Abstract:
An improved processing approach is disclosed in order to allow for different rates of stimulation to be used for different electrodes in a multi-electrode cochlear implant. When the incoming signal is processed by filter array ( 35 ), each channel is processed to determine amplitude ( 37 ) and to estimate the period of the signal in that channel ( 39 ). The amplitude and period information is used to determine which electrode is stimulated, and the timing of that stimulation.

Description:
TECHNICAL FIELD 
   The present invention relates to cochlear implant prostheses and in particular to an apparatus and method for applying stimulation to the neural structures of a cochlea in order to improve a subject&#39;s pitch and speech perception. 
   BACKGROUND ART 
   Cochlear implant systems are used to aid patients having a hearing deficiency. More particularly, these systems include a microphone receiving ambient sounds and converting the sounds into corresponding electrical signals, signal processing means for processing the electrical signals and generating cochlear stimulating signals and an electrode assembly for applying the cochlea stimulating signals to the cochlea of an implantee. In response to these electrical stimulations a perception of corresponding ambient sound is elicited in the implantee. 
   The inner ear of a normally hearing person includes hair cells which convert the displacement of the ear&#39;s basilar membrane in response to sound into nervous impulses. Different parts of the basilar membrane of the normal cochlea are displaced maximally by different frequencies of sound so that low frequency sounds maximally displace apical portions whereas higher frequency sounds cause displacement of more basal portions of the membrane. The nervous system is arranged so that a nervous impulse originating from a hair cell located adjacent an apical area of the membrane is perceived as a low frequency sound whereas a nervous impulse originating from a hair cell located adjacent a more basal position of the membrane is perceived as a higher frequency sound. This mapping of position to pitch is well known in the art as the tonotopic arrangement of the cochlea. 
   In a dysfunctional ear the hair cells may be damaged or absent so that no nervous impulses are generated. In such cases electrical stimulation impulses must be provided artificially to simulate the nervous activity of the hair cells in order to create a perception of sound. 
   With reference to  FIG. 1 , a typical cochlear implant is shown, which consists of an external component including a speech processor  1 , and an internal component including an implanted receiver and stimulator unit  6  and an intracochlear array  10 . The external component further includes a microphone  2  which is shown integral with the speech processor  1 . In this illustration the speech processor is arranged so that it can fit behind the ear with the microphone integral therewith. Alternative versions are also envisaged whereby the speech processor is worn on the body and separately attached to the microphone, and also where the speech processor and microphone are implanted in the patient. The present invention is applicable to all these alternatives. 
   In such cochlear implant devices, ambient sounds are detected by a microphone and a transduced signal is thereby generated, representative of the ambient sound. A processor unit then processes this transduced signal according to one of many strategies, (some of which will be explained further below) and based on this processing stimulation currents are applied between the electrodes of a coupled array. For example, in “monopolar” mode stimulation, stimulation currents may be caused to flow between an electrode of the electrode array  10  and an extracochlear electrode  115 . Nervous discharges elicited in the basilar membrane  8  are conveyed to the central nervous system of the wearer by auditory nerve  9 . 
   In the event that the stimulation current flows between an apical electrode such as electrode  111  and extracochlear electrode  115  then a lower pitched hearing sensation will be perceived by a wearer of the prosthesis than will be the case if stimulation current flows between basal electrode  107  and extracochlear electrode  115  because of the previously mentioned tonotopic arrangement of the cochlea. Further pitch information may be transmitted to the wearer corresponding to the rate at which stimulations are delivered. 
   Many possibilities exist as to the manner in which the signal processing means operates upon the electrical signals in order to produce stimulation signals. However it has been noted in the past that simultaneous stimulation of electrodes is not generally conducive to eliciting a perception of sound that is faithful to the actual acoustic signals being processed. This is because if electrodes are stimulated simultaneously, current paths between electrodes can interact, causing undesirable stimulation. Consequently most cochlear prosthesis stimulation strategies stimulate by means of only one electrode at a time. 
   In the past designers of cochlear implant stimulation strategies have striven to optimise the intelligibility of spoken words as perceived by the wearer of a cochlear implant. 
   One of the earliest sound processing and cochlear stimulation strategies is described in U.S. Pat. No. 4,532,930 to the present applicant. In that patent there is taught the use of a filter (F0) dedicated to extracting the voice pitch of a speech signal. The periodicity of the voice pitch is used to set the stimulation periodicity for two or three electrodes. A second, and possibly third, channel is analysed to determine periodicity and amplitude in a selected frequency band. 
   The periodicity extracted from the second filter, and possibly third filter, is used to select which electrode is stimulated for the second and third channels while the periodicity of stimulation of the channel is determined by the periodicity of the output signal from the F0 filter. 
   Another stimulation arrangement is described in U.S. Pat. No. 4,207,441. In that system there are n electrodes each coupled to one of n filters. Each electrode is stimulated once per analysis period, with an intensity corresponding to the amplitude of the corresponding filter channel. The analysis period of this system is predetermined and hence is not related to the signal on the filter outputs. 
   More recently in EP 0 745 363 there is described a stimulation system which takes into account the temporal behaviour of the cochlea. In an embodiment of the invention therein described a wavelet transformation is used to extract the temporal information with a view to using this information to determine the sequence of stimulation of the electrodes. The purpose of the invention is to improve the temporal resolution in response to a rapidly changing audio spectrum. 
