Abstract
A hearing assistance device includes an implantable part for electrically stimulating an auditory nerve of a user. The electrical stimulation of the cochlear nerve by a cochlear implant hearing assistance device is improved by providing an implanted part that comprises a) a current source generator; and b) an electrode array configured to be located inside one of the cochlear scala or adjacent to the auditory nerve. The hearing assistance device is configured to produce a time-varying waveform delivered by the current source generator, the time-varying waveform comprising a positively sloping positive pulse.
Claims
1. A method of operating a hearing assistance device, the hearing assistance device comprising an implantable part, the method comprising: providing an electrode array comprising one or more stimulation electrodes configured to be located inside one of the cochlear scala or adjacent to the auditory nerve, or at the auditory brainstem of a user of the hearing assistance device; providing stimulation current to generate electric stimulation pulses to one or more of said stimulation electrodes; and using said stimulation current to provide a parameterized time-varying waveform of said electric stimulation pulses to one or more of said stimulation electrodes, said parameterized time-varying waveform comprising a positively sloping positive pulse when plotted in terms of amplitude of current versus time, thereby creating auditory sensation for the user of the hearing assistance device, the auditory sensation corresponding to sound from a local environment of the user during normal operation of the hearing assistance device, wherein said stimulation pulses generated by said one or more stimulation electrodes are shaped as a positively sloping positive pulse when plotted in terms of amplitude of current versus time, according to said parameterized time-varying waveform.
2. A method according to claim 1 comprising providing a model of ionic currents present in a nerve or neuron, from which the temporal pattern of current to deliver can be computed so that a specific discharge probability and/or temporal accuracy can be obtained.
3. A method according to claim 1 comprising one or more of the following steps: computing a first passage time probability density using a model of said parameterized time-varying waveform, computing a set of fibres, which are activated by a single pulse of said parameterized time-varying waveform, computing interactions between sub-sequent pulses in a pulse train of said parameterized time-varying waveforms.
4. A method according to claim 1 comprising modulating the probability of discharging neurons of the cochlear nerve using a parameterized time-varying waveform specifically designed to limit the spread of excitation.
5. A method according to claim 1 comprising modulating the parameterized time-varying waveform by acting on either the spread of excitation or the discharge probability or the discharge latency.
6. A method according to claim 1 comprising a fitting procedure wherein the patient estimates the extent of the spread of excitation a specific electric stimulation pulse.
7. A method according to claim 1 comprising a fitting procedure, wherein a subjective measure related to the use of a masking paradigm is provided in which the patient is asked to detect the presence of a target stimulation in concurrence with a masker presented simultaneously or earlier.
8. A method according to claim 1 comprising a fitting procedure, wherein an objective measure for fitting the pulse-shape is provided, said objective measure being based on recording the nerve response after its stimulation or any evoked neural response produced by the stimulation.
9. A data processing system comprising a processor and program code means for causing the processor to perform the steps of the method of claim 1.
10. A fitting system configured to estimate the extent of the spread of excitation of different parameterized time-varying waveforms according to the method of operating a hearing assistance device defined in claim 6.
Description
BRIEF DESCRIPTION OF DRAWINGS
(1) The disclosure will be explained more fully below in connection with a preferred embodiment and with reference to the drawings in which:
(2) FIGS. 1A-1C shows a use case of a hearing assistance device comprising an implanted part according to the present disclosure, FIG. 1A schematically showing the head of a user wearing the device, FIG. 1B schematically showing a cross section of cochlea including a multi electrode array of the device, FIG. 1C schematically showing a perspective cross-sectional view of cochlea with the multi electrode array is mounted in scala tympani,
(3) FIGS. 2A-2C shows various partitions of a hearing assistance device according to the present disclosure; the embodiment in FIG. 2A comprising only an implanted part, the embodiment in FIG. 2B comprising an implanted part and an external part with a wireless communication link between them, and the embodiment in FIG. 2C comprising the same elements as the embodiment of FIG. 2B, but where the external part comprises an antenna part for establishing the wireless link to the implanted part and a processing part for processing an audio signal, and where the antenna and processing parts are connected by a wired link,
(4) FIGS. 3A-3E schematically shows elements of a neural response to an exemplary spatial stimulation pulse excited at a (single, mono-polar) location (L.sub.z) along the cochlear nerve, FIG. 3A showing the exemplary (prior art) bi-phasic square symmetric waveform stimulation pulse with a time delay between the positive and negative phases of the pulse, the positive and negative pulses having equal height and width, FIG. 3B illustrating stimulation of neurons along the cochlear nerve due to stimulation of a single specific electrode (E.sub.z) at location L.sub.z, FIG. 3C schematically showing exemplary waveforms of the stimulation pulses as seen by neurons located at various locations to both sides of the stimulated electrode (E.sub.z), FIG. 3D schematically illustrating a spatial current spread caused by the stimulation pulse at the stimulated electrode (E.sub.z), and FIG. 3E schematically illustrating a corresponding spatial spread of neuron excitation caused by the stimulation current,
(5) FIGS. 4A-4D schematically shows three examples of mono-polar and multi-polar stimulation schemes, FIG. 4A illustrating the (multi-array) stimulation electrode(s) spatially distributed along a cochlear nerve, FIG. 4B illustrating a mono-polar stimulation of electrode E.sub.z with a bi-phasic pulse, FIG. 4C illustrating a first multi-polar stimulation comprising stimulation of electrode E.sub.z with a positive pulse and neighbouring electrodes E.sub.z+1 and E.sub.z−1 with negative pulses, and FIG. 4D illustrating a second multi-polar stimulation comprising stimulation of electrode E.sub.z with bi-phasic pulse and neighbouring electrodes E.sub.z+1 and E.sub.z−1 with corresponding bi-phasic pulses of opposite phase,
