Apparatus And Prosthetic Device For Providing Electronic Correction Of Auditory Deficiencies For Aurally Handicapped Persons

Abstract

There is disclosed herein improved methods and apparatus for measuring and describing auditory deficiencies for purposes of prescribing compensatory amplification, and improved methods and apparatus for providing such compensatory amplification in a prosthetic device in practical wearable form, and particularly electronic circuits which may be used both in measuring and correcting such deficiencies.

Parent Case Text

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is directed to inventive concepts which are improvements over those described in copending U.S. Pat. application Ser. No. 133,229, filed Apr. 12, 1971, by William P. Stearns and entitled, "Method and Apparatus for Providing Electronic Sound Clarification for Aurally Handicapped Persons." The present application also is related to copending U.S. Pat. application Ser. No. 229,398 filed concurrently herewith in the names of William P. Stearns and Barry S. Elpern entitled, "Method for Providing Electronic Restoration of Speech Discrimination in Aurally Handicapped Persons," which describes and claims methods and apparatus disclosed in the present application.

Description

BACKGROUND OF THE INVENTION

This invention relates to the sound amplification arts, and to their application in the amelioration of auditory deficiencies resulting from damage to the sensori-neural structures of the human ear. It relates particularly to methods and apparatus for detecting and specifying deficiencies in a persons ability to perceive and to comprehend spoken language, and to methods and apparatus for correcting such deficiencies.

Sensori-neural hearing loss is generally considered to be the most prevalent type of auditory handicap found in the United States as well as in other civilized cultures. It constitutes a significant barrier to adequate communication in 5 to 10 percent of the total United States population, and in more than 50 percent of the population over 60 years of age. Furthermore, these proportions are expected to increase in conjunction with ongoing increases in ambient noise levels and life expectancy in our society.

Sensori-neural impairment may result from any one or more of a number of causes, including, but not limited to genetic and congenital factors, viral diseases, specific toxic agents, circulatory disturbances, specific physical trauma and excessive exposure to noise. Irrespective of the primary cause, however, sensory cells within the organ of hearing or their associated neural units suffer some degree of damage and are rendered partially or totally incapable of fulfilling their respective roles in the processing of auditory information. This form of damage cannot be repaired by means of currently known medical or surgical techniques, and the probability of discovery of effective techniques within the foreseeable future appears rather remote. Thus, in virtually all cases of sensori-neural hearing loss, amplification of incoming sounds represents the only possible means for restoring adequate hearing ability.

Hearing loss resulting from sensori-neural damage is usually irregular with respect to frequency, being selectively greater for particular portions of the audible frequency range. The ability to hear sounds in the range above 1,000 Hz is often affected more than the hearing of sounds below 1,000 Hz, although this is by no means a universal observation. The ultimate consequence of irregular hearing acuity for various portions of the audio frequency spectrum is distortion in the perception of complex sounds, i.e., sounds composed of a number of different frequencies.

A certain amount of distortion in complex sounds may be tolerable, but current information does not permit precise specification of the maximum amount of each type of distortion which may exist without interfering materially with accurate sound recognition. Many gross sounds, for example, do not demand a great deal of analytic power in the auditory system, so even a rather severely impaired system may function adequately in the interpretation of such sounds.

In audiologic parlance, the term "discrimination" denotes the capacity of the ear to analyze incoming acoustic patterns and interpret them appropriately. Analytic power may fail at any of several stages in the auditory process, commonly in the organ of hearing or first order neurons due to damage to these structures. Since the ear may be required to perform many degrees of discrimination, varying from extremely coarse to extremely fine, its analytic power may be measured through the use of tests which demand auditory discriminations of progressive difficulty until failure occurs.

Among the most difficult discriminations required of the human ear are those necessary for accurate interpretation of speech, particularly speech in the presence of noise. Because of the fundamental importance of spoken communication, it is obvious that chronic inability to understand what people say could profoundly influence an individual's social, economic, and cultural well-being. Tests of speech discrimination are commonly employed, therefore, to derive a realistic estimate of a person's everyday functional adequacy in hearing.

