US3784750A - Apparatus and prosthetic device for providing electronic correction of auditory deficiencies for aurally handicapped persons - Google Patents

Apparatus and prosthetic device for providing electronic correction of auditory deficiencies for aurally handicapped persons Download PDF

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US3784750A
US3784750A US00229322A US3784750DA US3784750A US 3784750 A US3784750 A US 3784750A US 00229322 A US00229322 A US 00229322A US 3784750D A US3784750D A US 3784750DA US 3784750 A US3784750 A US 3784750A
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amplifier
amplification
filter
filter means
test
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W Stearns
J Lauchner
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SHALAKO RESOURCE SYST INC US
SHALAKO RESOURCE SYSTEMS
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SHALAKO RESOURCE SYSTEMS
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    • H—ELECTRICITY
    • H03—ELECTRONIC CIRCUITRY
    • H03G—CONTROL OF AMPLIFICATION
    • H03G9/00—Combinations of two or more types of control, e.g. gain control and tone control
    • H03G9/02—Combinations of two or more types of control, e.g. gain control and tone control in untuned amplifiers
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R25/00—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception
    • H04R25/50—Customised settings for obtaining desired overall acoustical characteristics
    • H04R25/505—Customised settings for obtaining desired overall acoustical characteristics using digital signal processing
    • H—ELECTRICITY
    • H04—ELECTRIC COMMUNICATION TECHNIQUE
    • H04R—LOUDSPEAKERS, MICROPHONES, GRAMOPHONE PICK-UPS OR LIKE ACOUSTIC ELECTROMECHANICAL TRANSDUCERS; DEAF-AID SETS; PUBLIC ADDRESS SYSTEMS
    • H04R25/00—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception
    • H04R25/35—Deaf-aid sets, i.e. electro-acoustic or electro-mechanical hearing aids; Electric tinnitus maskers providing an auditory perception using translation techniques
    • H04R25/356—Amplitude, e.g. amplitude shift or compression

