JPH0367694B2 - - Google Patents

Description

Claim 1: A transmitting device operable to transmit a plurality of carrier signals, each along a separate transmission channel.
and a modulation device operable to modulate at least one of said carrier signals with a signal representing respective information content of audio frequencies; and a plurality of receiving channels each responsive to one of said carrier signals. a multi-channel receiving device for receiving each transmitted signal having a multi-electrode hearing aid; and a multi-electrode hearing aid for electrically stimulating the auditory nerve by said multi-electrode hearing aid when the device is in use.
Each receiving channel of the multi-channel receiving device is connected to at least one of the multi-electrode hearing aid devices.
and a means for connecting to one electrode, and the transmitting device has a plurality of channels each corresponding to a plurality of frequency bands within the audible frequency band, and each of the plurality of channels has one electrode.
a band waver for selecting one frequency band;
A signal processing circuit that is connected to the bandpass waveform generator and generates an output signal according to the frequency and amplitude of the signal from the bandpass waveform generator, and a transmission for modulating a carrier signal with the output signal of the signal processing circuit. A multifrequency device for electrical stimulation of the auditory nerve, characterized in that it includes a device for electrically stimulating the auditory nerve.

2. The multi-frequency device of claim 1, wherein each receive channel of the multi-channel receiver device has a respective tuned receive coil and a signal demodulator.

3. The multi-electrode hearing aid device comprises an elongated tissue-compatible molded body and a plurality of conductive wires disposed inside the molded body, each of the plurality of conductive wires terminating in one electrode contact, 3. The multifrequency device according to claim 1, wherein each electrode contact is attached to the surface of the molded body.

4. The method according to claim 3, wherein each of the plurality of conductive wires is made into a corrugated shape before the molding process of the molded body, in order to improve the flexibility of the molded body when introduced into the cochlea. Multi-frequency equipment.

5. A multi-frequency device according to claim 3 or claim 4, wherein each of said electrode contacts is positioned such that said electrode contacts can stimulate the cochlea according to its frequency response.

6 The signal processing circuit includes a pulse generator, a device connected to the bandpass generator for controlling the frequency of the pulse generator in response to the frequency of a signal from the bandpass generator, and the bandpass generator. and a device for controlling the pulses of said pulse generator in response to the amplitude of a signal passed through said transmitter, wherein said transmitter modulates a carrier signal by an output signal of said pulse generator. A multi-frequency device according to range 1.

7. The multifrequency device of claim 6, wherein said pulse generator is a monostable multivibrator, and wherein said device for controlling pulses in response to said signal amplitude controls pulse width.

8. The multi-frequency device according to claim 6, wherein the plurality of channels of the transmitting device correspond to bands of 0.25 to 0.5 KHz, 0.5 to 1.0 KHz, 1.0 to 2.0 KHz, and 2.0 to 4.0 KHz, respectively.

9 The multi-channel receiving device has four channels, each of which includes one coil, and the transmitting device has two coils, each of which has two coils.
two modulated carrier signals, two coils of the receiving device are inductively coupled to one coil of the transmitting device, and two other coils of the receiving device are inductively coupled to one coil of the transmitting device. 7. A multi-frequency device as claimed in claim 6, wherein the multi-frequency device is inductively coupled with.

10. The two coils of the receiving device are arranged in a superposition, and the other two coils of the receiving device are also arranged in a superposition, so that the mutual inductance of these coils is minimized. A multi-frequency device according to scope 9.

11 The transmitting device has at least one coil for transmitting the plurality of modulated signals, the receiving device has a plurality of coils corresponding to the number of channels, and the coil of the transmitting device 2. The multi-frequency device of claim 1, wherein: is inductively coupled to said plurality of coils of a receiving device.

12. The signal processing circuit includes a dynamic range compression device and a frequency treble dependence compensation device, and the oscillator modulates a carrier signal by the processed analog acoustic signal. 2. The multi-frequency device according to claim 1.

13 The multi-channel receiving device has four channels, each channel including one coil, and the transmitting device has two coils, each coil having two modulated channels. receiving a carrier signal, two coils of the receiving device being inductively coupled to one coil of the transmitting device, and two other coils of the receiving device being inductively coupled to another coil of the transmitting device; A multi-frequency device according to claim 12.

