoustic feedback
2.5.1. The feedback path
Feedback in a hearing instrument is the part of the hearing instrument output signal that
returns to the hearing instrument's input. If the hearing instrument mold fits very tightly into
the ear canal and has no other leakages only a small part of the output signal leaks out of the
ear canal and reaches the hearing instrument microphone. If earmold ventilations or other
leakages allow a larger portion of the feedback signal to reach the hearing instrument
Olsen, Measurement of DFS benefit
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microphone, the system might become unstable and behave like an oscillator (Kates, 1999).
The path from the receiver to the microphone is referred to as the feedback path. It consists of
the ear canal, the ear mold with ventilations and tubing, couplings between tubing and ear
mold (in a BTE), emission from the hearing instrument shell (in an ITE), and vibration and
acoustic transmission within the hearing instrument (Flack et al., 1995; Hellgren et al, 1999).
The feedback path is not static, but dynamic and is influenced by changes in ear canal shape
(i.e., when speaking or chewing)(Oliveira, 1997). Also, earwax in the ear canal can prevent
the earmold from being placed tightly in the ear canal (Dillon, 2001).
2.5.2. Loop gain
The loop gain is the total gain of sound traveling through the different parts of the hearing
instrument and back through the feedback path (Dillon, 2001). Feedback whistling will occur
only if the loop gain is greater than 0 dB. Loop gains above 0 dB occur because of an
acoustic leakage around the ear mold or shell or because of the use of a ventilation bore.
When increasing the ventilation through increased vent diameters, for example to decrease
the occlusion effect, the risk of acoustic feedback or whistling increases. Another requirement
for feedback whistling to occur is that the feedback signal adds in phase with the signal
already in the hearing instrument output (Dillon, 2001). If the feedback signal is in phase it is
called positive feedback. When the feedback signal is in the opposite phase to the already
existing signal it is called negative feedback. Negative feedback will not cause feedback
whistling since it cancels the incoming sound. The gain compensation for the change of ear
canal resonance normally used when fitting hearing instruments (see chapter 2.1) increases
the risk of having feedback whistling in the 3-kHz frequency region.
Olsen, Measurement of DFS benefit
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2.5.3. Artifacts
Before the hearing instrument actually starts to emit feedback whistling, feedback might
degrade the sound quality. Positive and negative feedback causes peaks and valleys in the
frequency response to build, and as a result the sound might get a ringing quality (Cox, 1982;
Dillon, 2001).
2.5.4. Algorithm interaction with feedback
Modern hearing instruments may often have one or two systems for feedback management
(see chapter 2.6), but often several algorithms are running simultaneously in the instrument
aiming at improving speech intelligibility, sound quality, and user comfort. Some of these
algorithms might also have an effect on the risk of having feedback whistling, and therefore
care should be taken to sort out any influence from other algorithms running in the hearing
instrument when evaluating the benefit from DFS algorithms. Examples algorithms are listed
below and the possible effect on feedback occurrence is described in each case.
2.5.4.1. Compression system
The hearing instruments compression system aims at maximizing speech audibility, speech
intelligibility and sound comfort. If the environment is very quiet, gain increases and as a
result feedback whistling might occur if the gain is increased to the level of instability. The
rather loud whistling makes the gain go down again. This sequence of on/off whistling can go
on as long as the hearing instrument user is in a quiet environment.
Olsen, Measurement of DFS benefit
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2.5.4.2. Noise reduction
Noise reduction algorithms are designed to reduce the gain in a specific frequency band if
noise is detected in that region. Detection is based on the degree of the modulation in the
signal since speech is a highly modulated signal while steady state noise has little
modulation. Feedback whistling is not modulated, and noise reduction algorithms might
therefore decrease gain in frequency bands containing feedback whistling.
2.5.4.3. Adaptive directionality
Some hearing instruments with dual microphone directional systems have algorithms that can
cancel out point-source noise by adapting the system’s null to the direction of the noise
source (adaptive directionality, AD). This algorithm might reduce feedback whistling, since
the feedback signal reaches both microphones approximately in phase and is therefore
suppressed just like a noise source from the side (90 degree azimuth).
2.6. Feedback management
2.6.1. Acoustic modifications and gain reduction
One way of reducing the risk of feedback whistle is to increase the attenuation in the
feedback path by tightening the ear mold. Reducing the ventilation size will increase
attenuation and the same can sometimes be achieved by making a slightly bigger earmold.
This might on the other hand lead to physical discomfort and increased complaints of
occlusion effects.
Increasing the distance between the hearing instrument microphone and receiver can increase
attenuation in the feedback path. Typically this will be achieved by using a BTE instead of an
Olsen, Measurement of DFS benefit
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ITE or a body-worn instrument instead of a BTE. Going from one hearing instrument type to
a bigger device will often meet resistance from the hearing instrument user due to the wish to
have a hearing instrument that is as invisible as possible. Bigger hearing instruments also
brings along other problems for instance with the physical fit behind the ear or with cables
creating noise in the body-worn instrument when touching the cloth.
Users often experience an increased occurrence of feedback whistles after some time of use.
This might for instance be because of changes of the ear canal size, because the earwax
pushes the earmold away from the ear canal wall, because the tubing is getting to loose a fit
on the earhook, because of splits in tube or hook, or because of increased hearing loss
resulting in increased volume setting (Dillon, 2001).
If the hearing instrument has a volume control the user will often solve the problem by
turning it down to just below the position that produce feedback whistle. At some point this
might result in so little gain that the user experiences poorer speech intelligibility.
If the audiologist fails to solve the problems with feedback whistling by increasing the
attenuation in the feedback path other measures must be taken. Often the gain can be reduced
separately in certain frequency bands, which can be used to reduce the risk of feedback
whistling without reducing gain too much in areas that are important for speech intelligibility.
In some fitting software programs notch filters are available. Such filters are designed to
reduce gain specifically at the frequencies where feedback whistle occurs.
Olsen, Measurement of DFS benefit
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Some hearing instruments even have systems that automatically reduces gain during use if
the hearing instrument starts to whistle. A risk introduced by using automatic whistle
detection and subsequent gain reduction (WDGR) in affected areas is that a whistling sound
not coming from acoustic feedback might trigger the gain reduction algorithm without being
needed.
All procedures of reducing gain to avoid feedback whistle may lead to decreased audibility of
important speech cues in the affected frequency regions. Therefore ways of reducing
feedback whistle without reducing gain are highly desirable.
2.6.2. Digital Feedback Suppression
Digital Feedback Suppression (DFS) is a technique that cancels feedback in the signal
without reducing hearing instrument gain and without compromising wearer comfort by
using tightly fitting earmolds. To achieve this, the precise transfer function of the feedback
path has to be determined through a measurement on the individual ear. A broadband signal
(Figure 1 a) is generated in the hearing instrument and delivered by the hearing instrument
receiver, through the tubing and earmold system to the ear canal. Some of the signal passes
out of the ear canal through leaks in and beside the earmold and reaches the hearing
instrument microphone (Figure 1 b). The signal returning to the microphone is analyzed to
determine the amplitude and phase characteristics across frequencies (Figure 1c) and a static
filter modeling these characteristics is calculated (Figure 1 d).
