referring to fig1 the eeg measurement system of the present invention is shown attached to a patient 10 . patient 10 is connected through at least two electrodes 20 a , 20 b to an impedance measurement circuit 50 through lead wires 40 . the output 55 of the impedance measurement circuit 50 containing the impedance measurement signal is sent to a digital signal processor 80 . the impedance measurement circuit 50 excites the electrodes 20 ( a ), 20 ( b ) by injecting a small current i 30 into one of the electrodes 20 a . the current i 30 passes through the patient 10 and out the other electrode 20 b and back into the impedance measurement circuit 50 . this current must be limited to less than 10 microamperes to maintain patient safety . in a preferred embodiment , the signal is approximately 1 nanoampere . at this current level the resulting voltage , according to ohm &# 39 ; s law , v = ir , is approximately 1 microvolt per kiloohm . the use of this very low excitation current facilitates keeping the resulting voltage very low compared to the biopotential signal being measured . in the case of eeg , such voltages range from hundreds of microvolts down to less than 1 microvolt . in the preferred embodiment , the excitation current is at a frequency just above the eeg band of interest , or 128 hz . referring now to fig2 the impedances of the two electrodes 20 a , 20 b are referred to herein as zp and zm . equal and opposite current sources 110 , 115 inject current into the electrode leads 20 a , 20 b such that the current flows out of the upper current source 110 , through the patient 10 and back into the lower current source 115 . instrumentation amplifier 60 has very high input impedance ( in a preferred embodiment about 50 megaohms ), and so only a negligible current passes through it . the resulting voltage across the inputs of the instrumentation amplifier 60 is equal to the value of the current times the combined impedances zp , zm of the two electrodes and the patient 10 . this voltage is amplified by the instrumentation amplifier 60 and sent to an analog to digital converter 70 . there , the signal is digitized at a sample rate at least exceeding twice the frequency of the excitation current . in a preferred embodiment , the analog to digital converter 70 runs at much higher sampling rate , requiring the use of an over - sampled analog to digital converter . the resulting digitized signal is forwarded to the digital signal processor 80 for computation . in a preferred embodiment , the computation consists of a fourier transform as described below . [ 0015 ] fig3 illustrates the synchronization of the frequency of the excitation current sources 110 , 115 with the analog to digital converter &# 39 ; s 70 sample rate and the digital signal processor &# 39 ; s 80 main clock 130 . the frequency of clock 130 is reduced by divider 200 , which outputs a pulse train to current sources 110 , 115 , which generate a current at a frequency that is exactly the clock rate divided by 2 14 or one part in 16 , 384 of the clock 130 . in a preferred embodiment , the clock rate is 2 . 097152 megahertz , resulting in a signal current that is 128 hz . the frequency of clock 130 is also reduced by divider 75 , which outputs a pulse train to the analog to digital converter 70 , which digitizes the impedance measurement signal obtained from the instrumentation amplifier 50 . in a preferred embodiment , the digitization occurs at 16 , 384 times a second , or at a rate of 1 / 2 7 or 128 th of the clock rate . the analog to digital converter 70 , therefore , is outputting a digitized version of the impedance test signals at a 16 , 384 sample per second rate to the digital signal processor 80 . the divider 75 also outputs a pulse train to the digital signal processor 80 in the preferred embodiment at 16 , 384 hz , again , 128 th of the clock signal fundamental frequency of 2 . 097152 megahertz . this enables the digital signal processor 80 to operate at a rate that is an even divisor of the system clock 130 , and is exactly synchronous with the frequency of excitation of the current into the electrodes 20 a , 20 b . the advantage of this technique is that the resulting processed signal is made to be exactly synchronous to the processor clock 130 , which enables the processor 80 to execute a form of synchronous detection . the benefits of this approach are that the detection can be done over a very narrow bandwidth ( the width of one bin in the discrete fourier transform dft ), thus filtering out noise over most frequencies and resulting in a more precise measurement . a further benefit is realized because the impedance test stimulus ( the excitation current ) and the detection ( dft ) are always synchronized ; thus , no adjustments are needed to make either circuit match the clock of the other . the impedance signal excitation current sources are depicted in fig4 . clock 130 outputs a pulse train with a frequency of 2 . 