Patent Description:
Hearing loss, which may be due to many different causes, is generally of two types, conductive and/or sensorineural. Conductive hearing loss occurs when the normal mechanical pathways of the outer and/or middle ear are impeded, for example, by damage to the ossicular chain or ear canal. Sensorineural hearing loss occurs when there is damage to the inner ear, or to the nerve pathways from the inner ear to the brain.

Individuals who suffer from conductive hearing loss typically have some form of residual hearing because the hair cells in the cochlea are undamaged. As such, individuals suffering from conductive hearing loss typically receive an auditory prosthesis that generates motion of the cochlea fluid. Such auditory prostheses include, for example, acoustic hearing aids, bone conduction devices, and direct acoustic stimulators.

In many people who are profoundly deaf, however, the reason for their deafness is sensorineural hearing loss. Those suffering from some forms of sensorineural hearing loss are unable to derive suitable benefit from auditory prostheses that generate mechanical motion of the cochlea fluid. Such individuals can benefit from implantable auditory prostheses that stimulate nerve cells of the recipient's auditory system in other ways (e.g., electrical, optical and the like). Cochlear implants are often proposed when the sensorineural hearing loss is due to the absence or destruction of the cochlea hair cells, which transduce acoustic signals into nerve impulses. An auditory brainstem stimulator is another type of stimulating auditory prosthesis that might also be proposed when a recipient experiences sensorineural hearing loss due to damage to the auditory nerve.

Certain individuals suffer from only partial sensorineural hearing loss and, as such, retain at least some residual hearing. These individuals may be candidates for electro- acoustic hearing prostheses.

<CIT>, <CIT>, <CIT> and <CIT> disclose known auditory prosthesis.

In one aspect, a method according to claim <NUM> is provided.

In another aspect, a auditory prosthesis according to claim <NUM> is provided.

Embodiments of the present description are described herein in conjunction with the accompanying drawings, in which:.

Embodiments of the present description are generally directed to the use of acoustic scene (environmental) analysis to determine the sound class of sound signals received at a hearing prosthesis and, accordingly, assess the estimated listening difficulty that the acoustic environment presents to a recipient of the hearing prosthesis. This difficulty of the recipient's listening situation can be used to adjust, adapt, or otherwise set the resolution (e.g., through multipolar widening and/or focusing) of the electrical stimulation signals delivered to the recipient to evoke perception of the sound signals. In other words, the resolution of the electrical stimulation signals is dynamically adapted based on the present acoustic environment of the hearing prosthesis. This dynamic adaption of the stimulation resolution may optimize the tradeoff between power consumption and hearing performance.

There are a number of different types of hearing prostheses in which embodiments of the present invention may be implemented. However, merely for ease of illustration, the techniques presented herein are primarily described with reference to one type of hearing prosthesis, namely a cochlear implant. However, it is to be appreciated that the techniques presented herein may be used in other hearing prostheses, such as auditory brainstem stimulators, electro-acoustic hearing prostheses, bimodal hearing prostheses, etc..

<FIG> is a schematic diagram of an exemplary cochlear implant <NUM> configured to implement embodiments of the present description. The cochlear implant <NUM> comprises an external component <NUM> and an internal/implantable component <NUM>.

The external component <NUM> is directly or indirectly attached to the body of the recipient and typically comprises an external coil <NUM> and, generally, a magnet (not shown in <FIG>) fixed relative to the external coil <NUM>. The external component <NUM> also comprises one or more sound input elements <NUM> (e.g., microphones, telecoils, etc.) for detecting/receiving input sound signals, and a sound processing unit <NUM>. The sound processing unit <NUM> includes, for example, one or more batteries (not shown in <FIG>) and a sound processor (also not shown in <FIG>). The sound processor is configured to process electrical signals generated by a sound input element <NUM> that is positioned, in the depicted embodiment, by auricle <NUM> of the recipient. The sound processor provides the processed signals to external coil <NUM> via, for example, a cable (not shown in <FIG>).

The implantable component <NUM> comprises an implant body <NUM>, a lead region <NUM>, and an elongate intra-cochlear stimulating assembly <NUM>. The implant body <NUM> comprises a stimulator unit <NUM>, an internal/implantable coil <NUM>, and an internal receiver/transceiver unit <NUM>, sometimes referred to herein as transceiver unit <NUM>. The transceiver unit <NUM> is connected to the implantable coil <NUM> and, generally, a magnet (not shown) fixed relative to the internal coil <NUM>.

The magnets in the external component <NUM> and implantable component <NUM> facilitate the operational alignment of the external coil <NUM> with the implantable coil <NUM>. The operational alignment of the coils enables the implantable coil <NUM> to transmit/receive power and data to/from the external coil <NUM>. More specifically, in certain examples, external coil <NUM> transmits electrical signals (e.g., power and stimulation data) to implantable coil <NUM> via a radio frequency (RF) link. Implantable coil <NUM> is typically a wire antenna coil comprised of multiple turns of electrically insulated single-strand or multi-strand platinum or gold wire. The electrical insulation of implantable coil <NUM> is provided by a flexible molding (e.g., silicone molding). In use, transceiver unit <NUM> may be positioned in a recess of the temporal bone of the recipient. Various other types of energy transfer, such as infrared (IR), electromagnetic, capacitive and inductive transfer, may be used to transfer the power and/or data from an external device to a cochlear implant and, as such, <FIG> illustrates only one example arrangement.

Elongate stimulating assembly <NUM> is configured to be at least partially implanted in cochlea <NUM> and includes a plurality of longitudinally spaced intra-cochlear electrical stimulating contacts (electrodes) <NUM> that collectively form a contact array <NUM>. Stimulating assembly <NUM> extends through an opening in the cochlea <NUM> (e.g., cochleostomy <NUM>, the round window <NUM>, etc.) and has a proximal end connected to stimulator unit <NUM> via lead region <NUM> that extends through mastoid bone <NUM>. Lead region <NUM> couples the stimulating assembly <NUM> to implant body <NUM> and, more particularly, stimulator unit <NUM>.