   A problem that has been faced by users of cochlear implants featuring prior art stimulation schemes is that while intelligibility of spoken words is often good the user&#39;s pitch perception, and in particular perception of music, is poor. Accordingly, it is an object of the present invention to provide an apparatus and method for use in a multi-channel cochlear implant which will improve a user&#39;s perception of pitch. 
   SUMMARY OF THE INVENTION 
   Broadly, the present invention seeks to use information about the periodicity of the signal in each filter channel as a factor in determining the rate of stimulation applied to a tonotopically placed electrode which corresponds to the relevant channel. This allows for an improved perception of pitch by the implant user. As a result, the rate of stimulation which occurs in practice for each electrode will be related to the periodicity of the signal in the filter channel corresponding to that electrode. 
   According to one aspect of the present invention there is provided a cochlear implant prosthesis of the type having a transducer for converting an acoustic signal to an electrical signal, a plurality of bandpass filtering means responsive to said electrical signal and operatively producing a plurality of bandpass filtered signals, signal processing means responsive to said plurality of bandpass filtered signals and operatively generating stimulation commands, electrode driving means responsive to said stimulation commands and an electrode array coupled to said electrode driving means for operatively delivering to a user of said cochlear implant prosthesis stimulations in accordance with said stimulation commands, said signal processing means including: 
   a) period estimation means, responsive to said filtered signals and operatively generating periodicity signals indicative of the periodicity of each of at least a number of said plurality of filtered signals; 
   b) amplitude estimation means responsive to said filtered signals and operatively generating magnitude signals indicative of the magnitude of each of said plurality of filtered signals; 
   c) selection means responsive to said magnitude signals arranged to select only one filtered signal of said plurality of filtered signals in each stimulation period, said selection means generating said stimulation commands including a command to stimulate by means of an electrode operatively tonotopically best corresponding to said one filtered signal, said command to stimulate further specifying a time for stimulation to occur dependent on a corresponding one of said periodicity signals. 
   According to another aspect, the present invention provides a processing device for a cochlear implant prosthesis, said prosthesis being of the type including electrode driving means responsive to stimulation commands and an electrode array coupled to said electrode driving means for operatively delivering to a user of said cochlear implant prosthesis stimulations in accordance with said stimulation commands, 
   said processing device being responsive to a transducer for converting an acoustic signal to an electrical signal and including a plurality of bandpass filtering means responsive to said electrical signal and operatively producing a plurality of bandpass filtered signals, signal processing means responsive to said plurality of bandpass filtered signals and operatively generating stimulation commands, said signal processing means including: 
   a) period estimation means, responsive to said filtered signals and operatively generating periodicity signals indicative of the periodicity of each of at least a number of said plurality of filtered signals; 
   b) amplitude estimation means responsive to said bandpass filters operatively generating magnitude signals indicative of the magnitude of each of said plurality of filtered signals; 
   c) selection means responsive to said magnitude signals arranged to select only one filtered signal of said plurality of filtered signals in each stimulation period, said selection means generating said stimulation commands including a command to stimulate by means of an electrode operatively tonotopically best corresponding to said filtered signal, said command to stimulate further specifying a time for stimulation to occur dependent on a corresponding one of said periodicity signals. 
   According to a further aspect of the present invention there is provided a method of operating a cochlear implant prosthesis of the type including a plurality of bandpass filters each having a characteristic centre frequency, said filters generating a corresponding plurality of filtered signals, said prosthesis further including stimulation delivery means coupled to an electrode array, said method including the steps of: 
   a) in each of a number of time intervals, determining the amplitude for each of said plurality of filtered signals and a periodicity value for at least some of said plurality of filtered signals; 
   b) selecting only one of said signals as a basis for stimulation in each stimulation period; and 
   c) applying a stimulation current by means of an electrode of said electrode array tonotopically closest to the centre frequency of the bandpass filter producing the signal determined in step b), said stimulation current being applied during a time interval determined from the periodicity value of the signal determined in step b). 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
       FIG. 1  depicts a typical cochlear implant device. 
       FIG. 1A  depicts a block diagram of the functional elements of a cochlear implant according to the present invention. 
       FIG. 2  depicts a dedicated hardware version of a cochlear implant prosthesis according to the present invention. 
       FIG. 3A  is a graph of a possible output of an amplitude estimator module of  FIG. 2 . 
       FIG. 3B  is a graph of a possible output of a period estimator module of  FIG. 2  of the same channel as the amplitude estimator of  FIG. 3A . 
       FIG. 3C  is a graph of stimulation currents generated via an electrode in accordance with the amplitude and period estimates of  FIGS. 3A and 3B . 
       FIG. 3D  is a graph of stimulation currents generated via a further electrode having a stimulation current occurring simultaneously with a stimulation current in the graph of  FIG. 3C . 
       FIG. 3E  is a graph of a single stimulation current. 
       FIG. 4  is a flowchart of procedural steps used in implementing the present invention by means of software on an apparatus of the type depicted in  FIG. 1A . 
       FIG. 5  is a flowchart of procedural sub-steps involved in one of the boxes appearing in  FIG. 3 . 
   

