Classifications

• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
  • A61N1/00—Electrotherapy; Circuits therefor
  • A61N1/18—Applying electric currents by contact electrodes
  • A61N1/32—Applying electric currents by contact electrodes alternating or intermittent currents
  • A61N1/36—Applying electric currents by contact electrodes alternating or intermittent currents for stimulation
  • A61N1/372—Arrangements in connection with the implantation of stimulators
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F11/00—Methods or devices for treatment of the ears or hearing sense; Non-electric hearing aids; Methods or devices for enabling ear patients to achieve auditory perception through physiological senses other than hearing sense; Protective devices for the ears, carried on the body or in the hand
  • A61F11/04—Methods or devices for enabling ear patients to achieve auditory perception through physiological senses other than hearing sense, e.g. through the touch sense
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
  • A61N1/00—Electrotherapy; Circuits therefor
  • A61N1/02—Details
  • A61N1/04—Electrodes
  • A61N1/05—Electrodes for implantation or insertion into the body, e.g. heart electrode
  • A61N1/0526—Head electrodes
  • A61N1/0541—Cochlear electrodes

Definitions

• the invention relates to methods, multi-channel electrode, multi-channel receiving device and multi-frequency system for electrical stimulation, in particular for the stimulation of nerves or muscles, such as to achieve hearing impressions in deaf or hearing impaired people; in particular, the invention relates to stimulation with electrical impulses.
• the implanted receiver is in direct contact with the bone, from which sound waves are transmitted to the inner ear using the bone conduction.
• the invention consists in that the system comprises: a transmission device for transmitting a plurality of carrier signals, each of which is modulated by a signal representing a frequency band, a multi-channel reception device for reception the transmission signals, each receiving channel being assigned to a signal representing a frequency band of the sound signal, and a device for applying the signal from each of the channels mentioned to at least one electrode contact of the multi-channel electrode, the multi-channel electrode being used for electrical stimulation.
• the present invention is preferably a multi-frequency system for improved electrical stimulation of the auditory nerve, which can have a plurality of transcutaneous transmission channels, the carrier frequencies of which are modulated onto signals which each represent a frequency band of the hearing range.
• a multi-channel receiving device consists in that it contains: a plurality of independent channels, each channel of which contains a coil which is tuned to receive a specific transmission signal, a demodulator connected to the coil for the purpose of demodulating the signal and a device for connection of the demodulated signal with the electrode contacts. These reception channels each receive one of the signals transmitted wirelessly through the skin.
• the invention further relates to a multichannel electrode for implantation in the cochlea, which according to the invention consists in the fact that it has an elongated, biocompatible casting body, inside which there are a plurality of wires, each of which ends in an electrode contact, the electrode contacts are attached to the surface of the body and the wires are crimped to achieve strain relief and flexibility of the electrode.
• the multi-channel electrode consists of the following process steps: a) production of a plurality of wires which have a contact at their end; b) waves of said wires, whereby the flexibility of the electrode body is increased; c) placing said wires in a mold, said contacts being held in position by vacuum suction; d) filling said mold with biocompatible material.
• a method for improving the electrical stimulation of the hearing consists of the following method steps: a) a plurality of modulated signals are emitted, each signal corresponding to a frequency band within the hearing range; b) said plurality of modulated signals are received; c) the modulated signals are demodulated; d) and the demodulated signals are applied to the multi-channel electrode which is implanted in the cochlea.
• FIG. 1 shows a schematic sectional view through a human auditory organ and the placement of the present invention
• Fig. 2 is an electrical block diagram of a preferred embodiment of the external sound processor transmitter as part of the multi-frequency system to improve the electrical stimu lation of hearing
• 3 shows an electrical circuit diagram of a preferred embodiment of a multi-channel receiving device as part of the multi-frequency system for improving the electrical stimulation of the hearing
• 4 shows a schematic illustration to illustrate two possible geometrical arrangements of receiving coils in accordance with an embodiment of the invention
• 5 shows an illustration of a multi-channel cochlear electrode as part of the multi-frequency system for improving the electrical stimulation of the hearing
• Outer ear 10 is guided to the eardrum 12, which is connected to the ossicles 14 of the middle ear and sets this in motion, thereby stimulating the worm 16.