Each of the phonic units of a spoken word is a complex sound, composed of several frequencies clustered in a more-or-less definable range. When the acuity of the ear has been selectively impaired in a specific frequency range, speech sounds or their components falling in that range may be heard at reduced intensity or not at all. Impairment in several frequency ranges compounds the difficulty and is probably responsible in large measure for the primary complaint of the individual with sensori-neural hearing loss, that he can hear a speaker's voice but cannot understand what is said. The mechanism for inhibiting such understanding may be the non-linear responses that result in intermodulation products and harmonics which could cause interference with the desired spectral components of speech.

On the basis of the foregoing information, it would seem quite reasonable to deal with sensori-neural hearing loss by selective spectrum amplification; that is, providing amplification only in those frequency ranges or bands in which acuity is deficient, and only in the amount of the deficiency. Thus, the ultimate value of selective spectrum amplification rests on the application of appropriate methods for measuring the degree of auditory deficiency as a function of various frequency bands, and also on the construction of a wearable device which is fully capable of producing amplification to compensate for the measured deficiencies. Because of existing inadequacies in both respects, the principle of selective amplification has fallen into disrepute, for the hearing aid industry has adopted the pure tone (single frequency) threshold audiogram as the criterion measurement, and has produced hearing aids with inadequate capabilities for providing proper acoustic output at each portion of the audio band.

The threshold audiogram curve represents an individual's measured absolute auditory threshold for a series of pure frequency tones, usually in the range of 250 Hz to 8,000 Hz sampled at octave intervals on the assumption that intra-octave tone thresholds follow the general audiogram contour. However, it is demonstrable that fairly marked departures from the overall pattern may exist at intermediate frequencies, i.e., frequencies between pure tones, one octave apart.

The rationale for utilizing threshold measurements is shrouded in history, but it is exceedingly interesting to note that the analogous procedure of measuring visual thresholds for monochromatic (single color) lights is never performed to measure the visual acuity of the eye or to prescribe eyeglasses. In fact, careful consideration of the types of measurements which are genuinely helpful in guiding the design of particular hearing aid features suggests that the pure tone threshold curve is virtually useless for several reasons:

A. under everyday circumstances, individuals react only to supra-threshold sounds, as these are the sounds of primary significance. For practical purposes, threshold sounds remain unnoticed.

B. the contour of an individual's threshold curve is observably different from the contour of his supra-threshold equal loudness curves or comfortable listening level curves.

C. an individual's recognition of complex phonic units or their combination into spoken words is essentially unrelated to his acuity for individual pure tones.

Control of acoustic output in current hearing aids is ordinarily achieved through manipulation of frequency response, which refers to the acoustic output of a sound transmission system at each of the frequencies within its pass band when the input level is maintained constant for all frequencies. A graphic representation of a system's frequency response is referred to as a response characteristic, curve or contour. Manufacturers commonly claim that they are able to build hearing aids to yield any required frequency response; but this does not appear to be the case in practice because there are definite limitations on the bandwidths and response curves available in prsent day aids. In practice, manufacturers use combinations of components which produce a limited choice of response patterns and simply select one which most closely corresponds to the criterion, which, as mentioned earlier, usually is a threshold audiogram curve.

One additional comment is relevant as a preface to the innovative concepts to which the present invention is particularly addressed. It is generally recognized that the ear with sensori-neural hearing loss is excessively susceptible to overloading, which is to say that, although it may be relatively insensitive to sounds of low or moderate intensity, it is hypersensitive to sounds of higher intensity (e.g., non-linear response characteristics). This condition restricts the useful operating range of the ear, referred to as the dynamic range; that is, the decibel difference between the lowest intensity at which a sound is reliably detected (absolute threshold) and the upper limit of comfortable loudness for that sound (discomfort threshold).

Whereas the dynamic range of the normal ear is of the order of 100 dB, the range of a sensori-neurally impaired ear may be as little as 10 or 15 dB, generally over a limited frequency spectrum range. Thus, for an impaired ear to function with any degree of adequacy, the full intensity range of the outside acoustic world must be restricted in some way to fit through an abnormally small sound window and such restriction must cause minimal intermodulation products, harmonics, and so forth which would result in distortion. Without such restriction, the ear is readily overloaded, leading to psychologic or physical annoyance and distortion of incoming acoustic patterns.