   DETAILED DESCRIPTION 
   The present invention will be described with reference to a specific implementation. However, it will be appreciated that the present invention can be implemented in various ways, with suitable modifications to suit the cochlear implant system in question. 
   With reference to  FIG. 1A , there is depicted a simplified digital hardware implementation of a cochlear prosthesis according to the present invention. Sound waves are transduced by microphone  11  and the electrical signal so produced is processed by a signal conditioning module  13 . Signal conditioning module  13  includes standard circuits for pre-amplifying and low pass filtering the signal prior to its processing by analog to digital converter  15 . Analog to digital converter  15  produces a 16 bit digital signal which is conveyed to microprocessor  17 . Microprocessor  17  operates according to a program stored in EPROM  19 . Microprocessor  17 , in accordance with its program operates upon the digital signal in order to generate a sequence of stimulation commands which are delivered to a switchable current source module  23 . The commands delivered to the switchable current source module  23  specify the amplitude of the current that is to flow from one or more electrode to one or more other electrodes, the timing of the stimulation current, and the mode of the stimulation. 
   The term ‘mode’ here refers to the selection of electrodes between which a stimulation current is to flow. Well known stimulation modes include bipolar, monopolar and common ground. Upon receiving commands specifying the parameters of the stimulation to be applied, switchable current source module  23  connects various electrodes of electrode array  21  to an internal controllable current source in order to generate the appropriate stimulation. The construction of a switchable current source is well known in the art and may be found in the applicant&#39;s U.S. Pat. No. 4,532,930. 
     FIG. 2  depicts a dedicated hardware implementation of the invention for purposes of explanation. While  FIG. 2  illustrates the invention as if individual tasks performed by microprocessor  17  were embodied in dedicated hardware, it remains the case that the invention is most readily practised according to the arrangement of  FIG. 1A . The invention will however be explained with reference to  FIG. 2  in order to most clearly impart an understanding of its operation to the reader. 
   Referring to  FIG. 2  it will be noted that the analog signal from pre-conditioning module  13  is first sampled at 16 kHz by sampler  31  thereby producing a sampled signal. The sampled signal is split 22 ways, each of the split signals providing an input to digital filters  35 A– 35 V which filter the signal into quarter octave frequency bands. It will be appreciated that the numbers of ways the signal is split, and the sampling rate, are matters of design choice appropriate to the system in which the present invention is implemented. 
   Digital filters  35 A– 35 V are bandpass and logarithmically spaced with the base frequency being typically at 150 Hz. Each filter is of a 6th order Chebychev Type I bandpass type implemented in three second order sections. The quarter octave frequency bands are as shown below: 
   