In general, the sound processor in sound processing unit <NUM> is configured to execute sound processing and coding to convert a detected sound into a coded signal that represents the detected sound signals. These encoded data are sometimes referred to herein as processed sound signals and are sent to the implantable component <NUM>. The stimulator unit <NUM> is configured to utilize the processed sound signals to generate electrical stimulation signals that are delivered to the recipient's cochlea via one or more stimulation channels. In this way, cochlear implant stimulates the recipient's auditory nerve cells, bypassing absent or defective hair cells that normally transduce acoustic vibrations into neural activity.

<FIG> illustrates an arrangement in which the cochlear implant <NUM> includes an external component. However, it is to be appreciated that embodiments of the present description may be implemented in cochlear implant systems having alternative arrangements. For example, <FIG> is a functional block diagram of an exemplary totally implantable cochlear implant <NUM> configured to implement embodiments of the present invention. Since the cochlear implant <NUM> is totally implantable, all components of cochlear implant <NUM> are configured to be implanted under skin/tissue <NUM> of a recipient. Because all components are implantable, cochlear implant <NUM> operates, for at least a finite period of time, without the need of an external device. An external device <NUM> can be used to, for example, charge the internal power source (battery) <NUM>. External device <NUM> may be a dedicated charger or a conventional cochlear implant sound processor.

Cochlear implant <NUM> includes an implant body (main implantable component) <NUM> and an implantable microphone <NUM>, an elongate intra-cochlear stimulating assembly <NUM> as described above with reference to <FIG>. The microphone <NUM> may be disposed in, or electrically connected to, the implant body <NUM>. The implant body <NUM> further comprises an internal transceiver unit <NUM>, a sound processor <NUM>, a stimulator unit <NUM> as described with reference to <FIG>, and the battery <NUM>.

The sound processor <NUM> is configured to execute sound processing and coding to convert received/detected sound signals (e.g., received by microphone <NUM>) into processed sound signals.

The transceiver unit <NUM> permits cochlear implant <NUM> to receive and/or transmit signals to external device <NUM>. For example, transceiver unit <NUM> may be configured to transcutaneously receive power and/or data from external device <NUM>. However, as used herein, transceiver unit <NUM> refers to any collection of one or more implanted components which form part of a transcutaneous energy transfer system. Further, transceiver unit <NUM> includes any number of component(s) which receive and/or transmit data or power, such as, for example a coil for a magnetic inductive arrangement, an antenna for an alternative RF system, capacitive plates, or any other suitable arrangement.

As noted above, <FIG> illustrates an embodiment in which the external component <NUM> includes the sound processor. As such, in the illustrative arrangement of <FIG>, processed sound signals are provided to the implanted stimulator unit <NUM> via the RF link between the external coil <NUM> and the internal coil <NUM>. However, in the embodiment of <FIG>, the sound processor <NUM> is implanted in the recipient. As such, in the embodiments of <FIG>, the processed sound signals do not traverse the RF link, but instead are provided directly to the stimulator unit <NUM>.

The human auditory system is composed of many structural components, some of which are connected extensively by bundles of nerve cells (neurons). Each nerve cell has a cell membrane which acts as a barrier to prevent intercellular fluid from mixing with extracellular fluid. The intercellular and extracellular fluids have different concentrations of ions, which leads to a difference in charge between the fluids. This difference in charge across the cell membrane is referred to herein as the membrane potential (Vm) of the nerve cell. Nerve cells use membrane potentials to transmit signals between different parts of the auditory system.

In nerve cells that are at rest (i.e., not transmitting a nerve signal) the membrane potential is referred to as the resting potential of the nerve cell. Upon receipt of a stimulus, the electrical properties of a nerve cell membrane are subjected to abrupt changes, referred to herein as a nerve action potential, or simply action potential. The action potential represents the transient depolarization and repolarization of the nerve cell membrane. The action potential causes electrical signal transmission along the conductive core (axon) of a nerve cell. Signals may be then transmitted along a group of nerve cells via such propagating action potentials.

<FIG> illustrates various phases of an idealized action potential <NUM> as the potential passes through a nerve cell. The action potential is presented as membrane voltage in millivolts (mV) versus time. The membrane voltages and times shown in <FIG> are for illustration purposes only and the actual values may vary depending on the individual. Prior to application of a stimulus <NUM> to the nerve cell, the resting potential of the nerve cell is approximately -<NUM> mV. Stimulus <NUM> is applied at a first time. In normal hearing, this stimulus is provided by movement of the hair cells of the cochlea. Movement of these hair cells results in the release of neurotransmitter into the synaptic cleft, which in return leads to action potentials in individual auditory nerve fibers. In cochlear implants, the stimulus <NUM> is an electrical stimulation signal (electrical stimulation).

Following application of stimulus <NUM>, the nerve cell begins to depolarize. Depolarization of the nerve cell refers to the fact that the voltage of the cell becomes more positive following stimulus <NUM>. When the membrane of the nerve cell becomes depolarized beyond the cell's critical threshold, the nerve cell undergoes an action potential. This action potential is sometimes referred to as the "firing" or "activation" of the nerve cell. As used herein, the critical threshold of a nerve cell, group of nerve cells, etc. refers to the threshold level at which the nerve cell, group of nerve cells, etc. will undergo an action potential. In the example illustrated in <FIG>, the critical threshold level for firing of the nerve cell is approximately -<NUM> mV. The critical threshold and other transitions may be different for various recipients and so the values provided in <FIG> are merely illustrative.