The consequences of overloading have been appreciated for many years, and output compression devices are widely used in today's hearing aids. Without exception, however, these devices operate on a broad frequency band, so that when any frequency component of a signal reaches a predetermined critical level, the entire pass band of the hearing aid is compressed. Consequently, the components which are not at a critical intensity are needlessly attenuated.

Our evaluation of relevant factors has led to the evolution of several innovative concepts concerned with improved methods and apparatus for measuring and describing auditory deficiency for purposes of prescribing compensatory amplification, and with improved methods and apparatus for providing such compensatory amplification in practical and wearable form.

SUMMARY OF THE INVENTION

While the present application and said co-pending appliation filed in the names of Stearns and Elpern include similar disclosures, and both teach the same inventive concepts for the sake of completeness, the claims of the present application are particularly directed to electronic circuits of the nature disclosed herein and equivalents thereof for enabling the objectives set forth herein to be accomplished. Accordingly, it is an objective of the present application to provide an electronic measurement system with the following capabilities:

a. division of the audible frequency spectrum into a series of adjacent frequency bands through the use of filter networks, the width of any band in the series being of such magnitude as may be deemed appropriate;

b. specific and individual intensity or volume control associated with each of the filter networks defined in (a) above;

c. specific and individually adjustable output compression associated with each of the filter networks defined in (a) above;

d. activation of the filter networks individually or in concert (all pass);

e. introduction and transmission of recorded material such as from tape, or electronically generated signals through the system of filter networks;

f. electro-mechanical transduction of electronically processed signals, such transduction occurring within the external auditory canal of the test subject;

g. pre-amplification and mixing of input signals and for broadband intensity control;

h. monitoring voltage and/or recording the swept frequency spectrum across the output transducer.

Through the use of the foregoing concepts, further objectives of this invention are:

a. to facilitate the use of narrow band signals as test stimuli inasmuch as they represent the most satisfactory compromise between the precise physical describability of pure tones and the complex acoustic composition of speech sounds;

b. to facilitate the use of continuous conversational speech as test stimulus;

c. to facilitate the measurement of comfortable listening level or other supra-threshold response to various frequency bandwidths as an index of auditory deficiency.

Another object of this invention is to enable the provision of sufficiently miniaturized hearing aid apparatus for wearing by aurally handicapped persons. Such miniaturization can be accomplished by electronic techniques, and the apparatus is intended to implement the amplification features determined by the electronic measurement techniques.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other objects and features of the present concepts will become better understood through a consideration of the following description taken in conjunction with the drawings in which:

FIG. 1 is a block diagram of test equipment used in testing the speech discrimination ability of a subject according to the present concepts;

FIGS. 2a and 2b are curves illustrating speech discrimination scores associated with various spectra and test conditions to be discussed later;

FIGS. 3 and 4 are curves illustrating the response characteristics of a master hearing aid after such response characteristics have been set to obtain the best speech discrimination ability for the patient under evaluation;

FIG. 5 is a block diagram of a wearable hearing aid according to the present invention;

FIG. 6 is a more detailed block diagram of a hearing aid of the nature of that illustrated in FIG. 5;

FIG. 7 is a curve indicating typical response of a single filter of the hearing aid of FIG. 6; and

FIGS. 8 through 11 are specific circuit diagrams of the aid of FIG. 6.

Briefly, in accordance with the present concepts, the format for measuring the auditory deficiency of a subject involves testing with a master hearing aid device in accordance with steps as set forth below.

1. The hearing aid receiver is inserted into the external canal of the test ear and secured with a packing of earmold impression material to provide an acoustic seal.

2. The non-test ear is occluded by an insert earplug and a circumaural muff to block any auditory perception by that ear.

3. The patient is seated comfortably opposite a loudspeaker within a sound-treated enclosure, the loudspeaker having been acoustically equalized to produce a flat frequency-response characteristic at the position of the patient's head. The test signals used to accomplish such equalization are narrow bands of noise.