     
       
             
             
             
           
             
             
             
           
         
             
                 
             
             
                 
               Lower Frequency Boundary 
               Upper Frequency Boundary 
             
             
               Filter Band 
               (Hz) 
               (Hz) 
             
             
                 
             
           
           
             
                 
             
           
        
         
             
               A 
               150.00 
               178.38 
             
             
               B 
               178.38 
               212.13 
             
             
               C 
               212.13 
               252.27 
             
             
               D 
               252.27 
               300.00 
             
             
               E 
               300.00 
               356.76 
             
             
               F 
               356.76 
               424.26 
             
             
               G 
               424.26 
               504.54 
             
             
               H 
               504.54 
               600.00 
             
             
               I 
               600.00 
               713.52 
             
             
               J 
               713.52 
               848.53 
             
             
               K 
               848.53 
               1009.10 
             
             
               L 
               1009.10 
               1200.00 
             
             
               M 
               1200.00 
               1427.00 
             
             
               N 
               1427.00 
               1697.10 
             
             
               O 
               1697.10 
               2018.20 
             
             
               P 
               2018.20 
               2400.00 
             
             
               Q 
               2400.00 
               2854.10 
             
             
               R 
               2854.10 
               3394.10 
             
             
               S 
               3394.10 
               4036.30 
             
             
               T 
               4036.30 
               4800.00 
             
             
               U 
               4800.00 
               5708.20 
             
             
               V 
               5708.20 
               6788.20 
             
             
                 
             
           
        
       
     
   