0917152 megahertz to divider 200 . divider 200 divides the clock pulse train by 2 14 , or 16 , 384 to provide a 128 hz clock signal . the digital signal from the divider 200 is output to switch 210 , which opens and closes at this 128 hz rate . this connects and disconnects the + 5 volts to resistor r 1 at 128 hz . when the + 5 volts is connected by the switch 210 to resistor r 1 , current flows forward through r 1 and into capacitor c 1 , causing it to build up voltage at the output of operational amplifier 230 . when switch 210 opens , the current through resistor r 1 and capacitor c 1 ceases . the charge on capacitor c 1 is now drained through resistor r 2 into the − 5 volt power supply . the resulting wave shape at the output of operational amplifier 230 is a triangle wave at a repetition rate of 128 hz . resistor r 3 is needed to maintain the bias for operational amplifier 230 . this triangle wave shape is inverted by the operational amplifier circuit comprised of the operational amplifier 220 and resistor network r 5 and r 4 , resulting in a wave shape at the output of operational amplifier 220 that is identical but whose amplitude is opposite in sign to that at the output of operational amplifier 230 . the resulting voltages at the outputs of operational amplifiers 230 and 220 are applied to resistors r 6 and r 7 , respectively . these resistors are of a value much greater than the impedances zp and zm being measured . in the preferred embodiment they are 4 . 7 megaohms , much greater than the impedances zp and zm , which are typically 0 to 100 kiloohms . the high value resistors r 6 and r 7 ensure that the excitation current from current source circuits 110 , 115 are largely independent of the values of impedances zp and zm . the current flowing through resistors r 6 and r 7 also passes through capacitors c 2 and c 3 . these capacitors serve to block any unwanted direct current and also serve to further increase the apparent output impedance of the current sources . the resulting currents flow through electrodes 20 a , 20 b , again as depicted in fig2 resulting in a voltage that is proportional to the combined impedance of zp and zm , appearing at instrumentation amplifier 60 . thus the voltage at the output of instrumentation amplifier 60 is proportional to the electrode impedances zp and zm . referring now to fig5 ( a )- 5 ( c ), a digital signal processing program executed on the digital signal processor 80 executes a fast fourier transform on the digitized signal . in fig5 ( a )- 5 ( c ), the fourier transform is plotted as voltage versus frequency . in fig5 ( a ) the background noise is relatively small , and the impedance signal can be easily seen as a spike arising from the rest of the signal at a frequency f 0 , which in the preferred embodiment is equal to 128 hz as described . in this instance its true value is relatively unaffected by the background noise . fig5 ( b ) depicts an instance where the impedance signal is beginning to be obscured by interference and background noise . in terms of the fast fourier transform of the digitized signal , wide band noise has the effect of raising the voltage across the entire frequency range . this causes the true value to be more effected by the background noise . in fig5 ( c ), the impedance signal is completely obscured by the background noise , causing its true value to be unmeasurable . that is , the fourier transform value at the excitation frequency is no longer primarily a function of the impedance value , but is more a function of the background noise . the digital signal processor 80 measures the electrode impedance by measuring the voltage amplitude of the impedance signal and multiplying it by a scale factor to convert the voltage into impedance in ohms . it can be seen in fig5 ( b ) and 5 ( c ) that in the presence of noise , this strategy may lead to instances where the derived impedance measurement is not solely a function of the impedance but is more a function of the interference and noise in the rest of the signal . in the present invention , the system discerns this situation by measuring the voltage level at a frequency f 1 very close to the impedance frequency f 0 . the digital signal processor then examines the difference between the voltage at the impedance signal , f 0 and the voltage noise signal , f 1 . if the signal f 0 is greater than the signal f 1 , the impedance measurement is said to have a positive signal to noise ratio . if the signal f 1 is greater than the signal f 0 , the impedance measurement is said to have a negative signal to noise ratio . if the signal to noise ratio is negative , or positive but less than a threshold , the measurement is deemed to be contaminated by interference and background noise . the digital signal processor 80 may alternatively use other methods to compute the voltage at frequencies f 0 and f 1 . for example , the digital signal processing may execute a discrete fourier transform ( dft ) or use other methods that are known to those skilled in the art . filters may also be used to measure the amplitude or power of the signal at frequencies f 0 and f 1 . such filters may be implemented in