Insufficient Gain

For patients with severe to profound hearing loss, sound amplification (or gain) is a primary concern. For a patient with air-conduction thresholds of 80 dB hearing level (HL) to perceive quiet sounds at a normal threshold of 0 dB HL, a hearing aid must amplify sound intensity by 80 dB, generating a 10,000-fold increase in sound pressure wave amplitude and a 100,000,000-fold increase in sound power intensity.² This represents the limit of existing traditional hearing aids. The maximum amplification at 1 kHz for a behind-the-ear (BTE) Phonak SuperFront PPCL4 power digital hearing aid is about 75 to 82 dB.³ Gain generally scales with hearing aid size, so that cosmetically less obtrusive hearing aids offer less gain. The maximum gains for digital in-the-ear (ITE), in-the-canal (ITC), and completely-in-canal (CIC) hearing aids currently are about 55 to 65 dB, 45 to 55 dB, and 35 to 50 dB.³

Acoustic Feedback

In practice, feedback often limits the useful gain of traditional hearing aids to less than the above-listed maximum gains. Acoustic waves from the hearing aid speaker leak through the air space between the hearing aid body and the external auditory canal wall back to the microphone, where (for a subset of frequencies) they add to existing microphone input and are amplified further. The resulting positive feedback loop causes the familiar squealing and hum that signify a poorly fitting or overamplified aid. Potential for feedback is worst for CIC aids, in which the microphone is closest to the speaker, and for ears with mastoid bowls, in which an air seal is difficult to obtain. At very high amplification, it is a problem even for BTE aids. Fitting aids tightly into the external auditory canal can decrease feedback, but at the cost of increased discomfort, and the risk of otitis externa, autophony, and blockage of natural sound input.

Distortion of Spectral Shape and Phase Shifts

Traditional hearing aids are limited in the frequency range over which they amplify sound. Most are optimized for performance in the speech band (500 to 2000 Hz) and cannot provide much amplification at less than 100 to 200 Hz or greater than 5000 to 6000 Hz. Although the most important spectral components for speech recognition are amplified, the loss of “bass” and “treble” can give the resultant sound percept an artificial character. Isolated severe hearing loss at low frequencies (e.g., in Meniere’s disease) or high frequencies (e.g., in presbycusis or ototoxic exposure) can be difficult to remediate with traditional hearing aids without overamplifying the midrange frequencies at which a patient may have normal hearing. Even in the 500- to 2000-Hz band, steep changes in audiometric threshold at neighboring frequencies (e.g., noise-induced hearing loss) cannot be perfectly fit because of inherent limitations in the rate of change of gain across frequencies. Too steep a change in amplifier gain across frequencies typically imparts phase shifts that distort sound percepts such as pitch and timbre. Digital signal processing technology has greatly improved spectral shaping options compared with analog aids, but these fundamental limitations persist.

Nonlinear Distortion

Similar to most other amplifiers, hearing aids are designed under assumption of linearity—that is, if a given sound input at the microphone leads to some output at the speaker (amplified and spectrally shaped by the aid), doubling the sound input’s magnitude should exactly double the output, without changes in spectral content or phase shifts. For high-intensity speaker output, this assumption breaks down as the speaker is driven into the range of movements for which it begins to saturate or clip. An input sinusoid can appear at the output as a waveform with blunted peaks. This nonlinear distortion imparts aberrant spectral components into the sound percept, giving it an artificial or robotic character. Although digital signal processing can mitigate distortion effects within an aid’s amplification circuitry, distortion produced by a speaker generating loud sounds in air remains a fundamental limitation of traditional hearing aids.

Occlusion Effects

To minimize feedback, most traditional hearing aids are fit to create an airtight seal with the external auditory canal wall, with the hearing aid’s speaker being isolated in the occluded ear canal. Canal occlusion has several undesirable effects. First, it can be uncomfortable because of pressure on canal skin. Second, it increases the likelihood of otitis externa because of disturbance of wax egress and air circulation. Some patients with chronic suppurative otitis media cannot tolerate hearing aids because of exacerbation of otorrhea. Third, it causes autophony and a sense of aural fullness that can worsen with changes in ambient barometric pressure. Fourth, it blocks the normal pathway for sound entry to the ear. Finally, canal occlusion disrupts the spectral shaping that normally occurs as a result of external auditory canal resonances.