   The bandpass filtered signal from each of the digital filters, for example  35 A, is connected to an amplitude detection module  37 A and a period estimation module  39 A. The output AMP[A] of amplitude detection module  37 A is a digital signal representing an estimation of the amplitude of the sampled signal from filter  35 A. The construction of module  37 A is straightforward, well understood by those skilled in the art, and will not be explained in detail other than to say that it could be based on a series of comparators. 
   Period estimation module  39 A counts sampling periods between positive zero crossings of the signal from filter  35 A. The output signal PERIOD[A] is scaled so that it is expressed in units of “timeslices”. 
   One “timeslice” is the time taken to deliver one stimulation pulse by means of an electrode. With reference to  FIG. 3E  an example of a stimulation current pulse waveform comprises a first ‘phase’  103  being a square wave of predetermined amplitude, an interphase gap  105  and a second phase  107  being a current square wave of the same magnitude and duration as the first phase but flowing in the opposite direction between an intra-cochlear electrode and (in mono-polar mode) an extra-cochlear electrode. Time periods  109  and  111  are present in which the system generating the stimulations has time to perform any necessary operations in order to configure for the next stimulation. The overall time taken to set-up, deliver and recover from application of a stimulation is one timeslice, in the present example a timeslice is of approximately 69 microseconds duration. 
   In the present implementation the preferred maximum stimulation rate is 8 kHz whereas the preferred sampling rate is 16 kHz. Accordingly PERIOD[A] is the number of samples occurring between positive-going zero crossings divided by two and rounded. The PERIOD[A] signal is updated at the end of each period. 
   The AMP[A], . . . , AMP[V] and PERIOD[A], . . . , PERIOD[V] signals contain magnitude and period information concerning the ambient sound picked up by microphone  11  for each of the frequency bands monitored by bandpass filters  35 . It is possible to simply stimulate via each corresponding electrode e[i] with a current intensity corresponding to AMP[i] at a time PERIOD[i] into the future in order to convey the information generated by amplitude detectors  37  and period estimators  39  to a wearer of the cochlear prosthesis. For example, with reference to  FIGS. 3A ,  3 B and  3 C a stimulation sequence via electrode e[A] is shown corresponding to amplitude and period values generated by amplitude detection module  37 A and period estimation module  39 A as shown plotted in  FIGS. 3A and 3B . Period[A] is equal to P 1  at time t=0 and Amp[A] is equal to a 1 . Accordingly at a time t=P 1  a stimulation current is delivered via electrode e[A] of electrical amplitude l(a 1 ) where l( ) is a loudness growth function which maps amplitude into the dynamic range of the wearer of the prosthesis. 
   The period P 2  and amplitude a 2  values are then obtained and a further stimulation is delivered at time t=P 1 +P 2  of amplitude l(a 2 ). This process is repeated continuously to produce the stimulation sequence of biphasic current pulses shown in the graph of  FIG. 3C . As previously mentioned, such a process could be simultaneously performed independently on all channels of the implant, (a “channel” as used here refers to a stimulation electrode, its corresponding filter and period and amplitude estimation modules). 
   In the system thus far described the period estimation module  39 A produces a period estimate which is simply the time delay between the last two positive-going zero crossings. While this system works well, any noise on the individual period estimates will degrade the performance of the system. To prevent this, it is desirable to calculate a smoothed period estimate. 
   The individual period estimates constitute a number sequence which is amenable to any of the methods of smoothing known to the art of digital signal processing. The smoothing may, for example, be implemented as a simple FIR or recursive digital filter—preferably a low-pass filter. Alternatively a rank-order filter, such as a median filter may be used. A rank-order filter has the advantage that it will completely remove any single data errors from the number sequence. A smoothed period estimate is thus produced by applying the sequence of period estimates to a digital filter, and taking the output of that filter. The smoothed period estimate is then utilised by taking the most recent output from the filter and using it in place of the (unsmoothed) period estimate. 
   With reference to  FIG. 3D  there is shown a stimulation sequence via electrode e[B]. It will be noted that stimulation pulse  102  occurs at exactly the same time as stimulation pulse l(a 2 ) of  FIG. 3C . Such coincidences occur more and more frequently depending on the number of channels used. 
   There are at least two problems associated with the above stimulation strategy by which stimulations may be delivered by two or more electrodes simultaneously. Firstly, as discussed above, it is well known that simultaneous, or very near simultaneous, stimulation produces a deterioration in the quality of the sound perceived by the user, due to the interaction of simultaneous current paths between the electrodes. 