(6) FIGS. 5A-5F shows six different exemplary bi-phasic stimulation ‘pulse’ waveforms according to the present disclosure and their modification at a neuron located spatially apart from the primary target neurons of the electrode emitting the stimulation pulse(s), FIG. 5A showing a stimulation pulse comprising a positive sloped pulse according to the present disclosure and an arbitrary negative pulse (or none), FIG. 5B showing a bi-phasic asymmetric waveform stimulation pulse comprising a positive sloped pulse according to the present disclosure and a square negative pulse, FIG. 5C showing a bi-phasic asymmetric waveform stimulation pulse as in FIG. 5B, but wherein the positive and negative pulses have different widths, FIG. 5D showing a bi-phasic sloped, symmetric waveform stimulation pulse according to the present disclosure, FIG. 5E showing a bi-phasic asymmetric waveform stimulation pulse comprising a positive positively sloped (triangular) pulse according to the present disclosure and a square negative pulse, and FIG. 5F showing a bi-phasic sloped (triangular), symmetric waveform stimulation pulse according to the present disclosure,
(7) FIGS. 6A-6B schematically shows in FIG. 6A an exemplary (step-like) relationship between the slope of a positively sloped (positive) pulse (cf. e.g. FIG. 5A) arriving at a neuron of the cochlear nerve and the probability of discharging the neuron (cf. FIGS. 7A-7C), while FIG. 6B illustrate exemplary stimulation pulse slopes having values below and above a threshold slope SL.sub.TH, respectively,
(8) FIGS. 7A-7C shows a combined illustration of the spatial range of excited neurons for two different stimulation pulse waveforms, a square waveform as shown in FIG. 3A and a positively sloped waveform according to the present disclosure as shown in FIG. 5D, FIG. 7A schematically illustrating waveforms of the positively sloped stimulation pulses as seen by neurons located at various locations to both sides of the (single, mono-polar) stimulated electrode (E.sub.z), FIG. 7B schematically illustrating a spatial current spread caused by the stimulation pulse at the stimulated electrode (E.sub.z), and FIG. 7C schematically illustrating a corresponding spatial spread of neuron excitation caused by the stimulation current, and
(9) FIG. 8 shows an embodiment of a hearing assistance device comprising an implanted part partitioned as schematically illustrated in FIG. 2C, and
(10) FIGS. 9A-9B schematically illustrates two further use cases of a hearing assistance device comprising an implanted part according to the present disclosure, both cases showing a bilateral (or binaural) fitting first and second hearing assistance devices, which may not (bilateral) or may (binaural) be in communication with each other, FIG. 9A showing a use case where each hearing assistance device comprises an implanted part according to the present disclosure, and FIG. 9B showing a use case where one of the hearing assistance devices comprises an implanted part according to the present disclosure, and where the other comprises an output transducer for mechanically or acoustically (e.g. a speaker) providing stimuli interpreted by the user as sound.
(11) The figures are schematic and simplified for clarity, and they just show details which are essential to the understanding of the disclosure, while other details are left out. Throughout, the same reference signs are used for identical or corresponding parts.
(12) Further scope of applicability of the present disclosure will become apparent from the detailed description given hereinafter. However, it should be understood that the detailed description and specific examples, while indicating preferred embodiments of the disclosure, are given by way of illustration only. Other embodiments may become apparent to those skilled in the art from the following detailed description.
DETAILED DESCRIPTION OF EMBODIMENTS
(13) FIGS. 1A-1C shows a use case of a hearing assistance device comprising an implanted part according to the present disclosure.
(14) FIG. 1A illustrates a monaural hearing assistance system comprising a single hearing assistance device of the cochlear implant type located at a right ear (ear1) of a user (U). Other embodiments may comprise a bilateral (or binaural, or hybrid solutions, e.g. comprising two implanted electrodes and one common processor) system wherein a hearing assistance device is located at each of the ears (ear1, ear2) of a user (U), the two hearing assistance devices being optionally in communication with each other in that they each comprise a transceiver for establishing a (wireless or wired) link between them allowing the transmission and reception of information to/from the other device. The hearing assistance device comprises an external part and an implanted part. Likewise, a hearing assistance system according to the present disclosure may additionally comprise any other multi-electrode-array stimulation, alone or combined with any other acoustic or vibrator-based stimulation on the same ear or the other. The external part is adapted to be located at or in an ear of the user and comprises in the embodiment of FIG. 1A a sound capture and processing part (BTE1) adapted to be located behind an ear of the user (U), and a communication part (COM1) in operational communication with the sound capture and processing part (BTE1), here via a wired connection. The communication part (COM1) is configured to communicate with the implanted part, including to transfer information about a current electric stimulus (e.g. representative of a current sound signal picked up by the sound capture and processing part (BTE1)) to be applied to the cochlear nerve (cochlear nerve). The cochlear nerves are connected to the auditory centre of the brain (the Primary Auditory Cortex, denoted PAC in FIG. 1A) as indicated by the bold dashed lines. The implanted part comprises a communication and stimulation unit (SP-TU1) and a multi-electrode array (mea1) in operational communication with each other. The communication and stimulation unit (SP-TU1) is configured to exchange information with the communication unit (COM1) of the external part, including to receive stimulation information, and to generate corresponding stimulation pulses, and to apply such pulses to electrodes of the multi-electrode array (mea1).