90
80
80
70
60
50
40
30
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a
Amplitude (dB SPL)
Amplitude (dB SPL)
90
0
100
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10000
b
Frequency (Hz)
90
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DFS filter
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Amplitude (dB SPL)
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Olsen, Measurement of DFS benefit
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1000
10000
Frequency (Hz)
Figure 1. A broadband signal is generated by the HI and delivered in the ear canal (a) .The signal is attenuated by the
feedback path and returns to the HI microphone (b). A static filter is fitted to model the feedback path attenuation (c), and
the filter is stored in the hearing instrument (d).
The function of digital suppression in the hearing instrument is depicted in Figure 2.
Numbers in the text refer to numbers on the figure.
Olsen, Measurement of DFS benefit
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DFS filter
7
6
Inversed, filtered signal
8
1
5
Feedback signal
+
DFS module
+
9
2
Hearing Instrument
3
Amplifier
Microphone
Feedback
path
Receiver
Filtering by
feedback path
4
Amplified signal
Figure 2. Digital feedback suppression in a hearing instrument. Please refer to text for details.
The incoming signal (1) is detected by the hearing instrument microphone (2). It is processed
and the resulting signal (3) is delivered in the ear canal by the hearing instrument receiver.
The amplified signal is attenuated by the tubing and earmold system (the feedback path (4))
on its way back to the hearing instrument microphone (2), where the feedback signal (5) and
the original signal adds up and might provoke feedback whistling in the hearing instrument.
To prevent feedback whistling the DFS algorithm is used in the following way: Before the
amplified signal reaches the receiver (3), it is lead through the DFS filter (6). After passing
through the DFS filter modeled on the feedback path, the filtered signal (7) is ideally very
Olsen, Measurement of DFS benefit
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similar to the feedback signal (5). When the inversed signal (8) is added to the microphone
signal (9) it therefore cancels the feedback part of the signal and whistling is avoided. As the
feedback path varies while the hearing instrument is used, the system must also include an
adaptive component, as described later in chapter 2.6.2.3.
2.6.2.1. Limitations to Digital Feedback Suppression
Under ideal conditions (i.e. when the feedback path model is perfect) the desired part of the
signal is not affected by the use of DFS (Kates, 1999). Using DFS in a hearing instrument
does not mean that it will always be possible to cancel feedback and thereby have a “whistlefree” hearing instrument. Hearing instruments have only limited space available for the
digital signal processor (DSP) circuitry and for the battery. DSP’s in hearing instruments are
therefore small, low-speed devices (Kates, 2003).
Inaccuracies in modeling the feedback transfer function and estimating the feedback is a
reason for limited DFS effectiveness. The inaccuracies are seen as a result of the limited
number of available bits for calculations in hearing instruments. The DFS filter is ‘semioptimally’ fitted for each hearing instrument user to make the best use of the available bits
(Perman, 2006, personal communication).
Narrowband signals can be cancelled out by adaptive feedback cancellation systems and
coloration artifacts can occur with such inputs because of a tendency for a mismatch between
the adaptive feedback path model and the actual feedback path for this type of signals (Kates,
2003). The result is instability and artifacts in the output signal.
Olsen, Measurement of DFS benefit
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Reverberation limits the effectiveness of feedback cancellation in hearing instruments (Kates,
2003). Reverberation typically lasts for a much longer time than the impulse responses of the
electro-mechanical components in the hearing instrument or the direct acoustic path from
receiver to the microphone. The short adaptive digital filter used to model the acoustic
environment in the ear canal and the surroundings of the head will therefore not be long
enough to model the room reverberation. Variations in reverberation due to change of
position in the room will change the feedback path, but the adaptive DFS filter will not be
able to model the changes.
The feedback path transfer function varies over time, and that means that the modeled
feedback path (in the adaptive DFS filter) might be inaccurate for the next block of samples if
the feedback path has changed. A fast updated model is more accurate than a slowly updated
model (Perman, 2006, personal communication) and the effectiveness of a slow DSP is
therefore less than in a fast acting DSP.
Non-linear distortion in the feedback path including the receiver is another limiting factor for
the DFS efficiency. If the signal gets distorted due to saturation of the receiver, the signal
filtered by the feedback path and the signal filtered by the DFS model are no longer similar,
and the feedback will therefore not be cancelled out (Perman, 2006, personal
communication).
2.6.2.2. Runtime DFS calibration
Bisgaard (1993) reported the first implementation of a DFS in a commercially available
hearing instrument (Danavox 145 DFS Genius). The hearing instrument was an analogue
Olsen, Measurement of DFS benefit
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BTE hearing instrument with a linear amplification scheme intended for use by profoundly
hearing-impaired subjects. In this hearing instrument a white noise test signal was used to
analyze the feedback path of the system. The analysis was going on continuously in order to
adjust the DFS system cancellation signal when the environment changed. This is referred to
as runtime calibration. The amount of extra whistle-free gain (headroom) provided by this
DFS system was in the order of 5-8 dB (Bisgaard, 1993).
One basic problem of the first DFS system was that the test signal was regularly presented to
update the feedback path model and this noise would be audible to many hearing instrument
users. This DFS system could therefore only be used by individuals with severe-to-profound
hearing losses.
2.6.2.3. DFS with initial calibration
Hearing instruments with DFS that is calibrated before use became available with the
introduction of fully digital hearing instruments. The feedback path is analyzed using a
broadband signal and the outcome of the analysis is stored in the hearing instrument.
Additionally an adaptive cancellation system corrects for any change of the feedback path
during use of the hearing instrument. This system analyzes changes in the environment using
the microphone signal and the receiver signal. The analysis results in continuous updates of
the feedback path model.
Early DFS versions in digital hearing instruments with initial calibration allowed an average
of 8 dB extra headroom (Pihl-Frank et al., 2000). Over the years DFS algorithms have been
Olsen, Measurement of DFS benefit
18
optimized to provide an average of up to 16 dB headroom (Merks et al., 2006) depending on
the hearing instrument and the DFS algorith.
In early implementations of DFS, the algorithm was mainly used to avoid degradation of the
sound quality due to positive and negative feedback (see chapter 2.5.3). It was not normally
recommended to use the DFS to extend the gain range that could be applied in the hearing
instrument without audible whistling. With the most recent developments it is possible to
keep devices whistle-free at gain levels that would be impossible without using the DFS
algorithm. This has made almost non-occluding fittings possible (Kiessling et al., 2003).
2.6.2.4. Implementations across brands
Until recently DFS algorithms were only implemented in a few brands of hearing instruments
that were equipped with the same programmable hardware platform, but over the last 1-2
years other manufacturers have launched hearing instruments with DFS implementations.