4. Tape recorded continuous discourse is delivered through the loudspeaker at a level approximating that of normal conversation, i.e., 60 dB SPL.

5. The patient is instructed to listen carefully to the tape recorded speech and to adjust the various filters and the main volume control to achieve maximum clarity and comprehension of the material.

6. When he feels he can attain no further improvement in clarity and comprehension of the recorded speech, the patient notifies the examiner and the recording is stopped.

7. Standardized tape recorded tests of speech discrimination ability are then delivered at the same level as the continuous discourse. The speech discrimination score is noted for each test given.

8. Relative performance following the present concepts may be assessed through comparison with unaided discrimination scores or scores obtained under identical listening conditions with other aids. Furthermore, the voltage level across the hearing aid receiver may be measured and recorded for each filter such as illustrated in FIGS. 2A-2B. These measurements when obtained for each filter network in the system, provide the response curve necessary to restore the subject's discrimination ability, and this response curve can then be used for prescribing the appropriate correction. If necessary, the response curve, which the subject has initially adjusted, may be revised by the subject or by the examiner to effect further improvement in speech discrimination ability. Additional standardized speech discrimination tests are then required to assess the effect of such revisions. Such adjustments may be based, for example, on whether the subject is missing vowels or consonants and so forth. Furthermore, the examiner may adjust the response to smooth large variations therein. Also, a pink or white noise test curve may be run and/or noise inserted during the test process.

Before giving further examples of measurement procedures and results, suitable test apparatus will be discussed. Turning to the drawings, FIG. 1 illustrates an organization of test equipment used in testing speech discrimination of a subject to the extent necessary to practice the method of the present invention. Speech from an audio tape recorder and playback apparatus 10 is applied to an audio mixer network 11. A pink or white noise generator 12 also may be coupled to the audio mixer network 11 to combine speech and noise. The speech and/or audio noise is thus fed to a filter network 13 which comprises a plurality of filter networks F.sub.1 through F.sub.n. Each of the filter networks F.sub.1 through F.sub.n has a discrete pass band and the entire combination 13 preferably covers a frequency range from approximately 125 Hz to at least 6,300 Hz. The filters F.sub.1 through F.sub.n divide the audible frequency spectrum into adjacent pass bands. The pass bands provided by each filter F.sub.1 through F.sub.n may be as wide or as narrow as desired in obtaining appropriate audio noise and/or speech recognition characteristics, and need not be related in an octave relationship as is frequently the case with filter networks. Filter set 13 may also be implemented with a combination of adjustable band reject filters arranged in a series configuration rather than parallel filters.

Furthermore, the signal amplitude variation is adjustable over a suitable range. The output of the filter set 13 is passed through an adjustable gain broadband audio amplifier 14, and hence to an audio receiver 15 positioned in a human ear 16. A vacuum tube voltmeter 17 can be used to measure the amplitude of the signal voltage impressed across the receiver 15.

By way of example, the noise generator 12 may be a HP (Hewlett Packard) 8057A precision noise generator, the tape recorder and playback 10 may be a Craig Model -2704 cassette recorder and playback unit, the mixer 11 may be a Shure .MG7 microphone mixer, the filter network 13 may be a HP 8056A, the audio amplifier 14 a McIntosh MC2505, the receiver 15 a Tibbetts model 102-10G hearing aid receiver, and the vacuum tube (RMS) voltmeter 17 a Ballentine model 320.

The response curves for two test subjects are shown in FIGS. 2a and 2b together with discrimination test scores. FIG. 2a pertains to one ear of one test subject, and FIG. 2b pertains to one ear of another test subject, the ordinate being logarithmic and indicating voltage across the receiver 15 in millivolts. The abscissa is frequency. Curve (a) in both figures represents the most comfortable listening level set by the subject for individual one-third octave bands of noise; curve (b) represents the comfortable listening levels set by each subject for maximum intelligibility of running speech; and curve (c) is an examiner's revision of curve (b) for example, to minimize response peaks. Curves (d) and (e) in FIG. 2a respectively represent frequency response of 3 dB per octave and 4 dB per octave. These latter two curves are simply a plot of frequency versus sound pressure level, but with a constant dB per octave change as compared with the random response plotted from the subject's own response setting, modified in some cases by the examiner as indicated above with respect to curve (c).