14. The two coils of the receiving device are arranged in a superposition, and the other two coils of the receiving device are also arranged in a superposition, so that the mutual inductance of these coils is minimized. A multi-frequency device according to range 13.

15. A multifrequency device according to claim 13, wherein said dynamic compression is performed by suitable non-linear electronic elements.

16. The multifrequency device of claim 15, wherein the nonlinear electronic element is a logarithmic amplifier.

17. The multifrequency device according to claim 13, wherein the dynamic range is compressed by driving a logarithmic amplifier with a signal whose frequency is shifted upward, and then shifting the frequency of the signal downward.

14. The multi-frequency device of claim 13, wherein the dynamic range is compressed by using non-linear elements arranged in a plurality of 18 octave wide bands.

19. The multifrequency device of claim 13, wherein the dynamic range is compressed by using a gain-controlled amplifier.

20. The multi-frequency device of claim 12, wherein the oscillator includes a base modulated, non-saturated output amplifier.

21. The multifrequency device of claim 12, wherein the oscillator includes an emitter modulated non-saturated output amplifier.

22. The multi-frequency device of claim 12, wherein the transmission device includes a band waver, the band waver comprising a transmitter tank and a tuned receiver circuit.

23. The multi-frequency device according to claim 22, wherein the bandpass filter is critically coupled.

24. The multi-frequency device of claim 13, wherein said transmitter is attached to an ear hook containing a transmitter coil.

Description The present invention relates to a method, a multichannel electrode, a multichannel receiving device, and a multifrequency device for electrically stimulating nerves or muscles so that a deaf or hard-of-hearing person can obtain hearing sensations, and in particular to Concerning giving more stimulation.

An improvement to a conventional prosthetic device (hearing aid) for implantation in the mastoid process that helps amplify sound waves is disclosed in U.S. Patent No.
Known from No. 3209081. In that case, the implanted receiver is in direct contact with the bone, from where the sound waves are guided to the inner ear by bone conduction.

Recently, devices that do not simply amplify sound waves but also convert sound waves into electrical pulses have become known. The pulses are used to electrically stimulate the auditory nerve and cause the sensation of hearing. U.S. Patent No. 3,449,768 includes
The use of encoded pulse trains to generate electric fields with desired gradients useful for stimulating the visual or auditory system has been described. Further, US Pat. No. 3,752,939 describes the use of an electrode having two conductive wires on its surface over the entire length of the electrode.

Journal of the Otolaryngology Society, Vol. 103, published December 1977, paper by Schindler et al. "Intracochlear implantation of multiple electrodes"
describes spatially localized stimulation of the auditory nerve in cats. Otolaryngology Annual Report 90/7, 1976, Clark and Hallworth's article ``Multiple electrode arrays for cochlear implantation'' describes a cochlear electrode with a thin film structure mounted on a flat flexible carrier material. . Other devices are known, such as thin wire bundles, which are used to provide electrical stimulation directly to the auditory nerve after being extracted from the cochlea.

European Patent Application No. 783005671 and German Publication No. 783005671
No. 2823798 describes an implantable multichannel hearing aid based on electrical stimulation of the auditory nerve, for which a circuit with active elements is used.

The invention comprises a device for transmitting a plurality of carrier signals, each modulated by a signal representing one frequency band, and a multi-channel receiver for receiving the transmitter signal, the device comprising one frequency of the acoustic signal. Each receiving channel is assigned to one signal representing a band, and the multichannel electrode further includes a signal application device for applying a signal from each of the receiving channels to at least one electrode contact of the multichannel electrode. The electrodes are used to provide electrical stimulation and are intended for multi-frequency devices for providing electrical stimulation via multi-channel electrodes.

The multi-frequency device for improving electrical stimulation of the auditory nerve according to the invention can have a plurality of transmission channels through the skin, and a signal representing each one frequency band of the audio waveband can be transmitted through those transmission channels. is superimposed on the carrier frequency of

A multi-channel receiver according to the invention comprises one channel receiver tuned to receive a certain predetermined oscillator signal.
a plurality of mutually independent channels each including a coil, a demodulator coupled to the coil for signal demodulation, and a coupling device for coupling the demodulated signal to the electrode contacts. Each of these receiving channels receives a respective signal transmitted wirelessly through the skin.