Some of these newly implemented DFS algorithms are calibrated during hearing instrument
fitting; others are calibrated during runtime utilizing environmental sounds as stimuli.
2.7. Evaluation of DFS
2.7.1. Currently used method
Hearing instrument users are able to appreciate whether a DFS system is efficiently dealing
with feedback and artifacts resulting from physical interaction with the hearing instrument
(e.g. getting dressed, putting on a hat, giving a hug, using the phone). Feedback and artifacts
Olsen, Measurement of DFS benefit
19
also occurs when leakages around the ear canal wall and the hearing instrument is introduced
by improper placement or by jaw movements (e.g. chewing).
There is at present no generally accepted method for testing DFS systems, but evaluation of
the headroom owing to the use of DFS has been done using the following procure (Merks et
al., 2006): The hearing instrument is set to a linear flat frequency-gain response and the
overall gain is increased until the maximum whistle-free gain (often called the maximum
stable gain, MSG) is reached. This is done with and without DFS.
The gain value found with activated DFS minus the value found with deactivated DFS is seen
as the DFS benefit. This method does not show DFS benefit as a function of frequency. The
method has been modified in the way that the amplification was adjusted up and down at
multiple frequencies to shape the maximum feedback-free linear amplification at individual
frequencies rather than in a frequency region. The modified method is however very time
consuming and if carried out on a human subject, unpleasant because of the ongoing
whistling. Also it is not possible to shape the MSG very accurately because of too few gain
handles in the fitting software.
2.8. Real-ear Measurements
Real-ear measurements (REM) can be performed to verify that the individually prescribed
hearing instrument gain has been realized at the plane of the tympanic membrane (Larsby and
Arlinger, 1988; Northern, 1992). This is to ensure that the components of the entire electroacoustical system, consisting of the hearing instrument, the tubing, acoustical filters, earmold
and the external ear, constructively interact to achieve the desired sound pressure in the ear
Olsen, Measurement of DFS benefit
20
canal as prescribed by the pre-selected prescription formula (Dillon, 2001). Gain prescription
for a certain hearing loss at specific frequencies is normally given as the real-ear insertion
gain (REIG), which is the difference in sound pressure at the tympanic membrane, with and
without amplification (Dalsgaard and Dyrlund-Jensen, 1976).
When performing REM, a known sound signal is presented to the aided and unaided ear and
the sound pressure level that is produced in the ear canal is measured. A loudspeaker delivers
the signal while the resulting sound signal is measured deep in the ear canal by means of a
probe tube that is connected to the measurement system microphone. The real-ear aided
response (REAR) is the sound pressure level (SPL) as a function of frequency measured in
the ear canal with the hearing instrument turned on (ANSI S3.46-1997; American National
Standards Institute, 1997). Two REM methods are often used:
2.8.1. The substitution method
When using the substitution method for REM, the sound level is first measured at the point
where the test subject’s head will be placed for the measurement but without the test subject
in place. This recording is stored in the measuring equipment memory as a reference for
subsequent measurements in the unaided and in the aided ear canal (ANSI S3.46-1997;
American National Standards Institute, 1997).
2.8.2. The modified pressure method
With the modified pressure method (ANSI S3.46-1997; American National Standards
Institute, 1997) no equalization of the test room is done, but the loudspeaker signal is
Olsen, Measurement of DFS benefit
21
continuously controlled to produce a constant sound pressure level throughout the
measurement, using a reference microphone placed near the subject’s ear canal entrance. Any
deviation from the preset sound pressure level will result in an adjustment of the loudspeaker
output.
The modified pressure method can be disturbed by sound leaking out of the ear canal when
using open fittings (Larsby and Arlinger, 1988). With conventional technology the impact
was limited since the obtainable feedback-free loop gain was small, but Dyrlund et al. (2006)
demonstrated that the problem increases with DFS algorithms allowing for relatively high
loop gains in open fittings without experiencing feedback whistling. Dyrlund and al. (2006)
showed that using the modified pressure method for REM on open fittings of hearing
instruments with DFS may result in underestimated REIG and that the more the feedback
suppression is utilized to provide stable gain not possible without feedback suppression, the
more the gain will be underestimated.
2.8.3. REM of DFS benefit
In a previous work, Olsen and Hernvig (2006) used modified REM equipment for
measurement of the DFS benefit in hearing instruments. While doing REM’s with the
modified pressure method (ANSI S3.46-1997; American National Standards Institute, 1997),
the sound pressure level was kept constant near the BTE microphone using a reference
microphone placed very close to the test hearing instrument microphone. When DFS allows
the loop gain to exceed 0 dB, the REM reference microphone used for the modified pressure
method will measure a sound pressure exceeding the preset stimulus value and hence the
equipment will reduce the test signal level to maintain the preset sound pressure level
Olsen, Measurement of DFS benefit
22
(Dyrlund et al. 2006). If the loop gain is high, the stimulus sound pressure will be reduced
with approximately as many decibels as the DFS allows the loop gain exceeding 0 dB.
Subsequently the aided sound pressure in the ear canal will drop with the same number of
decibels. Measurements with the substitution method (ANSI S3.46-1997; American National
Standards Institute, 1997) are unaffected by sound leaking out of the ear canal (Dyrlund et al.,
2006), and the substitution method therefore gives a representation of the “true” gain that can
be obtained with the DFS system activated. The difference between the outcomes obtained
with the two methods represents approximately the extra headroom provided by the DFS
system.
For a more exact estimation we must consider how the feedback signal and the test signal
interact at the reference microphone. Some of the sound reaching the reference microphone
comes directly from the test equipments loudspeaker, other parts of the sound are reflected
from walls, floor and ceiling in the test room, other sound is reflected from the head and
finally amplified sound leaks out of the ear canal. If the test signal is a warble tone or a noise
signal, with many frequency components, it is reasonable to assume that the feedback and test
signal power values are simply added. Then, denoting the actual test signal level with the
pressure method as Lin, the loop gain as Gloop, the nominal reference signal level as Lref, kept
constant by the test system, the total signal power registered at the reference microphone is
the sum of test-signal and feedback-signal power values:
Lin
(1)
10 10 10
Lin GIoop
10
This relation can be rewritten as
Lref
10 10
Olsen, Measurement of DFS benefit
(2)
1 10
GIoop
Lref Lin
10
10
10
23
REARdiff
10
10
The last equality is valid because the reduction in the test signal level (Lref-Lin) with the
pressure method is the same as the reduction in measured REAR. Thus the loop gain can be
estimated from the REAR difference as
REARdiff
(3)
Gloop
10 u log10 (10
10
1)
For example, if the loop gain is exactly 0 dB, the REARdiff would be 3 dB. At high loop gain
values, Gloop is approximately equal to REARdiff. For example, with REARdiff = 10 dB, the
loop gain would be Gloop 10 u log10 (9)
9.5 dB.