The course of the illustrative action potential in the nerve cell can be generally divided into five phases. These five phases are shown in <FIG> as a rising phase <NUM>, a peak phase <NUM>, a falling phase <NUM>, an undershoot phase <NUM>, and finally a refractory phase (period) <NUM>. During rising phase <NUM>, the membrane voltage continues to depolarize and the point at which depolarization ceases is shown as peak phase <NUM>. In the example of <FIG>, at this peak phase <NUM>, the membrane voltage reaches a maximum value of approximately <NUM> mV.

Following peak phase <NUM>, the action potential undergoes falling phase <NUM>. During falling phase <NUM>, the membrane voltage becomes increasingly more negative, sometimes referred to as hyperpolarization of the nerve cell. This hyperpolarization causes the membrane voltage to temporarily become more negatively charged than when the nerve cell is at rest. This phase is referred to as the undershoot phase <NUM> of action potential <NUM>. Following this undershoot phase <NUM>, there is a time period during which it is impossible or difficult for the nerve cells to fire. This time period is referred to as the refractory phase (period) <NUM>.

As noted above, the nerve cell must obtain a membrane voltage above a critical threshold before the nerve cell may fire/activate. The number of nerve cells that fire in response to electrical stimulation (current) can affect the "resolution" of the electrical stimulation. As used herein, the resolution of the electrical stimulation or the "stimulus resolution" refers to the amount of acoustic detail (i.e., the spectral and/or temporal detail from the input acoustic sound signal(s)) that is delivered by the electrical stimulation at the implanted electrodes in the cochlea and, in turn, received by the primary auditory neurons (spiral ganglion cells). As described further below, electrical stimulation has a number of characteristics/attributes that control the stimulus resolution. These attributes include for example, the spatial attributes of the electrical stimulation, temporal attributes of the electrical stimulation, frequency attributes of the electrical stimulation, instantaneous spectral bandwidth attributes of the electrical stimulation, etc. The spatial attributes of the electrical stimulation control the width along the frequency axis (i.e., along the basilar membrane) of an area of activated nerve cells in response to delivered stimulation, sometimes referred to herein as the "spatial resolution" of the electrical stimulation. The temporal attributes refer to the temporal coding of the electrical stimulation, such as the pulse rate, sometimes referred to herein as the "temporal resolution" of the electrical stimulation. The frequency attributes refer to the frequency analysis of the acoustic input by the filter bank, for example the number and sharpness of the filters in the filter bank, sometimes referred herein as the "frequency resolution" of the electrical stimulation. The instantaneous spectral bandwidth attributes refer to the proportion of the analyzed spectrum that is delivered via electrical stimulation, such as the number of channels stimulated out of the total number of channels in each stimulation frame.

The spatial resolution of electrical stimulation may be controlled, for example, through the use of different electrode configurations for a given stimulation channel to activate nerve cell regions of different widths. Monopolar stimulation, for instance, is an electrode configuration where for a given stimulation channel the current is "sourced" via one of the intra-cochlea electrodes <NUM>, but the current is "sunk" by an electrode outside of the cochlea, sometimes referred to as the extra-cochlear electrode (ECE) <NUM> (<FIG> and <FIG>). Monopolar stimulation typically exhibits a large degree of current spread (i.e., wide stimulation pattern) and, accordingly, has a low spatial resolution. Other types of electrode configurations, such as bipolar, tripolar, focused multi-polar (FMP), a. "phased-array" stimulation, etc. typically reduce the size of an excited neural population by "sourcing" the current via one or more of the intra-cochlear electrodes <NUM>, while also "sinking" the current via one or more other proximate intra-cochlear electrodes. Bipolar, tripolar, focused multi-polar and other types of electrode configurations that both source and sink current via intra-cochlear electrodes are generally and collectively referred to herein as "focused" stimulation. Focused stimulation typically exhibits a smaller degree of current spread (i.e., narrow stimulation pattern) when compared to monopolar stimulation and, accordingly, has a higher spatial resolution than monopolar stimulation. Likewise, other types of electrode configurations, such as double electrode mode, virtual channels, wide channels, defocused multi-polar, etc. typically increase the size of an excited neural population by "sourcing" the current via multiple neighboring intra-cochlear electrodes.

The cochlea is tonotopically mapped, that is, partitioned into regions each responsive to sound signals in a particular frequency range. In general, the basal region of the cochlea is responsive to higher frequency sounds, while the more apical regions of the cochlea are responsive to lower frequencies. The tonopotic nature of the cochlea is leveraged in cochlear implants such that specific acoustic frequencies are allocated to the electrodes <NUM> of the stimulating assembly <NUM> that are positioned close to the corresponding tonotopic region of the cochlea (i.e., the region of the cochlea that would naturally be stimulated in acoustic hearing by the acoustic frequency). That is, in a cochlear implant, specific frequency bands are each mapped to a set of one or more electrodes that are used to stimulate a selected (target) population of cochlea nerve cells. The frequency bands and associated electrodes form a stimulation channel that delivers stimulation signals to the recipient.

In general, it is desirable for a stimulation channel to stimulate only a narrow region of neurons such that the resulting neural responses from neighboring stimulation channels have minimal overlap. Accordingly, the ideal stimulation strategy in a cochlear implant would use focused stimulation channels to evoke perception of all sound signals at any given time. Such a strategy would, ideally, enable each stimulation channel to stimulate a discrete tonotopic region of the cochlea to better mimic natural hearing and enable better perception of the details of the sound signals. The present inventor has realized that, although focused stimulation generally improves hearing performance, this improved hearing performance comes at the cost of significant increased power consumption, added delays to the processing path, and increased complexity, etc. relative to the use of only monopolar stimulation. Additionally, the present inventor has realized that not all listening situations benefit from the increased fidelity offered by focused stimulation as different listening situations present varying levels of difficulty to cochlear implant recipients. For example, understanding speech in a quiet room is easier than understanding the same speech in a busy restaurant with many competing speakers. Accordingly, the present inventor has realized that recipients benefit more or less from the details of sound presented using increased stimulus resolution in different environments.