(15) The multi-electrode array (mea, mea1) may comprise a flexible, originally substantially linearly shaped carrier with a number of individually electrically accessible electrodes located along the length of the carrier. In an embodiment, the flexible carrier is configured to adapt to the form of cochlea when inserted. Alternatively, the multi-electrode array (mea, mea1) may (semi-rigidly) be pre-shaped to the form of cochlea.
(16) FIG. 1B schematically shows a cross section of cochlea including a multi electrode array of the hearing assistance device. The multi-electrode array (mea) is in the transversal cross-sectional view of cochlea of FIG. 1B located at an inner wall in the right side of scala tympani. It may, however, be located other places in the scala tympani (e.g. as indicated by electrodes (mea2, mea3, mea4) having a dotted outline). Further, the multi-electrode array may be located elsewhere in proximity of the cochlear nerve (e.g. in one of the other scala). The three scala of cochlea, Scala tympani, Scala media and Scala vestibuli, are schematically illustrated and denoted by the same names in FIG. 1B. The cochlear partition (Cochlear partition) hosting (a part of) the cochlear nerve (Cochlear nerve) and separating the Scala media and Scala vestibuli from the Scala tympani, is schematically indicated in FIG. 1B. The cochlear nerve comprises hair cells (Hair cell) reaching into Scala media.
(17) FIG. 1C schematically shows a perspective cross-sectional view of cochlea (Cochlea) with the (exemplary location of) multi electrode array (mea) being mounted in scala tympani (Scala tympani). The multi electrode array (mea) comprises a carrier (carrier) comprising a number of electrodes (electrode), e.g. 8 or more, distributed along its length (cf. dashed arrow denoted L (length) and indicating a Direction of helicotrema, where Scala tympani and Scala vestibuli meet). Each electrode (electrode) is configured to provide the option of electrical stimulation of a particular part of the cochlear nerve as indicated by the bold line denoted electrical connections in FIG. 1C. The electrical connections are operationally connected to the stimulation unit (SP-TU1) in FIG. 1A (or similarly to unit STU-MEU-PU-CONT in FIG. 8).
(18) For some clinical cases of profound, deafness with non-implantable cochlea (e.g., fully ossified, Mondini syndrome, etc.) or non-stimulable cochlear nerve (e.g., nerve cut following Neurofibromatosis acoustic tumor surgery), the hearing assistance device may comprise electrodes placed on the auditory brainstem, i.e. beyond the cochlear nerve and before the auditory cortex. The present disclosure comprises such embodiments where the hearing assistance device is an auditory brainstem implant.
(19) FIGS. 2A-2C shows various partitions of a hearing assistance device according to the present disclosure.
(20) FIG. 2A shows a hearing assistance device (HAD) in its most basic form comprising only a, preferably self-contained (e.g. battery driven, and comprising an input transducer, e.g. a microphone, and appropriate processing capability), implanted part (IMPp). FIG. 2B shows a hearing assistance device (HAD) comprising an implanted part (IMPp) and an external part (EXTp) with a wireless (e.g. inductive) communication link (Wireless link) between them. The external part (EXTp) may e.g. comprise an input transducer, e.g. a microphone, and a signal processing unit for enhancing a received electric input signal and possibly for preparing a scheme for stimulating electrodes of the implanted part (IMPp) in dependence of the current input signal. The external part (EXTp) may further comprise antenna and transceiver circuitry for transferring stimulation information (and possibly corresponding energy) to the implanted part (IMPp) (which comprises corresponding antenna and transceiver circuitry to allow reception of the transmitted signals and energy, to establish the Wireless link). Alternatively, the link from the external part (EXTp) to the implanted part (IMPp) may be based on a wired connection. FIG. 2C shows a hearing assistance device (HAD) as in FIG. 2B but where the external part (EXTp) comprises an antenna part (ANTp) for establishing the wireless link to the implanted part (IMPp) and a processing part (BTEp) for processing an audio signal, and where the antenna and processing parts are connected by a wired link (Wired link, e.g. a cable). In an embodiment, the processing part (BTEp) is configured to be located at an ear of the user. Alternatively, the processing part (BTEp) and the antenna part (ANTp) may be connected by a wireless link. This may be particularly relevant, if the processing part (BTEp) is located elsewhere than at an ear of the user.
(21) FIGS. 3A-3E shows elements of a neural response to an exemplary spatial stimulation pulse excited at a (single, mono-polar) location (L.sub.z) along the cochlear nerve.
(22) FIGS. 3A-3E schematically illustrates how excitation of neurons spreads around an electric stimulation intended to stimulate a specific location (area) of the cochlear nerve in an exemplary electrical stimulation of a cochlear implant, for which a multi-electrode array (mea in FIGS. 3B-3C) produces a stimulation (of bi-phasic symmetric, square pulse shape, cf. FIG. 3A).
(23) FIG. 3A shows an exemplary time-variant, (prior art) bi-phasic, square symmetric waveform stimulation pulse with a time delay (ΔT.sub.pn) between the positive (SQP.sub.a) and negative (SQP.sub.c) phases of the pulse, the positive (Positive pulse) and negative (Negative pulse) pulses having equal height (A.sub.s) and width (p.sub.wa=p.sub.wc) to conserve charge neutrality (as indicated by the areas enclosed by the respective pulses being equal: Area(AP)=Area(CP)). The time variant waveforms are drawn in an amplitude (Intensity, e.g. charge density or current (e.g. in units of A)) versus time (T) plot.