• the cochlea (snail) is a snail-shaped structure with about 2 1/2 turns. It contains an upper channel 18, the so-called Scala vestibuli, and a lower channel 20, the Scala tympani. Between the two scales is the cochlear duct 22. In the two liquid-filled
• Scales arise, caused by incident sound waves, Fluid waves, which trigger electrical impulses through the transducer function of the inner ear, which are passed on from nerve acusticus 24 to the brain and interpreted as hearing impressions.
• Fluid waves which trigger electrical impulses through the transducer function of the inner ear, which are passed on from nerve acusticus 24 to the brain and interpreted as hearing impressions.
• the stimulation system on which the present invention is based is used for direct electrical stimulation of the cochlea.
• the system includes a multi-channel sound processor / transmitter 30 that can be worn on the body.
• the transmitter is coupled to the implanted receiver.
• the coupling is preferably done inductively with the aid of coils 36 and 38, which are connected to sound processor / transmitter 30, and coils 32 and 34, which are part of the implanted receiving device.
• the sound processor / transmitter 30 generates a plurality of carrier signals, which are a series of signals from the audible
• Frequency range are modulated.
• the transmitted signals are picked up and demodulated by the implanted receiving device.
• the demodulated signal is connected through wires 42 and 44 to multichannel electrode 46, which is implanted in the cochlea.
• the surface of the multi-channel electrode contains a plurality of electrode contacts which serve for locally selective stimulation of the cochlea in accordance with the frequency arrangement thereof.
• the multi-frequency system contains four channels corresponding to four frequency bands into which the listening area is divided.
• each channel contains a bandpass filter 50, tuned to the desired frequency band (eg 0.25-0.5 KHz for channel 1).
• the filter 50 At the input of the filter 50 there is an audio signal which was picked up by the microphone 52 and amplified by a regulated amplifier 54. After the amplifier 54, the signal still encompasses a wide frequency range, as illustrated in FIG. 55. After passing through the bandpass filter 50, the frequency range is restricted, as indicated in FIG. 57. Circuits for signal delay can be installed in the channels for processing the lower frequencies, which serve to imitate the time available in the healthy ear, which the traveling wave requires over the length of the cochlea. The signal 57 is then through a comparator
• the limited signal 59 is passed to a frequency-voltage converter 60, which generates a variable DC voltage that is proportional to the frequency of the signal 59.
• the frequency voltage converter contains suitable modules, such as a monostable multivibrator, which is triggered by signal 59 in order to generate pulses of the same pulse width, which have the same repetition rate as the pulses from signal 59.
• the output of the monostable multivibrator is connected to a low-pass filter, which emits a variable DC voltage that is proportional to the pulse rate.
• the voltage signal at the output of converter 60 is then applied to a voltage-frequency converter 62, for example a voltage-controlled oscillator, the output signal 63 of which consists of a sequence of pulses which have a constant pulse width and whose pulse repetition frequency corresponds to the voltage controlling the oscillator.
• the pulse repetition frequency of the signal 63 can be within a certain range, for example between 40 and 400 Hz, while the bandpass filter can filter out a larger or smaller frequency range.
• the auditory nerve can best recognize repetitive frequencies in the stimulating signal below approximately 400 Hz. For this reason, the above-described transformation of the frequency band filtered out by the bandpass filter into the region of low frequencies best suited for the auditory nerve is carried out. This range is in many cases between 40 and 400 Hz, but can also be considerably larger.
• the signal at the output of the bandpass filter 50 is also sent to the series connection of a rectifier 66 and a logarithmic amplifier 68, which produces a DC voltage which logarithmically depends on the amplitude of the signal rectified by the rectifier 66.
• the monostable multivibrator produces output pulse sequences with pulses of variable pulse width, the repetition frequency of which varies between 40 and 400 Hz, corresponding to the frequency range which is particularly suitable for stimulating the cochlea. These pulse sequences are modulated onto 74 carrier signals in the transmitter.
• the multi-channel transmitter uses four carrier signals, two of which are at 12 MHz and the other two at 31 MHz.
• the pulse trains of channels 1 and 3 are used to modulate 12 MHz carrier signals; the pulse trains of channels 2 and 4 modulate 31 MHz carrier signals.
