In cases where individuals have experienced sensorineural deafness, the restoration of hearing sensations to such individuals has been achieved through the use of hearing aids and cochlear implants. Cochlear implants in particular have been in clinical use for many years. A wide variety of different speech processing strategies have been employed in order to process a sound signal into a basis for electrical stimulation via implanted electrode arrays. Some systems have focused upon extracting particular acoustic components of the detected sound signal, which are important to the user's understanding of speech, for example the amplitudes and frequencies of formants, and using these as a basis for generating stimuli. Other approaches have also attempted to utilise the generally tonotopic arrangement of the cochlea, so that each electrode corresponds generally to a particular frequency band.
One such approach, commercially used in speech processors sold by Cochlear Limited, is known as SPEAK. In the SPEAK system, the incoming sound signal is processed to provide an indication of the amplitude of the ambient sound signal in each of a predetermined set of frequency channels, and the channels with the largest amplitudes are selected as the basis for stimulation. In other approaches, the outputs of all channels are used to specify the stimulation patterns, rather than just the channels having the highest short-term amplitudes. The channels are defined by the partially overlapping frequency responses of a bank of band-pass filters. The filters may be implemented using a variety of analog or digital techniques, including the Fast Fourier Transform (FFT). The electrodes corresponding to those channels, determined by a clinical mapping procedure, are selected for activation in each stimulation period and are allocated to the channels according to the tonotopic organization of the cochlear. The rate of stimulation is preferably as high as possible subject to limitations imposed by the processing and power capacity of the external processor and implanted receiver/stimulator unit.
The range of electrical stimulus levels is usually determined by psychophysical measurement of threshold and comfortably loud levels on individual electrodes, using fixed-current pulse trains at the same rate as the stimulus cycle rate of the speech processor output. This may be described as per electrode loudness mapping. The problem with this method of loudness-mapping is that is does not take into consideration the effects of loudness summation when multiple electrodes are activated in quick succession, as they generally are in the output of speech processors.
Although most processing strategies activate a nominal fixed number of electrodes per stimulus cycle, it is important to realise that the actual number of electrodes stimulated in individual cycles is a variable subset of this number, depending on the level and bandwidth of the acoustic stimulus at each point in time. To illustrate this point, a low-level acoustic pure tone will lead to activation of a single electrode, and the electrical level on this electrode must be at least equal to the psychophysical threshold measured individually for that electrode to be audible. In contrast, a low level broad-band noise may activate (for example) eight electrodes in a stimulus cycle. If each of these eight electrodes are activated close to their individual psychophysical thresholds, as may occur with existing systems, then the resultant loudness will not be close to threshold loudness as intended, but will be closer to the maximum comfortable loudness.
This loudness summation leads to the situation that the output of the processor is too loud, even though the individual levels on each electrode do not exceed a comfortable loudness. Various practical methods have been employed to attempt to overcome this problem, including a global reduction of the upper level limit on each electrode, or the use of complex input signals to set the range of individual levels across electrodes. These methods, although alleviating the discomfort of implant users for loud sounds, do not address a second important issue, and that is the impact of loudness summation on speech perception.
Amplitude envelope fluctuations of a speech signal provide vital cues for speech perception, especially for those people who are less able to make use of spectral cues in the signal (for example, those with few active electrodes or poor electrode discrimination ability). Therefore it is important that the changes in acoustic intensity from moment to moment in a speech signal are accurately conveyed as the appropriate perceptual loudness changes to the implantee. The present loudness coding methods, whereby the acoustic output of a filter is mapped to a fixed range of electrical levels (however determined) on its corresponding electrode, lead inevitably to a perceptual distortion of the amplitude envelope shape because these methods do not take into account the variations from moment to moment of important aspects such as the number of electrodes activated in each stimulus cycle, and the relative loudness contributions from these other electrodes. In summary, the relative loudness of electrically stimulated hearing using present approaches does not accurately convey the relative loudness that a normally-hearing person would hear for the same acoustic input. As well as distorting the perception of the amplitude envelope of the acoustic signal, this effect will lead to narrow-band signals being masked by lower-level broad-band noise, thus disrupting the ability of implantees to understand speech in background noise.
Whilst the SPEAK approach has proven successful clinically, it is an object of the present invention to improve sound processing strategies so as to enhance intelligibility of speech and other sounds, for users of cochlear implants. It is a further object of the present invention to improve the perception of loudness provided to users of cochlear implants.