In an unpublished study using the proposed method and equipment for assessment of DFS
benefit the measurement was done five times on a KEMAR with the same hearing
instrument. The device, the probe tube and the reference microphone were removed and
replaced on the KEMAR head and the DFS was calibrated between each measurement. The
standard deviation of the differences between the extra headroom found at the five
measurements was up to 2.3 dB. This variance was considered to be too high since the
measured DFS benefit typically is in the order of 10 -15 dB.
2.8.4. Simulated real-ear measurements
Simulated real-ear measurements (sREM), e.g. insertion gain on a manikin with human
proportions may provide important information for the evaluation of hearing instrument
performance (ANSI S3.35-2004; American National Standards Institute, 2004). sREM of
DFS benefit allows control over the size of the auricle and the ear canal thus limiting the
variance between measurements carried out at different places and at different times. By
Olsen, Measurement of DFS benefit
24
using an occluded ear simulator it would be possible to avoid the use of a probe tube for the
measurement of sound pressure in the aided ear canal, which would also limit the inter test
variance.
3.
Material and methods
3.1. Equipment and test environment
The measurements in this study were made on the Knowles Electronics Manikin for Acoustic
Research (KEMAR) (Burkhard and Sachs, 1975). The KEMAR was equipped with a type
DB-066 rubber auricle, which is a medium ear size.
The Madsen Electronics Aurical with REM software module version 2.54 was used for the
measurement. Firmware versions were: Monitor version 1.00, program version 2.41 and DSP
version 2.41.
The normal reference microphone built into the REM headset was replaced by an external
microphone (Knowles MB6022ASC-2) that could be placed near the hearing instrument
microphone. In this way the equipment reduced the loudspeaker output according to the extra
feedback-free gain near the hearing instrument microphone. An occluded ear simulator (B&K
4157) mounted inside the KEMAR ear replaced the Aurical measuring microphone. In the
occluded ear simulator the microphone is located at the position that corresponds to the
tympanic membrane. The microphone signal is routed through a preamplifier (G.R.A.S. type
26AC) and a conditioning amplifier (G.R.A.S. type 12AA) to the Aurical system. The
Aurical test signal was delivered to an external loudspeaker (KEF Q85s).
Olsen, Measurement of DFS benefit
25
3.1.1. The test room
KEMAR was placed in the center of a single walled audiometric booth (C-A Tegnér AB)
measuring 2.41 m (width) x 3.03 m (length) x 2.17 m (height). The noise level in the room
was measured in 1/3 octave bands (NTI Minilyzer type ML1 connected to a measuring
microphone, NTI Mini SPL). The noise floor is depicted in Figure 3. The room has carpet on
the floor and sound absorption panels on all four walls. There is a window (0.71 m (width) x
0.915 m (height)) in one of the walls. The loudspeaker was mounted 0.5 m in front of the
KEMAR at head level and at an angle of 45° azimuth, which ensures the lowest variance
between repeated REM’s (Hawkins and Mueller, 1992; Dillon, 2001).
Figure 3. Noise floor in test room.
The researcher controlled the measurements from a position outside the test room.
Olsen, Measurement of DFS benefit
26
3.2. Calibration
The Aurical hearing instrument test (HIT) microphones were calibrated using the 94 dB SPL,
1 kHz B&K 3424 calibrator. After that the REM reference and probe microphones were
calibrated using the HIT microphones as reference. For a detailed description see the
appendix (chapter 8.1.)
3.3. Test hearing instrument and fitting software
For this experiment a micro BTE hearing instrument (ReSoundAIR) designed for open
fittings (Kiessling et al., 2003) was used. The hearing instrument is intended for use by
hearing impaired individuals with relatively good or normal hearing at low frequencies as it
does not amplify sound at low frequencies. (See the technical data sheet in appendix (chapter
8.2). The hearing instrument was programmed using the manufacturer’s fittings software
(ReSoundAIR, Aventa Version 2.05.6a, NOAH3 version 3.1.2 (1391D)). The test hearing
instrument provides no gain at low frequencies. At mid and high frequencies the gain for low
inputs (G50) is limited to 30 dB at 1 kHz and 35 dB at 2, 3 and 6 kHz and to 10 dB at 1 kHz
and to 15 dB at 2, 3 and 6 kHz for high inputs (G80). The gains are limited to avoid distortion
due to saturation of the receiver. The proposed method for evaluation of DFS benefit relies on
the fact that the input to the hearing instrument microphone decreases for loop gains
exceeding 0 dB, when REM’s are done with the modified pressure method. Decreased inputs
lead to higher gains, when the tested hearing instrument is in compression mode. Such
changes in the hearing instrument microphone input could influence the outcome of the DFS
evaluation and therefore a patch to the fitting software was made to allow increasing gain up
to a linear gain of 65 dB. Test stimulus levels were selected with caution so as to avoid
saturation of the hearing instrument receiver at high gain settings. In this study the noise
Olsen, Measurement of DFS benefit
27
reduction and the low-level expansion algorithms were switched off during testing in order to
avoid interaction (see chapter 2.5.4).
The NOAHLink interface was used for programming the test hearing instrument. The Aurical
software embedded in the Aventa fitting software was used to allow hearing instrument
handling and REM measurements from the same computer.
The manufacturer provides four different tube lengths with two different insertion lengths
that are combined with a silicone dome (earmold) that comes in three sizes to connect the
hearing instrument to the ear canal. The test hearing instrument was fitted using: short length
(type 1) tube, “short” tube insert length (type B) and “normal” dome size (diameter: 7 mm).
3.4. General procedures
Before each session the occluded ear simulator was taken out of the KEMAR ear and was
held close to the reference microphone in front of the loudspeaker and a calibration was
performed. The occluded ear simulator was put back in the KEMAR ear. The test hearing
instrument was placed on the left KEMAR ear and the NOAHLink interface was hanging
from the neck down the back of the KEMAR. The measuring equipments reference
microphone was taped to the KEMAR head as close to the hearing instrument microphone as
possible (Figure 4). The reference microphone was pointing forward. The hearing instrument
was connected to the NOAHLink left channel with a Flexconnect 2 CS 63 programming
cable.
Olsen, Measurement of DFS benefit
28
After the hearing instrument was connected to the computer, noise reduction and expansion
algorithms were deactivated and the DFS was calibrated and activated. The hearing
Reference microphone
Hearing instrument
microphone
Figure 4. Placement of hearing instrument and reference
microphone on KEMAR head.
instrument was set to the linear gain to be investigated and REAR was obtained using the
substitution method and after that with the modified pressure method. The results of the
measurements were exported to a spreadsheet for later analysis. The DFS benefit was
calculated as the difference between the REARs obtained with the substitution and the
modified pressure methods.