The present invention also includes an elongated tissue-compatible compact having a plurality of conductive wires disposed within the compact, each conductor terminating in an electrode contact, and a plurality of electrode contacts disposed within the compact. Mounted on the surface, the leads reside in a multi-channel electrode device for implantation in the cochlea, which is corrugated to make the electrode flexible and free from tensile loads.

The present invention also provides the following features: (a) manufacturing a plurality of conductive wires having one contact at the tip; (b) corrugating the conductive wire to increase the flexibility of the electrode body; and (c) making the contact point (d) placing a tissue-compatible material into the mold so as to keep it in place; Also covered is a method of manufacturing a multichannel electrode device for implantation in the cochlea.

A method of improving auditory electrical stimulation according to the present invention comprises: (a) transmitting a plurality of modulated signals, each corresponding to one frequency band within the audio frequency band; and (b) transmitting a plurality of these modulated signals. The method comprises: receiving a signal; (c) demodulating the modulated signal; and (d) applying the demodulated signal to a multichannel electrode implanted in the cochlea.

Next, further explanation will be given with reference to the drawings.

FIG. 1 is a cross-sectional view of a human ear, also showing the arrangement of the stimulation device according to the invention; FIG. FIG. 3 is an electrical block diagram illustrating a preferred embodiment; FIG.
5 is a perspective view of a multi-channel cochlear electrode as part of an auditory electrical stimulation device; FIG. The figure is a schematic diagram of the human ear to show pitch specificity, and shows the frequency (Hz) that causes maximum activity in the auditory nerve at a predetermined location along the cochlea. FIG. 8 is a perspective view showing a mold for manufacturing the cochlear electrode according to FIG. 7, FIG. 9a is an explanatory diagram showing a modification of the multi-channel cochlear electrode according to the present invention, Figure 9b is the 9th
Explanatory diagram showing the inserted state of the cochlear electrode in Figure a, No. 10
The figure is an electrical block diagram showing an external acoustic processing transmitter, FIG. 11 is an electrical wiring diagram of the dynamic range compression circuit shown in FIG. 10, and FIG. 12 is a preferred example of the transmitter shown in FIG. It is an electrical wiring diagram showing an example.

Sound waves are normally directed by the outer ear 10 to the tympanic membrane 12, which is coupled to and moves the auditory ossicles 14 of the middle ear, thereby energizing the cochlea 16. The cochlea 16 is a spiral formation with 2.5 turns. The cochlea 16 has an upper canal or scala vestibule 18 and an inferior canal or scala tympani 20. Between the scala vestibuli 18 and the scala tympani 20 is the cochlear canal 22 . In both floors 18 and 20 filled with liquid, the arriving sound waves generate liquid waves, and the waves generate electrical pulses by the transduction function of the inner ear, and these pulses are guided to the brain by the auditory nerve 24. , interpreted as the sense of hearing.

The inner ear of a completely deaf person cannot convert incoming sound waves into electrical signals that can be transmitted to the brain. Therefore, the multi-frequency stimulation device according to the present invention is adapted to electrically stimulate the cochlea directly.

The stimulator includes a body-worn, multi-channel acoustic processing and transmitting device 30. Acoustic processing transmitter 30 is coupled to an implanted receiver. The coupling is preferably effected by inductive action via the coils 36, 38;
is a coil 3 forming part of an implanted receiver.
2, 34 and the acoustic processing-transmission device 30.

Acoustic treatment transmitter 30, as detailed below.
generates a plurality of carrier signals onto which a series of signals in the audio frequency band are superimposed. The transmitted signal is received and demodulated by an implanted receiver. The demodulated signal is provided to multichannel electrode 46 via conductors 42 and 44. Electrode 46 is implanted in the cochlea. The electrode 46 has a number of electrode contacts on its surface, which electrode contacts are used to selectively stimulate the cochlea locally, corresponding to its frequency assignment.

According to a preferred embodiment, the multi-frequency device has four channels corresponding to four frequency bands, into which the listening area is divided.