(24) The solid line bi-phasic, symmetric stimulation pulse of FIG. 3A (denoted SQP.sub.a(L.sub.z) exhibits a square waveform of a given amplitude (A.sub.s, intensity, e.g. charge/phase) at the location (L.sub.z) of the neurons intended for receiving pulses from the electrode in question (cf. E.sub.z in FIG. 3B). When a given stimulation pulse arrives at neurons located a distance (L.sub.z+/−ΔL) away from the target neurons the amplitude (intensity) of the square pulse has been modified (decreased to A.sub.n, cf. dashed waveforms SQP.sub.a(L.sub.z+/−ΔL) and SQP.sub.c(L.sub.z+/−ΔL) in FIG. 3A for the positive and negative phases, respectively).
(25) FIG. 3B schematically illustrates stimulation of neurons (neurons) along the cochlear nerve (Cochlear nerve) due to stimulation of a single specific electrode (E.sub.z) at location L.sub.z. The current provided by the stimulation pulse(s) is indicated by ions (encircled +, − signs, respectively, in FIG. 3B). A multi-electrode array (mea) is schematically shown along the cochlear nerve with an accompanying length indication (arrow L) increasing towards the (rounded off) tip (tip) of the carrier. Electrodes E.sub.z−1, E.sub.z, E.sub.z+1, . . . , E.sub.Ne, where N.sub.e is the number of electrodes on the carrier (carrier), are spaced apart, e.g. according to a scheme adapted to a particular user, or to a general user. In an embodiment, the electrodes are regularly spaced by a predefined distance (Δd.sub.e). A reference electrode (Ref), e.g. to pick up charges during a mono-polar stimulation is shown. The reference electrode is preferably located outside cochlea.
(26) FIG. 3C schematically shows exemplary waveforms of the stimulation pulses as experienced by neurons located at various locations to both sides of the stimulated electrode (E.sub.z). The amplitude of the stimulation pulses is decreasing with increased distance from the stimulated electrode, as indicated by the graphs in dotted elliptical enclosures below the neurons (neurons) of the cochlear nerve in FIG. 3C. Corresponding amplitudes of the pulses are denoted A.sub.−2, A.sub.−1, A.sub.0, A.sub.+1, A.sub.+2, respectively.
(27) FIG. 3D schematically illustrates a spatial current spread (SCSP) caused by the stimulation pulse at the stimulated electrode (E.sub.z in FIG. 3B, 3C). The graph in FIG. 3D shows current versus distance (L) with a decrease in current to both sides of a maximum at the location L.sub.z of the stimulating electrode. FIG. 3E schematically illustrates a corresponding spatial spread of neuron excitation (Neural response) caused by the stimulation current. The graph in FIG. 3E shows neural response (NRES) versus distance (L) with a decrease in response to both sides of a maximum at the location L.sub.z of the stimulating electrode. A threshold value A.sub.TH indicates a level below which the neurons will not discharge. A corresponding spatial spread ΔL(A) around the location L.sub.z of the stimulating electrode is indicated. The threshold value A.sub.TH (and hence the spatial spread ΔL(A) of the neural response) depends on characteristics of the stimulation pulse, as indicated by the graphical insert of the bi-phasic, symmetric, square waveform associated with the dashed line indicating the threshold value A.sub.TH.
(28) In the example of FIGS. 3A-3E, a mono-polar stimulation using a single stimulation electrode (E.sub.z) and a reference electrode (Ref) is assumed. Further, a bi-phasic, symmetric square stimulation pulse is used for illustration. However, bi-polar (or multi-polar in general) stimulation and/or single phase (positive) or asymmetric stimulation may just as well be used (cf. FIGS. 4A-5F). According to the present disclosure, a positively sloped positive stimulation pulse is preferably used (cf. FIGS. 5A-5F). As indicated in FIG. 3C, the pulses experienced by neurons farther away from the stimulated electrode (E.sub.z in FIGS. 3A-3E) decrease in amplitude with distance from the stimulated electrode (cf. (A.sub.i, i=−2, −1, 0, +1, +2), but may still be large enough to be perceived by (i.e. to excite or fire) neurons at such locations, as illustrated in FIG. 3D, 3E.
(29) FIGS. 4A-4D shows three examples of mono-polar and multi-polar stimulation schemes. FIG. 4A illustrates the (multi-electrode array) stimulation electrode(s) (mea) spatially distributed (L) along a cochlear nerve comprising neurons (neurons) to be stimulated. A reference electrode (Ref) for use in mono-polar stimulation is further shown. FIG. 4B illustrates a mono-polar stimulation of electrode E.sub.z with a bi-phasic pulse (utilizing reference electrode Ref for the return current) in an amplitude A(E.sub.z) versus time t plot. FIG. 4C illustrates a first multi-polar (asymmetric) stimulation comprising stimulation of electrode E.sub.z with a positive pulse (A(E.sub.z) versus time t) and stimulating neighbouring electrodes E.sub.z+1 and E.sub.z−1 with negative pulses (A(E.sub.z+1) and A(E.sub.z−1), respectively, versus time t). To conserve charge balance, the total charge of the combined negative phases of stimulation pulses at electrodes E.sub.z+1 and E.sub.z−1 is equal to the charge of the positive stimulation pulse at electrode E.sub.z (as indicated by the corresponding (hatched) areas of the pulses). FIG. 4D illustrates a second multi-polar stimulation comprising stimulation of electrode E.sub.z with bi-phasic pulse and neighbouring electrodes E.sub.z+1 and E.sub.z−1 with corresponding bi-phasic pulses of opposite phase. Again, charge neutrality is intended as indicated by equality of the total areas of the positive phases and the total areas of the negative phases, respectively. The above stimulation schemes are only examples of mono-polar and multi-polar stimulation. Any mono-polar and multi-polar stimulation scheme (providing charge neutrality) may be used in combination with the stimulation waveform according to the present disclosure.