3.5. Experiments
3.5.1. Reduction of loudspeaker sound pressure
When using the modified pressure method for REMs the loudspeaker stimulus will be
reduced with approximately as many decibels as the DFS allows the loop gain to exceed 0 dB
(see chapter 2.8.3). Some of the reduction is, however, a result of the interaction between the
Olsen, Measurement of DFS benefit
29
stimulus signal and the aided signal leaking out of the ear canal. When the reduction is
known, the loop gain can be calculated using equation (3). In order to test the validity of the
equation, REARs with the substitution and the pressure methods were done with DFS
activated in the test hearing instrument. REARs were measured with the two methods in the
range 2 – 32 dB in two-dB steps. These gain settings were selected as the feedback influence
on the signal was negligible (See chapter 3.5.2). The stimulus level was kept constant at 50
dB SPL.
3.5.1.1. Results
Figure 5 shows two pairs of measurements obtained at gain settings 14 and 32 dB. At 32 dB,
where DFS is crucial for keeping the hearing instrument free from acoustic feedback, a large
difference between REARs measured with the two methods is seen. The difference is much
smaller at 14 dB, where DFS plays a smaller role.
Sound Pressure (dB SPL)
100
Subst 32
Press 32
Subst 14
Press 14
90
80
70
60
50
40
100
1000
10000
Frequency (Hz)
Figure 5. Two pairs of measurements (substitution and modified pressure methods) obtained at gain setting 14
and 32 dB. A large difference between the curves obtained with two methods is seen at 32 dB, while the
difference is smaller at 14 dB.
Olsen, Measurement of DFS benefit
30
Figure 6 shows the reduction of the loudspeaker sound pressure at the 16 gain settings as a
function of the hearing instrument gain setting. The sound pressure reduction was calculated
as the maximum difference between the REAR obtained with the modified pressure and the
substation methods found in the frequency range 2-8 kHz.
The theoretical reduction of the loudspeaker sound pressure for a certain loop gain can be
calculated from the following equation derived from equation (2):
G Ioop
(4)
REARdiff
Log10 (1 10
10
)
The thin line in the figure represents the REARdiff calculated for different loop gains using
equation (4). The theoretical function was fitted by eye to best correspond to the measured
value and has its 3-dB value at a hearing instrument handle setting close to 16 dB. This
corresponds to 0 dB loop gain. Without DFS the test hearing aid is whistling at a mean gain
of 22 dB, and feedback peaks in the REAR can be visually observed down to 9 dB.
Olsen, Measurement of DFS benefit
31
18
Measured reduction
16
Theoretical reduction
Diffence (Subst-press) (dB)
14
12
10
8
6
4
2
0
0
10
20
30
40
Gain setting (dB)
Figure 6. The reduction of the loudspeaker sound pressure as a function of the hearing
instrument gain setting (diamonds). The theoretical reduction of the loudspeaker sound
pressure as calculated from equation (1) is also shown (thin line).
3.5.2. Definition of maximum gain with acceptable feedback
The hearing instrument instability point is often defined as the maximum hearing instrument
gain that can be achieved without audible feedback whistling. Feedback whistling is,
however, not always heard at a fixed point. Between the gain level at which the hearing
instrument always whistles and the level where it never whistles, there are levels where the
instrument may whistle sometimes. Feedback is always present in the hearing instrument
signal, and at such gain levels the sound quality might be affected negatively by peaks and
valleys in the frequency response (Dillon, 2001). Cox (1982) demonstrated that peaks and
valleys are present in the hearing instrument frequency response at gain settings just below
Olsen, Measurement of DFS benefit
32
that required for producing audible oscillation. Such peaks and valleys get smaller with
reduced gain (Dillon, 2001). Based on these observations the following method for deciding
the presence of feedback in the signal was used:
x
DFS was activated
x
The hearing instrument was set to 0 dB linear gain at all frequency handles.
x
The hearing instrument gain was increased in 1-dB steps at all frequencies until
feedback whistling was audible.
x
The gain was then decreased at all handles by 1-dB steps until no feedback whistling
was heard.
x
The REAR was obtained with this gain setting for a 50 dB SPL warble tone input
using the substitution method (chapter 2.8.1).
x
The gain was decreased in 2-dB steps until the lowest possible gain setting, and
REAR was measured again at each gain step.
x
In each REAR the peak was identified and its gain value was depicted as a function of
the fitting software gain setting.
x
The amount of acoustical feedback in the signal was accepted if an increase of 2 dB
resulted in an increase of 2 dB ± 0.2 dB and if the peak frequency variation was less
than ±100 Hz. A spreadsheet was programmed to find the highest fitting software
gain setting at which these conditions were fulfilled. (See appendix, chapter 8.3).
x
The highest gain deemed free of feedback based on the requirements regarding stable
frequency and stable peak gain increments was selected as the “maximum feasible
gain”.
Olsen, Measurement of DFS benefit
33
This procedure was completed ten times with one single hearing instrument. Between each
measurement, the hearing aid was removed and replaced on the KEMAR ear and the DFS
was calibrated.
3.5.2.1. Results
In Figure 7 one of the ten sets of REARs obtained with linear gain settings (1 to 41 dB in 2dB steps) is seen for a 50 dB SPL input.
Simulated REM
(Input 50 dB SPL)
110
Output (dB SPL)
100
90
80
70
60
50
100
1000
10000
Frequency (Hz)
Figure 7. Simulated real ear aided responses with the test hearing instrument set
to linear gains ranging from 1 to 41 dB in 2-dB steps. Stimulus level was 50 dB
SPL.
Olsen, Measurement of DFS benefit
34
It is clearly seen that the measured gain-increase is more than 2 dB near the response peak for
high gain settings. In a signal free from acoustical feedback an increase of the hearing
instrument gain setting by 2 dB would result in a peak gain increase of 2 dB. The stretch of
gain settings for which this holds true (within ± 0.2 dB) was identified and the mean
difference between the REAR peak value and the hearing instrument gain handle setting was
calculated. Based on this mean value the expected peak value was calculated for all gain
settings. The expected value was subtracted from the measured value. The mean result (with
standard deviation) from the ten measurements can be seen in Figure 8 as a function of
hearing instrument gain handle setting. At the stretch from just above 10 to 30 dB gain setting
there is a close to linear relation between the gain setting and the peak value. Below 10 dB
the relation becomes nonlinear because of influence from internal hearing instrument noise
and above around 30 dB the relation becomes nonlinear due to acoustic feedback. At gain
settings below 30 dB the standard deviation is small, however above this limit the standard
deviation increases.
Olsen, Measurement of DFS benefit
35
Peak gain
6
Deviation (dB)
5
4
3
2
1
0
-1
0
10
20
30
40
50
Handle gain (dB)
Figure 8. The measurement shown in Figure 7 was repeated ten times. In the figure the mean peak increase of
the REARs found in ten the measurements are shown as a function of hearing instrument handle gain setting.
Also shown is the SD of the repeated measurements.
It is also seen in Figure 7 that the peak frequency changes much at high gains. The mean peak
frequency (with standard deviation) found in the ten measurements is depicted in Figure 9.
As seen the peak frequency is relatively stable up to just below 30 dB. Above that value there
are big changes in the peak frequency and the standard deviation increases.