   Consequently the application of stimulations on a number of channels simultaneously is undesirable. A further problem is that simultaneous stimulation requires very substantial processing power and so is not possible in the majority of cochlear implants presently available. 
   In light of the above problems the inventors have incorporated a preferred refinement for determining which information signals, coming from amplitude detectors  37  and period estimation modules  39  are most appropriate for acting upon in order to produce a high quality percept in a user. According to the invention, for any stimulation period i.e. “timeslice” t 0  the signals AMP[A,t 0 ], . . . , AMP[V,t 0 ] are ordered according to magnitude and a stimulation current is generated by means of the electrode which corresponds to the signal AMP[A,t 0 ], . . . , AMP[V,t 0 ] having the greatest magnitude. (It should be noted that, when implanted, electrodes e[A], . . . , e[V] are tonotopically mapped to filters fA, . . . , fV so that electrode e[A] is most apically placed whereas electrode e[v] is most basally placed.) For example, if it is found that AMP[F,t 0 ] has the greatest magnitude at a given timeslice then electrode e[F] is used to deliver the monopolar stimulation in the next timeslice t 1 . 
   A further variation to this scheme is that a number of the next largest magnitude signals are also determined in the same timeslot, for example AMP[G,t 0 ]&gt;AMP[B,t 0 ]&gt;AMP[K,t 0 ] might be determined to have the magnitudes next greatest to AMP[F,t 0 ]. In that case those values are assigned to AMP[G,t 1 ], AMP[B,t 1 ] and AMP[K,t 1 ] respectively. During the next timeslice, t 1 , the procedure is repeated and it may be that AMP[G,t 1 ] is selected as having the greatest magnitude so that electrode e[G] is selected for delivering a stimulation pulse of amplitude corresponding to AMP[G,t 1 ]=AMP[G,t 0 ]. By using this scheme it is possible that signals having a large magnitude, though not the greatest, are presented to the user after a short time delay. The inventors have found that most acoustic power is centred around the lower frequency bands which have longer periods associated with them whereas the higher frequency bands generally have less power associated with them as well as being of shorter period. Accordingly, it is predominantly higher frequency sounds which are delayed rather than lower frequency sounds. 
   A further refinement is that rather than calculate period estimates in respect of the outputs from filters centred at higher frequencies, for example for filters Fl, . . . , FV period estimators  391 , . . . ,  39 V simply generate a constant signal, or periodicity value, indicating a period of 1.25 ms i.e. a periodicity value towards the highest stimulation rate that the device is capable of supporting. 
   While the preceding description covers a system utilising period estimators on some or all of the bandpass filtered signals, it is possible to implement the system more simply. A stimulus could be requested each time a positive zero-crossing is detected on a filter output. Once per timeslice each channel is interrogated to see if it has a stimulation request pending. If there are no requests pending, then no action is required. If there is exactly one request pending, then a stimulus is generated corresponding to that request. 
   If more than one request is pending, then the following actions are taken. The requests are sorted according to the amplitudes of the corresponding bandpass filtered signals. A stimulus is generated corresponding to the bandpass filtered signal with the largest amplitude. The next N largest (preferably 5 largest) amplitude requests are delayed by one timeslice. Any remaining requests are cancelled. 
   This system is simpler to implement than that previously described. It has two main disadvantages, however. The previously described system utilising period estimates acts to limit the stimulation rate on higher frequency channels. This is directly beneficial in that excessive stimulation with little information content is avoided. More importantly, in the case of relatively large amplitude high frequency signals, the lower frequency signals will be completely lost in the simpler system. The rate-limiting effect of the period-estimation system will mean that the low-frequency signal will always get through. 
   The request generators  41 -A, . . . ,  41 -V and request arbitrator  43  operate to determine which electrode will be stimulated from the AMP[A], . . . , AMP[V] and PERIOD[A], . . . , PERIOD[V] signals. The operation of the request generators and the request arbitrator, in order to implement the aforedescribed scheme, will now be explained with exemplary reference to request generator  41 A. 
   The AMP[A] and PERIOD[A] signals are inputs to request generator module  41 A. Another input to the request generator is the CLK signal which corresponds to the present timeslice. The CLK signal is modulus  256  to avoid overflow problems. The last input to request generator  41 A is a command signal ARB_CMD[A] from request arbitrator  43 . 
   The outputs from request generator  41 A are TSLICE[A] and REQ_AMP[A]. 
   The TSLICE[A] represents the time at which it is proposed by generator  41 A that a stimulation be delivered having an amplitude value represented by REQ_AMP[A]. 
   The relationship between TSLICE[A] and PERIOD[A] and REQ_AMP[A] and AMP[A] is determined by the value of the ARB_CMD[A] signal. The ARB_CMD[A] signal can take one of three values REQUEST, DELAY, DISCARD. When ARB_CMD[A] takes the value: 
   REQUEST
         then REQ_AMP[A]:=AMP[A]; TSLICE[A]:=CLK+PERIOD[A]       