In accordance with the embodiments presented herein, a hearing prosthesis is configured to analyze received sound signals to determine the primary or main sound "class" of the sound signals. In general, the sound class provides an indication of the difficulty/complexity of a recipient's listening situation/environment (i.e., the environment in which the prosthesis is currently/presently located). Based on the sound class of the sound signals, the hearing prosthesis is configured to set the stimulus resolution of the electrical stimulation signals that are delivered to the recipient to evoke perception of the sound signals. The stimulus resolution is set in a manner that optimizes the tradeoff between hearing performance (e.g., increased fidelity) and power consumption (e.g., battery life). The hearing prosthesis uses higher resolution stimulation (i.e., stimulation that provides relatively more acoustic detail) in more challenging listening situations with increased expected listening effort, and uses lower resolution stimulation (i.e., stimulation that provides relatively less acoustic detail) in easier listening situations with lower expected listening effort. Since there is limited power available in a cochlear implant, it is therefore advantageous to adapt the stimulation resolution depending on the listening situation in order to optimize the stimulus resolution for the best overall hearing performance within the long-term power budget.

In accordance with the embodiments presented herein, the spatial resolution (i.e., the spatial attributes of the electrical stimulation) can be increased, for example, through use of focused stimulation. Conversely, the spatial resolution can be lowered, for example, through the use of monopolar, or wide/defocused, stimulation. These decreases in the stimulation resolution have the benefit of lower power consumption and lower complexity, but they also sacrifice listening fidelity (e.g., loss of sound details). In addition, the temporal resolution (i.e., the temporal attributes of the electrical stimulation) can be varied, for example, by changing the rate of the current pulses forming the electrical stimulation. Higher pulse rates offer higher temporal resolution and use more power, while lower pulse rates offer lower temporal resolution and are more power efficient.

Consequently, the stimulus resolution can be varied with differing associated power costs and, in certain situations, the techniques presented herein purposely downgrade hearing performance (e.g., speech perception) to reduce power consumption. However, this downgrade in hearing performance is dynamically activated only in listening situations where the recipient likely does not have difficulty understanding/perceiving the sound signals with lower stimulus resolution (e.g., monopolar stimulation, defocused stimulation, etc.) and/or does not need the details provided by high resolution (e.g., focused stimulation).

<FIG> is a schematic diagram illustrating the general signal processing path <NUM> of a cochlear implant, such as cochlear implant <NUM>, in accordance with embodiments presented herein. As noted, the cochlear implant <NUM> comprises one or more sound input elements <NUM>. In the example of <FIG>, the sound input elements <NUM> comprise two microphones <NUM> and at least one auxiliary input <NUM> (e.g., an audio input port, a cable port, a telecoil, a wireless transceiver, etc.). If not already in an electrical form, sound input elements <NUM> convert received/input sound signals into electrical signals <NUM>, referred to herein as electrical input signals, that represent the received sound signals. As shown in <FIG>, the electrical input signals <NUM> are provided to a pre-filterbank processing module <NUM>.

The pre-filterbank processing module <NUM> is configured to, as needed, combine the electrical input signals <NUM> received from the sound input elements <NUM> and prepare those signals for subsequent processing. The pre-filterbank processing module <NUM> then generates a pre-filtered output signal <NUM> that, as described further below, is the basis of further processing operations. The pre-filtered output signal <NUM> represents the collective sound signals received at the sound input elements <NUM> at a given point in time.

The cochlear implant <NUM> is generally configured to execute sound processing and coding to convert the pre-filtered output signal <NUM> into output signals that represent electrical stimulation for delivery to the recipient. As such, the sound processing path <NUM> comprises a filterbank module (filterbank) <NUM>, a post-filterbank processing module <NUM>, a channel selection module <NUM>, and a channel mapping and encoding module <NUM>.

In operation, the pre-filtered output signal <NUM> generated by the pre-filterbank processing module <NUM> is provided to the filterbank module <NUM>. The filterbank module <NUM> generates a suitable set of bandwidth limited channels, or frequency bins, that each includes a spectral component of the received sound signals. That is, the filterbank module <NUM> comprises a plurality of band-pass filters that separate the pre-filtered output signal <NUM> into multiple components/channels, each one carrying a single frequency sub-band of the original signal (i.e., frequency components of the received sounds signal).

The channels created by the filterbank module <NUM> are sometimes referred to herein as sound processing channels, and the sound signal components within each of the sound processing channels are sometimes referred to herein as band-pass filtered signals or channelized signals. The band-pass filtered or channelized signals created by the filterbank module <NUM> are processed (e.g., modified/adjusted) as they pass through the sound processing path <NUM>. As such, the band-pass filtered or channelized signals are referred to differently at different stages of the sound processing path <NUM>. However, it will be appreciated that reference herein to a band-pass filtered signal or a channelized signal may refer to the spectral component of the received sound signals at any point within the sound processing path <NUM> (e.g., pre-processed, processed, selected, etc.).

At the output of the filterbank module <NUM>, the channelized signals are initially referred to herein as pre-processed signals <NUM>. The number 'm' of channels and pre-processed signals <NUM> generated by the filterbank module <NUM> may depend on a number of different factors including, but not limited to, implant design, number of active electrodes, coding strategy, and/or recipient preference(s). In certain arrangements, twenty-two (<NUM>) channelized signals are created and the sound processing path <NUM> is said to include <NUM> channels.