circuitry or as digital filters in a computer or a dedicated digital signal processing integrated circuit . further , a wide range of potential signals at frequencies f 0 and f 1 may be used . referring to fig6 the process implemented by the digital signal processor 80 to measure bioelectric impedance is described . this process may also be executed by any other kind of processor or combination of processors . the digitized signals containing eeg and impedance measurement signals , as well as interference and background noise , are input to the digital signal processor . the impedance measurement is made in step 201 , as described earlier , by measuring the voltage at the excitation frequency f 0 , which is 128 hz in a preferred embodiment . the impedance measurement is comprised of the actual impedance voltage resulting from the excitation current , plus the contribution from noise . the apparent impedance will be the root of the sum of the squares of the actual impedance plus the noise contribution . the noise measurement is made in step 202 by measuring the voltage at a frequency f 1 or at a set of frequencies very close to the impedance measurement frequency . if a set of frequencies is used , an aggregation function is used to combine the voltages at each of the multiple frequencies . such aggregation function may be a mean , median , maximum , minimum or other such mathematical function well known in the art . in addition , the voltages at each of the multiple frequencies may be weighted to increase the contribution of certain frequencies over others . in the preferred embodiment , this noise measurement is the root mean square ( rms ) of the signal voltage in frequency range from 70 to 110 hz . this is beneficial since the rms voltage may be used elsewhere in the eeg system for detecting the presence of electromyogram ( emg ) signals , reducing the amount of computation required of the digital signal processor 80 . the biopotential measurement 203 in a preferred embodiment is the eeg . in a preferred embodiment , the eeg signal is used to calculate the depth of consciousness of a patient undergoing anesthesia using bispectral index monitoring . the measured impedance value 210 is divided by the measured noise value 212 in step 205 to form the impedance signal to noise ratio ( snr ) 209 . the ongoing biopotential signal is a contributor to the noise in the impedance measurement process . the measured impedance value 210 is compared against a threshold limit in step 204 . in a preferred embodiment , the threshold limit is 15 k ohms . if the measured impedance value 210 is less than or equal to the threshold , then it is known that the actual impedance is less than the threshold regardless of the background noise , and the biopotential measurement 203 is enabled in step 207 . if the measured impedance value 210 is greater than the threshold limit , as determined in step 204 , then the actual impedance may or may not be greater then the threshold . in this case the impedance signal to noise ratio 209 is tested against the snr threshold in step 206 to determine if background noise has increased the measured impedance to a value greater than that of the actual impedance value . if the impedance snr is greater than the snr threshold , the impedance measurement is deemed to be uncontaminated by noise . in this case , the biopotential measurement is blocked in step 208 as it is deemed that the impedance is too high to make an accurate biopotential measurement . in the case where the impedance snr is less than or equal to the snr threshold , then the impedance measurement is deemed to be invalid and , and the biopotential measurement continues to be enabled in step 207 . the impedance measurement is most susceptible to contamination by noise when the actual impedance is close to the threshold limit . this is the interval for which the smallest amount of noise may cause an acceptable impedance to appear to be unacceptably high . in the preferred embodiment , the value of the snr threshold is adjusted so that an actual impedance that is approximately 15 % below the impedance threshold would not be subject to contamination by noise that results in the blocking of the biopotential measurement . for example , in the preferred embodiment the snr threshold is set so an actual impedance of 13 kiloohm will not appear to be greater than the impedance limit of 15 kiloohm in the presence of noise . thus the noise voltage limit is set by the following equation : while the foregoing invention has been described with reference to its preferred embodiments , various alterations and modifications will occur to those skilled in the art . all such alterations and modifications are intended to fall within the scope of the appended claims .
a method for measuring bioelectric impedance in real time , in the presence of interference and noise is disclosed . a small electric current is injected into a biopotential electrode system , and then the measurement is tested for contamination by electrical interference or other noise sources .