36
5 Considerations in Modern Multichannel Nonlinear Hearing Aids FRANCIS K. KUK Introduction The field of hearing aids has witnessed a sig- nificant evolution in the last decade. Prior to the 1980s, over 80% of hearing aids manu- factured and dispensed were linear hearing aids with peak clipping as the only option for output limiting. Many of the nonlinear hearing aids were linear hearing aids with output compression limiting. The situation changed so that in the mid-1990s almost half of the hearing aids sold were nonlinear hear- ing aids (Mueller and Killion, 1996). Most of the nonlinear hearing aids were single chan- nel linear hearing aids with compression limiting, or single channel wide dynamic range compression (WDRC) hearing aids. The sale of nonlinear hearing aids steadily increased in the last few years, with an estimate that over 75% of hearing aids sold incorporated some types of nonlinear signal processing. Of these hearing aids, three- quarters were either programmable or true digital signal processing hearing aids (Strom, 2001). Many of these hearing aids are multi- channel devices utilizing complex process- ing algorithms that extend the benefits of conventional nonlinear hearing aids. Such rapid change in technology has led to an al- most inevitable consequence—our verifica- tion protocol has failed to keep pace with the advances such that many can neither be eas- ily evaluated nor fairly verified without spe- cial considerations. The interpretation that one makes with the results of verification may also be unclear. Failure to properly ver- ify a nonlinear hearing aid could compro- mise its potential benefit. Thus, this chapter focuses on the changes in current nonlinear hearing aids, explains how they may extend the benefits provided by conventional de- vices, and discusses how our current prac- tices on verification may need to be updated to reflect the changes in technology. Al- though the concept and characterization of compression are reviewed, readers are re- ferred to summary articles for a detailed dis- cussion on the rationale, classification, fit- ting, and efficacy of conventional nonlinear hearing aids (Braida et al, 1979; Walker and Dillon, 1982; Kuk, 1996, 2000; Dillon, 1996. Characterization of Conventional Compression Hearing Aids One way to study the operation of a hearing aid is to examine its static I-O curve. This curve can be determined with a steady-state sinusoid (ANSI S3.22–1996) or with a com- posite noise signal (ANSI S3.42–1992). Be- cause of its simplicity, most reports use a 178

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Page 1: Considerations in Modern Multichannel Nonlinear Hearing Aids · 2015-08-12 · 5 Considerations in Modern Multichannel Nonlinear Hearing Aids FRANCIS K. KUK Introduction The field

5Considerations in Modern Multichannel

Nonlinear Hearing Aids

FRANCIS K. KUK

Introduction

The field of hearing aids has witnessed a sig-nificant evolution in the last decade. Prior tothe 1980s, over 80% of hearing aids manu-factured and dispensed were linear hearingaids with peak clipping as the only optionfor output limiting. Many of the nonlinearhearing aids were linear hearing aids withoutput compression limiting. The situationchanged so that in the mid-1990s almost halfof the hearing aids sold were nonlinear hear-ing aids (Mueller and Killion, 1996). Most ofthe nonlinear hearing aids were single chan-nel linear hearing aids with compressionlimiting, or single channel wide dynamicrange compression (WDRC) hearing aids.The sale of nonlinear hearing aids steadilyincreased in the last few years, with anestimate that over 75% of hearing aids soldincorporated some types of nonlinear signalprocessing. Of these hearing aids, three-quarters were either programmable or truedigital signal processing hearing aids (Strom,2001). Many of these hearing aids are multi-channel devices utilizing complex process-ing algorithms that extend the benefits ofconventional nonlinear hearing aids. Suchrapid change in technology has led to an al-most inevitable consequence—our verifica-tion protocol has failed to keep pace with the

advances such that many can neither be eas-ily evaluated nor fairly verified without spe-cial considerations. The interpretation thatone makes with the results of verificationmay also be unclear. Failure to properly ver-ify a nonlinear hearing aid could compro-mise its potential benefit. Thus, this chapterfocuses on the changes in current nonlinearhearing aids, explains how they may extendthe benefits provided by conventional de-vices, and discusses how our current prac-tices on verification may need to be updatedto reflect the changes in technology. Al-though the concept and characterization ofcompression are reviewed, readers are re-ferred to summary articles for a detailed dis-cussion on the rationale, classification, fit-ting, and efficacy of conventional nonlinearhearing aids (Braida et al, 1979; Walker andDillon, 1982; Kuk, 1996, 2000; Dillon, 1996.

Characterization of ConventionalCompression Hearing Aids

One way to study the operation of a hearingaid is to examine its static I-O curve. Thiscurve can be determined with a steady-statesinusoid (ANSI S3.22–1996) or with a com-posite noise signal (ANSI S3.42–1992). Be-cause of its simplicity, most reports use a

178

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CHAPTER 5 ■ CONSIDERATIONS IN MODERN MULTICHANNEL NONLINEAR HEARING AIDS 179

2000-Hz sinusoid as the stimulus to illus-trate the static and dynamic properties of acompression hearing aid. Figure 5–1a showsan I-O curve of a simple compression hear-ing aid to illustrate the terminology.

Static Characteristics

gain (g)

Gain of a hearing aid is calculated as thedifference between the output level andthe input level prior to saturation (or the out-put level beyond which distortion productslike harmonic and intermodulation distor-tion product are present). For a compressionhearing aid, a fixed gain value may be dif-ficult to establish because its gain decreaseswith increasing input level. To meet qualityassurance purposes, most manufacturers re-port the maximum gain of the compressionhearing aid. This is usually the gain at thelow compression threshold. An I-O curve(ANSI, 1996), or a series of output frequencygain curves at different input noise levels(ANSI, 1992) will be needed to determinegain at various inputs. Figure 5–1a shows anI-O curve and Figure 5–1b shows an input-gain (I-G) curve. These figures show thatgain is 30 dB (90 dB � 60 dB = 30 dB) belowan input of 60 dB SPL. However, gain be-comes 25 dB at an input of 70 dB SPL and 20

dB at an input of 80 dB SPL. It is easy to seethat compression is similar to gain reduction.

compression threshold (ct)

CT represents the lowest input sound pres-sure level (SPL) at which the hearing aid re-duces its gain. This is frequently described asthe kneepoint on the I-O curve. For conven-tional compression hearing aids, linear pro-cessing is typically employed below the CT.In conventional input compression hearingaids where the volume control of the hearingaid is placed after the compressor (Kuk,1996), there is usually one CT that typicallyranges from 40 to 60 dB SPL. In Figure 5–1a,bthe CT is at 60 dB SPL. As indicated earlier,the CT represents the point of maximumgain on an input compression hearing aid.

compression ratio (cr)

CR represents the ratio of the change in inputlevels to the corresponding change in outputlevels when the SPL of the input is within thecompression range (i.e., CR = change ininput/change in output). The nominal CR iscalculated when the output reaches its steady-state value. The CR can also be calculated asthe inverse of the slope of the I-O curve withinthe compression range. In Figure 5–1a, the CRis calculated to be 2:1 [(90 � 60)/(105 � 90) = 2],

Figure 5–1. (a) A simple static input-output curve of a 2:1 input compression hearing aidwith the compression threshold at 60 dB SPL (left). (b) An input-gain curve showing thesame electroacoustic characteristics is shown on the right.

(a) (b)

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180 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

whereas the slope of the I-O curve is 0.5 [(105 �90)/(90 � 60) = 0.5].

compression range

Compression range represents the range ofinput at which the hearing aid is in compres-sion. In this case, it is 30 dB (60 to 90 dB SPL).

Dynamic Characteristics

attack and release times

Compression circuits operate through theuse of feedback loops. The time taken for thecompression hearing aid to change from lin-ear gain to within 3 dB of the compressedsteady state (i.e., reduced gain) is called theattack time. The time taken for the compres-sion hearing aid to return from the com-pressed state to within 4 dB of the linearsteady state (i.e., full gain) is called therelease time (ANSI, 1996).

The ANSI S3.22–1996 standard (ANSI,1996) specifies that attack and release timesbe measured with a sinusoid (typically 2000Hz) that abruptly changes from 55 to 90 dBSPL. The “duration of each of the levels shallbe sufficiently long so as to not influence themeasured attack and release times” (ANSI,1996, p. 11). The use of two signals with dif-ferent steady-state duration is required totest compression hearing aids with adaptiveand/or multiple release times. The assump-tion for using the 55 to 90 dB SPL signal levelis that 90 dB SPL is typically above the com-pression threshold and 55 dB SPL is typi-cally below the compression threshold. Theassumption for the low input level is nottrue for many digital nonlinear hearing aids.

Differentiating among ConventionalHearing Aid Circuits

Compression hearing aids are distinguishedfrom each other by differences in their staticand dynamic characteristics. As a preludeto the discussion on how technology has im-proved nonlinear hearing aids, it is importantto review how these compression parametersare chosen to achieve various functions. Dil-

lon (1996) provides an excellent tutorial onsuch classification as well as the efficacy foreach type of compression circuit.

Linear—No Compression-PreservingInput Intensity Difference

Linear processing refers to an amplificationscheme in which the same amount of gain isapplied to the range of inputs of interest. Fig-ure 5–1b shows that the range of input from 0to 60 dB SPL receives linear processing be-cause an input of 40 dB SPL receives thesame amount of gain (30 dB) as an input of60 dB SPL. A corollary is that for a givenchange in input range (e.g., 0 to 60 dB, or the60-dB range), the same change in outputrange occurs (e.g., 30 dB to 90 dB, or 60-dBrange for a 30-dB gain). Because the same in-tensity range is seen in the input and the out-put, a major advantage of linear processing isits preservation of the intensity difference (orthe waveform envelope) of the input signal.Such difference can be an important speechcue, especially for patients with more than amoderate degree of hearing loss (Boothroydet al, 1988; Dreschler, 1989; Van Tasell, 1993).

Gain on most linear hearing aids is ad-justed for a medium conversational (around65 dB SPL) input (Skinner, 1988). High-input-level sounds (>70 dB SPL), when added tothe fixed gain of a linear hearing aid, maysound uncomfortably loud. Although onecan limit the output of the hearing aid tobelow the wearer’s uncomfortable listeninglevel (UCL) using a peak clipping circuit,this method creates high distortion, whichmay affect sound quality and speech intelli-gibility. On the other hand, low-input-levelsounds (<50 dB SPL) might not be audible.Thus, a volume control (VC) has been madeavailable to allow gain adjustment in differ-ent environments.

Compression Limiting—High CT (>70 dB SPL), High CR (>8:1), Short Attack (<10 ms)—MaintainingComfort While Minimizing Distortion

One use of compression in hearing aids is tolimit the maximum output through the use

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CHAPTER 5 ■ CONSIDERATIONS IN MODERN MULTICHANNEL NONLINEAR HEARING AIDS 181

of compression limiting (CL) or output com-pression limiting (OCL). Instead of provid-ing the same fixed gain at all input levels, alinear hearing aid with CL provides thesame fixed gain up to a relatively high inputlevel (usually above an input of 70 dB SPL,or an input that results in an output levelthat is just below the wearer’s UCL). There-after, drastic gain reduction occurs (highcompression ratio, usually greater than 8:1)so that high input sounds will neither be-come too loud nor saturate the hearing aidto result in distortion. It should be notedthat CL is an approach to limit the output ofhearing aids. As such, it has a very short at-tack time (less than 10 msec) and a moder-ately long release time (around 100 to 200msec). Typically, it is used in linear hearingaids and is placed at the output stage of thehearing aids. Thus, it has also been called anautomatic gain control output (AGC-O) cir-cuit. Recently, such a circuit has also beenplaced at the input stage of hearing aids tominimize distortion at the input, as well asat the output stage of nonlinear hearing aidsusing WDRC. Thus, the term AGC-O is lessmeaningful to characterize a compression-limiting circuit than it once was.