The pre-processed signals <NUM> are provided to the post-filterbank processing module <NUM>. The post-filterbank processing module <NUM> is configured to perform a number of sound processing operations on the pre-processed signals <NUM>. These sound processing operations include, for example, channelized gain adjustments for hearing loss compensation (e.g., gain adjustments to one or more discrete frequency ranges of the sound signals), noise reduction operations, speech enhancement operations, etc., in one or more of the channels. After performing the sound processing operations, the post-filterbank processing module <NUM> outputs a plurality of processed channelized signals <NUM>.

In the specific arrangement of <FIG>, the sound processing path <NUM> includes a channel selection module <NUM>. The channel selection module <NUM> is configured to perform a channel selection process to select, according to one or more selection rules, which of the 'm' channels should be use in hearing compensation. The signals selected at channel selection module <NUM> are represented in <FIG> by arrow <NUM> and are referred to herein as selected channelized signals or, more simply, selected signals.

In the embodiment of <FIG>, the channel selection module <NUM> selects a subset 'n' of the 'm' processed channelized signals <NUM> for use in generation of electrical stimulation for delivery to a recipient (i.e., the sound processing channels are reduced from 'm' channels to 'n' channels). In one specific example, the 'n' largest amplitude channels (maxima) from the 'm' available combined channel signals/masker signals is made, with 'm' and 'n' being programmable during initial fitting, and/or operation of the prosthesis. It is to be appreciated that different channel selection methods could be used, and are not limited to maxima selection.

It is also to be appreciated that, in certain embodiments, the channel selection module <NUM> may be omitted. For example, certain arrangements may use a continuous interleaved sampling (CIS), CIS-based, or other non-channel selection sound coding strategy.

The sound processing path <NUM> also comprises the channel mapping module <NUM>. The channel mapping module <NUM> is configured to map the amplitudes of the selected signals <NUM> (or the processed channelized signals <NUM> in embodiments that do not include channel selection) into a set of output signals (e.g., stimulation commands) that represent the attributes of the electrical stimulation signals that are to be delivered to the recipient so as to evoke perception of at least a portion of the received sound signals. This channel mapping may include, for example, threshold and comfort level mapping, dynamic range adjustments (e.g., compression), volume adjustments, etc., and may encompass selection of various sequential and/or simultaneous stimulation strategies.

In the embodiment of <FIG>, the set of stimulation commands that represent the electrical stimulation signals are encoded for transcutaneous transmission (e.g., via an RF link) to an implantable component <NUM> (<FIG> and <FIG>). This encoding is performed, in the specific example of <FIG>, at the channel mapping module <NUM>. As such, channel mapping module <NUM> is sometimes referred to herein as a channel mapping and encoding module and operates as an output block configured to convert the plurality of channelized signals into a plurality of output signals <NUM>.

Also shown in <FIG> are a sound classification module <NUM>, a battery monitoring module <NUM>, and a stimulus resolution adaption module <NUM>. The sound classification module <NUM> is configured to evaluate/analyze the input sound signals and determine the sound class of the sound signals. That is, the sound classification module <NUM> is configured to use the received sound signals to "classify" the ambient sound environment and/or the sound signals into one or more sound categories (i.e., determine the input signal type). The sound classes/categories may include, but are not limited to, "Speech," "Noise," "Speech+Noise," "Music," and "Quiet. " As described further below, the sound classification module <NUM> may also estimate the signal-to-noise ratio (SNR) of the sound signals. In one example, the operations of the sound classification module <NUM> are performed using the pre-filtered output signal <NUM> generated by the pre-filterbank processing module <NUM>.

The sound classification module <NUM> generates sound classification information/data <NUM> that is provided to the stimulus resolution adaptation module <NUM>. The sound classification data <NUM> represents the sound class of the sound signals and, in certain examples, the SNR of the sound signals. Based on the sound classification data <NUM>, the stimulus resolution adaptation module <NUM> is configured to determine a level of stimulus resolution that should be used in delivering electrical stimulation signals to represent (evoke perception of) the sound signals. The level of stimulus resolution that should be used in delivering electrical stimulation signals is sometimes referred to herein as the "target" stimulus resolution.

The stimulus resolution adaptation module <NUM> is configured to adjust one or more operations performed in the sound processing path <NUM> so as to achieve the target stimulus resolution (i.e., adapt the resolution of the electrical stimulation that is delivered to the recipient). The stimulus resolution adaptation module <NUM> may adjust operations of the filterbank module <NUM>, the post-filterbank processing module <NUM>, the channel selection module <NUM>, and/or the mapping and encoding module <NUM> to generate output signals representative of electrical stimulation signals having the target stimulus resolution.

The stimulus resolution adaptation module <NUM> may adjust operations of the sound processing path <NUM> at a number of different time scales. For example, the stimulus resolution adaptation module <NUM> may determine the target stimulus resolution and make corresponding processing adjusts in response to a triggering event, such as the detection of a change in the listening environment (e.g., when the sound classification data <NUM> indicates the cochlear implant <NUM> is in a listening environment that is different from the previous listening environment). Alternatively, the stimulus resolution adaptation module <NUM> can determine the target stimulus resolution and make corresponding processing adjusts substantially continuously, periodically (e.g., every <NUM> second, every <NUM> seconds, etc.,), etc..

<FIG> illustrates an arrangement in which the cochlear implant <NUM> also comprises a battery monitoring module <NUM>. The battery monitoring module <NUM> is configured to monitor the charge status of the battery/batteries (e.g., monitor charge level, remaining battery life, etc.) and provide battery information <NUM> to the stimulus resolution adaptation module <NUM>. In addition to the sound classification data <NUM>, the stimulus resolution adaptation module <NUM> may also use the battery information <NUM> to determine the target stimulus resolution and make corresponding processing adjusts to the sound processing path operations. For example, if the battery information <NUM> indicates that the cochlear implant battery/batteries are below a threshold charge level (e.g., below <NUM>% charge), the stimulus resolution adaptation module <NUM> can switch the sound processing path <NUM> to a power saving mode that uses lower resolution (e.g., monopolar stimulation or defocused stimulation only) to conserve power.