Poor Appearance

Many patients reject hearing aids because of the social stigma of being seen as old or infirm. Eight percent of hearing aid candidates wearing the less noticeable ITC hearing aids listed poor cosmetic appearance as a major contributing factor in their decision not to use a hearing aid.⁴ Hearing aids are difficult to conceal in patients who are balding or who wear their hair short. Miniaturization of electronics continues to improve hearing aids in this regard, and CIC hearing aids are essentially invisible to the casual observer. Miniaturization comes at the cost of lower gain, more feedback, and higher cost, however. Ultimately, battery size becomes the limiting factor, with smaller batteries costing more and requiring more frequent replacement.

Poor Transduction Efficiency

Loss of energy because of impedance mismatching and transduction losses is another inherent drawback of conventional hearing aids. The mechanical impedance (change in pressure for a given displacement flow) of the air-filled external auditory canal differs from that of the fluid-filled cochlea. Without a middle ear mechanism, most of the acoustic energy striking the stapes footplate reflects back into the air. When the tympanic membrane and ossicular chain are functioning normally, they act as an impedance-matching transformer by virtue of the relative areas of the tympanic membrane and stapes footplate, and by the lever action of the ossicular chain.⁵ The relatively large-displacement, low-pressure movements of air against the tympanic membrane are transduced to the relatively small-displacement, high-pressure movements of the footplate. Except for bone-conducting hearing aids, all traditional hearing aids use a speaker to output an (amplified) acoustic wave into the air of the external auditory canal. When this middle ear apparatus is malfunctioning, as occurs in otosclerosis, tympanic membrane perforation, or a canal wall down mastoidectomy, traditional hearing aids must overcome the impedance mismatch. The result is reduced effective gain, increased distortion, or both.

Even when the ossicular chain is functioning normally, the transduction of acoustic energy from air at the input of a traditional hearing aid to stapes footplate motion is imperfect. Whenever a signal flows from one physical realm (e.g., electric current in a speaker) to another (e.g., acoustic waves in air), some energy is lost (e.g., to heat) and noise or distortion can add to the desired signal. There are several transduction steps for a traditional hearing aid—from acoustic waves in ambient air, to electric current in a microphone, to a larger electrical current in a speaker wire or piezoelectric driver, to acoustic waves in air within the external auditory canal, to movement of the tympanic membrane and ossicular chain, to acoustic waves in perilymph, to hair cell stereociliary deflection and depolarization. Most steps are unavoidable (although cochlear implants bypass these steps through direct cochlear nerve stimulation). Directly coupling a hearing aid actuator movement of the ossicular chain can delete some of these transduction steps, however, boosting gain and reducing distortion. Nearly all implantable hearing aids make use of this approach.

Recruitment and Compression of Dynamic Range

A primary source of frustration of patients with SNHL is reduced speech intelligibility. The common complaint, “I can hear, but I cannot understand,” underscores this point. This problem relates to abnormal frequency resolution and aberrant patterns of growth in loudness, each of which reduces speech intelligibility in challenging listening conditions such as noisy surroundings.^6,7

SNHL also imposes severe constraints on the dynamic range of perceived sound. For normal listeners, the dynamic range from sensing soft sounds to the loudest tolerable noise is more than 100 dB. Within this wide dynamic range of hearing lies an approximately 35-dB dynamic range of conversational speech. In contrast, the dynamic range of patients with SNHL is often narrowed by an increase in the threshold of audibility and a lowering of the ceiling of tolerance to high-intensity sounds. This compaction of dynamic range leads to “recruitment,” an abnormal growth in loudness as sound intensity increases.^8,9 What sounds normal for someone with normal hearing may be too soft for someone with recruitment, whereas what is too loud for someone with normal hearing is also too loud for a patient with recruitment. In effect, there is a much narrower range of sound intensity that a patient with recruitment can tolerate. Further adding to the difficulty, recruitment is observed in frequencies that are most impaired, in the high frequencies, which also carry crucial information for speech understanding.