Hearing aids with compression limitinghave been used commercially since the1950s (Davis et al, 1947; Hudgins et al, 1948;Lynn and Carhart, 1963). The general re-search suggests that compression limiting ispreferable over peak clipping circuits (e.g.,Dillon, 1988; Dawson et al, 1991; Hawkinsand Naidoo, 1993). This is probably becauseof its ability to reduce distortion while main-taining the temporal and spectral integrityof the signal most of the time (because of lin-ear processing in the hearing aid). Dillon(1988) suggested that patients with a mild-to-moderate degree of hearing loss wouldmost likely enjoy such benefits. Addition-ally, patients with a severe-to-profoundhearing loss who have sufficient residualfrequency selectivity may also benefit fromreduced distortion as long as sufficient out-put is available from the hearing aid. Killionand Fikret-Pasa (1993) also recommendedcompression limiting for their type III pa-

tients who have a severe-to-profound senso-rineural hearing loss.

Wide Dynamic Range Compression(WDRC)—Low CT (<60 dB SPL), Low CR(<4:1), Short Attack (<10 msec) and Release(<50 msec)—Ensuring Audibility andComfort Without VC Adjustment

Another difficulty with linear amplificationis that soft sounds may not be audible all thetime unless the wearer actively adjusts theVC on the hearing aids. This could happenwith a change in vocal effort of the speaker(Pearsons et al, 1977), or when the speaker–listener distance increases. Although an adultlistener may be able to adjust the VC on thelinear hearing aid to compensate for thechange in input levels, it is unreasonable toexpect a very young child to be able to dothe same. Unfortunately, it is more likely tofind substantial changes in the input levelsof a child’s environment than of an adult’s(Stelmachowicz et al, 1993).

To ensure audibility for soft sounds whilemaintaining intelligibility and comfort formedium level and loud sounds, a hearingaid must provide sufficient gain so that softsounds, when amplified, exceed the thresh-old of the wearer. The same amount of gainmust be reduced for louder sounds to mini-mize discomfort at the higher input levels.Thus higher gain (than linear) should beprovided for soft sounds, and gain reduction(or compression) should occur above thelevel of soft sounds so that audibility, intelli-gibility, and comfort can be ensured at allinput levels. Nonlinear hearing aids withsuch characteristics are often called wide dy-namic range compression (WDRC), full dy-namic range compression (FDRC), or en-hanced dynamic range compression (EDRC)hearing aids. Typically, these devices pro-vide more gain than linear hearing aids forsoft sounds. In addition, the compressionthreshold occurs at a relatively low inputlevel (typically below 60 dB SPL) and at alow to modest rate (i.e., low compressionratio below 3–4:1). Because gain reductionoccurs at a low input level, the risk of satura-

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182 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

tion distortion at high input/output is alsominimized. One should note, however, thatsaturation distortion can still occur at veryhigh inputs (>85 dB SPL). Thus, some high-end programmable and digital WDRC hear-ing aids also incorporate CL at its outputstage to further minimize distortion.

The choice of attack time and release timeon a WDRC hearing aid is critical. If one de-signs a WDRC hearing aid to model after thephysiology of the cochlea, the attack timeand release time should be short becausephysiologic data of the outer hair cell (OHC)functions and basilar membrane motion sug-gest a very fast, active compressive mecha-nism in the healthy cochlea (Brownell et al,1985; Ruggero and Rich, 1991). Furthermore,studies comparing the growth of loudnessperception of short versus long sounds alsosuggest a relatively fast compressive mecha-nism (Buus, 1999).

A relatively short attack time and short re-lease time (or fast-acting compression) arealso logical if the purpose of the WDRC is toprovide automatic gain adjustment to followclosely the intensity variation within thespeech signal. For example, a compressionhearing aid designed to follow the intensityvariation at a phoneme level should use re-lease time that is less than 80 msec (Plomp,1983). Finally, a short attack time and a shortrelease time are necessary to achieve thestated static compression ratio on the input-output curves (Kuk, 1998b). To provide theextra audibility for soft speech sounds, thesehearing aids typically use a short attack time(less than 10 msec) and a short release time(less than 100 msec).

Research data on the efficacy of WDRChave been mixed. Braida et al (1979), basedon a review of research data from 1952 to1976, concluded that there was no conclusiveevidence to suggest this type of compressionis beneficial for patients with sensorineuralhearing loss. Walker and Dillon (1982), basedon a review of the literature up to that time,reported that there may be some “tentativesupport” for WDRC. Hickson (1994), basedon a review of later research findings, re-

ported more positively on the efficacy ofsuch a circuit design. Dillon (1996) surveyedthe literature up to that time and concludedthat WDRC improved speech recognition inquiet over a linear hearing aid at a low inputlevel and when subjects were not allowed toadjust the VC on the hearing aids.

The difference in performance betweenWDRC and linear amplification may be ex-plained from the standpoint of audibility.Figure 5–2 compares the amount of gainprescribed for an impaired ear between aWDRC hearing aid (dotted line) and a linearhearing aid (solid line). To restore “normal”loudness, the WDRC hearing aid suppliesmore gain at lower input levels and less gainat higher input levels when gain at a conver-sational level (65 dB SPL) is similar for bothhearing aids. If no volume adjustment wereallowed, the higher gain from the WDRChearing aid at a low input level wouldensure better audibility, and thus higherspeech recognition scores. Because less gainis available to a WDRC hearing aid than to alinear hearing aid at high input levels, awearer of such device may find loud soundsmore bearable with WDRC. This may ac-

Figure 5–2. Hypothetical input/output curvescomparing the difference in output between a lin-ear hearing aid (solid line) and a wide dynamicrange compression (WDRC) hearing aid (dottedline) when both are adjusted to have the sameoutput to a conversational input (65 dB SPL).

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CHAPTER 5 ■ CONSIDERATIONS IN MODERN MULTICHANNEL NONLINEAR HEARING AIDS 183

count for the increased comfort associatedwith the use of WDRC hearing aids in dailyenvironments (e.g., Moore et al, 1992).

Automatic Volume Control (AVC)—Moderate CT (60–75 dB SPL), Low-Moderate CR (4–6:1), Long Attack (>20 msec) and Release (>1 s)—Preserving the Input Range Within theWearer’s Residual Dynamic Range

One negative consequence of providing moregain to low input sounds and less gain tohigh input sounds at a fast rate is a reduc-tion of the intensity contrasts inherent in thespeech signal. This effect is termed smearing.Its effect can be examined in both the tempo-ral domain and the spectral domain. Thiscould affect intelligibility, especially of intel-ligibility of speech in noise. It would bedesirable to receive the audibility and com-fort benefits offered by a WDRC hearing aidwithout altering the intensity relationship ofthe input signals. This can be achievedthrough the use of a longer attack time and alonger release time.

effective compression ratio

The static or nominal compression ratio that isdetermined by measuring the steady-state re-sponse of the compression circuit to sinusoids(i.e., nominal CR) is only achievable whenvery short attack times (less than 5 msec) andrelease times (less than 30 msec) are used. Theeffective (or real-life) compression ratio is af-fected by factors that include the precom-pressor gain on the hearing aid, the peak-to-valley ratio of the input signals, the overallinput level, the time constants of the com-pression circuit, and the intervals betweendifferent acoustic peaks (Blesser, 1969; Stoneand Moore, 1992). Because daily acousticsignals (including speech) fluctuate greatlyin intensity levels, peak-to-valley ratios, andinterstimulus intervals, such factors can in-teract with the time constants of the com-pression hearing aid to yield a lower effec-tive compression ratio than the nominalvalue. Kuk (1996, 1998a) provides a quanti-

tative illustration of how the release timechanges the nominal compression ratio. Thelonger the release time (for the same fixedintersyllabic interval), the more likely theeffective compression ratio regresses towardone (1:1, or linear).

subjective preference for longerrelease time

The “linearization” of compression hearingaids with a long release time raises potentialconflicts with the rationale for WDRC. How-ever, subjective data on the effect of releasetimes seem to support their use. Neuman etal (1995) showed that hearing-impaired sub-jects preferred a longer release time (1000msec) in some noise backgrounds in a singlechannel compression system. Hansen (2001)also reported that both normal hearing andhearing-impaired subjects preferred a longerrelease time (4000 vs. 40 msec) on a -octaveband multichannel compression hearing aid.Verschuure et al (1996) also showed similarfindings. If the supposition that a longer re-lease time could result in a lower effectivecompression ratio were true, these studiessuggest that hearing-impaired people prefera lower compression ratio than a higher one.It is of interest to note that Neuman et al(1994) reported that hearing-impaired sub-jects preferred compression hearing aidswith a lower compression ratio (1.3) than ahigher one (>3).

need for the nominal or staticcompression ratio

The concept of the effective compressionratio raises the question of whether theamount of compression is adequate for thewearer if a longer release time were used.After all, the static CR is estimated to “com-press” the range of acoustic stimuli into theresidual dynamic range (DR) of the hearing-impaired patient. For example, if one as-sumes that the range of input is 100 dB andthe hearing-impaired patient has a dynamicrange of 50 dB at all frequencies, a CR of atleast 2:1 is necessary. A compression ratio

13

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184 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

smaller than 2:1 would mean some of theinput would be beyond the wearer’s DR. Asmaller CR is only acceptable (in theory) ifthe range of sounds in the environment isequal to or smaller than the residual DR ofthe hearing-impaired wearer.

Although environmental sounds canrange from a very low level (lower than 0 dBSPL) to a very high level (over 100 dB SPL),the range of intensity in any single listeningenvironment is smaller than the assumed100 dB SPL. Indeed, Pearsons et al (1977)measured the range of intensity variationsin many listening environments (inside ahome, in a department store, etc.) and for thesame persons speaking at different vocal ef-forts. They showed that the intensity rangewas typically 30 to 40 dB in almost all ofthe measured conditions. This suggests thatrather than assuming a 100-dB range of inputlevels, a more typical range should be 30 dBin any listening environments. Other re-searchers also advocated similar ranges(e.g., Boothroyd et al, 1988; Skinner, 1988).

The residual DR of the hearing impairedpatient is dependent on the wearer’s degreeof hearing loss and the individual’s toler-ance limit, i.e., UCL. Although individualdifferences exist, Pascoe (1988) showed thaton average, UCL increased as the hearingloss increased. Although the residual DR de-creased with increasing hearing losses, theresidual DR was typically greater than 30 dBfor hearing losses smaller than 80 to 90 dBHL. Figure 5–3 shows the relationship be-tween thresholds, most comfortable listen-ing levels (MCLs), and UCLs.

These two pieces of evidence suggest thatfor the average patient with less than a 80 to90 dB hearing loss, wide dynamic range com-pression, while capable of achieving extra au-dibility for soft sounds and comfort for loudsounds, is not necessary to transpose therange of environmental sounds into the pa-tient’s residual dynamic range. This is be-cause the residual DR is typically larger thanthe range of acoustic inputs in a single listen-ing environment. Although AGC is necessaryto achieve audibility of the softer sounds(e.g., less than 50 dB SPL) and comfort for the

louder sounds (e.g., higher than 75 to 80 dBSPL) without VC adjustment, it may not benecessary to use fast attack and release timesin order to preserve the nominal or staticcompression ratio. Indeed, it may not even besubjectively desirable (Neuman et al, 1994,1995). Once the overall gain level of the hear-ing aids has been automatically adjustedbased on the input level, it can actually re-main in a “linear processing” mode so thatthe intensity range of sounds from the spe-cific listening environment is maintained.Kuk (1998b) provides a detailed explanationfor the rationale of using long attack and re-lease times in WDRC hearing aids.