   DELAY
         then TSLICE[A]:=TSLICE[A]+1       

   DISCARD
         take no action.       

   The principle behind these rules is that in the event that request arbitrator  43 , whose operation will shortly be described, determines that a stimulation pulse should be applied corresponding to the output from filter  35 A then by sending an ARB_CMD[A] signal having the value REQUEST to request generator  41 A the amplitude and timing of the stimulation pulse is made available at the next timeslice. Alternatively, if ARB_CMD[A] takes the value DELAY then the corresponding TSLICE[A] variable is incremented. Construction of the request generator, in order to implement the above rules is readily accomplished according to established synchronous digital design techniques. 
   Request arbitrator module  43  takes as its input the signals TSLICE[A], . . . , TSLICE[V] from each of the request generators  41 A- 41 V, REQ_AMP[A], . . . , REQ_AMP[V] and the CLK signal. Arbitrator module  43  generates a signal P_CHAN which identifies which of electrodes e[A], . . . , e[V] of the electrode array is to be used to apply a stimulation. 
   The arbitrator module also generates a signal P_AMP which takes a value REQ_AMP[A], . . . , REQ AMP[V] which is used, after scaling as will subsequently be described, to determine the amplitude of the signal to be used when applying stimulation. Request arbitrator module  43  operates according to the following rules: 
   1. Find all TSLICE[i] with a value equal to CLK. 
   2. Find N channels of those determined in Step 1 which have the largest value of REQ_AMP[j]. The channel with the largest value of REQ_AMP[j] and TSLICE[j] as determined in step 1 is found and P_CHAN set to j and P_AMP set to REQ_AMP[j]. So that a stimulation is directed via electrode e[j] with amplitude scaled from the value P_AMP=REQ_AMP[j]. This is accomplished by setting the ARB_CMD[j] signal to REQUEST. 
   3. The channels having the next N−1 largest amplitude values REQ_AMP[ ] are delayed by one timeslice for consideration during set up for the next stimulation. This is achieved by sending an ARB_CMD[ ] signal to the corresponding N−1 request generators to DELAY. 
   4. The remaining channels, which were selected in step 1 but not in step 2 are discarded. This is achieved by sending the corresponding request generators an ARB_CMD[ ] signal of value DISCARD. 
   5. If any of the request generators is sending a specific “no pulse request” then the corresponding ARB_CMD[ ] signals are set to REQUEST. 
   Once the P_CHAN and P_AMP values have been determined they are passed to Loudness growth function module  47 . The growth function module takes into account the predetermined comfort and threshold levels of the user of the cochlear prosthesis in order to map the P_AMP values into the listeners dynamic range. Such mapping is known in the prior art and the reader is referred to U.S. Pat. No. 4,532,930 to the same applicant for further details. 
   The invention is most conveniently practised, in accordance with the embodiment of  FIG. 1 , by programming a SPrint speech processor, available from Cochlear Limited of 14 Mars Road, Lane Cove, NSW 2066, Australia, in order to perform the operations described in relation to  FIG. 2 . The SPrint speech processor is used in conjunction with a CI24M cochlear implant receiver-stimulator from the same vendor. 
   With reference to  FIG. 4  there is depicted a block diagram of the overall operational procedure for implementing the present invention in software. At block  201 , for each sample period, the sound signal is filtered into the required number of channels. At block  202 , the signal in each channel is analysed to determine its amplitude, and the period of the signal. The latter may be performed by determining the time between successive zero crossings, as described above. Based upon the values for the amplitude and period for each channel, block  203  selects which channel signal is to be used as the basis for stimulation, and hence the electrode to be stimulated. Loudness mapping block  204  performs the function of mapping the desired amplitude stimulus within the dynamic range for the selected electrode. The latter step is well known to those skilled in the art. 
     FIG. 5  further details the operational steps performed in box  203  of  FIG. 4 . At initialisation, in block  210 , the stimulus selector updates all inputs. The inputs are the values for amplitude and period as described previously. At block  211 , each input channel is checked to determine if a stimulation is being requested for the next period, based upon the period value, and all such channels are sorted according to amplitude. At block  212 , all channels but the largest amplitude channel are delayed to a later stimulation period. In block  214 , the largest amplitude channel is selected as the basis for stimulation, and the inputs for that channel are updated to reflect that a stimulus will be delivered that period. Block  213  completes the process by discarding the remaining channels by updating their inputs, and returning to the process of block  210  for the next period, time j. 
   While the invention has been described with reference to preferred embodiments, it is to be understood that these are merely illustrative of the application of principles of the invention. Accordingly, the embodiments described in particular should be considered exemplary, not limiting, with respect to the following claims.