Olsen, Measurement of DFS benefit
36
Peak Frequency (kHz)
Peak frequency
4,1
4,0
3,9
3,8
3,7
3,6
3,5
3,4
3,3
3,2
0
10
20
30
40
50
Handle gain (dB)
Figure 9. The measurement shown in Figure 7 was repeated ten times. In the figure the mean peak frequency of
each REAR is shown as a function of the gain setting. Also shown is the SD of the repeated measurements.
The procedure for identifying the maximum feasible gain (MFG, chapter 3.5.2) was done ten
times. The mean MFG found with a 0.2 dB gain deviation limitation was 32.4 dB (SD 1.6).
The MFG found with a 0.1 kHz frequency deviation limitation was 31.6 (SD 2.3). If the
lowest value of the two was selected the mean MFG was 31.2 (SD 2.0).
3.5.3. sREM test-retest variance.
After the feasible gain (32 dB) had been selected (see chapter 3.5.2) the procedure for sREM
of DFS described in chapter 3.4 was completed ten times. The input level was 50 dB SPL and
the gain was set to 32 dB. One single hearing instrument was used for the measurements.
Between each measurement, the hearing aid was removed and replaced on the KEMAR ear
and the DFS was calibrated. The reference microphone was also removed and replaced on the
KEMAR head between measurements.
Olsen, Measurement of DFS benefit
37
3.5.3.1. Results
Figure 10 shows the mean curves obtained with the modified pressure method and the
substitution methods.
Output (dB SPL)
Simulated REAR
100
95
90
85
80
75
70
65
60
55
50
100
Substitution
Pressure
1000
10000
Frequency (Hz)
Figure 10. The mean REARs obtained with the modified pressure method and the substitution method (solid
line). The measurement was done ten times on KEMAR with 32 dB hearing instrument handle gain and a 50 dB
SPL warble tone sweep as input signal.
The mean difference between the two curves is depicted in Figure 10. Also seen is the
standard deviation (SD) of the differences between the curves obtained at the ten
measurements.
Olsen, Measurement of DFS benefit
38
DFS benefit
(10 measurements)
18
Mean
16
+ SD
Benefit (dB)
14
- SD
12
10
8
6
4
2
0
-2
100
1000
10000
Frequency (Hz)
Figure 11. The mean difference (solid line) between the REAR obtained with the modified pressure method and
the substitution method. The measurement was done ten times on KEMAR with a hearing instrument handle 32
dB gain and a 50 dB SPL warble tone sweep as input signal. The standard deviation (SD) of the differences
obtained at the ten measurements is also shown (thin line).
The SD as a function of frequency is shown in Figure 12. The maximum SD in the amplified
frequency area is around 2 dB. The same measurements were done with a gain setting of 27
dB. With this setting the maximum difference between the two REARs was around 8 dB and
the max SD up to 1.2 dB.
The mean difference between the two measurements in Figure 10 is peaking at 4.25 kHz at
which frequency the loudspeaker sound pressure has been reduced by 14.3 dB owing to the
modified pressure control system. Using equation (3) the loop gain at that frequency is
calculated to 14.1 dB, which is the DFS headroom or in other words the benefit achieved by
using the DFS algorithm.
Olsen, Measurement of DFS benefit
39
Standard deviation
2,5
SD (dB)
2,0
1,5
1,0
0,5
0,0
100
1000
Frequency (Hz)
10000
Figure 12. The standard deviation (SD) of the differences obtained at the ten measurements of the differences
between the REAR found with the modified pressure method and the substitution method. The measurement
was done ten times on KEMAR with 32 dB gain and with a 50 dB SPL warble tone sweep as input signal. The
hearing instrument and the reference microphone was removed and replaced on the KEMAR.
3.5.4. Test-retest variance with the currently used method
The currently used method for estimation of DFS headroom (chapter 2.7.1) was carried out
ten times on the same test hearing instrument used in chapter 3.5.3. Between the
measurements the hearing instrument was taken out of the KEMAR, replaced and the DFS
was calibrated. An audiologist who was not able to see the hearing instrument handle gain
setting while setting the gain carried out the test. The gain setting just below the setting where
feedback whistling was heard was stored as the MSG with and without DFS.
3.5.4.1. Results
The mean MSG was 22 dB (SD: 0.7) without DFS and 40 dB (SD: 0.7) with DFS. The DFS
benefit was calculated to 18 dB (SD: 0.4 dB).
Olsen, Measurement of DFS benefit
4.
40
Discussion
The results depicted in Figure 6 support the assumption made for equation (1). Hence
equation (3) can be used for calculation of the loop gain when the difference between the
REAR obtained with the substitution and the modified pressure method is known.
Feedback-influence on the hearing instrument signal was investigated by testing the stability
of the peak frequency and of the sound pressure increase as a result of 2-dB increments of the
gain handle-setting. The mean MFG found with the two methods separately and with the
combination of the two methods was nearly the same: around 32 dB with this test hearing
instrument. The very small changes in peak frequencies and gain increments might not be
audible for a user of hearing instruments, but apparently the changes can still disturb the
repeatability of the sREM (Figure 12), since the repeatability of sREMs was better with a
gain setting of 27 dB. Still the full DFS benefit might be shown better at a 32 dB gain setting.
A SD of about 2 dB seems to be to much for a measurement DFS that yields a benefit of
around 14 dB. To improve the accuracy, the DFS headroom could be calculated based on the
mean result of several sREMs.
To be a valid method for measurement of DFS headroom, sREM must accurately show how
much the DFS allows the gain to be increased in a hearing instrument without having audible
whistle or artifacts resulting from feedback. The present method is valid in the sense that the
loop gain can be accurately calculated from the difference found between sREM with the
substitution and the modified pressure methods. Problems arise from the fact that at gain
levels where changes of the peak sound pressure and frequency are inaudible to hearing
Olsen, Measurement of DFS benefit
41
instrument users, they still cause a too high test/retest variance. It is known that narrowband
signals can be cancelled out by adaptive feedback cancellation systems (Kates, 2003). The
extent to which the narrow band signal (in this case a warble tone) is cancelled out from one
trial to another may vary and cause some of the test/retest variance. The use of narrow band
noise signals instead of warble tones might help getting a lower variance.
The mean DFS headroom found with the currently used method was somewhat higher (18
dB) and the SD of the inter test differences (0.4 dB) was much smaller than found with the
proposed new method (mean headroom: 14 dB, SD: 2 dB). It is, however, a serious drawback
with the currently used method, that headroom is measured at gain just below the point where
the hearing aid is going to whistle. At this gain level the sound quality is very poor, and the
higher value found with the currently used method for this reason does not show the useable
DFS benefit because of distortion and because the hearing instrument would easily start
whistling even with minor changes of the feedback path. In contrast the new method offers
measurement of DFS benefit at gain levels where the signal is not distorted by feedback.