Compression hearing aids with a longattack time and a long release time have beencalled automatic volume control (AVC) hear-ing aids in the past. These hearing aids havea moderately high CT (around 60 to 70 dBSPL) and a moderate compression ratio (lessthan 5:1). Even though the release times areclassified as long, many are around 1 secondin duration. Recently, WDRC hearing aids

Figure 5–3. Relationship between thresholds,most comfortable listening level (MCL), and un-comfortable listening level (UCL) expressed in dBHL. [From Pascoe D. (1988). Clinical measure-ments of the auditory dynamic range and their re-lation to formulas for hearing aid gain in presbya-cousis and other age related aspects. In: Jensen JH,ed. Proceedings of the 13th Danavox Symposium.Danavox, Denmark: pp. 129–147. With permissionfrom Danavox Jubilee Foundation, Denmark.

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CHAPTER 5 ■ CONSIDERATIONS IN MODERN MULTICHANNEL NONLINEAR HEARING AIDS 185

with extra long attack and release times havealso become available. To distinguish fromthe traditional AVC hearing aids, which typi-cally use an intermediate compression thresh-old, these hearing aids are called slow-actingwide (or enhanced) dynamic range compres-sion (SA-WDRC) hearing aids to reflect thelower CT in these hearing aids.

Relatively few studies were conductedto examine the efficacy of AVC hearing aids.Those that did, supported the use of AVC,especially at a high noise input (King andMartin, 1984; Neuman et al, 1994). Becausefast-acting multichannel compression sys-tems introduce temporal and spectral smear-ing (Boothroyd et al, 1988; Van Tasell, 1993)and patients with a hearing loss need consis-tent audibility and as many speech cues aspossible over a wide range of input withoutadjusting the VC, Byrne (1996) suggested theuse of slow-acting compression instead offast-acting WDRC to minimize any potentialsmearing while allowing for automatic gainadjustment.

Results of Recent Advances inNonlinear Technology

More Complex Input-Output Curves

The use of digital signal processing (DSP) inhearing aids allows the implementation of

various forms of I-O functions in currentnonlinear hearing aids. Although many DSPhearing aids still use I-O curves that suggestbasic functions, many use advanced algo-rithms that result in complex I-O curves.They have several functional units and mayinclude many that were described in the pre-vious sections. One example of a complexI-O curve is discussed below. There may beother complex I-O curves as well.

Figure 5–4 shows the I-O curve at the 2000-Hz channel of a 15-channel DSP hearing aid.Instead of having the linear-compression-peak clipping segments that were describedin Figure 5–1a, this I-O curve has five seg-ments, each with a different compressionratio and each serving a different function.From an input level of 0 to 2.5 dB SPL, theoutput of the hearing aid increases from 42.5to 47.5 dB SPL. This represents a gain changefrom 42.5 dB at an input of 0 dB SPL to a gainof 45 dB at an input of 2.5 dB SPL. A regionof the I-O curve that shows an increase ingain as input level increases is termed an ex-pansion region. In this case, the region from 0to 2.5 dB SPL illustrates expansion. If oneuses Walker and Dillon’s (1982) approach tocalculate the compression ratio in this region(change in input range divided by thechange in output range), one obtains a com-pression ratio of 2.5/5 or 1:2. In other words,a compression ratio less than 1 describes an

Figure 5–4. Example of a com-plex input-output curve show-ing expansion, linear, and com-pression functions at differentinput levels. The range of com-pression (from 7.5 to 90 dB SPLor 82.5 dB) is larger than the typ-ical 40 to 50 dB range seen inconventional compression hear-ing aids.

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186 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

expansion circuit. On the I-O curve, an ex-pansion circuit also has a slope (change inoutput divided by change in input) that isgreater than 1. Although an expansion cir-cuit can be achieved using analog means,currently only DSP hearing aids utilize suchcircuitry.

As one moves from 2.5 to 7.5 dB SPL, onenotes that the output changes from 47.5 to52.5 dB SPL. The sameness in input range (5dB) and output range (5 dB) defines linearprocessing. This region is included to pro-vide a smooth gain transition as input in-creases from 2.5 to 7.5 dB SPL.

As the input increases from 7.5 to 52 dBSPL (average conversational level at 2000Hz), the output increases from 52 to 70 dBSPL. Because the change in input range (44.5dB SPL) is greater than the change in outputrange (18 dB SPL), compression is indicatedin this input region. In this case, the com-pression ratio is 2.5:1. This is called low-level compression (LLC) to designate thefact that the CT is significantly lower thanconventional CT (at around 40 dB SPL orhigher), and the fact that the same CR is ap-plied only up to the conversational level.

As one increases the input from 52 to 90 dBSPL, output increases from 70 to 100 dB SPL.This translates to a compression ratio of 1.3:1.This is called high-level compression (HLC)to distinguish from compression limiting,which typically starts at a high input levelabove 60 to 70 dB SPL. In this case, HLCstarts at the conversational level and contin-ues until a high input level is reached. Thecompression ratio is closer to 1 and is smallerthan the compression ratio for the LLC. Thisis to approximate the freedom from recruit-ment seen in normal loudness growth func-tions at high input levels (Hellman andMeiselman, 1990) and to preserve the inten-sity contrasts in the high input signals.

The output above an input of 90 dB SPLdoes not change significantly as the input in-creases. Indeed, the output increases to 101dB as the input increased above 100 dB SPL.This corresponds to a CR of 10:1. This is thecompression limiting (CL) circuit and it

serves the purpose of minimizing saturationdistortion while maintaining comfort.

The complex I-O curve in Figure 5–4 re-veals several advances in nonlinear technol-ogy over the simpler designs. First, a com-pression threshold as low as 7.5 dB SPL ispossible. This is in contrast to the typical 40dB SPL seen in conventional designs. Sec-ond, an expansion region is available belowthe compression threshold. Typically in con-ventional design, a linear circuit is used in-stead. Third, the range of compression (fromCT to compression limiting) is significantlyincreased. In this case, it is 82.5 dB SPL (from7.5 to 90 dB SPL) instead of the 40 (from 50 to90 dB SPL) to 50 (from 40 to 90 dB SPL) dBSPL range seen in the conventional compres-sion hearing aids. Fourth, more than onecompression ratio, instead of one fixed CR,is seen in the compression range.

The advantages of having more complexI-O characteristics can be evidenced in atleast three areas, as described in the follow-ing subsections.

audibility with minimal increase incircuit noise—expansion

The rationale for expansion (or gain in-creases as input increases) is to minimize theamount of amplification for sounds belowthe compression threshold. Typically, com-pression hearing aids utilize linear process-ing below the compression threshold. Inputsbelow the CT will be amplified maximallyand may become audible to wearers whohave normal hearing or a mild hearing loss.

Figure 5–5 illustrates the effect of an ex-pansion circuit on the audibility of circuitnoise. In most hearing aids, the audible cir-cuit noise originates from the microphonethat is used. Because the microphone isplaced at the front end of the amplifier cir-cuit, any noise that it generates will be am-plified by the gain of the hearing aid. If themicrophone has a noise level of 20 dB SPL, ahearing aid with 40 dB gain at and below theCT will amplify the noise level to 60 dB SPL.On the other hand, if a 1:2 expansion circuit

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is used below the CT, a gain of 20 dB will beat the 20 dB SPL input instead. The ampli-fied microphone noise will be 40 dB SPL in-stead. This represents a 20-dB reduction ofmicrophone noise level. Depending on thehearing acuity of the hearing aid wearer, thisnoise level may not be audible.

Because an expansion circuit allows theaudibility of soft sounds without the associ-ated circuit noise that may be common withlinear processing, an associated advantageof expansion circuit is to allow hearing aidmanufacturers to lower the CT further withminimal artifacts from circuit noise. As mi-crophone noise level decreases in the future,one may be able to achieve a lower CT withor without an expansion circuit.

more specificity in intensity processingand fine-tuning

Nonlinear WDRC hearing aids are fine-tuned differently than linear hearing aids.Because linear hearing aids provide thesame gain at all input levels, adjustment ofthe gain parameter would affect sounds atall input levels. In contrast, nonlinear com-pression hearing aids offer gain adjustment

in at least three input levels separately: soft,normal (or conversational), and loud. Thepresence of these parameters allows fine-tuning at specific input levels and minimizesthe occurrence of side effects. Kuk (2001a)provides a summary of examples of howthese parameters may affect subjective pref-erence and how subjective comments can berelated to electroacoustic parameters on athree-channel compression hearing aid.

For example, the I-O curve shown in Figure5–6 illustrates the specificity of adjustment of-fered by current multichannel compressionhearing aids. Figure 5–6a shows the effect ofadjusting the gain parameter for low-level in-puts. As the gain parameter is increased (dot-ted line), the effect is seen at the very lowinput level. The effect decreases as input in-creases toward the conversational level. Thissuggests that adjustment of this parameterwill affect only soft sounds, and will not af-fect conversational and loud sounds. Figure5–6b shows the effect of adjusting the gainparameter for loud sounds. Its effect is pri-marily restricted to input levels above theconversational level. Its effect increases asinput level increases. Figure 5–6c shows theeffect of adjusting the gain parameter for nor-mal conversational input. In contrast to theparameter for soft and loud sounds, the effectis different depending on the extent of adjust-ment. As the gain parameter is increased ordecreased slightly, its effect is concentratedmostly around the conversational input lev-els. Gain at the higher and lower input levelsremained the same or decreased (dashedline). As the parameter is adjusted further,however, gain across the whole input range isdecreased (dotted line) or increased (notshown). Such diverse adjustment and in-creased specificity represents the complexprocessing used in some of the current multi-channel compression hearing aids.

lower compression threshold for evenbetter audibility at low inputs

One of the advantages of using DSP in cur-rent multichannel nonlinear hearing aids is

Figure 5–5. Two input-output curves that differin the type of processing below the compressionthreshold. The linear circuit is expected to yieldan output of 60 dB SPL for a microphone noiselevel of 20 dB SPL. In contrast, an output of 40 dBSPL is expected of the expansion circuit.

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188 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

Figure 5–6. Effect of adjustingvarious gain parameters on theoutput. (a) Effect of gain parame-ter for soft sounds. (b) Effect ofgain parameter for loud sounds.(c) Effect of gain parameter forconversational sounds.

(a)

(b)

(c)

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that a lower CT can be achieved with mini-mal artifacts like whistling as a result of feed-back. Kuk (1999) describes the rationale forcurrent DSP hearing aids and how they canachieve a low CT despite potential artifacts.

There are several advantages to a low CT.One immediate advantage is an increasein gain (and thus output) at inputs belowthe compression threshold over the same cir-cuit with a higher CT. Indeed, Kuk (1998a)showed that a lower CT is associated withhigher gain for inputs below the CT. Thiswould mean a lower (or better) aided thresh-old in sound-field measurement (Kuk et al,2001) as well as improved speech recognitionscores for low (50 dB SPL) input presentation(Gabbard and O’Grady, 2000; Kuk et al, 2001).

Lee et al (1998) also demonstrated thata lower CT may overcome some of the limi-tations of hearing aids with directional mi-crophones. These investigators studied thelimitation in audibility of directional micro-phones to sounds presented from the backand the speech in noise improvement of-fered by them. A total of 30 subjects with amoderate degree of sensorineural hearingloss participated. In the first study, subjectswere presented with soft speech (50 dB SPL)in quiet at 180 degrees. They wore a digitalhearing aid with a fixed directional micro-phone and a programmable hearing aidwith switchable directional microphones—omnidirectional and directional modes ofoperation. When the programmable hearingaid was worn in the directional mode, amean score of 13% correct was achieved.When it was worn in the omnidirectionalmode, a mean score of 37% correct wasachieved. A mean difference of 24% wasseen in the speech score. In contrast, a meanscore of 44% was achieved with the digitalhearing aid in the fixed directional modeunder the same test condition. The investi-gators attributed the difference in speechscores to the compression threshold used inthe hearing aids (50 to 55 dB SPL in the pro-grammable hearing aid versus 20 dB HL inthe digital hearing aid).