The percentages indicated at the upper right edge of the curves of FIGS. 2a and 2b indicate speech discrimination scores achieved by the respective two subjects. This refers to a test wherein the subject is provided with hearing aid apparatus having a response curve set according to the respective curves indicated in FIGS. 2a and 2b and the subject listens to speech supplied to the input of the hearing aid apparatus. The test material consists of standardized phonetically balanced lists of words, each list comprising 50 words to which the subject must respond by repeating each word immediately after it is presented by a tape recorded speaker. Considering FIG. 2a in this regard, it will be seen that the speech discrimination score was only 4 per cent (4 percent) for response curve a, that is, merely a comfortable listening level set by the subject. On the other hand, a speech discrimination score of 76 per cent was obtained when the hearing aid apparatus was adjusted to response curve ba significant improvement; and increased to 92 per cent for response curve c (the examiner's revised curve b). As shown by FIG. 2b, the original score was 20 .sub.per cent for curve a, but increased to 54 per cent for curve b, and to 88 per cent for curve c. It will be understood that another test subject's response may be significantly different from that shown in either FIGS. 2a or 2b and, in fact, each ear of a subject may differ significantly in its uniqueness.

It should be noted that the final response adjustment (curve c) for optimum discrimination was made by an examiner using a spectrum shaping technique that has been developed as part of this invention, and principally involves smoothing of response peaks of curve b as noted earlier. In the case of FIG. 2a, the individual's ability to distinguish the spoken words of the speech discrimination tests increased to 9 percent (curve c), which is a significant improvement over the 76 percent for curve b and a significant improvement over the scores for the constant dB per octave change curves d and e.

Several specific examples and results in accordance with the present test method are described below. Subject X was tested with three hearing devices. In each instance the left ear was tested and the right ear was covered. Test No. 1 was made while the subject used her personal, commercially purchased, hearing aid. The test involved use of Comm Tech Auditory Test N-1 sentences with a competing signal of two female talkers in the background. This is a relatively difficult discrimination test as compared to mere use of a word list. The resulting speech discrimination score (SDS) was 22 percent. Test No. 2 involved a wearable hearing aid with adjustable filters using the circuit shown in FIG. 6 hereof and the same test as in Test No. 1, and the speech discrimination score was 61 percent after adjustment of the wearable hearing aid in accordance with steps 5 and 6 of the test noted earlier. Test No. 3 was conducted several days later and involved the use of CID Auditory Test W-22, List 4D. The speech discrimination score with the subject's personal aid was 68 percent. Test No. 4 involved CID Auditory Test W-22, List 2F, and the use of a master hearing aid with a circuit as illustrated in FIGS. 5 and 6. The wearable aid and master aid are the same electrically and use essentially the same components, but the master aid is physically larger and has more easily adjustable knobs for filter settings. The speech discrimination score in this case was 88 percent. FIG. 3 illustrates the frequency response of the master hearing aid after it was set by Subject X in the test.

Test No. 5 involved an audiological evaluation of Subject Y conducted by a university Speech and Hearing Clinic. The phonetically balanced speech discrimination score was 70 percent. This same subject was tested in accordance with the present method after the subject set the response of the master hearing aid. The subject was tested with CID Auditory Test W-22. List 4D, and had a speech discrimination score of 92 percent in the same ear. A similar test of Subject Y was conducted with CID Auditory Test W-22, List 2F, wherein the subject adjusted the response of the master hearing aid and then the response was trimmed by an audiologist in the manner noted earlier, and the speech discrimination score was improved to 96 percent. FIG. 4a is an oscillographic waveform, similar to the curve of FIG. 3, indicating the response of the master hearing aid after being set and trimmed by the examiner.