Recruitment remains one of the principal challenges of hearing aid rehabilitation, and is responsible for a common phenomenon most clinicians dealing with hearing loss have witnessed or experienced: At average speaking levels, an individual with recruitment may ask a speaker to talk more loudly, yet with even a slight increase in intensity the speech becomes intolerably loud, and the speaker is told not to shout. Although many hearing aids can be programmed to prevent sound from amplifying into a range that is uncomfortable, even the most advanced aids cannot fully replicate the complex, nonlinear response patterns of a healthy cochlea.⁹

Disordered Perception of Pitch and Sound Localization

Another challenge faced by conventional hearing aids relates to impaired pitch resolution and sound localization. Although simple amplification strategies can compensate for this loss of sensitivity, they do little to restore the exquisite selectivity of the cochlea in organizing pitch information. Much of the benefit provided by the normal auditory system is based on the presence of two ears that independently sample the acoustic environment and send this information to the brain for comparison and analysis. Sound detection is an important first step in hearing, and can be performed reasonably well by one sensitive ear. There is significant additional auditory benefit, however, when sound input occurs through both ears. When only one ear functions, or when there is a drastic difference in hearing levels between ears, the ability to locate the source of sound decreases, as does suppression of background echoes and the ability to “block out” background noise.

All of the aforementioned factors play a role to various degrees in patients, and combine to increase the noncompliance rate in hearing aid users. For most patients, these are the principal reasons why hearing aids are unacceptable.¹⁰ With all these challenges concerning conventional hearing aids, it is not surprising that many hearing aid users feel frustrated when hearing aids fail to perform to expectations.

Transducer Design

Conventional hearing aids function by receiving acoustic energy through a microphone, processing and amplifying the signal (in either an analog or a digital fashion), and transmitting the signal through a speaker adjacent to the eardrum. This amplified sound travels through the normal auditory pathway of the tympanic membrane and ossicles to the inner ear. Implantable middle ear hearing aids also employ a microphone that senses sounds and a signal processor that amplifies and alters the sound signal, just as in a conventional hearing aid. From here, they differ from conventional hearing aids. In one of several mechanisms unique to each device, the implantable middle ear hearing aids convert the electric signal into a mechanical energy that is coupled directly to the ossicular chain (see Fig. 157-1).The critical components of these systems are the transducer, which enables the device to output a signal that can be received by the ossicles, and the mechanism by which the transducer is coupled to the ossicles, which can either be a direct or indirect contact.¹⁵ Variations in these critical components distinguish the various implant designs (see Fig. 157-2). Goode and colleagues¹¹ previously outlined the history of the development of implantable middle ear hearing aids.

Two basic types of transducers have been incorporated into middle ear implantable hearing aids: piezoelectric and electromagnetic. Piezoelectric implantable hearing aids function by connecting the ossicles to an amplifier using a piezoelectric crystal vibrator. Piezoelectric materials are dielectric materials with coupled electrical and mechanical properties. Applying a voltage across an appropriately designed piezoelectric rod causes it to bend or lengthen, with a predictable change in deflection based on the voltage applied. Within the implant, sounds picked up by a microphone are converted by the signal processor to an electrical signal, and sent to the piezoelectric rod. As this rod vibrates in response to the converted auditory signal, it lies in direct contact with the ossicles. In this manner, sound waves are transferred directly to the ossicles, which travel along the normal auditory pathway. A critical feature of implants that use a piezoelectric transducer is direct contact between the piezoelectric unit and the ossicles.

Implantable middle ear hearing aids may also incorporate an electromagnetic transducer. These transducers generate a magnetic field using a coil carrying a current that encodes the microphone output. This magnetic field from the coil can cause motion in a ferromagnetic substance that is nearby, causing it to vibrate.¹⁵ As applied to middle ear implants, sound received from a microphone is converted by the signal processor, amplified, and sent to the electromagnetic transducer. The electromagnetic field that is set up vibrates a ferromagnetic unit coupled directly to the ossicles. In contrast to the piezoelectric unit that must be in direct contact with the ossicles, electromagnetic transducers need only to be near the ossicles, which have attached to them the ferromagnetic device that transmits the auditory signal to the ossicles.