Some may question that the enhancedaudibility of speech presented from the back

could reduce the effectiveness of the direc-tional microphone. In a second part ofthe study, the effectiveness of the directionalmicrophones was compared. In this case,speech at 65 dB SPL was presented in thefront and a party noise was presented in theback at a signal-to-noise ratio (SNR) of �5dB. Rather than scoring lower with the digi-tal directional hearing aid, subjects obtainedsimilar scores between the digital fixed di-rectional hearing aid (with a low CT at 20 dBHL) and the programmable hearing aid inits directional mode (53% for the digital and45% for the programmable). This confirmedthat the digital directional hearing aid wasat least similar to, if not better than, theprogrammable hearing aid in its speech-in-noise improvement while minimally affect-ing audibility of soft sounds presented fromthe back. The difference in the presentationlevels (50 dB SPL for soft speech and 70 dBSPL for noise) and the spectra of the stimulimay account for such observations.

how low should the ct be?

If CT affects the audibility of soft sounds, alogical question to ask is how low should theCT be to ensure best audibility. In theory, ifthe purpose of a WDRC hearing aid is to en-able patients with a hearing loss to hear like anormal hearing person at 0 dB HL, then theideal CT should be at 0 dB HL. The degree ofhearing loss or the configuration of the hear-ing loss should not be factors to consider. Thisis also the position argued by Dillon (1996)and Cornelisse et al (1995). There are, how-ever, practical factors to suggest otherwise.

microphone noise floor

Because the microphone is the first activecomponent in a hearing aid, any noise fromit will be amplified by the gain on the hear-ing aid to become audible. This suggests thatpatients with a mild degree of hearing lossmay hear the thermal noise from the micro-phone and react negatively. Thus, a low CTis acceptable only if the microphone noisewill not be amplified above the wearer’sthresholds.

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There are two potential solutions to thisproblem. First, if the noise floor of the micro-phone is reduced, then its amplified level maybe below the wearer’s threshold and remainsinaudible. The typical microphones todayhave an overall noise level of around 20 to 25dB SPL, with the spectral level of around 10 to15 dB SPL in the mid-frequencies. This placesa lower limit for the CT of around 15 to 20 dBSPL in the mid-frequencies. Recent advancesin microphone technology have reduced suchnoise level to an overall level of around 10 to15 dB SPL and the spectral level under 5 dBSPL in the mid-frequencies (higher in thelows). This could potentially allow for a lowerCT in modern nonlinear compression hearingaids.

Another approach to achieve a lower CTwith minimal circuit noise is the use of an ex-pansion circuit. This has the effect of provid-ing less gain to microphone noise to result ina lower, inaudible output. Obviously, theneed for an expansion circuit is more criticalwhen the amplifier is used with a micro-phone that has a high noise floor. The needfor such circuit may lessen with the newer,quieter microphones. The limitation of themicrophone noise floor may no longer be anissue that prohibits the use of a low CT.

useful information

Despite the practical ability to achieve a CTbelow 10 dB SPL, one has to question thepractical value of having such a low CT.Clearly, the lower the CT, the lower the inputlevel one can hear. If the purpose of a nonlin-ear hearing aid is to allow the perception ofnormal speech, then the required CT wouldnot need to be lower than 20 dB HL1 across

frequencies. This is because the softest portionof normal speech is around 20 dB HL acrossfrequencies when such data are displayed onan audiogram format (Byrne et al, 1994).However, if the goal is to hear the softest por-tion of soft speech, a CT that is significantlylower than 20 dB HL should be desirable.

Although the actual value for a CT has notbeen agreed upon, the general consensus isthat compression threshold should be as lowas possible (Sweetow, 1998). In choosing acompression hearing aid, it may be worth-while to select one with the lowest CT andmake adjustment to it if necessary accordingto the subjective comments of the patientsafter they have worn the hearing aid forsome time (Kuk, 2001a).

not everyone favors a low ct

Although a low CT has the advantage ofproviding increased audibility, the extra au-dibility may not always be desirable. Barkerand Dillon (1999) studied subject preferencefor a single channel, 2:1, fast-acting WDRChearing aid that allowed variation of the CT.Subjects wore the hearing aid at a CT of 40dB SPL and a CT of 65 dB SPL for a total pe-riod of 1 month before indicating their pref-erence. Linear processing was used belowthe CT. The results of the study showed thatthe majority of the subjects preferred thehearing aid set at the higher CT (65 dB SPL)than at the lower CT (40 dB SPL). Distractionfrom the ambient noise associated with thelower CT was reported as the major com-plaint. It is conceivable that the instructionsgiven to the subjects, the type of hearing aidused (single channel, fast-acting versus slow-acting, multichannel), and the type of pro-cessing below the CT (linear versus expan-sion) contributed to the outcome of thestudy. Nonetheless, it is clear that, althougha low CT can be achieved, how it is imple-mented on a compression hearing aid candirectly influence its acceptance. It is not suf-ficient to conclude on the merit of a featurebased on its theoretical considerations; itshould be considered in the context of thewhole hearing aid design.

1This is expressed in dB HL and not dB SPL. dB HLand dB SPL are linked by the minimal audible field(MAF) or minimal audible pressure (MAP) functions. Inessence, 0 dB HL is the dB SPL needed to reach a thresh-old sensation. It is also expressed as the audiometriczero. Because the sound pressure level needed to reachthreshold varies with frequencies, expressing theneeded CT in dB HL has the convenience of the samevalue across frequencies and not discrete value for eachfrequency.

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More Specificity in Frequency Processingwith More Processing Channels

Another major change in modern nonlinearhearing aids is the availability of more in-dependent compression channels. Today’smultichannel nonlinear hearing aids varyfrom as few as two channels to as many as 20.channels. This has the advantage of increas-ing the specificity of processing in the fre-quency domain.

differences between bands and channels

A band in a hearing aid refers to a frequencyregion where gain adjustment is made. Aclose analogy of the “band” concept is theequalizer where one can make gain (or sen-sitivity) adjustment in many frequency re-gions (or bands). The nature of the signalprocessing (e.g., compression ratio) of all thebands, however, remains the same.

A channel in a hearing aid refers to a fre-quency region where the same signal pro-cessing takes place. The relationship be-tween “bands” and “channels” is illustratedin Figure 5–7, which shows a hypotheticalhearing aid that has 10 frequency regionswhere gain adjustment can be made (i.e.,bands) and three broader frequency regionswhere different ratios of compression can beapplied (i.e., channels). In this case, bands 1to 4 contribute to the “low” channel, bands 5to 7 to the “mid” channel, and bands 8 to 10

to the “high” channel. Although gain in eachof the four bands in the “low” channel maybe quite different, the change in gain (orcompression ratio) of all four bands is iden-tical because they are in the same channel.This being the case, one may conclude thatthe number of channels in a hearing aidis equal to or smaller than the number ofbands in it.

achieving desired output/gain withatypical hearing loss configurations

Because gain provided by a processing chan-nel is the same for all inclusive frequencies,the need for more channels increases asthe hearing loss configuration becomes moreirregular (e.g., an “inverted V” graph). Thisis necessary so that each frequency regionmay receive independent gain adjustmentfor speech to be delivered close to the opti-mum level. This goal can be achieved with amultiband system as well.

recruitment compensation

Because most patients with a hearing loss ex-hibit different rates of loudness growth atdifferent frequencies, different rates of gainadjustment (i.e., compression ratio) will benecessary for different frequency regions torestore normal loudness. Consequently, therequirement for independent multiple pro-cessing channels increases as the hearing lossconfiguration departs from a flat configura-tion. In theory, the more threshold changesacross frequencies, the greater is the need formore independent processing channels.

monitoring one’s own voice

When a person speaks, his/her voice mea-sured at the ears has a different spectrumthan when it is measured at a conversationaldistance of 1 m. It is about 15 to 20 dB SPLhigher in the low-to-middle frequency re-gion and about 5 dB SPL lower in the highfrequency region than that measured at aconversational distance. This is due to a com-bination of distance factor and radiationcharacteristics of vocal production (Dunn

Figure 5–7. Theoretical distinction between aband and a channel.

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192 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

and Farnsworth, 1939). If the patient wears asingle channel linear hearing aid, the exces-sive low frequency from the patient’s ownvoice could result in upward spread of mask-ing on the high-frequency consonants andrender them inaudible. However, if thespeaker wears a single channel nonlinearhearing aid, the excessive low-frequencyinput would reduce the gain of the hearingaid across the whole spectrum. Although thethreat of upward spread of masking is mini-mized, the gain of the high-frequency soundsis also reduced. The end result in both casesis that the high-frequency sounds may be in-audible. If speech production is a function ofhow well the speaker monitors his/her ownvoice, a single channel device may compro-mise the speaker’s speech production skill.

A multiple channel hearing aid is useful inthat different frequency regions can be am-plified using different compression settings.For example, one may maintain the samegain for both the high- and low-frequencyregions. However, one may increase thecompression ratio (even more gain reduc-tion at high inputs) for the low-frequency re-gion and decrease the compression ratio inthe high-frequency region to minimize thespread of masking of the low-frequencyinput signal on the high-frequency input sig-nal, and to provide separate processing be-tween the two frequency regions. The abilityof a multiple channel hearing aid in allow-ing the wearers to monitor their own voicemay be one of the reasons why improve-ment in speech production skills was seen inchildren wearing a three-channel digitalhearing aid (e.g., Kuk et al, 1999c).

occlusion management

Along the same line, the higher input of thespeaker’s voice to his/her own hearing aidswould also result in a higher output in theperson’s ear canal. This could contribute tothe perception of “hollowness” or “barreleffect” that is commonly known as theocclusion effect. In addition to many otherapproaches like venting and increasing thecanal length of the hearing aid, one commonapproach to minimize this effect is to reduce

the amount of low-frequency gain (Kuk,1990; Sweetow and Valla, 1997).

Gain reduction in the low frequency, al-though effective in minimizing the occlusionperception, may lead to undesirable results ifone fails to control the precise frequencies atwhich gain reduction occurs. Fletcher (1953)reported on the relative contribution of eachfrequency region to overall “intelligibility”and “overall loudness.” He showed that fre-quencies below 500 Hz contribute substan-tially (60%) to overall loudness perception (orthe “boomy” sensation), whereas the samefrequencies contribute less than 5% to overallintelligibility. That is, if one can reduce gainbelow 500 Hz successfully without affectinggain at higher frequencies, one could reducethe perception of hollowness with minimaleffect on intelligibility. The difficulty is theease of reducing gain below 500 Hz withoutaffecting gain above that frequency region.

In a single channel system, reducing gainbelow 500 Hz may also reduce gain above500 Hz. Indeed, the more gain reductionbelow 500 Hz, the more gain reduction at thenearby mid-frequencies. This is also a prob-lem with a dual-channel system that uses afixed crossover frequency. For example, ifthe crossover frequency between the twochannels is fixed at 1500 Hz, any gainchange in the low-frequency channel wouldaffect frequencies up to 1500 Hz. The fre-quency region between 500 and 1500 Hz,however, contributes as much as 55% tospeech intelligibility and 36% to overall loud-ness. By lowering gain in the low-frequencychannel, one would have minimized the hol-lowness perception, but speech intelligibilitymay suffer also. The ideal solution is to haveenough independent channels so that gainmodification at 500 Hz and below does not af-fect gain in the nearby mid-frequencies. Alter-natively, movable crossover frequencies maybe used to vary the bandwidths of each pro-cessing channel.

noise reduction algorithm—analysisand implementation

Many DSP hearing aids today include noisereduction algorithms that are capable of (1)

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identifying “speech” and “noise”, and (2)differentially amplifying “speech” and“noise.” Although there are many differentalgorithms that can achieve these goals,their effectiveness depends in part on howwell they can distinctly identify speech andnoise signals, and how well noise is attenu-ated without affecting speech. The numberof channels in the DSP hearing aid sets alimit on its effectiveness.