<FIG> also illustrates a specific arrangement that includes one sound classification module <NUM>. It is to be appreciated that alternative embodiments may make use of multiple sound classification modules. In such embodiments, the stimulus resolution adaption module <NUM> is configured to utilize the information from each of the multiple sound classification modules to determine a target stimulus resolution and adapt the sound processing operations accordingly (i.e., so that the resulting stimulation has a resolution that corresponds to the target stimulus resolution).

Although <FIG> illustrates a cochlear implant arrangement, it is to be appreciated that the embodiments presented herein may also be implemented in other types of hearing prosthesis. For example, the techniques presented herein may be used in electro-acoustic hearing prostheses that are configured to deliver both acoustical stimulation and electrical stimulation to a recipient. In such embodiments, the prosthesis would include two parallel sound processing paths, where the first sound processing path is an electric sound processing path (cochlear implant sound processing path) similar to that is shown in <FIG>. In such arrangements, the second sound processing path is an acoustic sound processing path (hearing aid sound processing path) that is configured to generate output signals for use in acoustically stimulating the recipient.

<FIG> is a flow diagram illustrating further details of the techniques presented herein. For ease of illustration, <FIG> will be described with reference to the arrangement of <FIG>.

The flow of <FIG> begins at <NUM> where the cochlear implant <NUM> receives input sound signals for analysis by the sound classification module <NUM>. At <NUM>, the sound classification module <NUM> determines the sound class of the input sound signals. <FIG> illustrates example sound classes <NUM> that include: "Speech," "Music," "Noise," and "Quiet. " However, it is to be appreciated that additional sound classes are also possible.

In the example of <FIG>, the stimulus resolution adaption module <NUM> is configured to adapt the sound processing path <NUM> differently depending on the determined sound class. The different adaptions are generally shown in <FIG> at blocks <NUM>(A), <NUM>(B), and <NUM>(C). More specifically, <FIG> illustrates that for the "Music" and "Noise" sound classes, a higher target stimulus resolution is preferred since these are more difficult listening situations for a recipient. Accordingly, at <NUM>(B), the stimulus resolution adaption module <NUM> implements one or more adjustments to the sound processing path <NUM> to set a higher target resolution for the electrical stimulation delivered to the recipient. For the "Quiet" sound class, at <NUM>(C) the stimulus resolution adaption module <NUM> implements one or more adjustments to the sound processing path <NUM> to set a lower target stimulus resolution.

For the "Speech" sound class, a signal-to-noise ratio (SNR) of the input sound signals is used to control the resolution. The SNR provides a measure of how much speech compared to noise is present in the input sound signals. Therefore, at <NUM>, the SNR of the input sound signals is estimated and, at <NUM>(A) the SNR is used to determine the target stimulus resolution. In general, the higher the SNR, the lower the target stimulus resolution and, conversely, the lower the SNR, the higher the target stimulus resolution. The SNR can be used to select one of a number of discrete stimulation resolution levels, or could be applied across a continuum. For example, in one embodiment, the stimulus resolution adaption module <NUM> uses only two (<NUM>) stimulus resolution levels for the "Speech" sound class. These two levels comprise a low resolution setting for SNRs greater than a certain threshold (e.g., greater than <NUM> dB) where speech is much stronger than the noise, and a high resolution setting for SNRs below the threshold (e.g., below <NUM> dB). In other embodiments, the stimulus resolution adaption module <NUM> may make use of a large number of different resolution levels that each correlate to different SNR ranges. In these embodiments, the determined SNR is mapped to one of the number of different ranges and, accordingly, used to select the corresponding stimulus resolution.

Also shown in <FIG> is the battery information <NUM> which may be an optional input to the resolution adaptation blocks <NUM>(A)-<NUM>(C). As noted, the battery information <NUM> represents the state of charge of the battery/batteries and can be used to modify the target resolutions at <NUM>(A)-<NUM>(C). For example, if the battery charge is low, for example below <NUM>%, then a target higher resolution at <NUM>(A) or <NUM>(B) might be lowered so that battery life is given more priority. The battery information <NUM> could also be used to modify the SNR threshold(s) for the speech class. For example, in a low battery situation (e.g., charge below a threshold level), the SNR threshold(s) for selecting between two resolution steps could also be lowered, for example from <NUM> dB to <NUM> dB, to favor the use of lower resolution and, accordingly, conserve power.

As described elsewhere herein, the stimulus resolution adaption module <NUM> may set or adjust various operations of the sound processing path <NUM>, such as the operations of the filterbank module <NUM>, the post-filterbank processing module <NUM>, the channel selection module <NUM>, and/or the mapping and encoding module <NUM>, to set the stimulus resolution of the delivered electrical stimulation signals. In one embodiment, the spatial/spectral attributes of the stimulus resolution are set by switching between different channel/electrode configurations, such as between monopolar stimulation, wide/defocused stimulation, focused (e.g., multipolar current focusing) stimulation, etc.. <FIG> are a series of schematic diagrams illustrating exemplary electrode currents and stimulation patterns for five (<NUM>) different channel configurations. It is to be appreciated that the stimulation patterns shown in <FIG> are generally illustrative and that, in practice, the stimulation current may spread differently in different recipients.

Each of the <FIG> illustrates a plurality of electrodes shown as electrodes <NUM>(<NUM>)-<NUM>(<NUM>), which are spaced along the recipient's cochlea frequency axis (i.e., along the basilar membrane). <FIG> also include solid lines of varying lengths that extend from various electrodes to generally illustrate the intra-cochlear stimulation current <NUM>(A)-<NUM>(E) delivered in accordance with a particular channel configuration. However, it is to be appreciated that stimulation is delivered to a recipient using charge-balanced waveforms, such as biphasic current pulses and that the length of the solid lines extending from the electrodes in each of <FIG> illustrates the relative "weights" that are applied to both phases of the charge-balanced waveform at the corresponding electrode in accordance with different channel configurations. As described further below, the different stimulation currents <NUM>(A)-<NUM>(E) (i.e., different channel weightings) results in different stimulation patterns <NUM>(A)-<NUM>(E), respectively, of voltage and neural excitation along the frequency axis of the cochlea.