Test No. 6 involved Subject Z whose unaided phonetically balanced speech discrimination score was 66 percent. He was tested in accordance with the present method after he had adjusted the response of the master hearing aid. The master hearing aid in this case differed from that previously employed in Test No. 5, in the following respect: six adjacent filter networks divided the overall speech spectrum into unequal bandwidths, whereas, for previous tests, the filter bands were each one octave in width. The unequal bandwidths were selected on the basis of their relative contribution to overall speech intelligibility. Such bandwidths are often referred to as "equal intelligibility bands." With the aforesaid master hearing aid adjusted by the subject (See FIG. 4b) for optimum speech intelligibility, he achieved a speece discrimination score of 96 percent on CID Auditory Test W-22, List 3-D. A control test under identical acoustical and procedural conditions with a conventional hearing aid yielded a speech discrimination score of 84 percent on CID Auditory Test W-22, List 3-F.

Turning now to an exemplary aid, a block diagram of a wearable hearing aid is shown in FIG. 5. The Figure illustrates the practical miniaturized circuitry for a hearing aid which can be adjusted to duplicate the response curve obtained with the test apparatus shown in FIG. 1. A microphone and FET (field effect transistor) amplifier 21 feeds the received input signals to a broadband audio IC (integrated circuit) amplifier 22 which has a volume (amplitude) control 23. Driver amplifier 24 provides a low impedance source for filter network 25, including plural amplitude controls 26 and plural active IC bandpass filters 27, the outputs of which are fed to a summation, or all-pass, network 28. As will be apparent to those skilled in the art, the bandpass filters 27 each have a bandpass and amplitude control 26 suitable for closely providing or approximating the desired response curve (e.g., curve c in FIGS. 2a and 2b) and thus are selected to provide suitable speech recognition. These filters 27 accordingly may each provide a portion of the total pass band of the filter network 25. The pass band provided by each filter may be as wide or as narrow as required ot obtain optimal speech discrimination and need not be related in any octave relationship or fractional combination thereof. An integrated circuit amplifier 29 having a pass band commensurate with that of filter network 25 provides the final signal amplification prior to the signal being applied to a receiver 30. Automatic saturation elimination control 31 provides signal compression when the signal exceeds a predetermined level. The foregoing hearing aid configuration offers the following advantages: Independent control of pass band amplitude for each of several portions of the spectrum; separate response control for each ear (binaural); ease of readjustment as the patient's requirements change with time; this may be accomplished by the replacement of filter elements having different pass bands and adjusted for different amplitudes. In desired cases a narrow band notch rejection filter may be added after the summation network to alleviate narrow band resonance problems observed in some patients. The concept is readily adaptable to MSI (medium scale integration) integrated circuit techniques. This permits a substantial size reduction in hearing aid models. Rechargable or long life batteries may be used as desired. Ease of repair, ruggedness, and waterproof sealing of the electronic circuits can be readily accomplished. Attractive and compact packaging can be provided.

FIG. 6 is a detailed block diagram of a hearing aid of the nature of that illustrated generally in FIG. 5, and may be manufactured in a miniaturized wearable form. Also, the circuit of FIG. 6 can be used in the master hearing aid which, as noted earlier, is preferably a larger test instrument having larger and more readily adjustable knobs for varying the response characteristics thereof during testing. The wearable aid may be as small as practical. A prototype aid having thumbwheel adjustments for the filter circuits has been constructed and packaged with outside dimensions of five by three by one and one-eighth inches, but obviously smaller sizes can be manufactured. An exemplary master aid has been constructed with outside dimensions of 15 by 10 by 4 and 1/2 inches. The basic hearing aid shown in FIG. 6, includes an integrated microphone/low-noise FET amplifier stage 50 followed by a low-noise amplifier section 52 which drives a bank of parallel and independently adjustable bandpass filters indicated generally at 54. The filters are side-by-side in frequency and are adjustable in gain only after initial frequency alignment. FIG. 7 is a scope trace showing a typical single filter response at a center frequency, fo, of 1 KHz. A summation circuit 56 adds all of the filter outputs on a common bus in a linear summation. The summed signal then is applied to a linear amplifier and drive circuit 58 which, in turn, drives a miniature magnetic receiver 60 of the hearing aid.