• the carrier signals, to which the signals of channels 1 and 2 are modulated, lie on a transmitter coil, and the carriers to which the signals of channels 3 and 4 are modulated, lie on a second transmitter coil.
• the two carrier frequencies applied to a transmitter coil are therefore different, and mutual interference between the channels can be avoided.
• the use of only one transmitter coil for two channels is a simplification. For other reasons, however, it can prove beneficial to use a separate transmitter coil for each channel.
• FIG 3 shows the construction of a preferred embodiment of the multi-channel receiving device for four channels.
• Each channel includes a coil 81-84, the coils 81 and 82 being inductively coupled to the transmitter coil 76 and the coils 83 and 84 being inductively coupled to the transmitter coil 78.
• Each of the receiving coils 81-84 has a capacitor 85 connected in parallel, which produces resonant circuits with a resonance frequency of 12 or 31 MHz.
• the signal picked up by the coil 81 passes to a demodulator consisting of the diode 86, the capacitor 87 and the resistor 88. If pulse width modulation and demodulation are used, a zener diode can be installed in parallel with the resistor 88, thereby limiting the voltage at the detector output becomes. This allows voltage fluctuations to be minimized by changing the coupling between the transmitting and receiving coils.
• the voltage at the output 90 of the detector is preferably between 0 and 3 volts and has a frequency between 40 and 400 Hz in accordance with the modulation signal.
• the following Mete ⁇ ode can be used advantageously: In order to by using a plurality of space requirements arising from receiving coils, the receiving coils are overlapped and arranged in groups.
• each of the receive coils within a group is tuned to its own frequency, the mutual inductance of the two or three receive coils would lead to an unacceptably high crosstalk if the coils were simply arranged one above the other. If the coils are arranged in such a way that their mutual magnate flows compensate one another, as shown in FIG. 4, then their mutual inductance disappears. In this way, two or three independent channels with negligible crosstalk are obtained, which only require a little more space than a single channel. 4, the coils 81 and 82 and 83 and 84 are arranged in two separate groups, the two groups being intended to illustrate different possibilities for arranging two coils within a group to compensate for the mutual inductance.
• Each of the coils has a diameter of the order of 1.5-2 cm and the distance between the two groups of coils is approximately 3 cm to prevent crosstalk between the groups.
• each receiving channel Since the coils 81 and 82 are tuned to different frequencies (for example 12 and 31 MHz), each receiving channel only receives and demodulates the signal in the associated transmission channel.
• the demodulated signal from each of the receiving channels is passed to a multi-channel electrode, as is shown, for example, in FIG. 5.
• Each channel can be connected to one or more electrode contacts 92, which are arranged along the body of the multi-carial electrode, so as to be able to stimulate specific locations along the cochlea for targeted pitch perceptions.
• bipolar stimulation, unipolar stimulation against a remote ground electrode or stimulation against a distributed common ground can be used.
• the multi-channel electrode consists of an elongated cast body made of silicone elastomer, e.g. Silastic, in which a plurality of wires, indicated by 91, are embedded. Each wire ends in a contact ball 92 that lies on the surface of the electrode body 90.
• the arrangement of the contact bodies on the surface of the electrode enables selective stimulation of individual cochlear regions after implantation of the electrode in the cochlea.
• the response areas for high frequencies are in the basal region, those for lower frequencies in the apical region of the cochlea.
• desired pitch perceptions can be produced according to the placement of the electrode contacts 92 within the cochlea.
• a pitch continuum can be achieved by the additional variation of the stimulation frequency.
• the schematic section through the multi-channel electrode in FIG. 7 illustrates the placement of the wires 93 and 94 within the electrode. For simplicity, only two wires are shown. Each of the wires is crimped for strain relief, which also facilitates bending of the electrode as it is inserted into the cochlea.
• the wires are made of teflon-coated platinum (90%) - iridum (10%) with a diameter of 25 ⁇ m.
• the balls at the ends of the wires are 0.3 mm in diameter and are obtained by melting the wire in a flame, forming a small ball.
• the contact balls are arranged in pairs in two rows opposite one another along the electrode body.
• the largest diameter of the electrode is 0.9 mm and decreases to 0.5 mm towards the tip.
• the total length of the electrode must correspond to an insertion of the electrode into the cochlea over a length of 20 - 25 mm.