Definitions

  • PATENTED JAN 8 7 SHEU 6 BF 8 m sum PATENTED JAN 81974 SHEET 8 0f 8 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. 7
  • This invention relates to the sound amplification arts, and to their application in the amelioration of auditory deficiencies resulting from damage to the sensorineural structures of the human'ear. It relates particularly to methods and apparatus for detecting and speci- .fying 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. lt constitutes a significant barrier to adequate communication in to percent of the total United States population, and in more than percent of the population over 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 doesnot 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 adefirst order neurons due to damage to these structures.
  • 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.
  • Each of the phonic units of a spoken word is a complex sound, composed of several frequencies clustered ina more-or-lessdefinable 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.
  • the threshold audiogram curve represents an individuals 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.
  • Control of acoustic output in current hearing aids is ordinarily achieved through manipulation of frequency response, which refers to the acoustic output ofa sound transmission system at each of the frequencies within its pass band when the input level is maintained constant for all frequencies.
  • frequency response refers to the acoustic output ofa 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.
  • lt 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).
  • the range of a sensori-neurally impaired ear may be as little as 10 or 15 dB, generally over a limited frequency spectrum range.
  • 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.
  • 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.
  • 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 blockdiagram of a wearable hearing aid according to the present invention.
  • FIG. 6 is a'mo're 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.
  • FIGS. 8 through 11' are specificcircuitdiagrams of the aid of FIG. 6; f
  • 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.
  • 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.
  • the non-test ear is occluded by an insert earplug and a circumaural muff to block any auditory perception by that ear.
  • 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 patients head.
  • the test signals used to accomplish such equalization are narrow bands of noise.
  • FIG. 1 illustrates an organization of test equipment used in testing speech discrimination of a subject to the extent necessary to practice the methodof the present invention.
  • Speech from an audio tape recorder and playback apparatus 10 is applied to an audio mixer network 11.
  • a pink or whitenoise 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 through F,,.
  • Each of the filter networks F through F has a discrete pass band and the entire combination 13 preferablycovers a frequency range from approximately Hz to at least'6,300 Hz.
  • the filters-F, through F divide the audible frequency spectrum into adjacent pass bands.
  • each filter F, through F 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.
  • 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 l4, 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.
  • 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 l02-lOG hearing aid receiver
  • the vacuum tube (RMS) voltmeter 17 a Ballentine model 320.
  • FIGS. 20 and 2b The response curves for two test subjects are shown in FIGS. 20 and 2b together with discrimination test scores.
  • FIG. 2a pertains to one ear of one test subject
  • 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;
  • curve (0) is an examiners 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.
  • 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.
  • the test material consists of standardized phonetically balanced lists of words, each list comprising 50 words to whichthe subject must respond by repeating each word immediately after it is presented by a tape recorded speaker.
  • 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.
  • 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. l, 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 subjects personal aid was 68 percent. Test No.
  • FIG. 4 illustrates the frequency response of the master hearing aid after it was set by Subject X in the test.
  • Test N0. 5 involved an audiological evaluation of Subject Y conducted by a university Speech and Hearing Clinic. The phonetically balanced speech discrimination score was 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 car. 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 thespeech 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 trimmedby 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 eachone 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.
  • FIG. 5 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 re.- sponse 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 lowimpedance 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, orall-pass, network 28.
  • the bandpass filters 27 each have a bandpass and amplitude control 26 suitable for closelyproviding or approximating the desired response curve (e.g., curve 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 beas wide or. as narrow as required 0t obtain optimal speech discrimination and need not be related in any octave relationship orfractional 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 ofthe, spectrum; separate response control for each ear (binaural); ease of readjustment as the patients requirements change with time; this maybe accomplished by the replacement of filter elements having different pass bands and adjusted for different amplitudes.
  • a narrow band notch rejection filter may be added after the summation network to alleviate narrow band resonance problems-observed in some patients.
  • Theconcept 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.
  • the circuit of FIG. 6 can beused 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 l0.by 4 and /5 inches.
  • the basic hearing aid shown in FIG. 6, includes an integrated microphone/low-noise FET amplifier stage 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 ingain only after initial frequency alignment.
  • FIG. 7 is a scope trace showing a typical single filter response at a center frequency, f0, of l KI-Iz.
  • 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.
  • an automatic saturation elimination (ASE) circuit 62 provides a gain controlled loop back to the front end circuits.
  • TheASE feedback signal can either be sensed at the signal line 64 to the filter bank 54 or at the receiver drive point66 in the hearing aid output.
  • a volume control 63 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 8+ 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 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.
  • 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.
  • Additionalfront end gain is provided by two operational amplifiers 82 following the low noise preamplifier and volume control 68. These amplifiers may incorporate very low current drain lC 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. I i
  • FIG. 9 An exemplary filter bank is illustrated in FIG. 9 and includessix 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.
  • each filter band is made up of three operational amplifiers 90-92 inactive 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 (Rs) and capacitors (Cs) 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.
  • 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.
  • 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 8+ for the FET amplifier in the microphone assembly.
  • a large signal at the input of theASE circuit 62 results in a drop in the amount of voltage supplied to the microphone PET 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 uf 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 lN9l4 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
  • 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 aiddepending on the patients 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.
  • 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. receivingsignals therefrom and each of said filter means providing output signals,
  • 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
  • summation means for combining the output signals after the output signals have been acted on by said plurality of compression means.
  • each of said active filter means includes variable amplitude control means coupled with said active filter means.
  • said input circuit means includes microphone means coupled with an amplifier, said amplifier being coupled with said amplifier means.
  • 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.
  • Figure 5 should appear as shown on the attached heet.

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  • Health & Medical Sciences (AREA)
  • General Health & Medical Sciences (AREA)
  • Neurosurgery (AREA)
  • Otolaryngology (AREA)
  • Physics & Mathematics (AREA)
  • Engineering & Computer Science (AREA)
  • Acoustics & Sound (AREA)
  • Signal Processing (AREA)
  • Measurement Of The Respiration, Hearing Ability, Form, And Blood Characteristics Of Living Organisms (AREA)
  • Tone Control, Compression And Expansion, Limiting Amplitude (AREA)
US00229322A 1972-02-25 1972-02-25 Apparatus and prosthetic device for providing electronic correction of auditory deficiencies for aurally handicapped persons Expired - Lifetime US3784750A (en)

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JP (1) JPS4898698A (enrdf_load_stackoverflow)
CA (1) CA984460A (enrdf_load_stackoverflow)
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Cited By (32)