In order to prevent overdriving the receiver into the nonlinear region, an automatic saturation elimination (ASE) circuit 62 provides a gain controlled loop back to the front end circuits. The ASE feedback signal can either be sensed at the signal line 64 to the filter bank 54 or at the receiver drive point 66 in the hearing aid output. A volume control 68 ahead of the filter bank 54 permits the overall hearing aid gain to be set at any desirable quiescent value.

Low pass filter circuits 70 and 72 are used for B+ and B- noise filtering and decoupling at various points throughout the system as desired. The hearing aid is designed to operate from hearing aid batteries providing a balanced plus and minus voltage with respect to the common bus 64.

Turning to a more specific discussion of the system of FIG. 6, a miniature ceramic microphone with a built-in low-noise FET amplifier may be used as the input transducer stage 50. Units similar to the Knowles BL-1671 may be employed. This unit has a response from less than 100 Hz to greater than 8,000 Hz as measured by standard hearing aid microphone measurement techniques. With the system as shown in FIG. 6, a 1.3 volt dc supply voltage is used to power the build-in FET amplifier of the input transducer. Control of this voltage to lower levels is one way of controlling the front end gain of the hearing aid by such means as the ASE control loop 62. A dc voltage on the output leads, in combination with the audio signal, requires a decoupling capacitor prior to feeding the low noise preamplifier in the hearing aid front end.

To minimize front end noise contribution, a dual Darlington connected amplifier pair (such as a Motorola 2N5089 NPN low-level, low noise device) operating at low current levels and with a large input current limiting resistor, is included following the microphone circuit as more specifically illustrated in FIG. 8 in the low noise two stage preamplifier 80. The dual low-noise amplifier is connected to the volume control potentiometer 68 that sets the quiescent gain of the overall hearing aid.

Additional front end gain is provided by two operational amplifiers 82 following the low noise preamplifier 80 and volume control 68. These amplifiers may incorporate very low current drain IC operational amplifiers, such as the Solitron UC 4252 dual unit. Feedback resistors around each operational amplifier permit the gain to be set at any desired value within the operating range. A complementary pair driver stage 84, including devices such as Motorola 2N5089 and 2N5087 transistors, provides a push-pull drive signal to the signal line 64 which feeds the filter bank circuits 54. This same bus is an alternate source for feeding the ASE automatic gain control feedback loop 62 as noted earlier.

An exemplary filter bank is illustrated in FIG. 9 and includes six parallel filter networks numbered 1-6 of adjacent frequency bands, and each has independent gain control. However, it is to be noted that different numbers and types of filter networks may be used as desired. Active three pole filters are included which incorporate operational amplifiers such as the Solitron UC 4253C triple operational amplifier, in integrated circuit configuration. Each amplifier draws microamperes of current, which is of prime importance in minimizing battery drain for longer operating life.

As illustrated in FIG. 9, each filter band is made up of three operational amplifiers 90-92 in active filter circuit configurations. The first filter section 90 is a low pass filter followed by a high pass filter 91 and then a band pass filter 92. Selection of the proper resistors (R's) and capacitors (C's) determines the center frequency, band pass, ripple and gain of each three pole filter section. A gain potentiometer 94 is included at the input to each filter section to provide independent gain control for the particular frequency band represented.

The selection of the band limits is flexible during the initial alignment. Possible alignments include octave bands, one-third octave bands, unequal bands adjusted for optimum speech discrimination, and band with frequency gaps in special areas for selective sound elimination purposes. The operational amplifiers operate between a balanced positive and negative battery supply with a quiescent output level of zero volts. This permits maximum voltage swing of the output waveform prior to reaching saturation, as well as minimum quiescent current drain during absence of signal. The six filter outputs are linearly summed in a resistive summation network 56 prior to feeding the post amplification circuits 58 of the hearing aid.