FIG. 2 is an electrical block diagram of the acoustic processing transmitter 30.
It has four channels corresponding to 0.5-1.0KHz, 1.0-2.0HHz and 2.0-4.0KHz. The configuration of each channel is shown in the form of a block diagram for channel 1. Each channel has a desired frequency band (for example, 0.25-0.5KHz for channel 1)
It is equipped with a band wave generator 50 tuned to . An audible signal is supplied to the input of the transducer 50 and is received by a microphone 52 and amplified by a controlled amplifier 54 . After the amplifier 54 the audible signal still contains another frequency band, for example the frequency band 55, but after passing through the waver 50 it is limited as shown by the frequency band 57.

A signal delay circuit may be placed in the low frequency processing channel to simulate the time that a traveling wave across the length of the cochlea would exist in a healthy ear.

Signal 57 is then fed to comparator 58, which produces a limited signal 59. This signal 59
The zero crossings of signal 57 are matched with the zero crossings of signal 57.

The limited signal 59 is transferred to a frequency to voltage converter 60
guided by. Converter 60 generates a variable DC voltage proportional to the frequency of signal 59. converter 60
has a suitable structural element, for example a monostable multivibrator, which receives the signal 5
9 and generates a pulse of the same pulse width with the same repetition frequency as the pulse from signal 59. The output of the multivibrator is connected to a low frequency generator that generates a variable DC voltage proportional to the pulse repetition frequency.

The voltage signal at the output of converter 60 is fed to a voltage-frequency converter 62, which is a voltage-controlled oscillator. The output signal 63 of the converter 62 consists of a series of pulses having a constant pulse width and whose pulse train frequency corresponds to the voltage controlling the oscillator.

The pulse train frequency of the signal 63 may be within a certain band, e.g.
0 may leave out larger or smaller frequency bands.

As explained further below, the auditory nerve is best able to discriminate repetition frequencies in the stimulation signal below about 400 Hz. This is why the frequency band waved by bandpass generator 50 is converted as described above into a lower frequency band that is best suited to the auditory nerve. This band is often 400-400Hz, but may be larger.

The signal at the output of the bandpass filter 50 is passed through the rectifier 66
and a logarithmic amplifier 68, which produces a DC voltage that depends on the logarithm of the amplitude of the signal rectified by the rectifier 66.

All channels have identical circuit elements behind different bandpass generators. The monostable multivibrator in each channel corresponds to a frequency range of 400-400Hz, which is particularly suitable for cochlear stimulation.
generates an output pulse train consisting of pulses of variable pulse width whose repetition frequency varies within a range of . These pulse trains are superimposed on the carrier signal at the transmitter 74.

In this four-channel embodiment, the multi-channel oscillator 74 uses four carrier signals, two at 12 MHz and the other two at 31 MHz.
The pulse trains of channels 1 and 3 are used to modulate a 12 MHz carrier signal, and the pulse trains of channels 2 and 4 modulate a 31 MHz carrier signal. Channel 1,
2 signals are superimposed on it. The carrier signal is applied to one transmitter coil and the carrier wave, on which the signals of channels 3 and 4 are superimposed, is applied to a second transmitter coil. Both carrier frequencies applied to the oscillator coil are therefore unequal and interaction between channels is avoided.

However, by using this transmitter coil for two channels, the structure is simplified. However, using a dedicated transmitter coil for each channel is also advantageous for other reasons.

FIG. 3 shows a preferred configuration example of a multichannel receiving apparatus for four channels.

Each channel has coils 81-84, with coils 81 and 82 inductively coupled to transmitter coil 76 and coils 83 and 84 inductively coupled to transmitter coil 78. A capacitor 85 is connected in parallel to each of the coils 81 to 84, so that 12M
An oscillation circuit with a resonant frequency of Hz or 31MHz is formed.

The signal received by coil 81 is fed to a demodulator consisting of diode 86, capacitor 87 and resistor 88. In the case of pulse width modulation and demodulation, the voltage at the detector output is limited by connecting a zener diode in parallel with resistor 88. Voltage fluctuations due to changes in the coupling between the transmitter and receiver coils can thereby be minimized.