These two types of amplification have different advantages for use in an implantable hearing aid. Electromagnetic transducers do not directly contact the ossicular chain, but rely on electromagnetic transmission to a ferromagnetic unit that is attached to the ossicles. Electromagnetic transducers can be packaged in a smaller housing, an important factor in the ear. In contrast, piezoelectric devices are larger and couple directly to the ossicular chain. The advantage of this type of transducer is its ability to deliver more distortion-free amplification directly to the ossicular chain.¹⁵ In either event, to accommodate patients with moderate to severe SNHL in the 55- to 90-dB hearing level range, any implant, whether electromechanical or piezoelectric, should be able to deliver a maximum mechanical stimulation output equivalent to 120 to 130 dB sound pressure level (SPL).^16,17 How each implant achieves this output level is described subsequently.

Ossicular Coupling

As noted by Huttenbrink,¹⁸ to obtain optimal efficiency in implantable hearing aids, they must adapt to the mechanics of the middle ear and ossicular chain. As the third ossicle in the chain, and the one in continuity with the inner ear, the motion of the stapes is crucial in sound transmission. The stapes moves in a piston-like fashion.¹⁹ The ideal motion of the stapes by a prosthesis would be in this piston-like fashion. Any implant that directly attached to the stapes would function best by mimicking this motion.^20,21

Various implants employ numerous ways to couple vibration stimuli to the inner ear. Some employ a piezoelectric transducer to contact either the stapes head directly (e.g., St. Croix Envoy/Esteem-Hearing Implant) or the incus body (e.g., Implex TICA). Others use an electromagnetic transducer typically attached to the incus (e.g., Med-El Vibrant Soundbridge, Otologics Carina/MET), but sometimes adapted to contact the stapes capitulum, stapes footplate, or round window membrane. Others clip to the incudostapedial joint (e.g., Soundtec Direct Drive System). Each of these systems is described in greater detail individually.

Total versus Partially Implantable Hearing Aids

Implantable middle ear hearing devices may be either totally or partially implantable. Partially implantable devices consist of an external microphone and speech processor that are connected to a transmitter with an external coil that transmits electrical energy transcutaneously to the internal device, in much the same way cochlear implants are now powered. The battery-powered system is contained within the external device, decreasing the size of the implanted component. The internal device consists of an internal receiving coil, which provides electric energy to the mechanical driver connected to the ossicular chain.

In contrast, the fully implantable systems house all of the above-mentioned components within the implanted portion of the device, including the battery pack. Although the size and complexity of the implanted components are increased, the visibility of the device is reduced, a feature that is desirable among hearing aid users. Because of the finite life of the rechargeable batteries, the totally implantable designs employed in clinical settings so far require reoperation at approximately 5-year intervals to exchange the spent device with one containing a new battery.

Vibrant Soundbridge

To date, the implantable middle ear hearing aid with the most clinical data is the Vibrant Soundbridge, manufactured by Symphonix Corporation until 2002, when Symphonix was dissolved and the Vibrant technology was bought and later re-released by a subsidiary of the Med-El Corporation (Innsbruck, Austria), which resumed Vibrant Soundbridge sales in 2004. Because it was the first commercially available, FDA-approved middle ear hearing aid, a thorough review of this device is instructive.

The Vibrant Soundbridge is a semi-implantable hearing aid that was first introduced in 1997 (Fig. 157-3). Since its introduction in 1997, more than 2300 patients have been implanted worldwide.¹⁴ The device was approved for commercial availability in Europe in 1998, and in the United States in 2000. The electromagnetic transducer uses a proprietary vibrating ossicular prosthesis (VORP), which typically is attached to the long process of the incus. The VORP contains a rare earth magnet surrounded by a receiving coil, a demodulator package, a conductor link, and a proprietary floating mass transducer (FMT).^22,23 The transducer was developed by Ball, and incorporated prior conceptual work performed by Goode and colleagues.^11,22,24 The VORP is electrically connected to the internal receiver.²⁵ These internal components are coupled to the audio processor via a telemetry link, similar to a cochlear implant. The processor, which is worn behind the ear, contains the microphone, audio processing electronics, magnet, telemetry transmission equipment, and room for a standard 675 zinc battery to power the implant.