(30) From a general stimulation point of view (spatial and temporal definitions), the polarity will change due to both temporal and spatial definition. I.e., a positive temporal waveform definition can be inversed with polarity inversion of the electrode from a spatial definition criterion, in other words, physically inversed.
(31) FIGS. 5A-5F shows six different exemplary bi-phasic stimulation ‘pulse’ waveforms according to the present disclosure and their modification at a neuron located spatially apart from the primary target neurons of the electrode emitting the stimulation pulse(s).
(32) FIG. 5A shows a parameterized time-varying waveform (Intensity versus Time) comprising a positively sloping positive pulse stimulation pulse according to the present disclosure and an arbitrary negative pulse (or none).
(33) The (optional) negative pulse is shown in the dotted box denoted Negative pulse. A number of purely exemplary waveforms of the (optional) negative pulse are indicated in dashed line in the dotted box, including a passive discharge waveform. The parameterized time-varying waveform of the positive pulse (Positive pulse, SLP.sub.a(L.sub.z), solid line waveform in FIG. 5A) comprises a positively sloped waveform defined by heights of the rising (A.sub.s1) and falling (A.sub.s2) edges of the positive pulse (resulting in slope angle α) at the location (L.sub.z) of stimulation. The positive pulse has a width in time of p.sub.wa. When a given positive stimulation pulse arrives at neurons located a distance (ΔL) away from the target neurons the amplitudes and the slope of the sloped pulse has been modified (both decreased, to (A.sub.n1, A.sub.n2) and β, respectively, cf. dashed line waveform SLP.sub.a(L.sub.z+/−ΔL) in FIG. 5A). In general, the width of the positive pulse (p.sub.wa) is adapted to the current application (and possibly dynamically adapted to the current need for stimulation in the frequency range aimed at by a particular electrode). In an embodiment, the width of the positive pulse is of the order of tens of microseconds. In an embodiment, the width of the positive pulse is larger than 5 μs. In an embodiment, the width of the positive pulse is smaller than 100 μs. The negative pulse(s) is preferably configured to maintain charge neutrality together with the positive pulse(s). The negative pulse(s) may—together with the positive pulse—form part of a biphasic pulse in a mono-polar stimulation configuration, or may be applied to another electrode in a multi-polar stimulation configuration (cf. e.g. FIGS. 4A-4D).
(34) The parameterized time-varying waveform stimulation pulse of FIG. 5B is bi-phasic and asymmetric in that it comprises different positive and negative pulse waveforms. The positive stimulation pulse (Positive pulse, SLP.sub.a(L.sub.z), solid line waveform) exhibiting a positively sloped waveform is equal to that of FIG. 5A. The negative stimulation pulse (Negative pulse, SQP.sub.C(L.sub.z), dashed line waveform) is a square pulse. The pulse width (p.sub.wc) of the negative phase is equal to the pulse width (p.sub.wa) of the positive phase. Preferably, the area (charge) of the positive and negative pulses are equal (or a difference is otherwise compensated for, e.g. by multi-polar stimulation). The positive and negative phases are separated by a time delay ΔT.sub.pn. In an embodiment, the time delay is zero (see e.g. FIG. 5C).
(35) FIG. 5C shows a bi-phasic asymmetric waveform stimulation pulse as in FIG. 5B, but wherein the positive and negative pulses have different widths (p.sub.wa<p.sub.wc) and where the time delay between the positive and negative phases is minimal (e.g. intended to be zero). Again, preferably, the area (charge) of the positive and negative pulses are equal.
(36) FIG. 5D shows a bi-phasic sloped, symmetric waveform stimulation pulse according to the present disclosure comprising an arbitrary time delay (ΔT.sub.pn) between the positive and negative phases of the bi-phasic pulse. The positive pulse is as shown in FIGS. 5A, 5B and 5C, and the negative pulse is a symmetrically generated version thereof (e.g. mirrored around a horizontal axis). Hence, the area (charge) of the positive and negative pulses are equal, thereby preserving charge neutrality.
(37) FIG. 5E shows a bi-phasic asymmetric waveform stimulation pulse comprising a positive positively sloped (triangular) pulse TRP.sub.a(L.sub.z) according to the present disclosure and a square negative pulse SQP.sub.c(L.sub.z) comprising an arbitrary time delay (ΔT.sub.pn) between the positive and negative phases of the bi-phasic pulse. The triangular pulse is a special case of the parameterized time-varying waveform stimulation pulse of FIG. 5A, where the vertical rising edge is absent (A.sub.s1=A.sub.n1=0). Otherwise, it behaves as previously described, e.g. in connection with FIG. 5A. As in FIG. 5C, the positive and negative pulses have different widths (p.sub.wa<p.sub.wc). Again, preferably, the area (charge) of the positive and negative pulses are equal.