The voltage of the detector output section 90 is preferably 0 to 3V.
40 in response to the modulating signal.
It has a frequency of ~400Hz.

For devices for tissue stimulation with a small number of mutually independent channels activated by multiple signals at the same time, especially devices with 2 to 9 channels, the following method may be advantageous. It is a method of arranging receiver coils in stacked groups to minimize the space demands created by the use of multiple receiver coils.

Even if each receive signal coil in a group is tuned to one particular frequency, simply arranging the coils in a superposed manner will not allow the signal to be heard due to the mutual inductance of the two or three receiver coils. Stronger crosstalk may occur. If the coils 81 and 82 are arranged as shown in FIG. 4 so that the opposing magnetic fluxes are compensated, mutual inductance disappears. As a result, two or three mutually independent channels can be obtained, which occupy less space than a single channel and have negligible crosstalk. Corresponding to this concept, coils 81, 82 and coils 83, 84 are arranged in two mutually separate groups in FIG. These two groups represent different possibilities for arranging two coils in a group to compensate for mutual inductance. 85, 86, and 87 are three coils grouped together to compensate for mutual inductance. The diameter of the coils 85 to 87 is approximately 1.5 to 2 cm, and the interval between the two coil groups is as follows:
To prevent crosstalk between them, the width should be approximately 3 cm.

The coils 81 and 82 have different frequencies (e.g.
Since the channels are tuned to 12 and 31 MHz, each receive channel receives and demodulates only the signal of the transmit channel to which it is assigned. The demodulated signals from each receive channel are directed to multichannel electrodes as shown in FIG. One or more electrode contacts 92 can be coupled to each channel, and these electrode contacts can be used to stimulate multiple locations along the cochlea for the desired pitch perception. It is arranged along the electrode body.

When using various connection connections, bipolar stimulation, unipolar stimulation to remote ground electrodes or stimulation to a common distributed ground can be used.

The multi-electrode is made of a molded body made of silicone elastomer, such as "Silastick", and a large number of conductive wires 91 are embedded inside this molded body. Each lead 91 terminates in a spherical electrode contact 92 on the surface of electrode body 90 . The arrangement of contact bodies on the electrode surface allows for selective stimulation of each cochlear region following implantation of the electrode in the cochlea. As can be seen from FIG. 6, which is a schematic diagram of the human cochlea, the high frequency response area is in the root area and the low frequency response area is in the apical area. Thereby the electrode contacts 92 in the cochlea
A desired pitch perception can be obtained in accordance with the arrangement of . Pitch continuity is achieved by additional stimulation frequency changes.

FIG. 7, which is a schematic cross-sectional view of the multi-channel electrode, shows the arrangement of conductive wires inside the electrode.
For simplicity of illustration, only two conductors 93,9 are shown.
Only 4 are shown. The conductors 93, 94 are corrugated to remove tensile loads, but this facilitates bending of the electrodes when introduced into the cochlea.

According to a preferred embodiment, the conductors 93, 94 are
Teflon-insulated platinum (90%) with a diameter of 26 μm
It is an iridium (10%) wire. The spheres at the tips of the conductors 93, 94 have a diameter of 0.3 mm.
It is formed by flame melting.

The electrode contacts 92 are arranged in pairs along the electrode body 90 in two rows facing each other.

According to a preferred embodiment, the diameter of the electrode body 90 is 0.9 mm, decreasing to 0.5 mm towards the tip. The total length of the electrode body 90 is 20 to 25 mm into the cochlea.
It must be compatible with the introduction of mm electrodes.

FIG. 8 shows one mold half 96 of two identical mold halves for controlling multi-channel electrodes. The semi-molding mold 96 has a groove 97 that gradually becomes narrower in a desired electrode shape, and the groove 97
A large number of through holes 98 are formed therein, all communicating with one vacuum conduit 99. When manufacturing multi-channel electrodes, the conductive wire 9 is inserted into the groove 97 of the semi-molded mold.
3,94 will be installed. At this time, the spherical tips of the conductive wires 93 and 94 are positioned in the through hole 98 of the half mold 96 by vacuum action. Thereafter, the mold 96 is assembled and the shirastick material is pressed into the groove 97.
Vacuum suction through the through holes 98 precisely positions the electrode contacts 92 on the surface of the electrode body 90 .