The number of channels in the hearing aidcan limit the resolution of the noise analysisand its treatment. In general, increasing thenumber of channels (thus narrowing band-width) increases the precision of the analy-sis. When a broad bandwidth is used (fewerfilters), details of the speech signal may bemissed. When a narrower filter is used(many filters), details of the signal would beretained for a more precise analysis.

The precision of the speech/noise treat-ment can be improved by increasing thenumber of independent processing chan-nels. Imagine the same dual-channel hearingaid with a crossover frequency at 1500 Hz.Let’s assume that a narrow band of noisethat is centered at 500 Hz occurs at the sametime as a female talker who has a peak in herspeech spectrum around 1000 Hz. The resultof the analysis calls for a gain reduction inthe low-frequency channel. However, themagnitude of the 500 Hz noise and the desir-able speech around 1000 Hz will be attenu-ated. The result is a diminution of noiseas well as speech. Thus the SNR may notimprove.

However, in a 15-channel system, only the500-Hz channel would incur gain reduction.This will spare the speech information at1000 Hz. In addition, the overall loudness ofthe output is lowered because of gain reduc-tion at 500 Hz. This would increase listeningcomfort while preserving intelligibility. Forindividuals who may be easily affected bythe spread of masking, this may even im-prove their speech recognition ability as thespread of masking is minimized throughgain reduction.

Despite the many potential advantages of-fered by multiple channel hearing aids, oneinherent limitation of having many indepen-

dent channels is the increased risk of tempo-ral and spectral smearing. Such risks increasedramatically as the number of independentchannels increases, and especially in a fast-acting (i.e., short attack time and releasetime) WDRC hearing aid.

Figure 5–8 shows the effect of the numberof channels on the spectral level of the vowel/i/. The formants of the vowels are indi-cated by the peaks in the figure. The inten-sity difference between the formants and be-tween the peak and the “trough” serves ascues for speech identification. As one caneasily see, a two-channel compression sys-tem causes a small reduction of the spectraldifference. As the number of channels in-creases to 10, however, significantly more re-duction of peak-to-trough difference is seen.That is, spectral smearing increases as thenumber of channels increases.

Figure 5–9 shows the correspondingsmearing (or contrast reduction) in the tem-poral or intensity domain. Again, the inten-sity difference between the peak and the val-ley of the waveform (i.e., envelope) couldserve as cues for speech understanding. Inthis case, the reduction (or smearing) wors-ens as the number of channels increases. Be-cause patients with more than a moderate

Figure 5–8. Example of spectral smearing of thevowel /i/ caused by a 2-channel and a 10-channelWDRC hearing aid with a short release time(50 msec). The original formant is included forcomparison.

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degree of hearing loss rely more heavily onthe temporal contrasts for speech under-standing, such smearing could also affectspeech understanding (Van Tasell et al, 1987).

Several studies have examined the effectof the number of compression channelson speech recognition. Many showed a de-crease in speech recognition scores inthe multichannel device or as the numberof channels increases in the multichannelhearing aid (Bustamante and Braida, 1987;Plomp, 1988; Drullman et al, 1996). How-ever, several authors (e.g., Mangold andLeijon, 1979; Moore et al, 1985; Yund andBuckles, 1995a–c) showed improvement inspeech recognition scores with multichannelcompression processing. For example, Yundand Buckles (1995b) showed that speechrecognition score improved significantly(48% vs. 52%) as the number of channelsin a compression system increased from fourto eight. Interestingly, the same authors(Yund and Buckles, 1995c) in a related studyshowed that speech recognition for an eight-channel compression system was not signifi-cantly different from that for a linear hearingaid (50.4% vs. 51.6%). Clearly, the findingson the impact of the number of channels onspeech recognition are mixed at this time.Although there may be theoretical advan-tages with using more channels, one may

encounter diminishing returns as the num-ber of channel increases. Worse yet, negativeartifacts (e.g., smearing), may become morenoticeable if careful consideration is notpaid to the design process. The evaluationprotocol, the subject population, and thedesign of the multichannel compressionhearing aid (e.g., slope of filters, channelcoupling, etc.) may lead one to different ob-servations and conclusions.

More Sophisticated Noise Reduction Algorithms

Conventional compression hearing aids havebeen used for noise reduction purposes (Dil-lon, 1996). By definition, noise reductionhearing aids are those that are capable ofchanging their frequency gain characteristicsas a result of the speech/noise analysis. Inessence, noise reduction hearing aids arecompression hearing aids that involve notonly an overall gain change, but also a fre-quency response change. Examples of con-ventional compression hearing aids thatserve as noise reduction hearing aids are thebass increase at low levels (BILL) and treble in-crease at low levels (TILL) circuits, which de-crease the low-frequency gain (for BILL) andhigh-frequency gain (TILL) as input level in-creases (Killion et al, 1990). In these two ex-

Figure 5–9. Examples of tem-poral (or intensity envelope)smearing as a consequence ofchannel effect. The original sig-nal is indicated for comparison.The middle display is the resultof a single-channel linear hearingaid. The bottom display is the re-sult of a 10-channel WDRC hear-ing aid.

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amples, the frequency response as well asthe overall gain of the hearing aid change asthe input level changes. The assumption for“noise” is the high overall input level.

Instead of using the overall level of theinput to signal gain reduction, many currentdigital hearing aids use sophisticated algo-rithms to estimate the nature of the inputsignal. One approach that is used by severalcommercial digital hearing aids is the use ofthe modulation index of the input as a basisfor speech/noise estimation. The rationaleis that speech modulates greatly over itscourse (e.g., Plomp, 1983) whereas a typicalnoise does not vary greatly in its intensityover time. This is seen on the left of Figure5–10, which shows the waveform envelopesof a speech segment (top left) and a car noise(bottom left). It is clear that over the time pe-riod sampled, the intensity of the speech

segment varied significantly, whereas thenoise did not show much variation. Thus,speech has a high modulation index, whereasnoise has a low modulation index. By deter-mining the modulation indices of the inputsignals, one can determine if it is speech ornoise so that appropriate gain adjustmentmay be made afterward. In general, gain ad-justment (typically reduction) beyond thatfrom compression will be made in the fre-quency channels where a noise is detected.No gain changes beyond that from compres-sion will be made if speech is detected.

An alternative approach to estimate the na-ture of the input (i.e., speech or noise) is to de-termine its level-distribution function, which isa plot of the percentage of time sounds withina specific frequency channel occur at a partic-ular intensity (Ludvigsen, 1997). For a stimu-lus that does not vary greatly in its intensity

Figure 5–10. Examples of waveforms showing high modulation (speech, top left) andlow modulation (car noise, bottom left). On the right are the corresponding level distribu-tion functions of the same listening environments. Note the bimodal distribution of speechand unimodal distribution of car noise.

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over time (i.e., noise), it will appear to have aunimodal distribution on a level distributionfunction. This is the same as saying that a par-ticular intensity level has a high frequency ofoccurrence. By the same token, because aspeech signal shows intensity fluctuation overtime, its level distribution function will have amultimodal distribution to reflect that differ-ent intensity levels occur in the input or thatsignificant fluctuations in intensity occurred.For the most part, the use of modulation indexor level distribution function yields similar es-timation of speech or noise. The right panel ofFigure 5–10 shows the level distribution func-tions for the speech and noise inputs.

The use of the results of a statistical detec-tor instead of a level detector (peak-to-peakor root-mean-square) to signal the onsetof gain adjustment at different frequenciesmeans that the static I-O curve is only par-tially reflective of the actual performanceof the hearing aids in real life. More specifi-cally, it is restricted only to speech condi-tions and not to noise conditions. In additionto a different real-life gain, the concept ofcompression threshold (onset of gain reduc-tion) would need to be modified by the na-ture of the input level as well as its overalllevel. This has a significant impact on howsuch hearing aids are evaluated today, aswill be discussed in the following sections.

Need to Reconsider Current VerificationStrategies with Nonlinear Hearing Aids

The advances in multichannel nonlinearhearing aid technology raised issues on itsoptimal verification and validation. Specifi-cally, because many of the evaluative toolsthat are used today have their origins fromsingle channel linear hearing aids, difficultiesmay arise when they are applied directly tocurrent multichannel nonlinear hearing aids.

distinction between validation andverification

Validation involves an assessment of thewearer’s performance with the hearing aids.For example, speech recognition tests, dailyquestionnaires, and subjective quality judg-ment in the form of either category rating or

paired comparison are validation tools. Thesensitivity and reliability of these validationtools to evaluate today’s multichannel non-linear hearing aids, although an importantconsideration, will not be discussed in thischapter. The reader is referred to Chapter 5of the companion text, Strategies for Selectingand Verifying Hearing Aid Fittings, 2nd edi-tion (Thieme Medical Publishers) for a de-tailed description of these tools.

Verification refers to the physical measure-ment of the output of a hearing aid. The goal isto determine if it matches some predefinedstandards. An example of verification is mea-suring the electroacoustic output of a hearingaid to determine if it agrees with the data onthe specification sheet. Other tools or indices ofverification include measuring coupler outputto determine if it matches a prescriptive out-put/gain target, or measuring the real-ear out-put of a hearing aid to determine if it matches aspecific target for gain/output. Some also in-clude the determination of the aided thresholdas a verification metric. Regardless of wherethe measurement is performed (i.e., real ear orcoupler), a decision on the suitability of thehearing aid is made as a consequence. Adjust-ment will be made to those hearing aids tomeet the criteria. Verification tools do not yieldinformation on the wearer’s performance withthe recommended hearing aid settings. Many,however, assume the contrary and expect satis-faction when the verification criteria are met.The reader is referred to Chapters 2 and 3 of thecompanion text, Strategies for Selecting and Veri-fying Hearing Aid Fittings, 2nd edition (ThiemeMedical Publishers) for a detailed descriptionof these tools.

The availability of current nonlinear hear-ing aids has created two areas of concernabout verification efforts: the ability to reli-ably measure the output of such hearing aids,and the interpretation on the adequacy ofthe output relative to prescriptive targets.

Hearing Aid Factors That Affect Reliabilityof Measure

Reliability of the measure, whether matchingthe output to specification data accordingto ANSI protocol (e.g., ANSI, 1996), or match-

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ing it to a prescriptive target using either cou-pler measures or real-ear approaches, can beaffected by the characteristics of the hearingaid processing. Although many nonlinearhearing aids use digital signal processing, notall such hearing aids create the same difficul-ties in measurement because of differences intheir processing features. Special precautionsare needed to account for nonlinear process-ing. The special processing is addressed inthe following subsections.

noise reduction algorithms

Many current multichannel nonlinear hear-ing aids include noise reduction algorithmsthat use statistics to estimate the nature ofthe incoming signal to determine if addi-tional gain adjustment across frequencies isnecessary. As was discussed earlier, manyuse either the modulation index or the leveldistribution function to make such deci-sions. Although the basis of the speech/noise decision is fairly similar among manu-facturers, the aids do differ at the input levelat which noise reduction is active (highinput level only or all input levels), the max-imum amount of gain adjustment across fre-quencies, as well as the speed at which the

time-varying filter changes its frequency re-sponse characteristics.