(38) FIG. 5F shows a biphasic sloped (triangular), symmetric waveform stimulation pulse according to the present disclosure comprising an arbitrary time delay (ΔT.sub.pn) between the positive and negative phases of the bi-phasic pulse. The positive pulse is a triangular pulse as shown in FIG. 5E and the negative pulse is a symmetrically generated version thereof. Hence, the area (charge) of the positive and negative pulses are equal, thereby preserving charge neutrality.
(39) To summarize FIGS. 5A-5F: According to the present disclosure, an important property of the stimulation pulse is the temporal shape of the positive phase. A possible time lag between the positive and negative phases and the waveform of the negative phase are of minor importance. An advantage of the present, sloped stimulation pulse scheme is that it allows to use a variation of slope to code for intensity. Square pulses allow intensity coding too. Preferably, a combination of a square pulse and a triangular pulse (here termed a ‘sloped pulse’, cf. FIG. 5B, 5C, 5D) can be used. A slope could be a constantly rising current as shown in FIGS. 5A-5F, or a fast succession of flat and rising current like a stair, or any other appropriate increase of the intensity from a lower start value to a higher end value.
(40) The goals are: to reduce the spatial current spreading (improve spatial selectivity) to improve intensity coding (e.g. have same amplitude coding with less energy using sloped instead of square pulses).
(41) For sloped pulses (cf. FIGS. 5A-5F), a neuron located a distance from the neuron(s) intended for stimulation sees a smaller pulse. However, in addition to a smaller amplitude of the pulse, the slope of the pulse (stimulating current) has also decreased.
(42) It is assumed that neurons in the auditory system are sensitive to the rate of depolarization. This means that they will discharge only if the slope of stimulation is higher than a certain value. This is assumed to be due to the presence of a fast activating sub-threshold potassium channel.
(43) Because of this rate threshold, using a pulse with ramp (a sloped pulse) in its temporal profile will allow to reduce the stimulation spatial selectivity (as illustrated in FIGS. 7A-7C).
(44) FIGS. 6A-6B shows in FIG. 6A an exemplary (step-like) relationship between the slope of a positively sloped (positive) pulse (cf. e.g. FIG. 5A) arriving at a neuron of the cochlear nerve and the probability of discharging the neuron (cf. FIGS. 7A-7C), while FIG. 6B illustrate exemplary stimulation pulse slopes having values below and above a threshold slope SLTH, respectively. It is believed that the observed property of neurons in the auditory system to be dependent on the slope of the stimulation pulses is linked to the presence of the (low voltage activated) potassium (K+) current I.sub.KLVA.
(45) FIGS. 7A-7C is a combined illustration of the spatial range of excited neurons for two different stimulation pulse waveforms, a square waveform as shown in FIG. 3A and a positively sloped waveform according to the present disclosure as shown in FIG. 5D. FIG. 7A (corresponding to FIG. 3C dealing with the same issue but for a prior art, square waveform) schematically illustrates waveforms of the positively sloped stimulation pulses as seen by neurons located at various locations to both sides of the (single, mono-polar) stimulated electrode (E.sub.z). As also indicated in FIGS. 5A-5F, the amplitudes (A.sub.i, i=−2, −1, 0, +1, +2) AND slopes (β.sub.i, i=−2, −1, 0, +1, +2) of the pulses decrease with increasing distance from the stimulation electrode (E.sub.z). FIG. 7B schematically illustrates a spatial current spread (SCSP) caused by the stimulation pulse at the stimulated electrode (E.sub.z). The graph in FIG. 7B shows current (SCSP) versus distance (L) with a decrease in current to both sides of a maximum at the location L.sub.z of the stimulating electrode. FIG. 7C schematically illustrates a corresponding spatial spread of neuron excitation (Neural response) caused by the stimulation current by illustrating a probability of neuron excitation along a length of the cochlear nerve centred around a location (L.sub.z) of a stimulated electrode (E.sub.z) for two different bi-phasic simulation pulse waveforms (as illustrated in FIGS. 3A and 5D, respectively, and indicated by inserts associated with the two different neural response curves), resulting in a different spread ΔL of excitation of neurons for the two waveforms. The threshold value A.sub.TH indicating a level below which the neurons will not discharge (dashed line). Corresponding spatial spreads ΔL(SQ) and ΔL(SL) around the location L.sub.z of the stimulating electrode is indicated for each the two stimulation pulse waveforms, square (SQ) and sloped (SL), respectively. As indicated in FIG. 7C, the spatial spread ΔL(SQ) of the neural response of the square stimulation pulse waveforms is larger than the spatial spread ΔL(SL) of the neural response of the sloped stimulation pulse waveforms.
(46) FIG. 8 shows an embodiment of a hearing assistance device comprising an implanted part partitioned as schematically illustrated in FIG. 2C. FIG. 8 illustrates a ‘normal operation scenario’, where electrodes (E.sub.z, z=1, 2, . . . , N.sub.e) of a flexible multi-electrode array (mea) of the implanted part (IMPp) (inserted into one of the scala of cochlea, e.g. scala tympani, and having its electrodes distributed along the extent of the cochlear nerve). The individual electrodes (E.sub.z) are stimulated in dependence of an acoustic input signal (AlnS) picked up by a microphone of an external part (EXTp) of the system (cf. FIG. 2B or 2C, here external part BTEp, e.g. adapted for being located behind an ear of a user). In the embodiment of FIG. 8, the relevant current stimulation scheme generated in the external BTEp part and the accompanying necessary electric energy are transferred to the implanted part via a communication link (Com-Link) between the implanted part (IMPp) and an external antenna part (ANTp).