Electrode body 100 according to the modification shown in FIG. 9a
has a curvature corresponding to that of a spiral.

Figure 9b shows the method of introducing the electrode body 100 into the cochlea. A straight rod, for example a steel wire, is inside the electrode body 100 and
0 is gradually withdrawn while being inserted into the cochlea. Thereby, the electrode body 100 returns to its original shape.

Structural forms of externally provided or internally implanted subsystems other than those described above are as follows.

(b) Four or more channels of electrodes introduced into the cochlea, combined with four channels of internal equipment and one channel of external sound processing and transmitting equipment.

In this case, one electrode channel for stimulation is selected or the desired number of electrode contacts are coupled together. In the latter case, different threshold values for the separate electrode contacts can be compensated for by corresponding adaptations.

(b) A single channel electrode at or near a round window, fixed in combination with a large earth electrode.
A ground electrode can also be attached to the outside of the cochlea. Single channel electrodes can be used with single channel internal devices and single channel external acoustic processing and transmitting devices.

The device in (b) has already proven its worth in several deaf volunteers, who can use electrical stimulation alone to learn 60 to 60 minutes of normal conversation about unknown words or sentences, without the need for supplementary lip reading. 70% understanding was obtained. This makes it clear that this prosthesis can be seen as a useful hearing aid for all deaf people.

The device (b) is mainly used for people who are not completely deaf, that is, people with hearing loss and children with hearing problems.

Whether the external acoustic processing transmitter is a single channel system or a multi-channel system,
Instead of converting sound waves into pulse trains, analog signals can be used. In this case, suitable dynamic range compression and compensation of the frequency dependence of the sound intensity become very important characteristics of the acoustic processing transmitter. These characteristics are important for the following reasons. In other words, the dynamic range between the stimulus values required for marginal hearing and excessive hearing is much narrower than when a person with normal hearing hears, and, to make matters worse, in many cases , the perception of the threshold sound and the sound intensity that exceeds the threshold is highly dependent on the stimulus frequency.

Non-linearity is particularly utilized for dynamic range compression. This non-linearity may be a logarithmic or power function, it may consist of an intermittent linear function, or it may take any other suitable form.

Nonlinearity is obtained by using suitable differential or operational amplifiers coupled to diode-wired transistors or diode networks.

Nonlinearity can be controlled by a frequency-shifted signal so that dynamic range compression does not produce undesired frequencies. In this case, the even-numbered distorted wave components can be removed by a narrow band transducer in the frequency-shifted signal, and then the signal can be mixed down-shifted back into the original audio frequency range.

Another possibility to reduce the distorted wave content is to use multiple nonlinearities within an octave wide band.

Controlled amplifiers with sufficiently low response or fall time constants of 2-10 msec to 100-300 msec may be used. Appropriate nonlinearities can be added to the signal path to give the desired shape to the amplitude characteristics of the controlled amplifier.

The dynamic range compression circuit may precede the frequency response compensation circuit. Although the amount of frequency response compensation required in this case is relatively small,
This compensation must be done very precisely.

FIG. 10 shows a block diagram of a single channel or multichannel acoustic processing and transmitting device.
A multichannel stimulator consists of a plurality of substantially identical channels, each with a specific radio frequency transmission section. Each channel shares a predetermined audio frequency band that is waved by a respective bandpass waver. For single channel stimulators,
Since only one channel is used, the bandpass filter 103 is not required. The signal received by the electrate microphone 104 is 80 dB
It has the above dynamic range. Converting this large dynamic range to a range of permissible stimulus values of about 10 dB is accomplished by either a dynamic compression circuit 107, described in detail later, or an input-controlled inverse control acting on the amplifier 105. It is done by both. Compared to dynamic compression using nonlinear elements, this control has the advantage of less time-induced nonlinear distortion, but due to the final response time, the sudden occurrence of a high-frequency signal can cause Since the disturbance peaks are also passed through, control as well as dynamic compression is generally used.