The types of stimuli that are routinely usedfor verification purposes (real ear or coupler)could be problematic with these current hear-ing aids. This is because the typical stimuli, bethey sinusoids or broadband noise, are sta-tionary in nature. That is, their intensity leveldoes not change during the course of thestimulus presentation. Although any of thesesignals would be acceptable for conventionalhearing aids, all of these signals would be re-garded as “noise” by current nonlinear hear-ing aids that include noise reduction algo-rithms. Figure 5–11 shows the gain of athree-channel noise reduction hearing aid to abroadband noise with the noise reduction al-gorithm deactivated and activated. As muchas 12-dB gain reduction is seen when thenoise reduction algorithm is activated. SeeChapter 3 in the companion text, Strategies forSelecting and Verifying Hearing Aid Fittings,2nd edition (Thieme Medical Publishers) for adetailed description of issues related to real-ear measures of DSP hearing aids.

The amount of gain reduction variesacross manufacturers. In addition, the SNRof the input as well as the duration of thestimulus could affect the measured gainlevel. In general, more gain reduction is as-

Figure 5–11. Frequency-gainresponse of a digital hearing aidto a broadband speech noisewhen (a) the noise reductionalgorithm is not active, and (b)the noise reduction algorithm isactive.

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sociated with inputs that have lower SNRs.The duration of the stimulus interacts withthe estimation time used by the noise reduc-tion algorithm to yield varying output. Ifthe stimulus duration is significantly shorterthan the noise estimation time, no gain ad-justment is made because the noise reduc-tion algorithm has insufficient time to makean estimation. If the duration of the stimulusis about equal to the estimation time, gainreduction occurs. Gain reduction continuesto the maximum for the specific SNR as longas the stimulus is present. For example, athree-channel digital hearing aid that theauthor has experience with takes approxi-mately 10 seconds to be fully activated. Thepotential effect of the noise reduction algo-rithm will not be seen if a stimulus is pre-sented for only 5 seconds. If, however, thestimulus is presented for 15 seconds, gainfrom the hearing aid could be reduced by asmuch as 12 dB.

This observation suggests that one mustcontrol the duration of the stimulus presen-tation during verification efforts. Otherwise,the same hearing aid settings could result indifferent output. If the purpose is to examinethe gain of the hearing aid without theaction of the noise reduction algorithm, thenthe stimulus must be a broadband stimulusthat is presented for a sufficiently long timeto activate the compression circuit but notlong enough to activate the noise reductionalgorithm. The exact time would be depen-dent on the specific manufacturer. Typically,a stimulus that is about 3 to 5 seconds in du-ration is acceptable for many hearing aids.If, however, the purpose is to examine thegain with the noise reduction fully activated,stimulus duration of over 15 seconds will benecessary.

There are simpler methods to bypass theaction of the noise reduction algorithm dur-ing verification. Most fitting software has theoption to allow deactivation of the specialfeatures (e.g., noise reduction) during verifi-cation. This would allow reliable testing ofthe hearing aid output. Another option is touse a specially designed broadband stimu-lus, the International Collegium of Rehabili-

tative Audiology (ICRA) signal for testing.The ICRA signal is a speech-shaped noisethat is randomly modulated so it appearsas speech to current hearing aids with noisereduction algorithms (see Chapter 1). Thiswould prevent the hearing aid from lower-ing its gain and avoid variation of the outputeven with different stimulus duration. Sev-eral manufacturers of testing equipment alsoincluded this stimulus as an option.

attack and release time

The release time of nonlinear hearing aidsmay affect the reliability of the coupler mea-sure (Revit, 1994). The problem is aggra-vated by some current nonlinear hearingaids that use even longer attack and releasetimes (more than 2 seconds) than were con-sidered by the manufacturers of hearing aidtest systems. This is especially true whenautomatic test sequences are allowed. Theoutput of the hearing aid varies dependingon when the test stimulus is presented inrelation to the recovery phase of the hearingaid. For example, Figure 5–12 shows the fre-quency gain curves at input levels from 50 to80 dB SPL in 10-dB steps from a hearing aidthat has a maximum release time of 20 sec-onds. Figure 5–12a shows that no gain dif-ference is seen across the four input levelswhen the delay between stimulus presenta-tion is 0.5 second. One may have concludedthat the hearing aid is a mild gain linearhearing aid. Figure 5–12b shows the resultswith a 2-second delay between presenta-tions, whereas Figure 5–12c shows the sameat a delay of 5 seconds. Significant deviationin the gain of the hearing aid is seen acrossthe three graphs. This shows that if the delaybetween presentations is shorter than the re-lease time of the hearing aid, one may obtainvariable results depending on when the sub-sequent stimulus is presented during thegain recovery phase of the nonlinear hearingaid. In general, the delay between stimuluspresentations must be longer than the maxi-mum release time of the hearing aid to mini-mize variability in output. For hearing aidtest systems that provide automatic stimulus

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presentation, it would be important to setthe automatic test sequence so that sufficientdelay between presentations is introduced.

low compression threshold

The test protocol recommended by the ANSIworking group to test the attack time and

the release time of a nonlinear hearing aidis to use a stimulus that changes its intensitylevel from 55 to 90 dB SPL. The choice ofsuch intensity levels assumes that the non-linear hearing aids being evaluated are com-pression limiting or automatic volume con-trol devices. These types of compressionhearing aids typically have a high compres-

Figure 5–12. Frequency gaincurves for inputs from 50 to 80dB SPL (in 10-dB steps) at vari-ous delays between presenta-tions: 0.5 second (a), 2 seconds(b), and 5 seconds (c).

(a)

(b)

(c)

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sion threshold (above 60 to 65 dB SPL). Theuse of this stimulus to measure the dynamiccharacteristics of a WDRC hearing aid maybe questionable. This is because the thresh-old of WDRC hearing aids, even the conven-tional ones, is often below 55 dB SPL (from40 to 55 dB SPL). This suggests that the hear-ing aid is in compression even for the 55 dBSPL portion of the test signal. This is eventruer with current digital signal processinghearing aids where many have CTs below30 to 40 dB SPL. A different signal level isneeded if one wishes to characterize the dy-namic aspects of current compression hear-ing aids with low CT (below 40 dB SPL). Theissue is complicated even more when manyof these multichannel devices have differentdynamic characteristics and expansion pro-cessing in each channel. More than one testfrequency may be needed in the future.

multiple channels

The choice of stimulus type (sinusoid orbroadband noise) makes minimal differenceon the obtained frequency-gain curves whentesting a linear hearing aid at typical inputlevels. For example, Figure 5–13 shows alinear power hearing aid tested with a pure-tone sweep, an ANSI composite noise asimplemented on the Frye 6500 test system(ANSI, 1992), and an ICRA noise imple-mented on the Frye 6500 system. Notethat identical frequency response curves areobtained.

Testing nonlinear compression hearingaids, even single channel devices, requiresspecial consideration. Preves (1996) cau-tioned about the artificial effects of “bloom-ing” in the output frequency response whena sweep-tone stimulus is used as the test sig-nal instead of a broadband noise. Because ofsuch artifacts, it has been suggested that astationary broadband noise with definedcharacteristics be used to measure the fre-quency response of nonlinear hearing aids(ANSI, 1992).

The availability of multichannel nonlinearhearing aids brings even more variability inthe measured outcome when different stim-uli are used. One such stimulus is the ICRA

noise that was discussed previously. Figure5–14a shows the gain of a three-channelWDRC hearing aid measured with a 60 dBSPL sweep-tone (ANSI, 1996). The samehearing aid is also tested with a 60 dB SPLcomposite noise specified by ANSI (1992)(Fig. 5–14b) and a 60 dB SPL ICRA noise(Fig. 5–14c). In addition to the blooming ef-fect seen in Figure 5–14a (over Fig. 5–14b),significant difference in the gain values isobserved between Figure 5–14b and 5–14caround the 1000-Hz region. This is becausethe ANSI noise and the ICRA noise differ intheir frequency composition. In a multichan-nel hearing aid where each channel can haveits own compression threshold and com-pression ratio, a difference in the input lev-els among frequencies would suggest dif-ferent gain. This variability in output withstimulus types raises the issue of (in)accu-racy of our target match across clinical sitesor within the same site that uses more thanone type of test signal as offered by sometest systems. This calls for a more standard-ized noise for ease of data comparison acrossfacilities and for verifying the output/gainof multichannel nonlinear hearing aids. At aminimum, the type of stimulus used in veri-fication must be reported.

Factors Affecting Interpretation of Target Match

A criterion that is commonly used in the fit-ting of hearing aids (linear and nonlinear) isto determine if the output (and/or gain) ofthe hearing aid matches a prescribed target.This measurement can be obtained via probe-microphone measurement or through cou-pler measurement with appropriate real-ear-to-coupler corrections. It is often assumedthat real-world satisfaction is guaranteedwhen the output (real ear or coupler) of thehearing aid meets a specific gain target. Al-though many clinicians recognize that stricttarget matching is impossible to achieve andmay not even be desirable, there are addi-tional reasons why one may need to be evenmore cautious when it comes to current mul-tichannel nonlinear hearing aids. The authoris not ignoring the many benefits provided

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by prescriptive formulae. Rather, the pointof this discussion is to challenge our currentthinking so that suitable adjustment to ourcurrent practice can be made to reflect thestate of technology.

assumptions of prescriptive targets

Most prescriptive formulae for linear hearingaids are based on the assumption of amplify-ing normal conversational speech (with an

Figure 5–13. Frequency gain curves of a linear power aid for different stimuli: (a) pure tonesweep; (b) Frye 6500 composite speech noise; (c) Frye 6500 International Collegium of Reha-bilitative Audiology (ICRA) noise. Note the same frequency response curve in all situations.

(a)

(b)

(c)

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overall level of 65 to 70 dB SPL, in quiet) to thewearer’s most comfortable listening (MCL)level. Suitable adjustment to the low- andhigh-frequency gain is made to minimize am-plifying too much room noise (in the low fre-

quency) and the occurrence of feedback in thehigh frequency (Skinner, 1988). Differencesexist among various formulae from the differ-ence in the assumption of the speech spec-trum, as well as modification based on empir-

Figure 5–14. Frequency gain curves of a nonlinear power aid for different stimuli: (a) puretone sweep; (b) Frye 6500 composite speech noise; (c) Frye 6500 ICRA noise. Note the differ-ence in frequency response especially in the low to middle frequency region in all situations.

(a)

(b)

(c)

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ical evidence. Regardless of the formula cho-sen, frequency gain at one input level (i.e.,conversational level) is recommended fordaily use in all listening environments.

The fitting philosophies for nonlinearhearing aids are more diverse. Some advo-cate the use of individual loudness scaling toset the individual compression parameterson the nonlinear hearing aids. Examples arethe Independent Hearing Aid Fitting Forum(IHAFF) approach (Valente and Van Vliet,1997) and the loudness growth by octaveband (LGOB) approach (Allen et al, 1990).These approaches aim at restoring normalloudness in the impaired ear. Some areprescriptive approaches that use the ratio ofthe wearer’s residual dynamic range tothe dynamic range of normal hearing lis-tener to define the compression characteris-tics. An example is the desired sensationlevel (DSL[i/o]) approach formulated bySeewald and his colleagues (Cornelisse et al,1995). Some take an eclectic approach to pre-scribe gain (and thus compression ratio) atdifferent input levels (FIG6, 1996). Many ofthese approaches use a loudness normaliza-tion philosophy (Byrne, 1996). Still othersprescribe gain by attempting to ensure equalloudness at each frequency region (e.g., Na-tional Acoustic Laboratories, 1999). SeeChapter 1 in the companion text, Strategies forSelecting and Verifying Hearing Aid Fittings, 2ndedition (Thieme Medical Publishers) for a de-tailed description of each method. Unlikethe prescriptive targets for linear hearingaids, frequency gain recommendations atthree input levels are provided for a calcula-tion of the desired compression thresholdand ratio. But similar to the linear targets, allthe targets assume listening in quiet and thatthe recommendation is applicable to allhearing aids (i.e., they are generic).

should there be only a single target?