(47) The external BTEp part comprises a forward signal path comprising: a microphone (or microphone system, e.g. for providing directionality in a specific DIR-mode), an A/D converter (A/D) for converting an analogue input signal to a digital signal by sampling the analogue input signal with a sampling frequency f.sub.s, a pre-emphasis filter (PEF) (e.g. a FIR filter) for adapting the input levels to a loudness perception of a normally hearing person (psychoacoustic adaptation), an analysis filter bank (A-FB) for converting a single time variant input signal to time-variant signals in a number p of frequency bands (I.sub.1:I.sub.p). The analysis filter bank may e.g. comprise a 128 point FFT providing p=64 frequency bands (or alternatively a filter bank followed by an envelope detector), a regrouping unit (REGR) for allocating p frequency bands to a number q of channels (CH.sub.1:CH.sub.q) equal to the number of electrodes used, e.g. q=20, configurable based on user data (cf. unit User specific data), e.g. based on the Bark scale or ‘critical bands’), a noise reduction algorithm (NR, with settings based on User specific data) adapted to attenuate signal components that are judged not to be part of a target signal, the noise reduction algorithm e.g. working independently on signals of each channel (CH.sub.1:CH.sub.q), a compression scheme (COMP, with settings based on User specific data) adapted to provide a level dependent compression of an input signal of each channel (CH.sub.1:CH.sub.q), a stimulation generator (STG) for generating a representation of the stimuli corresponding to a given intensity in a given frequency range at a given point in time (reflecting the current input audio signal) to be applied to corresponding electrodes of the implanted part, a local energy source (BAT), e.g. a battery, such as a rechargeable battery for energizing components of the hearing assistance device (BTEp, ANTp, IMPp), and a stimulus data coding unit (COD-PLS, with settings based on User specific data) for generating a scheme (incl. providing energy for stimulating each of the (active) electrodes (E.sub.z, max q electrodes, typically less) of the implanted part (IMPp), and forwarding stimuli (or coded stimuli) and energy via a cable to the antenna part (ANTp).
(48) The unit User specific data) may represent user data stored in a memory of the BTEp part or user data read into the various algorithms during a fitting session (or a combination of the two). Such data may include frequency dependent hearing thresholds and uncomfort levels (related to electric stimulation of the individual electrodes). The user specific data may include age, gender, etc.
(49) In an alternative embodiment, the components of the external part (BTEp) are included in the implanted part (IMPp), whereby the hearing assistance device is self-contained (cf. FIG. 2A). In such an embodiment, only a communication link to an external fitting system is necessary.
(50) In the embodiment of FIG. 8, a cable (denoted Cable to ANTp, and Cable from BTEp, in the BTEp- and ANT-p-ends, respectively) connects the BTE-part (BTEp) to the antenna part (ANTp). The cable provides separate digital data and power (denoted Stimuli-data+power) to the antenna part (ANTp).
(51) The antenna part (ANTp) is adapted for being located at the ear of the user allowing a communication link (Com-link) to be established with the implanted part (IMPp). The antenna part comprises: a power and data mixing unit (e.g. incl. a crystal oscillator) forming part of an inductive transmitter (and backlink receiver), (TX(Rx)) and antenna coil (Ant).
(52) The implanted part (IMPp) comprises: an inductive antenna coil (Ant) and receiver (and backlink transmitter), (RX(Tx)), a multi-electrode array (mea) comprising a (typically flexible) carrier (e.g. of silicone rubber) with a multitude of electrodes (E.sub.z) (of a corrosion resistant, e.g. noble, metal), each being individually connectable to a current source of a stimulation unit (STU) and preferably a voltage measurement unit for capturing a nerve response by a capacitor: a stimulation unit (STU) comprising a data extraction circuit, for extracting configuration data and stimuli data a current generator for generating a stimulus current (based on the extracted stimulus data) to be applied to the electrodes (E.sub.z), an interface to the electrodes (E.sub.z) comprising capacitors and switches (SW) for switching between individual electrodes and their connection to the stimulation unit (STU) and to a measurement unit (MEU), an operational amplifier (e.g. forming part of the measurement unit MEU) and preferably a processing unit (PU) for processing and identifying nerve response measurements (e.g. eCAPs), and a control unit (CONT) configured to control the timing and waveform of the application of stimulation signals in a stimulation time period and the coupling (via switch unit (SW)) of a relevant stimulation electrode to the stimulation unit (STU) and the optional measurement of a resulting response in a measurement time period and the optional coupling (via switch unit (SW)) of a relevant recording electrode to the measurement unit (MEU).