• Fig. 8 one of two identical mold halves 96 is shown for producing the multi-channel electrode. It contains a narrowing groove 97 of the desired electrode geometry. There are a plurality of holes 98 in the groove, all of which are connected to a vacuum channel 99.
• the wires are first positioned in the channel of the mold half by holding the spherical ends of the wires in place in the holes 98 of the mold halves with the aid of vacuum suction. The mold is then assembled and into channels 97
• FIG. 9a An alternative embodiment of the electrode 100 is shown in FIG. 9a, the curvature of which is made in accordance with the cochlea.
• a straight rod 102 such as a steel wire, is located inside the electrode and is slowly pulled out while the electrode is inserted into the cochlea. As a result, the electrode 100 returns to its original shape.
• either an electrode channel for stimulation or any one can be selected Number of electrode contacts are interconnected. In the latter case, different thresholds for different electrode contacts can be compensated for by appropriate adaptation.
• a single-channel electrode attached to or near the round window in connection with a larger ground electrode, which is also attached outside the cochlea, can be used together with a single-channel implant and an external single-channel sound processor / transmitter.
• the system described in 1) has already been tested on some selected completely deaf volunteers, and an open speech understanding of 60 - 70% for unknown words or sentences can be achieved without additional lip reading, only by means of the electrostimulation alone. This means that this prosthesis can already be regarded as a valuable hearing aid for the completely deaf.
• the system in 2) is primarily intended for deaf people who are not completely deaf, but for hearing impaired children and for hearing impaired children. Both for the single-channel and for the
• Multi-channel version of the external switching processor / transmitter it is possible not to convert the sound waves into pulse chains, but to use the analog signal.
• suitable dynamic range compression and compensation of the dependence of the volume on the frequency are very important properties of the sound processor / transmitter. Both properties are so important because the dynamic range between the stimulus intensity, which is necessary for a threshold hearing impression and that which is necessary for a too loud hearing impression, is much smaller than in the acoustic case for normal hearing, and aggravatingly, in several cases, a strong dependency of the Threshold and above-threshold volume sensations from the stimulation frequency can be recorded.
• Nonlinearities are preferably used for dynamic range compression. These nonlinearities can be logarithmic, obey a power function, consist of a piecewise linear function or be of another suitable form.
• the non-linearity is preferably realized with the aid of suitably connected differential amplifiers or operational amplifiers in connection with diode networks or transistors connected as diodes.
• the non-linearity can be controlled by a frequency-shifted signal.
• the even-numbered distortion products can be eliminated by a narrowband HF bandpass filter in the up-shifted signal before the signal is mixed down again into the original audible frequency range.
• Another way to reduce distortion products is to use multiple nonlinearities within octave-wide bands.
• controlled amplifier that has sufficiently low response and decay time constants of 2 - 10 ms or 100 - 300 ms.
• the amplitude characteristic of this controlled amplifier can be given the desired shape by inserting suitable non-linearities into the signal path.
• the circuit for dynamic range compression can also precede the circuit for frequency response compensation. In this case, the extent of the frequency response compensation required is smaller, but it must be carried out very precisely.
• Fig. 10 shows the block diagram of a single or multi-channel sound processor including transmitter.
• a multi-channel stimulator consists of several essentially identical ones Channels with their own high-frequency transmission. Each channel is responsible for a specific audio frequency band, which is filtered out by the respective band filter. In the case of a single-channel stimulator, only one of these channels is built up and the bandpass filter 103 can therefore be omitted.
• the signal picked up by an electret microphone 104 has a dynamic range of over 80 dB.
• the transformation of this large dynamic range to the range of the permissible stimulation intensity of approximately 10 dB takes place with the help of dynamic compressors 107 described in more detail below and / or an input-controlled backward regulation which acts on the control amplifier 105.
• the control has the advantage of the low temporal non-linear distortion, but due to the finite response time, spikes are passed through when loud signals suddenly set in, so that both dynamic compression and control are generally used.
• the circuit 106 for adapting the frequency response to the previously measured frequency dependency of the patient's iso-loudness curves contains frequency-dependent components such as RC or LC elements. It is constructed in a conventional manner. In principle, it can also be installed after the dynamic compression circuit 107, in which case only slight but very precise frequency influences are required; the signal to be transmitted is transmitted to the implanted, tuned receiving circuit 110 and the demodulator via an amplitude-modulated transmitter 108 with the output circuit 109 111 and from there to the electrode 112.