To meet the prescriptive requirement, a lin-ear hearing aid has to meet the target gainfor conversational level ~65 db SPL only. It isassumed that the use of the volume control(VC) on the hearing aid would compensatefor the overall level difference for the low

and high input levels (Byrne, 1996). The as-sumption is that the frequency response re-quirement for low and high inputs is similarto the medium input.

Several studies that examined the preferredfrequency responses under different listen-ing conditions raised doubts about the as-sumption of a fixed frequency response forall input levels. Kuk et al (1994) measuredthe preferred frequency responses in hearing-impaired subjects across listening conditionsthat differed in the SNR while maintaining thesame overall input level. Their results showedthat most subjects preferred a broad fre-quency response when the SNR of the listen-ing environment was positive. As the SNR be-came poorer, however, subjects preferred lesslow-frequency gain as well as lower overallgain from the hearing aid. This suggests thatthe preferred frequency gain on a hearing aidis determined not only by the overall inputlevel but also by the SNR of the input signal.

Keidser et al (1995, 1996) also showedthat patients with a hearing loss preferred fre-quency gain characteristics that are differentfrom the National Acoustic Laboratory’s (-re-vised) (NAL-R) (Byrne and Dillon, 1986) tar-get in situations other than quiet. In situa-tions where low-frequency noise or babblenoise dominates, a steeper response thanthe NAL-R target is preferred. In situationswhere high-frequency noise is present, a flat-ter response than the NAL-R target is pre-ferred. Because real-life listening situations ofhearing aid wearers will necessarily includemany types of noise, these studies questionthe limited value of having a fixed frequency-gain response for all environments. Fixed tar-get at one level or the same frequency re-sponse target at multiple input levels may notbe optimal. Targets must account for not onlythe input level effect but also the SNR effectin daily situations. That is certainly an impe-tus for the design of the so-called noise reduc-tion hearing aids that respond not only to theoverall input level, but also to the SNR of thelistening environments.

Nonlinear fitting targets, especially thosemeasured on an individual basis, have theappearance of higher face validity in that thelevel effect is considered. The effect of SNR,

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204 HEARING AIDS: STANDARDS, OPTIONS, AND LIMITATIONS

however, is still not considered. Further-more, the fitting procedures of these ap-proaches require the patients with a hearingloss to perform loudness scaling in whichtheir perceived loudness of discrete, station-ary tone bursts or narrow band noise is usedto define the static compression characteris-tics of a hearing aid (Kuk, 2000). Althoughloudness scaling is certainly one approach tostudy loudness perception, its use in fitting anonlinear compression hearing aid, espe-cially modern multichannel nonlinear hear-ing aids, may be questionable.

normalization for all frequencies?

Loudness normalization is based on thepremise that the manner of amplificationwould restore normal loudness perceptionin the patient with a hearing loss. Becauseof the difference in the degree of recruit-ment among frequency regions (other thanperhaps a flat hearing loss configuration),compression hearing aids would necessarilyrequire many independent processing chan-nels so that each can perform independentprocessing for a restricted frequency region.A compression hearing aid that can truly re-store loudness normalization would requireas many channels as there are independentauditory processing units. Even if one ap-proximates each critical band as one inde-pendent processing unit, it would require atleast 15 to 20 independent channels in a non-linear compression hearing aid to achieveloudness normalization at all frequencies.Many multichannel compression hearingaids do not have enough channels to qualifyfor this requirement.

is loudness normalization critical?

The importance of loudness normalizationwas questioned by Byrne (1996), who ar-gued that loudness normalization does notguarantee optimal speech intelligibility forthe hearing aid wearer. Rather, he arguedthat the process of loudness normalizationmay have placed too much emphasis on thefrequency-intensity domain and that unde-sirable side effects may occur to offset any

potential advantages of loudness normaliza-tion. This may be the case when the nonlin-ear hearing aids use fast-acting compressionin many channels. One such negative effectis the potential of smearing in the temporaland spectral domains. This could obliteratesome of the cues available in the input forspeech understanding (Van Tasell, 1993).

Byrne (1996) also questioned the “unimpor-tance” of normal loudness relationship. He in-dicated that the long-term spectrum of speechand the short-term spectra of sounds varyconsiderably among talkers and listening en-vironments. Yet, as long as they are audibleand occur within the dynamic range of the lis-tener, most people do not have any problemunderstanding them. This directly questionsthe need to maintain a “normal” loudness re-lationship as advocated by the loudness nor-malization fitting approaches. Other investi-gators also echoed similar opinions throughindependent studies (Horwitz et al, 1991; VanTasell, 1993; Van Buuren et al, 1995). The opin-ions of these researchers suggest that strictloudness matching in a nonlinear hearing aidmay not be critical to provide optimal results.

the effect of multiple channels has notbeen considered in the targets

Many of the generic prescriptive targetswere generated based on theoretical con-structs (e.g., DSL [i/o], Cornelisse et al, 1995)or empirical evidence derived from existinghearing aid wearers (e.g., NAL-RP [includ-ing profound loss], Byrne et al, 1990). It isreasonable to expect that the generated tar-gets reflect the dominant hearing aid tech-nology of the time. Given that assumption,it would follow that many of them are ap-plicable only to single channel linear hearingaids. Issues related to multichannel com-pression hearing aids would probably nothave been considered in the formulation ofthese targets.

Kuk and Ludvigsen (1999a) provide a de-tailed illustration on the output effect of thenumber of channels in a multichannel com-pression hearing aid. In summary, these in-vestigators showed that two compression

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hearing aids that have the same static I-Ocurve (i.e., same gain, compression threshold,and compression ratio) but differing in thenumber of channels would yield differentoutput to a broadband stimulus in real life.This would be the case when one matches theoutput of both hearing aids to the same gaintargets during verification with a tonal stimu-lus. The one with more channels wouldsound louder than the one with fewer chan-nels to complex stimuli in real life. The outputdifference between the two hearing aids in-creases as the difference in the number ofchannels between them increases. The differ-ence also increases as the compression ratioof the two hearing aids becomes higher. Lin-ear hearing aids, on the other hand, do notshow this output difference even if they differin the number of channels.2

Kuk and Ludvigsen’s (1999a) observationshave several implications. First, it providesanother piece of evidence to suggest that tonalstimulus is not a good stimulus to use to ad-just gain on compression hearing aids, espe-cially ones with multiple compression chan-nels. Second, it suggests that current genericfitting targets, even if they truly predict real-life performance, must be updated to accom-modate the difference in the number of chan-nels seen in today’s hearing aids. Specifically,this means that one must modify the targetvalues to include hearing aids with multiplechannels. As of date, only the NAL-NL1 (Na-tional Acoustic Laboratories, 1999) prescrip-tive target has considered the effect for up tofour channels. Such limitation in current pre-scriptive targets should caution one againststrictly matching the output of a multichannelhearing aid to a prescriptive target.

the effect of noise reduction has notbeen considered in the targets

Another area of concern in gain verificationis the availability of noise reduction algo-

rithms in many current hearing aids. Previ-ously it was discussed how this mechanismcould make the verification/measurementof output variable. This section questionsthe validity of the gain targets for these hear-ing aids, and examines how such hearingaids should be verified.

The noise reduction algorithm in mostmultichannel nonlinear hearing aids worksby reducing or modifying the prescribedgain for a quiet listening situation. For suchhearing aids to provide optimal listening inquiet and in noise, their manufacturers musthave some knowledge of optimal gain re-quirement in these situations before theyspecify the amount of gain adjustment to beprovided by the noise reduction algorithm.Indeed, several manufacturers of digitalmultichannel hearing aids have developedproprietary fitting formulae based on exten-sive in-house research and clinical data.Their proprietary gain targets have consid-ered the dynamic changes in gain require-ments beyond what would have been con-sidered by generic prescriptive formulae.

A clinician who unknowingly adjusts theparametric settings on such hearing aids tomatch a generic target could risk compro-mising the hearing aid fit for the wearer—inquiet, in noise, or in both types of environ-ments. Consequently, some manufacturersrecommend against a strict target match.This may be the same reason why some re-searchers reported that the preferred gain ofadult hearing aid wearers was closer to themanufacturer’s recommendation than to ageneric target’s (Stelmachowicz et al, 1998).Because not all manufacturers of digitalhearing aids have gone through this re-search effort, clinicians need to know the ra-tionale behind the hearing aid gain targetprior to deciding if a generic fitting targetshould be matched.

An equally complex issue is whether thenoise reduction algorithm should be deacti-vated or should remain active during targetmatching (if one decides to match its outputto a generic target). If one assumes thatall prescriptive targets are generated fortypical noise-free listening conditions, then

2A “multichannel linear” hearing aid is essentially anequalizer and not a true multichannel hearing aid. Atrue multichannel hearing aid refers to one with at leasttwo channels of independent compression.

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the noise reduction algorithm should be by-passed during verification/target matching.The caveat is that the output may be “in-flated” relative to a generic target if themanufacturer’s proprietary target has con-sidered the effect of the noise reduction (as-suming that the generic target has correctedfor channel and release time effects also). Inthat case, one should guard against loweringthe output/gain to meet the target. If onematches the target with the noise reductionalgorithm activated, one may be examiningthe impact of noise on the hearing aidperformance. One may receive less outputthan the prescriptive recommendation. Ifone proceeds to increase the gain setting tomatch a target, it could result in too muchgain for everyday situations. These are com-plex issues that need further deliberation.Bearing in mind that a basic objective in anyacceptable hearing aid fitting is to ensureaudibility of the softest sound and comfortfor the loudest sound, it would seemappropriate that the noise reduction algo-rithm be deactivated during verification toexamine if the criterion of audibility is firstand foremost achieved. Testing with noisereduction activated may be desirable to ex-amine the comfort issue when loud soundsare present, but it should be a secondary ob-jective at this point. Kuk and Ludvigsen(1999b) discuss these issues.

the effect of attack time and releasetime has not considered in the targets

It was discussed earlier that the static char-acteristics on nonlinear hearing aids are se-lected based on either a prescriptive ap-proach (e.g., FIG6, DSL[i/o], NAL-NL1) oran individual loudness rating approachusing short, steady-state tone bursts or nar-row bands of noise (e.g., LGOB, IHAFF).None of these approaches considered the ef-fect of time constants (attack time and re-lease time) on the subjective perception ofthe output of a nonlinear hearing aid.

It was discussed earlier that the effectivecompression ratio of a compression hearingaid is affected by the time constants of thehearing aid and the input acoustic signal.

Real-life signals fluctuate greatly in their in-tensity levels. As one would predict, theshorter the time constants (or faster attackand release), the better is the nonlinear cir-cuit in following the intensity variations ofthe input signal to ensure audibility of theweaker sounds. It also suggests an effectivecompression ratio that approximates moreclosely the nominal (or static) compressionratio. The intensity difference among speechcomponents, however, will be decreasedand ambient noise more noticeable duringpauses in the speech input. In contrast, alonger release time (greater than 1 second orso) would preserve the intensity relationshipof the speech components better. The effec-tive compression ratio, on the other hand,would regress toward 1 (or linear) (Verschu-ure et al, 1996). Figure 5–15 shows the differ-ence in waveform for the same speech inputprocessed by two different release times.A linear representation of the same signalis shown for comparison. In the first casewhere a release time of 200 msec is used,clear reduction in the intensity difference isnoted. In the second case where the releasetime is 10 seconds, the intensity difference ismostly retained. Subjectively, hearing aidswith a longer release time sound clearerand more natural than hearing aids with ashorter release time (Neuman et al, 1995;Hansen, 2001). Spectrally, a longer releasetime results in an increase in the mid-frequency output over a shorter release time.