(53) An inductive, preferably bi-directional, communication link (Com-link) (e.g. comprising a 4 MHz carrier) is established by the inductive coils (Ant) of the antenna part (ANTp) and the implanted part (IMPp) when the two are located in an operational position (e.g. near the ear, on each side of the skin of a person). A back-link from the implanted part to the antenna- (and BTE-) part can e.g. be based on ‘load communication’. Due to the inductive coupling between the two antenna coils, any draw of current in the implanted part can be sensed in the antenna part. Thereby data-messages can be transmitted to the processor of the BTE-part (e.g. implant status signals (e.g. power level), electrode measurement data (impedances, and eCAPs). The backlink data can e.g. be coded in the signal using pulse width modulation (PWM) or amplitude modulation (AM). Alternatively, a digital coding scheme can be applied
(54) The external parts (BTEp and ANTp) can be partitioned in any other appropriate way than shown in FIG. 8. In an embodiment, the outputs of the BTE part (BTEp) are a) digitally coded data representing the electrode stimuli and b) a battery voltage, whereas the antenna part (ANTp) comprises a 4 MHz crystal oscillator whose output is mixed with the coded data to provide an on-off-coded signal, which is transmitted to the implant receiver via the inductive link. In an embodiment, all non-implanted parts of the hearing assistance device are located in a single external device (EXTp) and a communication link (Wireless link) between the implanted and external parts allowing the necessary exchange of information between the two parts (and possibly between the implanted (and/or the external) part and a fitting system), see e.g. FIG. 2B.
(55) In a fitting situation or during operation, nerve responses (e.g. eCAPs) and/or electrode impedance measurements are communicated to a fitting system for setting up the hearing assistance device according to a user's particular needs, either directly via the antenna part (ANTp) or via the BTE part (BTEp).
(56) The analogue electric signal representing an acoustic signal from the microphone is converted to a digital audio signal in the analogue-to-digital converter (A/D). The analogue inputs signal is sampled with a predefined sampling frequency or rate f.sub.s, f.sub.s being e.g. in the range from 8 kHz to 48 kHz (adapted to the particular needs of the application) to provide digital samples x.sub.n (or x[n]) at discrete points in time t.sub.n (or n), each audio sample representing the value of the acoustic signal at t.sub.n by a predefined number N.sub.s of bits, N.sub.s being e.g. in the range from 1 to 16 bits. A digital sample x has a length in time of 1/f.sub.s, e.g. 50 μs, for f.sub.s=20 kHz. In an embodiment, a number of audio samples are arranged in a time frame. In an embodiment, a time frame comprises 64 audio data samples. Other frame lengths may be used depending on the practical application.
(57) In an embodiment, the analysis filter bank (A-FB) comprise(s) a TF-conversion unit for providing a time-frequency representation of an input signal. In an embodiment, the time-frequency representation comprises an array or map of corresponding complex or real values of the signal in question in a particular time and frequency range. In an embodiment, the TF conversion unit comprises a filter bank for filtering a (time varying) input signal and providing a number of (time varying) output signals each comprising a distinct frequency range of the input signal. In an embodiment, the TF conversion unit comprises a Fourier transformation unit for converting a time variant input signal to a (time variant) signal in the frequency domain. In an embodiment, the frequency range considered by the hearing assistance device from a minimum frequency f.sub.min to a maximum frequency f.sub.max comprises a part of the typical human audible frequency range from 20 Hz to 20 kHz, e.g. a part of the range from 20 Hz to 8 kHz, e.g. 400 Hz to 6 kHz.
(58) FIGS. 9A-9B schematically illustrates two further use cases of a hearing assistance device comprising an implanted part according to the present disclosure, both cases showing a bilateral (or binaural) fitting first and second hearing assistance devices (which may not (bilateral) or which may (binaural) be in communication with each other). FIG. 9A shows a use case where each hearing assistance device comprises an implanted part according to the present disclosure. The functional parts of the individual first and second hearing assistance devices are discussed in connection with FIG. 1A and FIG. 8. FIG. 9B shows a use case (a so-called bimodal configuration) where one of the hearing assistance devices (the second) comprises an implanted part according to the present disclosure, and where the other hearing assistance device (the first) comprises an output transducer (OT1) for mechanically or acoustically (e.g. a speaker) providing stimuli intended to be interpreted by the user as sound. The output transducer (OT1) of the first hearing assistance device in FIG. 9B is shown as a (loud)speaker for generating acoustic stimuli, but may alternatively or additionally comprise a vibrator for mechanically exciting bones of the user (e.g. the skull). In an alternative embodiment, one of the hearing aid devices may comprise a speaker as well as an implanted part comprising a multi-electrode array. The first hearing assistance device may comprise a normal air conduction type hearing assistance device. The functional parts of the second hearing assistance device (comprising an implanted part) are discussed in connection with FIG. 1A and FIG. 8. The first and second hearing assistance devices may be configured to be able to exchange information between them. In an embodiment, first and second hearing assistance devices each comprises transceiver units allowing a wired or wireless link to be establish between them. An advantage of using a hearing assistance device according to the present disclosure in a bimodal fitting is that is that an improved frequency resolution of the implanted device can be provided to better match the frequency resolution of the corresponding air conduction hearing device.
(59) The invention is defined by the features of the independent claim(s). Preferred embodiments are defined in the dependent claims. Any reference numerals in the claims are intended to be non-limiting for their scope.
(60) Some preferred embodiments have been shown in the foregoing, but it should be stressed that the invention is not limited to these, but may be embodied in other ways within the subject-matter defined in the following claims and equivalents thereof.
REFERENCES
(61) U.S. Pat. No. 4,207,441 U.S. Pat. No. 4,532,930 [Clark; 2003] Graeme Clark, Cochlear Implants, Fundamentals and Applications, AIP Press, Springer Science+Business Media, Inc., New York, N Y, 2003. [Bal & Oertel; 2001] Ramazan Bal and Donata Oertel, Potassium Currents in Octopus Cells of the Mammalian Cochlear Nucleus, Journal of Neurophysiology, Vol. 86, pp. 2299-2311, Published 1 Nov. 2001