• FIG. 11 shows the circuit used for dynamic compression. It is based on the integrated circuit 112 (TL441), which essentially consists of four differential amplifiers, the outputs of which are connected in parallel and controlled via voltage dividers with different attenuation.
• the exit voltage which is the difference voltage between points 122 and 123 (y or ) is available, depends logarithmically on the input voltage at 121.
• This input voltage is supplied to the input 124 of the circuit 112 via the capacitor 113, which is used for blocking possibly existing DC voltages.
• Resistor 120 is used to determine the DC voltage level at input 124 of the circuit.
• the differential voltage between points 122 and 123 is converted into a single-ended voltage in a conventional manner with the aid of an operational amplifier 114, which forms a subtraction circuit with resistors 115 to 118.
• the trim potentiometer 119 is used to adjust the offset voltage.
• Fig. 12 is the circuit diagram of the amplitude modulatable transmitter.
• the transmitter consists of the oscillator 120, which generates the carrier frequency of 12 MHz, and the power output stage.
• the output circuit 126 forms, together with the implanted receiving circuit 127 tuned to the same frequency, a band filter.
• a current-controlled band filter that is to say a high-impedance control at the input, shows a maximum of the voltage induced in the secondary circuit for a certain coupling, the so-called critical coupling. In the vicinity of this stationary point, the induced secondary voltage will therefore only depend slightly on the coupling.
• the tolerance indicator shafts with regard to a shift in the position of the transmitting coil will be very favorable given the distance between the transmitting and receiving coils leading to critical coupling.
• a "critical distance" of approximately 10 to 12 mm is obtained by suitably selected circular qualities. With a transmitter coil diameter of 23 mm and a shift of ⁇ 10 mm, the secondary voltage only changes by -5%.
• the collector can be used for amplitude modulation modulation are not used, but it must be worked either with emitter current modulation or with base modulation (as in the example shown here).
• the modulation voltage is supplied to the base via resistor 133, coupling winding 134 and resistor 130.
• Resistor 130 is selected in accordance with the required output power and enables precise metering of the RF drive power without an otherwise necessary but cumbersome change in the number of turns of coupling winding 134.
• Capacitors 131 and 132 serve to connect the modulation input and the operating voltage supply in accordance with RF with mass.
• the Schottky diode 127 prevents the occurrence of instabilities which could give rise to interference vibrations.
• the output circuit 126 of the transmitter is located on a plexiglass earpiece and can be positioned exactly above the implanted receiving coil in this way
• the entire transmitter is designed in miniature, so that it can also be accommodated on this ear piece.
• This has the advantage of very low radiation of electromagnetic waves, since all high-frequency lines are shorter than 2 cm and therefore hardly act as an antenna.
• the described method of electrostimulation of the hearing using a multi-frequency system according to the present invention provides improved hearing in completely deaf and severely impaired hearing.
• the use of a division of the sound signal into bands and the selective stimulation of different locations within the cochlea improve the quality and intelligibility of the hearing impressions. Because the implanted receiver circuit contains only passive electronic components, no supply power is necessary, and it only has to the stimulation signals themselves are transmitted through the skin into the interior of the body.
• the hearing prosthesis including the multi-channel cochlear electrode can be easily manufactured, and the manufacturing method ensures an exact positioning of the electrode contacts to achieve the desired pitch impressions when stimulating the cochlea. While pulse width modulation is used in the preferred embodiment, other modulation strategies such as amplitude or frequency modulation can also be used.
• Analog stimulation signals after suitable electronic processing can be used as well as pulsed or digital signals. While in the described digital version the audio frequency bands are transformed into corresponding signal bands between 40 and 400 KZ, the corresponding signal bands can have the same frequency range as the audio frequency bands in both the digital and the analog circuit design.

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• 1981
  • 1981-03-23 JP JP56500943A patent/JPH0367694B2/ja not_active Expired
  • 1981-03-23 WO PCT/AT1981/000005 patent/WO1982000760A1/en unknown
• 1989
  • 1989-01-26 JP JP1017391A patent/JPH01244750A/ja active Pending
  • 1989-01-26 JP JP1017392A patent/JPH01244751A/ja active Pending

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