Referring first to <FIG>, shown is the use of a monopolar channel configuration where all of the intra-cochlear stimulation current <NUM>(C) is delivered with the same polarity via a single electrode <NUM>(<NUM>). In this embodiment, the stimulation current <NUM>(C) is sunk by an extra-cochlear return contact which, for ease of illustration, has been omitted from <FIG>. The intra-cochlear stimulation current <NUM>(C) generates a stimulation pattern <NUM>(C) which, as shown, spreads across neighboring electrodes <NUM>(<NUM>), <NUM>(<NUM>), <NUM>(<NUM>), and <NUM>(<NUM>). The stimulation pattern <NUM>(C) represents the spatial attributes (spatial resolution) of the monopolar channel configuration.

<FIG> illustrate wide or defocused channel configurations where the stimulation current is split amongst an increasing number of intracochlear electrodes and, accordingly, the width of the stimulation patterns increases and thus provide increasingly lower spatial resolutions. In these embodiments, the stimulation current <NUM>(A) and <NUM>(B) is again sunk by an extra-cochlear return contact which, for ease of illustration, has been omitted from <FIG>.

More specifically, in <FIG> the stimulation current <NUM>(B) is delivered via three electrodes, namely electrodes <NUM>(<NUM>), <NUM>(<NUM>), and <NUM>(<NUM>). The intra-cochlear stimulation current <NUM>(B) generates a stimulation pattern <NUM>(B) which, as shown, spreads across electrodes <NUM>(<NUM>)-<NUM>(<NUM>). In <FIG>, the stimulation current <NUM>(A) is delivered via five electrodes, namely electrodes <NUM>(<NUM>)-<NUM>(<NUM>). The intra-cochlear stimulation current <NUM>(A) generates a stimulation pattern <NUM>(A) which, as shown, spreads across electrodes <NUM>(<NUM>)-<NUM>(<NUM>). In general, the wider the stimulation pattern, the lower the spatial resolution of the stimulation signals.

<FIG> illustrate focused channel configurations where intracochlear compensation currents are added to decrease the spread of current along the frequency axis of the cochlea. The compensation currents are delivered with a polarity that is opposite to that of a primary/ main current. In general the more compensation current at nearby electrodes, the more focused the resulting stimulation pattern (i.e., the lower the width of the stimulus patterns increase and thus increasingly higher spatial resolutions). That is, the spatial resolution is increased by introducing increasing large compensation currents on electrodes surrounding the central electrode with the positive current.

More specifically, in <FIG> positive stimulation current <NUM>(D) is delivered via electrode <NUM>(<NUM>) and stimulation current <NUM>(D) of opposite polarity is delivered via the neighboring electrodes, namely electrodes <NUM>(<NUM>), <NUM>(<NUM>), <NUM>(<NUM>), and <NUM>(<NUM>). The intra-cochlear stimulation current <NUM>(D) generates a stimulation pattern <NUM>(D) which, as shown, only spreads across electrodes <NUM>(<NUM>)-<NUM>(<NUM>). In <FIG>, positive stimulation current <NUM>(E) is delivered via electrode <NUM>(<NUM>), while stimulation current <NUM>(E) of opposite polarity is delivered via the neighboring electrodes, namely electrodes <NUM>(<NUM>), <NUM>(<NUM>), <NUM>(<NUM>), and <NUM>(<NUM>). The intra-cochlear stimulation current <NUM>(E) generates a stimulation pattern <NUM>(E) which, as shown, is generally localized to the spatial area adjacent electrode <NUM>(<NUM>).

The difference in the stimulation patterns <NUM>(D) and <NUM>(E) in <FIG>, respectively, is due to the magnitudes (i.e., weighting) of opposite polarity current delivered via the neighboring electrodes <NUM>(<NUM>), <NUM>(<NUM>), <NUM>(<NUM>), and <NUM>(<NUM>). In particular, <FIG> illustrates a partially focused configuration where the compensation currents do not fully cancel out the main current on the central electrode and the remaining current goes to a far-field extracochlear electrode (not shown). <FIG> is a fully focused configuration where the compensation currents fully cancel out the main current on the central electrode <NUM>(<NUM>) (i.e., no far-field extracochlear electrode is needed).

As noted, <FIG> collectively illustrate techniques for adjusting the spatial resolution (i.e., adjusting the spatial attributes of the electrical stimulation) in accordance with embodiments presented herein. However, also as noted, it is to be appreciated that other methods for altering the stimulus resolution could be used in combination with, or as an alternative to, adjustments to the spatial resolution enabled by different stimulation strategies. For example, another technique for adapting the stimulus resolution includes varying the temporal resolution via pulse rate (i.e., higher pulse rates for higher temporal resolutions and lower pulse rates for lower temporal resolutions). In general, changes to the temporal resolution may be implemented in the post-filter bank processing module <NUM> (e.g., during calculation of the channel envelope signals) and/or in the mapping and encoding module <NUM> (e.g., selection of the pulse rate).

Another technique for adapting the stimulus resolution includes varying the instantaneous spectral bandwidth of the stimulation by changing the number of maxima in the channel selection. For example, the instantaneous bandwidth can be increased by increasing the number of channels selected by the channel selection module <NUM> and decreased by decreasing the number of channels selected by the channel selection module <NUM>.