In this study the goal was to find ways of measuring the benefit provided by DFS, defined as
how much extra useable gain that can be achieved by using the specific DFS algorithm. In
order to study the “clean” DFS effect on feedback, the test hearing instrument was placed on
a KEMAR, set to linear amplification and noise reduction and expansion algorithms were
deactivated. By doing so the test hearing instruments was tested in a mode in which it is not
generally used by the hearing instrument user. Furthermore the measurements in this study
were all done under static conditions, i.e. no attempts were done to measure changes in DFS
performance resulting from changes in the feedback path.
Olsen, Measurement of DFS benefit
42
Little or nothing can therefore be said on how DFS works in daily life based on the sREM
proposed in this study. The sREM can be used to compare the effectiveness of different DFS
algorithms and implementations in different devices, but the measurements tell nothing about
how the algorithm interacts with other algorithms running in the hearing instrument, when
worn on a living ear canal.
Algorithm developers should not tune DFS algorithms to work to perfection on KEMAR to
achieve high performance when compared to competing systems. Hearing instruments should
always be developed with end user benefit in mind, - not to manikins.
5.
Conclusion
x
A method for simulated real ear measurements showing frequency dependent
information on DFS benefit is proposed
x
The new method yields a more true value for DFS benefit, than the currently used
method does
x
The test/retest variance of the objective method for measurements of DFS headroom
is higher than that of the currently used method, but this can be solved by using the
mean from repeated measurements.
Olsen, Measurement of DFS benefit
6.
43
References
American National Standards Institute. (1997) Methods of measurement of real-ear
performance characteristics of hearing instruments. (ANSI S3.46-1997). New York:
Acoustical Society of America.
American National Standards Institute. (2004) Method of measurement of performance
characteristics of hearing instruments under simulated real-ear working conditions. (ANSI
S3.35-2004). Melville: Acoustical Society of America.
Bisgaard N. (1993) Digital feedback suppression - Clinical experiences with profoundly
hearing impaired. In: Beilin J, Jensen GR, eds. 15th Danavox Symposium. Recent
Developments in Hearing Instrument Technology. Copenhagen: Stougaard Jensen, 371-384.
Burkhard M, Sachs R. (1975) Anthropomorphic manikin for acoustic research. J Acoust Soc
Am 58:214-222.
Courtois J. (1988) Open molds. In: Jensen JH, ed. 13th Danavox symposium. Hearing
Instrument Fitting. Theoretical and Practical Views. Copenhagen: Stougaard Jensen, 175200.
Cox RM. (1982) Combined effects of earmold vents and suboscillatory feedback on hearing
instrument frequency response. Ear Hear 3:12–17.
Olsen, Measurement of DFS benefit
44
Dalsgaard S, Jensen OD. (1976) Measurements of the insertion gain of hearing instruments. J
Audiol Tech 15:170.
Dillon H. (2001) Hearing Aids. New York: Thieme.
Dyrlund O, Olsen SØ, Haastrup A, Lantz J. (2006) REIG measurements on open fittings
using digital feedback suppression. In: Rasmussen AN, Poulsen T, Andersen T, Larsen CB,
eds. 21st Danavox Symposium on Hearing Aid Fitting. Brønby: Centertryk, 167-171.
Flack L, White R, Tweed J, Gregory DW, Qureshi MY. (1995) An investigation into
attenuation by earmold tubing. Br J Audiol 29:237–245.
Hawkins DB, Mueller HG. (1992) Procedural considerations in probe-microphone
measurements. In: Mueller HG, Hawkins DB, Northern JL, eds. Probe Microphones
Measurements. Hearing Instrument Selection and assessment. San Diego: Singular
Publishing Group, 67-89.
Heide VH. (1994) Nicolet Project Phoenix, Inc. 1984 - 1989. The Development of a
Wearable Digital Signal Processing Hearing Aid. In: Sandlin RE, Ed. Understanding
Digitally Programmable Hearing Aids. Needham Heights: Allyn and Bacon, 123-150.
Hellgren J, Lunner T, Arlinger S. (1999) System identification of feedback in hearing
instruments. J Acoust Soc Am 105:3481-3496.
Olsen, Measurement of DFS benefit
45
Herbst G. (1989) The digitally programmable instrument. Part 2: The PHOX. Hear Instr 40
(3):36-40.
Kates JM. (1999) Constrained adaptation for feedback cancellation in hearing instruments. J
Acoust Soc Am 106:1010-1019.
Kates JM. (2003) Adaptive feedback cancellation in hearing instruments. In: Benesty J,
Huang Y eds. Adaptive Signal Processing: Application to Real-World Problems. NewYork:
Springer, 23-58.
Kiessling J, Margolf-Hackl S, Geller S, Olsen SØ. (2003) Researchers report on a field test of
a non-occuding hearing instrument. Hear J 56(9):36-41.
Killion MC. (1993) The K-AMP™ hearing aid: An attempt to present high fidelity for the
hearing impaired. In: Beilin J, Jensen GR, eds. 15th Danavox Symposium. Recent
Developments in Hearing Instrument Technology. Copenhagen: Stougaard Jensen, 167-230.
Larsby B, Arlinger S. (1988) Comparison of different types of equipment for insertion gain
measurement. In: Jensen JH, ed. 13th Danavox symposium. Hearing Instrument Fitting.
Theoretical and Practical Views. Copenhagen: Stougaard Jensen, 39-48.
Mangold S, Eriksson-Mangold M, Israelsson B, Leijon A, Ringdahl A. (1990) Multiprogrammable hearing aid. Acta Otolaryngol Suppl 469:70-75.
Olsen, Measurement of DFS benefit
46
Merks I, Banerjee S, Trine T. (2006). Assessing the effectiveness of feedback cancellers in
hearing aids. Hear Rev 13 (4), 53-57.
Northern JL. (1992) Introduction to computerized probe-microphone real-ear measurements
in hearing instrument evaluation procedures. In: Mueller HG, Hawkins DB, Northern JL, eds.
Probe Microphones Measurements. Hearing Instrument Selection and Assessment. San
Diego: Singular Publishing Group, 1-19.
Oliveira RJ. (1997) The active ear canal. J Am Acad Audiol 8:401–410.
Olsen SØ, Hernvig LH. (2006) Objective evaluation of maximum stable gain level (MSG)
and of the performance of digital feedback suppression (DFS) in hearing instruments. In:
Rasmussen AN, Poulsen T, Andersen T, Larsen CB, eds. 21st Danavox Symposium 2005 on
Hearing Aid Fitting. Denmark: Centertryk, 399-406.
Olsen SØ, Perman S, Hansen KH, Pedersen, BD. (2003) Experiments with a new generation
digital feedback suppression system. Paper presented at 6th EFAS Congress, Crete, Greece,
May 9-12.
Pihl-Frank L, Leijon A, Adolfsson L, Shahabi J.(2000) Danalogic 163D field trial. Paper
presented at the Tema Hörsel conference, Karlstad, Sweden, April 4-6.