The adaptation circuit 106, which adapts the frequency response to the pre-measured frequency dependence of the user's iso-pitch curve, comprises a frequency-dependent element, for example an RC element or an LC element, connected in the usual manner. Ru.
The adaptation circuit 106 has a structure that in principle follows the dynamic compression circuit 107, but in this case very small but very accurate frequency effects are required. The signal to be transmitted is output circuit 1
09, a tuned implant receiving circuit 110 and a demodulator 111.
and from there it is further transmitted to the electrode 112.

The circuit used for dynamic compression, shown in FIG. The output parts of these differential amplifiers are connected in parallel with each other. The output voltage obtained as the differential voltage between points 122, 123 (y,) depends logarithmically on the input voltage at point 121. This input voltage is led via a capacitor 113 to a point 124 at the input of the integrated circuit 112 to suppress any DC voltage that may be present. Resistor 120 is circuit 1
It is used to fix the DC voltage level at 12 points 124. The differential voltage between points 122 and 123 is, as is conventionally used, by an operational amplifier 114 forming a subtraction circuit with resistors 115-118.
Can be changed to unbalanced voltage. Trimmer potentiometer 119 is used to balance the offset voltages.

FIG. 12 shows a circuit diagram of an amplitude modulation oscillator. The oscillator is an oscillator 120 producing a carrier frequency of 12 MHz.
and an output stage. The output circuit 126 is
Implanted receiver circuit 12 tuned to the same frequency
Together with 7, it forms a bandpass filter. Current-controlled (and therefore high input ohmic resistance) bandpass generators, at a given coupling, the so-called critical coupling,
Indicates the maximum value of secondary induced voltage. Therefore, around this fixed point, the secondary induced voltage hardly depends on the coupling. With the inductive coupling used in this case, therefore, the tolerance characteristics regarding the positional movement of the transmitting coil are very loose at the distance between the transmitting coil and the receiving coil that causes critical coupling. The critical distance can be set between 10 and 10 by choosing the quality of the circuit appropriately.
It is kept at 12m. Transmitter coil diameter 23mm, movement ±
In the case of 10 mm, the amount of variation in the secondary voltage is only -5%.

Due to the high circuit quality required for the oscillator circuit and the need to avoid saturation of the output transistor 128, collector modulation cannot be used for amplitude modulation, and emitter current modulation or (as in the illustrated example) ) must be based on base modulation. The modulating voltage is led to the base via resistor 133, coupling coil 134 and resistor 130.
The resistor 130 is selected to correspond to the required output magnitude, allowing accurate values of the high frequency control output and eliminating the need for cumbersome changes in the number of turns of the coupling coil 134. Capacitors 131 and 132 are used for coupling the modulation input section and operating voltage supply section to the ground according to high frequencies. The Schottky diode 127, upon unexpected saturation of the output transistor,
This prevents the occurrence of instabilities that cause disruptive oscillations.

The transmitter output circuit 126 is located on the Plexiglas earpiece so that it can be precisely positioned via the implanted receiver coil.

The entire transmitter is small and can also be mounted on an ear piece. In this case, the total length of the high-frequency circuit is 2 cm or less, and it hardly acts as an antenna, so it has the advantage of less radiation of electromagnetic waves.

The above-described method of electrically stimulating hearing using the multi-frequency device of the invention improves the hearing of deaf and hard of hearing individuals.

By dividing the audible signal into bands and selectively stimulating various locations within the cochlea, the quality and intelligibility of the auditory sensation is improved. Since the implanted receiver circuit consists only of passive electronic elements, no power supply is required and only the stimulation signal needs to be transmitted through the skin and into the body.

A hearing prosthesis containing multi-channel cochlear electrodes is easily manufactured, and the manufacturing method described above provides precise positioning of the electrode contacts to obtain the desired pitch sensation during stimulation of the cochlea.
Although pulse width modulation is used in the example described above, other modulations may be used, such as amplitude modulation or frequency modulation.

The analog stimulation signal may be used like a pulsed signal or a digital signal after appropriate electronic processing. In the above example using a digital signal, the audible frequency band is converted to a corresponding signal band of 40 to 400 Hz, but the corresponding signal band is the same frequency range as the audible frequency not only in digital circuits but also in analog circuits. It may have.