Figure 157-3. A-H, Med-El Vibrant Soundbridge semi-implantable hearing aid. A and B, External microphone couples via an inductive transcutaneous link to the actuator, a vibrating ossicular prosthesis (VORP) that couples to the incus long process (B). C, Implanted component of device. D, VORP actuator. E, Programming unit and external circuitry. F-G, Inductive link is implanted against retromastoid cortical bone, similar to a cochlear implant (F and H). VORP is clipped to the incus via a facial recess approach (G).

The function of the externally worn processor is similar to a conventional hearing aid. Sound is picked up by the microphone and processed by the external receiver by the electronics. The signal is sent via telemetry to the internal receiver. From here it is sent to the demodulator, which provides a safety feature to limit maximum power output and prevent overstimulation. The demodulator generates a current encoding the sound, and this current travels to the FMT, causing the FMT to vibrate through inductive interaction with the magnet. Because the FMT is firmly attached to the incus using a titanium clip, amplified sound is transmitted directly to the ossicular chain. Because the FMT includes a magnetic component, implantation of this device prevents a patient from undergoing magnetic resonance imaging (MRI) without device removal.

The Vibrant Soundbridge is indicated in patients with moderate to severe SNHL and desire for an alternative to an acoustic hearing aid (Fig. 157-4).²³ Studies have indicated that it may be suitable for patients with hearing loss of 70 dB pure-tone average (PTA) hearing thresholds.²⁶ When first introduced, the externally worn audio processor was an analog device that incorporated a wide dynamic range compression. Because of inadequate gain of the device, particularly at lower frequencies and with low sound levels,²⁷ a newer, more powerful digital audio processor was introduced in 1999. Studies done shortly after its introduction failed to show superior hearing results compared with conventional aids, although patients preferred the device, in part because of ability to wear it for a much longer time without discomfort.²⁶

Figure 157-4. Audiologic selection criteria for the Soundbridge. The Soundbridge is indicated in patients with moderate to severe sensorineural hearing loss of 70 dB pure-tone average hearing thresholds. The shaded area of the figure corresponds to the pure-tone audiometric implant criteria.

The device typically is implanted using a standard transmastoid, facial recess approach to the middle ear. The VORP is crimped into the long process of the incus. The internal receiver is placed in a bony trough in the retrosigmoid bone several centimeters behind the ear, in a location similar to the receiver placed during cochlear implantation. One potential drawback of this design concerns the size of the VORP. Because of the confined space between the eardrum and the incus, the dimensions and mass of the FMT are limited, restricting its output.²⁸ A second potential drawback is erosion of the incus long process, owing to ischemia at the clip attachment site. Modifications of the typical surgical approach allow treatment of mixed hearing losses resulting from otosclerosis or ossicular erosion or agenesis through direct placement of the VORP on the stapes superstructure, round window, or oval window.^29–32

Owing to its primacy in regulatory approval processes and patient availability, to date more clinical data are available for the Vibrant Soundbridge than for all other middle ear implants. Results from the phase III clinical trial of 53 patients were sufficiently favorable to merit FDA approval of the device (phase I trials typically test feasibility and safety in a small group of subjects, phase II trials typically test effectiveness in a slightly larger group, and phase III trials assess safety and effectiveness in a large group).²³ Patients in the study were analyzed with conventional analog hearing aids, digital hearing aids, and the Vibrant Soundbridge. Standard audiologic tests included pure-tone air-conduction and bone-conduction thresholds (250 to 8000 Hz), functional gain, speech recognition tests, the Profile of Hearing Aid Performance, and the Hearing Device Satisfaction Scale.

Although one patient had a shift of 18 dB after device placement, the effect of implantation on residual, native hearing was believed not to be significant (<10 dB PTA at 500 to 2000 Hz). This finding corroborated earlier studies that showed little to no effect of the presence of the FMT on hearing sensitivity.³³

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