Ringdahl A, Eriksson-Mangold M, Israelsson A, Lindkvist A, Leijon A, Mangold S. (1988)
Evaluation of a programmable hearing aid with multiple memories. In: Jensen JH, ed. 13th
Olsen, Measurement of DFS benefit
47
Danavox symposium. Hearing Instrument Fitting. Theoretical and Practical Views.
Copenhagen: Stougaard Jensen, 175-200.
Rosen S, Howell P. (1997) Signals and systems for speech and hearing. London: Academic
press.
Stypulkowski PH. (1994) 3M programmable hearing instruments. In: Sandlin RE, ed.
Understanding Digitally Programmable Hearing Aids. Needham Heights: Allyn and Bacon,
97-121.
Westermann S. (1994) Principles and application of the Widex multiprogrammable hearing
aid system. In: Sandlin RE, ed. Understanding Digitally Programmable Hearing Aids.
Needham Heights: Allyn and Bacon, 41-66.
Yates GK. (1995) Cochlear structure and function. In: Moore BCJ, ed. Hearing. Handbook of
Perception and Cognition. 2nd edition. London: Academic Press, 41-74.
Olsen, Measurement of DFS benefit
7.
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Acknowledgements
The idea of developing an objective repeatable method for measuring DFS benefit came up
when I worked with Ole Dyrlund, Astrid Haastrup (GN ReSound, Copenhagen) and Johannes
Lantz (GN Otometrics, Copenhagen) on describing the measurement errors occurring in
REM of hearing instruments equipped with DFS. We found that the degree of error increased
the more the DFS algorithm was put at work to keep the hearing instrument stable, and I felt
that this in some way could be used to measure the headroom made possible using DFS.
People from several countries worked with me with this study:
My mentors were Arne Leijon (Royal Institute of Technology, Department of Signals,
Sensors and Systems, Stockholm) and Ingrid Lennart (Lund University, Department of
Clinical Sciences, Logopedics, Phoniatrics and Audiology). Anders Jönsson (Lund
University) volunteered to be my adviser during this project. Peter Kossek (GN Otometrics,
Taastrup) did the modifications of the Madsen Aurical equipment used for the measurements
in this study, and Arne Hutzflesz (GN Otometrics, Taastrup) helped with software updates
and solving problems whenever they occurred. Jesper Rye Nielsen (GN ReSound, Taastrup)
calibrated the system and Tim Buchwald Christensen (GN ReSound, Taastrup), patched the
Aventa fitting software so hearing instruments could be set to linear gain. The sREMs were
made with Jim Luther (Algorithm Group, GN ReSound, Taastrup) and the test/retest with the
currently used method was done by Lotte Hagen Hernvig (Research Audiology, GN
ReSound, Taastrup). Maïa Weddin, Shawn Perman and Lars Spicker Olesen (Algorithm
Group, GN ReSound, Taastrup) helped me understanding how DFS works and how it
interacts with other algorithms working in modern digital hearing instruments. John Nelson
(Director of Research Audiology Group, GN ReSound, Taastrup) revised earlier version of
Olsen, Measurement of DFS benefit
49
this manuscript and supported with ideas and suggestions. Jeff Bondy and Maureen Coughlin
(Auditory Research Laboratory, GN ReSound, Glenview, USA) helped with advices,
comments and revision of the manuscript.
Jan Grenner (Consultant at the department of otorhinolaryngology, University Hospital,
Lund) was examiner and my colleague Anna-Pia Pehrsson was opponent at the defense
which took place June 2nd 2006. Both gave me valuable input that improved the thesis
significantly.
Thanks to all these advisers, colleagues and friends for their kind cooperation.
Steen Østergaard Olsen
Copenhagen, June 2006
Olsen, Measurement of DFS benefit
8.
50
Appendix
8.1. Calibration of the Aurical System
Calibration of the Aurical hearing instrument test (HIT) microphones
x
Start the Aurical HIT module
x
Choose Calibration from the setup menu.
x
Select the “Ref. microphone” box in the calibration window
x
Connect 94 dB SPL 1 kHz B&K 3424 calibrator to the reference microphone and turn
it on.
x
Click “Calibrate” to start calibration.
x
Select the “Meas. Microphone” in the calibration window
x
Connect 94 dB SPL 1 kHz B&K 3424 calibrator to the measure microphone and turn
it on.
x
Click “Calibrate”
x
Exit calibration and save data
Calibration of the Aurical real-ear measurement (REM) microphones
Calibration of REM reference microphone:
x
Start the Aurical REM module
x
Select “Calibration” from the setup menu.
x
Select “Calibrate with HIT microphone”
x
Select the “REM Probe 1” box and “Internal chamber” box
x
Place the HIT reference microphone in the HIT chamber center spot
Olsen, Measurement of DFS benefit
x
51
Place the sound inlet of REM reference microphone as close as possible to the HIT
microphone sound inlet
x
Close the HIT box
x
Click on “Calibrate” and hold the HIT reference microphone and the REM reference
microphone steady until the calibration has finished
x
Close “Calibration” and save the calibration.
Calibration of REM probe microphone:
x
Take out the B&K 711 coupler of the KEMAR ear.
x
Hold the REM reference microphone in front of the Aurical internal speaker, and as
close as possible to the sound inlet of B&K 711 coupler.
x
Click “Tube Calibrate” and hold the REM reference microphone and the B&K 711
coupler steady until the sweep has finished.
x
The calibration is automatically stored
Olsen, Measurement of DFS benefit
8.2. ReSoundAIR technical data sheet
52
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53
Olsen, Measurement of DFS benefit
54
8.3. Method for selection of “maximum feasible gain” (Excel sheet)
Find highest gain with linear relation between fitting software gain setting and realized peak
gain.
x
Find REAR maximum value (PEAK.SPL) for each fitting software gain setting
(SWgain).
x
Find difference between PEAK.SPL and SWgain for all SWgains (PEAKdiff).
x
Find minimum PEAKdiff (PEAKdiffMin).
x
Calculate mean value of PEAKdiff for which PEAKdiff < PEAKdiffMin + 0.5
(PEAKlinMean)
x
Add PEAKlinMean to all SWgains (PEAKlin)
x
Calculate absolute value of the difference between PEAK and PEAKlin for all
SWgains (PEAKnonlin).
x
Find the maximum SWgain for which PEAKnonlin < 0.2 (StableSPL)
Find highest gain with stable peak frequency.
x
Find the frequency corresponding to PEAK.SPL for all SWgain (PEAK.FRQ).
x
If more than more PEAK.FRQ is found the mean frequency is used.
x
From the PEAK.FRQ found at each SWgain subtract the PEAK.FRQ found at
SWgain – 2 (PEAK.FRQ.INCREASE).
x
Find the maximum SWgain for which PEAK.FRQ.INCREASE < 0.1 (Stable.FRQ)
Find maximum feasible gain
x
The lowest value of StableSPL and Stable.FRQ is selected as the maximum feasible
gain.