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US3989904A (en) * 1974-12-30 1976-11-02 John L. Holmes Method and apparatus for setting an aural prosthesis to provide specific auditory deficiency corrections
US3995114A (en) * 1974-05-09 1976-11-30 Dahlberg Electronics, Inc. Ultra low current amplifier
US4025723A (en) * 1975-07-07 1977-05-24 Hearing Health Group, Inc. Real time amplitude control of electrical waves
US4099035A (en) * 1976-07-20 1978-07-04 Paul Yanick Hearing aid with recruitment compensation
US4139730A (en) * 1977-08-10 1979-02-13 Barbara Franklin Method of testing human auditory responses
US4224468A (en) * 1978-10-05 1980-09-23 Calder Jr Howard B Masking level difference adaptor for audiometers
US4276781A (en) * 1978-10-09 1981-07-07 U.S. Philips Corporation Method of and arrangement for adapting a hearing aid
FR2497592A1 (fr) * 1981-01-07 1982-07-09 Gougelot Louis Marie Dispositif pour la reeducation de l'audition et procede pour la realisation de ce dispositif
EP0071845A3 (en) * 1981-08-06 1983-04-20 Siemens Aktiengesellschaft Apparatus for compensating hearing deficiencies
US4390748A (en) * 1979-12-21 1983-06-28 Siemens Aktiengesellschaft Electro-acoustical measuring device and method
US4484345A (en) * 1983-02-28 1984-11-20 Stearns William P Prosthetic device for optimizing speech understanding through adjustable frequency spectrum responses
WO1987007464A1 (en) * 1986-05-27 1987-12-03 Voroba Technologies Associates Patient controlled master hearing aid
US4764957A (en) * 1984-09-07 1988-08-16 Centre National De La Recherche Scientifique-C.N.R.S. Earpiece, telephone handset and headphone intended to correct individual hearing deficiencies
WO1989008353A1 (en) * 1988-02-23 1989-09-08 Resound Corporation Improved multi-band programmable compression system
WO1989008371A1 (en) * 1988-02-23 1989-09-08 Resound Corporation Improved low voltage programmable compressor
US4941179A (en) * 1988-04-27 1990-07-10 Gn Davavox A/S Method for the regulation of a hearing aid, a hearing aid and the use thereof
US5278912A (en) * 1991-06-28 1994-01-11 Resound Corporation Multiband programmable compression system
US5434924A (en) * 1987-05-11 1995-07-18 Jay Management Trust Hearing aid employing adjustment of the intensity and the arrival time of sound by electronic or acoustic, passive devices to improve interaural perceptual balance and binaural processing
US5500902A (en) * 1994-07-08 1996-03-19 Stockham, Jr.; Thomas G. Hearing aid device incorporating signal processing techniques
US20050008177A1 (en) * 2003-07-11 2005-01-13 Ibrahim Ibrahim Audio path diagnostics
US6885752B1 (en) 1994-07-08 2005-04-26 Brigham Young University Hearing aid device incorporating signal processing techniques
US20050111683A1 (en) * 1994-07-08 2005-05-26 Brigham Young University, An Educational Institution Corporation Of Utah Hearing compensation system incorporating signal processing techniques
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US20070009130A1 (en) * 2001-08-10 2007-01-11 Clear-Tone Hearing Aid BTE/CIC auditory device and modular connector system therefor
US20070064966A1 (en) * 2001-08-10 2007-03-22 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US20070185710A1 (en) * 2004-03-11 2007-08-09 Rion Co., Ltd. Apparatus and method for preventing senility
EP2579619A1 (en) * 2011-10-07 2013-04-10 Starkey Laboratories, Inc. Audio processing compression system using level-dependent channels
EP2898705B1 (en) 2012-09-18 2017-08-23 Sonova AG Cic hearing device
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EP0071845A3 (en) * 1981-08-06 1983-04-20 Siemens Aktiengesellschaft Apparatus for compensating hearing deficiencies
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US5434924A (en) * 1987-05-11 1995-07-18 Jay Management Trust Hearing aid employing adjustment of the intensity and the arrival time of sound by electronic or acoustic, passive devices to improve interaural perceptual balance and binaural processing
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AU611273B2 (en) * 1988-02-23 1991-06-06 Resound Corporation Improved multi-band programmable compression system