There are several implications from thestudy of the effect of release time on the ef-fective compression ratio (and subjectivepreference). First, if two nonlinear hearingaids are identical in all but their time con-stants, the one with a longer attack and re-lease time would sound more natural andpleasant than the one with a shorter at-tack/release. This suggests that matchingthe static I-O curve is no guarantee of subjec-tive acceptance. Second, because the releasetime changes the overall loudness of the am-plified sound, modification to the prescrip-tive target is necessary to account for theloudness effect resulting from a difference inrelease time. Such has not been consideredin any generic prescriptive targets. Third,

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if release time changes the dynamic I-Ocharacteristics of the hearing aid, one mayquestion the value of measuring loudnessgrowth function for the fitting of compres-sion parameters. After all, the effective com-pression ratio will not be the same as thestatic compression ratio that is determinedthrough loudness growth procedures. And ifthe goal is to determine a general long-termcompression ratio, several studies have con-firmed that average loudness growth func-tions (and static compression ratios) can bepredicted accurately and efficiently withthreshold data (Hellman and Meiselman,1993; Jenstad et al, 2000).

Reconsidering the Concept and the Valueof the Aided Threshold

meaning of the aided threshold

A frequently used verification tool, that ofmeasuring the functional gain and/or theaided threshold, should be reconsidered.Functional gain is the gain provided by thehearing aid when it is worn over the un-aided condition. It is obtained as a differencebetween the aided and the unaided sound

field thresholds under the same test condi-tions. This index is not new. It has been usedwith linear hearing aids as a way to quantifythe improvement in auditory sensitivity.However, many may have refrained fromsuch claims because as one adjusts the vol-ume control position on the linear hearingaid (as would be typical in real life), themagnitude of the aided threshold (and func-tional gain) changes. This significantly di-minishes the meaningfulness of the aidedthreshold. When used with linear hearingaids, it was often cited that the magnitude ofthe functional gain was similar to the inser-tion gain that the wearer received for con-versational input (Pascoe, 1975; Hawkinsand Haskell, 1982).

In reality, one should recognize that theaided threshold represents the lowest inputlevel that one can hear with a hearing aid atthe specific gain setting. Because one wouldlikely adjust the VC on a linear hearing aidin real life, the aided threshold could changefrom the value determined during sound-field threshold measures. Thus, not manywould claim that the aided threshold repre-sents the softest sound that a wearer can

Figure 5–15. Effect of release time on the peak-to-valley envelope difference. The samecompression ratio (CR) (2.5:1) is used for both release times (200 msec and 10 seconds).SNR, signal-to-noise ratio.

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hear with a linear hearing aid. If one can ac-cept, however, that the design of a nonlinearhearing aid should free its wearer from leveladjustment (via the VC), the aided thresholdwould represent the softest sound that thewearer hears with the nonlinear hearingaid. Such assumption is reasonable becausethe signal processing in many modern DSPnonlinear hearing aids has considered thechanges in gain requirement across variousinput levels and SNR conditions (throughnoise reduction algorithm). Thus, the needfor a VC is minimal (although some studiesreported that a percentage of wearers stillprefer to have a VC on their nonlinear hear-ing aids for safety and control reason). Theaided sound-field threshold that is mea-sured in the clinic should represent the soft-est sound that the wearer can hear with thenonlinear hearing aid. Knowing the softestaudible sound can be extremely useful fordocumentation of hearing aid benefit (im-provement in audibility/sensitivity or di-minishment of impairment), or as a counsel-ing tool to set proper expectations.

For example, one way to evaluate if a non-linear fitting is optimal is to compare thevalue of the aided threshold with the aver-age speech spectrum displayed on an audio-gram. In a well-calibrated sound field, aidedthresholds that are around 20 dB HL acrossfrequencies would suggest that the softestpart of the speech spectrum (which occursaround 20 dB HL across frequencies) is audi-ble. A higher (or poorer) aided thresholdwould suggest that some of the speech spec-trum may not be audible. In light of the av-erage speech spectrum level, one may usethe absolute value of the aided threshold asa target. For listening to the softest part ofnormal speech, an acceptable target wouldbe an aided threshold of 20 dB HL across fre-quencies. An aided threshold that is lowershould be achieved if better audibility is de-sired. Note, however, that this discussion isapplicable only to nonlinear hearing aids.The same conclusion must not be applied tolinear hearing aids, especially when fittingsomeone with a severe-to-profound hearingloss (Kuk, 2001b).

One can compare the results of the fittingof new nonlinear hearing aids with thewearer’s own linear hearing aids. If the out-put of the linear hearing aids and the newnonlinear hearing aids are matched to a con-versational speech input, one would expectthat the nonlinear hearing aids yield a lower(or better) aided threshold than the linearcase. This has been demonstrated in Figure5–2. If the aided thresholds for both hearingaids are the same, then either the linear hear-ing aids have given the wearer more gainthan required, or the nonlinear hearing aid isnot prescribing enough gain. In both cases,the wearer will complain of insufficient loud-ness from the nonlinear hearing aids.

The value of the aided threshold may serveas a clue for troubleshooting. It is not uncom-mon for experienced linear hearing aid wear-ers to complain that sounds are not loudenough with their new nonlinear hearingaids. Such complaints may indeed reflect in-sufficient gain at all input levels, or simply areduction in gain at a high input level relativeto the linear hearing aid experience. One ap-proach to decide if further gain adjustment isneeded is to examine the value of the aidedthresholds. Wearers who have aided thresh-olds around 20 dB HL would have been en-sured of adequate audibility and are probablyreacting to the reduced loudness of nonlinearhearing aids to high input sounds. In such acase, counseling or adjustment of gain forhigh input sounds would be necessary. On theother hand, if the aided thresholds with thenonlinear hearing aid are similar to or poorerthan the aided threshold of the linear hearingaid, adjustment of overall gain would be nec-essary. See Kuk (2001a) for a detailed descrip-tion of the considerations in troubleshootingmodern nonlinear hearing aids.

additional variables affecting theaided threshold

One criticism of the concept of sound-fieldthreshold measurement concerns the vari-ability of the measured index (Humes andKirn, 1990). Walker (1995) summarizes theenvironmental factors that may confound

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the results of sound-field measures. This sec-tion concentrates on the additional consider-ations when one attempts to measure theaided threshold of a modern nonlinear hear-ing aid. The key factor affecting the accuracyand reliability of the measure is the timeconstants of the nonlinear hearing aids, andespecially for devices with a low compres-sion threshold (below 40 dB SPL).

First, because a compression hearing aiddecreases its gain as input level increases,the typical bracketing approach in thresholdestimation may yield more variable resultswhen the real aided threshold is at or abovethe compression threshold of the nonlinearhearing aid. In this case, the input stimulusthat is just above the CT receives lower gainthan it would otherwise if it were below theCT. This may artificially raise the value ofthe aided threshold and/or add to the vari-ability of the measurement. An alternativeapproach is to present the stimulus in an as-cending manner once the aided threshold isapproximated with a bracketing approach.This could minimize variability in the aidedthreshold measure.

Recall that the attack time is the time ittakes a nonlinear hearing aid to settle intothe reduced gain state when the input isabove the CT (or compression active). Con-sequently, if the stimulus level is below thecompression threshold, the duration of thestimulus would have no effect on the gain(and output) of the hearing aid. However,when the stimulus level is above the CT, theduration of the signal could interact with theattack time and affect the output. Whenthe stimulus duration is significantly longerthan the attack time, the initial part of thestimulus may receive the full gain of the hear-ing aid and the remaining part of the stimu-lus the reduced gain. If the initial part is sig-nificantly above the wearer’s threshold, thewearer may indicate a threshold response.Otherwise, the stimulus level has to beraised until the reduced gain part is abovethe wearer’s threshold. The drawing on theleft of Figure 5–16 illustrates the situation.

In contrast, in hearing aids using a long at-tack time, the same stimulus would receive

the same gain from the hearing aid. How-ever, a higher output may be possible (re-sulting from a longer duration of lineargain), and a lower aided threshold resultsfor the same audiometric threshold. Thus,two hearing aids with identical I-O charac-teristics could have different aided thresh-olds if they are significantly different in theirattack times. Thus, knowledge of the attacktime used by a compression system is useful.Given that most WDRC hearing aids use arelatively short attack time (less than 10msec), it is foreseeable that the aided thresh-old is measured with the hearing aids in areduced gain state unless the hearing aid hasa high CT. Otherwise, the obtained thresh-old may be higher than the real threshold.For WDRC systems using a low CT, the useof a long attack time may slow the gain re-duction and circumvent the problem ofachieving a higher aided threshold. Consid-ering the effect of temporal integration, CT,and attack time of most hearing aids, a stim-ulus duration of approximately 1 or 2 sec-onds should be used for reliable measure.

As discussed in the previous section oncoupler and real-ear measures, the delay be-tween stimulus presentation could also af-fect the reliability of the aided threshold. Formost commercial WDRC hearing aids that

Figure 5–16. Hypothetical illustration of the ef-fect of attack time on the output waveform. Forthe same audiometric threshold, a nonlinearhearing aid with a long attack time (right) couldyield a better aided threshold than one with ashorter attack time (left).

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use short attack and short release times, nospecial considerations are necessary to pacethe stimulus presentation accordingly be-cause the typical delay between presenta-tions is longer than the release time of thehearing aid (less than 200 msec). However,the delay between presentations can bemore critical in a slow-acting WDRC systemwhere the release time can be several sec-onds. If possible, one should wait for theperiod of the release time before the nextpresentation to minimize variability frompresentation at various gain levels. But ifone assumes that the difference betweenpresentations is typically 5 dB, the minimumrequired delay between presentations can beas brief as 1 or 2 seconds, even for a slow-acting WDRC hearing aid.

In summary, the attack time and the re-lease time on a nonlinear hearing aid mayaffect the reliability of the obtained aidedthreshold measure. Typically, an ascendingapproach with stimulus duration of about 1or 2 seconds and delay between presenta-tions of 1 or 2 seconds is appropriate to en-sure a reliable outcome.

Conclusion: What to Do with CurrentNonlinear Hearing Aids

Nonlinear hearing aids are different fromlinear hearing aids in that different gain val-ues are assumed at different input levels.This provides significant advantages overlinear hearing aids in meeting the wearer’slistening needs. At the same time, variousaspects of its processing have not beenconsidered in conventional verification ap-proaches. Even for the simplest single chan-nel WDRC hearing aids, their verificationand validation would require special consid-erations beyond that of a linear hearing aid.The use of multichannel nonlinear hearingaids further complicates existing verificationand validation efforts. The effects of thenumber of channels, time constants, com-pression threshold, and the presence of spe-cial features like noise reduction require amodification of our current verification prac-

tice to accommodate these advances. Moreadvances are forthcoming to result in evenmore modifications to our current protocol.

These changes require clinicians to recon-sider the appropriateness of their verifica-tion protocols for the chosen hearing aid.Knowledge of the static input-output curvesof the selected hearing aid is needed, but isnot enough. How the dynamic characteris-tics such as time constants and the noise re-duction feature relate to the temporal char-acteristics of the hearing aid/real-ear testsystem should be thoroughly understood.

It is critical that hearing aid manufactur-ers provide explanations of the use of theirhearing aids. How their hearing aids can befitted accurately and evaluated fairly shouldbe clearly articulated. Specific fitting andverification (and validation) protocols, ifthey deviate from the generally acceptedpractice, should be communicated to theclinicians who fit the hearing aids.

It is time that the profession reevaluates itsprotocol for fitting and verification of nonlin-ear hearing aids. The current approaches, al-though arguably appropriate for linear hear-ing aids, are not appropriate for use withmodern multiple channel nonlinear hearingaids without significant revision and modifi-cation. Although clinicians must verify andvalidate the results of their clinical fittings,care must be applied in interpreting the re-sults of their efforts. The profession must eval-uate and agree on a common set of stimulusand test protocols so that some level of agree-ment in verification and validation is shared.

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