David A. Preves

Programmable multichannel hearing aid with adaptive filter

A hearing aid is programmable with dual-tone multiple-frequency signals, received through the hearing aid microphone, to adjust operating coefficients of signal conditioning circuitry in the aid. A DTMF receiver filters and detects DTMF tone pairs into digital words provided to a controller for decoding, some of the digital words representing programming instructions and others representing data. In accordance with the instructions, the controller conveys the data to memory operatively associated with a plurality of control ports to the signal conditioning circuitry, with operating coefficients of the conditioning circuitry determined by the contents of the memory.

Frequency selection circuit for hearing aids

A signal processing circuit for hearing aids includes a broadband peak detector for generating a control voltage based upon the sound pressure level of an incoming acoustical signal over its entire frequency spectrum. The control signal is used to determine the cut-off frequency of a voltage controlled adaptive high-pass filter. An amplified electrical signal, corresponding to the acoustical signal, also is provided to the high-pass filter. In setting the cut-off frequency, the control voltage causes the high-pass filter to selectively suppress the low frequency portion of the signal, generating a modified signal in which the noise component is reduced.

Hyperacusis: Case studies and evaluation of electronic loudness suppression devices as a treatment approach

Hyperacusis, as defined here, is a relatively rare condition in which the patient, with or without hearing loss, experiences severe loudness discomfort to everyday environmental sound levels. The case studies of 14 patients with severe hyperacusis are described; all wore passive attenuators (earplugs and/or earmuffs) in an attempt to alleviate their discomfort, frequently producing communication difficulties. These subjects were fitted binaurally with experimental electronic loudness suppression devices housed in in-the-ear casings. The devices supplied low-level amplification followed by an extreme form of amplitude compression for moderate or high-level inputs in an attempt to reduce loudness discomfort without reducing audibility. Many of the subjects were found to function with a wider dynamic range with the active devices compared with passive attenuators or the unoccluded ear, and most reported that they benefited from the devices in at least some listening situations.

A Single-chip Hearing Aid With One Volt Switched Capacitor Filters

An integrated circuit has been designed and fabricated in BiCMOS technology which contains three switchedcapacitor fourth-order filters and one continuous-time fourth-order filter, as well as all other circuitry needed for a modern hearing aid. The circuit operates from a single hearing aid battery and functions down to 1.1 Volt, while the switched-capacitor filters are designed to allow power supplies as low as 1.0 Volt.

A new technique for quantifying temporal envelope contrasts.

A new technique has been developed for precisely quantifying the temporal contrasts that exist between two sound samples. This technique is based on envelope subtraction, and generates an Envelope Difference Index that may be used to help clarify whether alteration of the natural speech envelope via amplification improves or degrades speech intelligibility. The Envelope Difference Index method may also be used to assess hearing aid saturation, and may have other applications as well. The technique is applicable whenever a precise quantitation of the difference between two temporal envelopes is required, regardless of stimulus duration.

Strategies for Enhancing the Consonant to Vowel Intensity Ratio With In the Ear Hearing Aids

Numerous investigators have suggested that increasing the consonant to vowel intensity ratio (CVR) may improve speech intelligibility. This investigation was designed to determine the extent to which analog circuits. small enough to fit into in the ear hearing aids, can increase the CVR, and whether CVR enhancement is of benefit to hearing-impaired listeners. Real ear CVRs, calculated from real ear recordings of nonsense syllables, were obtained from eight hearing-impaired listeners. Recordings from each listener were obtained through each of four hearing aid circuits: (1) an adaptive high-pass filter; (2) a faster acting adaptive high-pass filter; (3) the fast-acting adaptive high-pass filter with expansion; and (4) an infinite amplitude clipper. The amount of CVR enhancement was compared to performance of the subjects with a NST speech recognition task. Subjects also ranked the four circuits for amount of consonant emphasis provided. Results indicated that the four hearing aid circuits increased the real ear CVR by 4 to 6 dB, relative to unaided. Aided CVR varied, however, across circuits and between fricative and stop consonants. Performance on the NST recognition task was generally consistent with the amount of CVR increase provided. Rank ordering for consonant emphasis was consistent with aided CVR for stop consonants, but not for fricatives.

Progress achieved in setting standards for hearing aid/digital cell phone compatibility