In order to prevent distortion in the hearing aid during the presence of large audio signal levels, an automatic saturation elimination circuit (ASE) 62 is included as noted earlier. This circuit samples the audio signal either at the signal line 64 of the filter band 54 or at the drive point 66 for the hearing aid receiver. As illustrated in FIG. 10, the audio signal is detected in a voltage doubler circuit 96, passes through a low pass filter 97, and then feeds an NPN transistor common emitter driver stage 98. The latter stage 98 incorporates a device such as the Motorola 2N5089 transistor in a low current drain circuit. The output of this stage supplied B+ for the FET amplifier in the microphone assembly.

A large signal at the input of the ASE circuit 62 results in a drop in the amount of voltage supplied to the microphone FET and thus reduces the gain of the signal into the receiver front end. The response time of the circuit 62 is in the range of a few milliseconds and can be adjusted to other values if desired. A large filter capacitor 1 .mu.f at the collector of the driver transistor in stage 98 minimizes the noise applied to the FET amplifier B+ supply and also provides a time constant needed in the ASE loop to prevent loop oscillations. The diodes used in the doubler 96 may be types such as 1N914 low cost silicon units available from several manufacturers. The doubling action permits a smaller signal to activate the ASE loop 62 without the addition of transistor gain stages and the corresponding power dissipation.

A dual operational amplifier and complementary pair transistor driver similar to the corresponding circuits ahead of the filter bank 54 are used as a post amplifier circuit 58 to drive the miniature magnetic receiver assembly 60. This post amplification circuit is illustrated in FIG. 11. Like circuit components are used, including the dual integrated circuit operational amplifier 102 (like amplifier 82 of FIG. 8) and an NPN/PNP complementary driver transistor circuit 103 (like driver 84). The gains of the operational amplifiers are set by means of feedback resistor networks. Typical gain values of 10 dB per amplifier may be used in the post amplifier stages. Current setting resistors in the operational amplifier circuits permit quiescent operation with microamperes of drain. The driver stage (complementary pair) have the biases adjusted to provide a minimum current needed for driving the receiver. A balanced positive and negative power supply with respect to the signal line 64 permits low quiescent current drain in the absence of a signal. An optional output signal connection to the gain control loop 62 as described earlier permits the gain control sensing to be supplied at the receiver input.

Finally, the system of FIG. 6 preferably employs a miniature magnetic receiver. Various miniature magnetic receivers can be connected to the driver circuit of the hearing aid depending on the patient's requirements. For persons requiring more volume, larger diaphragm receivers can be used. Smaller receivers capable of being placed entirely within the ear canal can also be driven by the same driver stage.

The present embodiments of this invention are to be considered in all respects as illustrative and not restrictive, the scope of the invention being indicated by the appended claims rather than by the foregoing description, and all changes which come within the meaning and range of equivalence of the claims therefore are intended to be embraced therein.

Claims

What is claimed is:

1. Apparatus useful for measuring human auditory deficiency and/or providing compensatory amplification for aurally handicapped persons comprising:

input circuit means for receiving complex phonic signals to be selectively amplified over a plurality of pass bands,

selective audio amplification means comprising a plurality of independently adjustable filter means having adjacent pass bands for enabling independent adjustment of amplification within each of a plurality of adjacent audio pass bands, said amplification means being coupled with said input circuit means for receiving signals therefrom and each of said filter means providing output signals,

a plurality of compression means, each compression means coupled to the output of a corresponding filter means for providing output signal compression when the output signals from said corresponding filter means exceeds a predetermined level, and

summation means for combining the output signals after the output signals have been acted on by said plurality of compression means.

2. Apparatus as in claim 1 wherein said filter means are active filter means.

3. Apparatus as in claim 2 wherein each of said active filter means includes variable amplitude control means coupled with said active filter means.

4. Apparatus as in claim 3 wherein said input circuit means includes microphone means coupled with an amplifier, said amplifier being coupled with said amplifier means.

5. Apparatus as in claim 3 wherein each of said active filter means includes operational amplifier low pass, high pass, and band pass filters, said operational amplifier filters being formed of integrated circuits.

6. Apparatus as in claim 1 wherein the complex phonic signals are continuous discourse.