US4941179A (en) * 1988-04-27 1990-07-10 Gn Davavox A/S Method for the regulation of a hearing aid, a hearing aid and the use thereof
US5278912A (en) * 1991-06-28 1994-01-11 Resound Corporation Multiband programmable compression system
US5500902A (en) * 1994-07-08 1996-03-19 Stockham, Jr.; Thomas G. Hearing aid device incorporating signal processing techniques
US5848171A (en) * 1994-07-08 1998-12-08 Sonix Technologies, Inc. Hearing aid device incorporating signal processing techniques
US8085959B2 (en) 1994-07-08 2011-12-27 Brigham Young University Hearing compensation system incorporating signal processing techniques
US6885752B1 (en) 1994-07-08 2005-04-26 Brigham Young University Hearing aid device incorporating signal processing techniques
US20050111683A1 (en) * 1994-07-08 2005-05-26 Brigham Young University, An Educational Institution Corporation Of Utah Hearing compensation system incorporating signal processing techniques
US20070064966A1 (en) * 2001-08-10 2007-03-22 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US20070009130A1 (en) * 2001-08-10 2007-01-11 Clear-Tone Hearing Aid BTE/CIC auditory device and modular connector system therefor
US8050437B2 (en) 2001-08-10 2011-11-01 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US8976991B2 (en) 2001-08-10 2015-03-10 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US8094850B2 (en) 2001-08-10 2012-01-10 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US7606382B2 (en) 2001-08-10 2009-10-20 Hear-Wear Technologies LLC BTE/CIC auditory device and modular connector system therefor
US20090296969A1 (en) * 2001-08-10 2009-12-03 Hear-Wear Technologies, Llc Bte/cic auditory device and modular connector system therefor
US9591393B2 (en) 2001-08-10 2017-03-07 Hear-Wear Technologies, Llc BTE/CIC auditory device and modular connector system therefor
US20100226520A1 (en) * 2001-08-10 2010-09-09 Hear-Wear Technologies, Llc BTE/CIC Auditory Device and Modular Connector System Therefor
US20060167681A1 (en) * 2002-12-19 2006-07-27 Hiroshi Rikimaru Diagnosis device and diagnosis method
US7421392B2 (en) 2002-12-19 2008-09-02 Rion Co., Ltd. Diagnosis device and diagnosis method
US20050008177A1 (en) * 2003-07-11 2005-01-13 Ibrahim Ibrahim Audio path diagnostics
US8223982B2 (en) * 2003-07-11 2012-07-17 Cochlear Limited Audio path diagnostics
US20070185710A1 (en) * 2004-03-11 2007-08-09 Rion Co., Ltd. Apparatus and method for preventing senility
US7729907B2 (en) 2004-03-11 2010-06-01 Rion Co., Ltd. Apparatus and method for preventing senility
US8861760B2 (en) 2011-10-07 2014-10-14 Starkey Laboratories, Inc. Audio processing compression system using level-dependent channels
EP2579619A1 (en) * 2011-10-07 2013-04-10 Starkey Laboratories, Inc. Audio processing compression system using level-dependent channels
US9736583B2 (en) 2011-10-07 2017-08-15 Starkey Laboratories, Inc. Audio processing compression system using level-dependent channels
EP2898705B1 (en) 2012-09-18 2017-08-23 Sonova AG Cic hearing device
US10433089B2 (en) * 2015-02-13 2019-10-01 Fideliquest Llc Digital audio supplementation
US11134867B2 (en) * 2016-10-19 2021-10-05 Mimi Hearing Technologies GmbH Method for accurately estimating a pure tone threshold using an unreferenced audio-system
US20180270590A1 (en) * 2017-03-17 2018-09-20 Robert Newton Rountree, SR. Audio system with integral hearing test
US10375489B2 (en) * 2017-03-17 2019-08-06 Robert Newton Rountree, SR. Audio system with integral hearing test
US10848877B2 (en) 2017-03-17 2020-11-24 Robert Newton Rountree, SR. Audio system with integral hearing test

Also Published As

Publication number Publication date
GB1419851A (en) 1975-12-31
DK138149B (da) 1978-07-17
DK138149C (da) 1978-12-18
CH560038A5 (enrdf_load_stackoverflow) 1975-03-27
JPS4898698A (enrdf_load_stackoverflow) 1973-12-14
FR2173316B1 (enrdf_load_stackoverflow) 1978-02-10
DE2309026B2 (de) 1976-08-19
FR2173316A1 (enrdf_load_stackoverflow) 1973-10-05
DE2309026A1 (de) 1973-09-06
CA984460A (en) 1976-02-24

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