A still other technique for adapting the stimulus resolution includes varying the frequency resolution. The frequency resolution of the filterbank module <NUM> can be increased by, for example, in an FFT filterbank using a higher-point FFT. The frequency resolution of the filterbank module <NUM> can be decreased by, for example in an FFT filterbank using a lower-point FFT.

Shown below is a table (Table <NUM>) illustrating different types of stimulus attributes and associated resolution adaptions in accordance with embodiments presented herein.

The embodiments presented herein have been primarily described with respect to the use of the sound class as the mechanism for determining a target stimulus resolution. It is to be appreciated that other techniques for determining the target stimulus resolution may be used in accordance with embodiments presented herein. For example, a direct measure of the degree of listening difficulty (listening effort) could be used to determine the target stimulus resolution. The listening effort may be determined, for example, using electroencephalography (EEG) (i.e., an electrophysiological monitoring method to record electrical activity of the brain), pupil dilation, or a physiological measure of stress (e.g., blood pressure, cortisol level, etc.). In these embodiments, if the listening effort is determined to be relatively high (with reference to a baseline), then the resolution can be increased. Conversely, if the listening effort is determined to be relatively low (with reference to a baseline), then the resolution can be decreased.

In a further embodiment, the target stimulus resolution can be determined based on a direct input from the recipient. The input may indicate the recipient's listening effort or directly indicate a desired stimulus resolution. In certain such embodiments, the system could be trained so as to, over time, automatically adjust stimulus resolution based on previously received recipient inputs. For example, the hearing prosthesis may be preconfigured with certain thresholds that cause changes between different stimulus resolutions. Over time, the stimulus resolution adaption module can adjust these thresholds using the recipient inputs.

<FIG> is a schematic block diagram illustrating an arrangement for a sound processing unit, such as sound processing unit <NUM>, in accordance with an embodiment of the present description. As shown, the sound processing unit <NUM> includes one or more processors <NUM> and a memory <NUM>. The memory <NUM> includes sound processor logic <NUM>, sound classification logic <NUM>, battery monitoring logic <NUM>, and stimulus resolution adaption logic <NUM>.

The memory <NUM> may be read only memory (ROM), random access memory (RAM), or another type of physical/tangible memory storage device. Thus, in general, the memory <NUM> may comprise one or more tangible (non-transitory) computer readable storage media (e.g., a memory device) encoded with software comprising computer executable instructions and when the software is executed (by the one or more processors <NUM>) it is operable to perform the operations described herein with reference to the sound processor, sound classification module <NUM>, battery monitoring module <NUM>, and stimulus resolution adaption module <NUM>.

<FIG> illustrates software implementations for the sound processor, the sound classification module <NUM>, and the stimulus resolution adaption module <NUM>. However, it is to be appreciated that one or more operations associated with the sound processor, the sound classification module <NUM>, the battery monitoring module <NUM>, and the stimulus resolution adaption module <NUM> may be partially or fully implemented with digital logic gates in one or more application-specific integrated circuits (ASICs).

Merely for ease of illustration, the sound classification module <NUM> and the stimulus resolution adaption module <NUM> have been shown and described as elements that are separate from the sound processor. It is to be appreciated that the functionality of the sound classification module <NUM> and the stimulus resolution adaption module <NUM> may be incorporated into the sound processor.

<FIG> is a flowchart illustrating a method <NUM> in accordance with embodiments presented herein. Method <NUM> begins at <NUM> where a hearing prosthesis receives input sound signals. At <NUM>, the hearing prosthesis determines a sound class of the input sound signals and, at <NUM>, the hearing prosthesis generates, for delivery to a recipient of the hearing prosthesis, electrical stimulation signals that are representative of the input sound signals. At <NUM>, a resolution of the electrical stimulation signals is set based on the sound class of the input sound signals.

<FIG> is a flowchart illustrating a method <NUM>, which does not form part of the present invention, in accordance with embodiments presented herein. Method <NUM> begins at <NUM> where a hearing prosthesis located in an acoustic environment receives sound signals. At <NUM>, the hearing prosthesis assesses the acoustic environment based on the sound signals. At <NUM>, the hearing prosthesis generates electrical stimulation representative of the sound signals at a stimulus resolution that is set based on the assessment of the acoustic environment. At <NUM>, the electrical stimulation signals are delivered to a recipient of the hearing prosthesis.

As described in detail above, presented herein are techniques that analyze the acoustic scene/environment of a hearing prosthesis and, accordingly, adjust, adapt, or otherwise set the resolution of electrical stimulation based on the acoustic environment (e.g., based on an estimated listening difficulty that the acoustic environment presents to a recipient of the hearing prosthesis). The techniques presented herein leverage the idea that there are many listening situations that are not difficult for recipients and, in such situations, power should not be wasted to transmit the most accurate neural representation possible. Likewise, in more challenging listening situations that are more taxing for recipients, it may be beneficial to use more power in order to create a more accurate neural activation pattern that lessens the listening burden on the recipient. Accordingly, the techniques presented optimize power consumption and hearing performance based on the listening situation.

It is to be appreciated that the above described embodiments are not mutually exclusive and that the various embodiments can be combined in various manners and arrangements.

Claim 1:
An auditory prosthesis, comprising:
one or more input elements (<NUM>, <NUM>) configured to receive sound and/or acoustic signals; and
a sound processing path (<NUM>) configured to convert the sound and/or acoustic signals into one or more output signals for use in delivering electrical stimulation to a recipient via one or more stimulation channels;
the auditory prosthesis further comprising:
a stimulus resolution adaption module (<NUM>) configured to select, based on a environment of the sound and/or acoustic signals, an electrode configuration for use in delivering the electrical stimulation to the recipient via the one or more stimulation channels, wherein the electrode configuration controls, for a given stimulation channel, a degree of current spread associated with the electrical stimulation signals in response to stimulation signals delivered via the given stimulation channel.