This paper traces a brief history of the development of standards for assessing the compatibility of hearing aids and telephones. In particular, we will discuss standards for measuring the immunity of hearing aids to digital cell telephone (DCT) interference and the harmonization of the IEC and ANSI hearing aid immunity standards. To prevent acoustic feedback and to suppress environmental noise, hearing aid wearers have traditionally used an induction coil or telecoil in their instruments to pick up the magnetic signals emanating from telephones and inductive loops. The amount of magnetic “leakage” signal produced by a telephone is a function of the specific receiver design. Early telephone receivers, such as the U-type, had sufficient magnetic leakage to be sensed effectively by telecoils. However, frustration among hearing aid wearers attempting to use the telephone via telecoil began to arise in the mid-1960s after telephone companies “improved” their receiver designs to reduce their magnetic emissions.1 The telephone companies wanted the hearing aid industry to change its designs to use acoustic rather than inductive coupling to telephones. The hearing aid industry rejected this suggestion, producing a stalemate between the two industries on hearing aid-telephone compatibility. Led by David Saks in the early 1970s, consumer groups took action, which led to Senate hearings and a directive to the two industries to solve the problem together. Some of the first meetings between engineers from the two industries occurred in 1982-1983 when an ad hoc technical committee addressed the problem of poor performance for hearing aid wearers attempting to use a telephone with their telecoils via induction. This produced a greater understanding of the nature of the magnetic leakage signal emanating from the telephone. Ultimately, this committee produced the EIA/HIA RS-504 standard Magnetic Field Intensity Criteria for Telephone Compatibility with Hearing Aids.2 This standard was later incorporated as a labeling criterion in the Hearing Aid Compatibility Act (HAC) of 1988, which required all corded and cordless landline phones to be hearing aid compatible. However, the Federal Communications Commission (FCC) exempted cell phones from this requirement. Meanwhile, as use of DCTs increased, hearing-impaired people began to complain of a buzzing interference generated in their hearing aids that did not occur with analog cellular phones. This problem is caused by the DCT turning the RF signal on and off periodically at a low frequency rate that lies within the audio frequency band. Hearing aid circuitry may unintentionally demodulate this periodic signal, which results in a low-frequency buzz. Detailed descriptions of the nature of this interference can be found in Victorian3 and Preves.4 Early studies regarding this type of interference were published in Australia5,6 and Europe.7 This work led to formulation of IEC 60118-13 Electromagnetic Compatibility (EMC) – Product Standard in 1997, which is concerned with simulated far-field measurements of hearing aid immunity in a gigahertz transverse electromagnetic (GTEM) cell. At first, this standard addressed only “bystander” interference to hearing aids, but it is now being modified so it also includes hearing aid immunity measurements for a hearing aid wearer using a DCT.8

Assessing the feasibility of Bluetooth in hearing rehabilitation

In the midst of a variety of other wireless alternatives, the growth of Bluetooth products since the introduction of the protocol in 1994 has been phenomenal. Beyond its initial use in cell phones, Bluetooth technology is an accessory in many automobiles, MP3 players, wireless watches, jewelry, and toys, just to name a few.1 In fact, Bluetooth technology has been incorporated in the NOAHlink programming device (HIMSA) to eliminate the bothersome wires that traditionally have connected the host computer to the hearing aids that are being programmed. With its inherent features of security, portability, robustness, low power, and low cost, Bluetooth is also an obvious candidate for applications in hearing instruments. As this paper will make clear, however, practical limitations with the current Bluetooth standard are limiting its application to hearing aids, and Bluetooth technology will not become accessible to the majority of hearing aid wearers until these limitations are resolved.

Hearing aid having a supply source providing multiple supply voltages

A system includes a number of electronic devices that uses multiple voltage sources and a supply source to provide the different supply voltages without up-converting a voltage level or down-converting a voltage level. In an embodiment, the supply source is realized using a battery having multiple voltage taps, where the battery provides the multiple voltage sources. In an embodiment, the system is a hearing aid. A single battery includes a common substrate on which a number of battery regions is formed, where each battery region provides a supply voltage at a rated voltage level different than the other battery regions. The common substrate may be configured as a rigid platform, a flexible platform in a folded configuration, a flexible platform in a rolled configuration, or other platform configurations that provide for multiple battery regions on a single platform.

Satisfying patient needs with nine fixed acoustical prescription formats

INTRODUCTION Conventional hearing aid fittings are usually accomplished by providing a set of predetermined electroacoustic characteristics to the patient’s ear based on some fitting prescription or rationale. The desired characteristics may then be preset in a hearing aid and verified by 2-cc coupler measurements or programmed while in the patient’s ear. The prescription can then be verified by insertion or functional gain measures. In both procedures, the hearing aids may be adjusted further to meet any subjective loudness, quality, or clarity requirements of the patient. This approach may lead to confusion on the part of the dispensing professional when attempting to adjust programmable or trimmer settings regarding what effect each adjustable parameter has on performance. There are differing opinions on how best to compensate for a specific type and severity of hearing loss. Some fitting prescriptions attempt to predict desired gain functions from threshold data.1 Others calculate target characteristics from direct loudness scaling from individual patients.2 The rationale behind some prescriptive techniques is to ensure loudness equalization (amplifying components of speech to contribute equally to loudness at different frequencies), whereas the goal of others is to perform loudness normalization (to restore the overall loudness of sounds to normal loudness and to restore the relative loudness of different frequency components). Because of the inherent differences among these fitting procedures, if they are applied to a single patient, the electroacoustic outcome of each may sound different to the patient depending on the formula selected by the fitter. Although there has been no single formula sufficiently compelling to be embraced by the majority of clinicians, the consensus is that most are beneficial. In addition to the acoustic variances resulting from the selection of one fitting formula versus another, there is considerable interaction between the prescription used and subjective performance and preference. Further, sound quality and clarity are not simply subjective in nature, but are influenced by factors associated with the pathologic changes in the individual’s ear, such as loss of hair cells and impaired function of the peripheral and central neural pathways. These factors cannot be compensated for by using any single fitting formula, nor can the considerable differences in subjective sound quality, clarity, or loudness be accurately predicted. Thus, an important step in hearing aid fittings is fine-tuning or changing electroacoustic characteristics to improve performance or subjective preference. Recent developments of non-linear fitting formulas, along with improved compression circuitry, have resulted in more natural perception of soft, medium, and loud sounds by patients wearing hearing aids. As with linear formulas, each of the non-linear hearing aid prescriptive formulas is based on a particular fitting rationale which results in differences in the final fitting.