MAGNETOMETER FOR MEDICAL USE

Abstract

A method of using a magnetometer system (30) to analyse the magnetic field of a region of a subject's body is provided. The method comprises using one or more detectors (60) to detect the time varying magnetic field of a region of a subject's body, using a digitiser (42) to digitise a signal or signals from the one or more detectors (60), each signal that is digitised including noise and a periodic signal produced by one or more of the one or more detectors (60) due to the time varying magnetic field of the region of the subject's body, and averaging the digitised signal or signals over plural periods. The magnetometer system (30) is configured such that the noise in each signal provided to the digitiser for digitisation is greater than about 25% of the interval between digitisation levels of the digitiser (42).

Claims

1. A method of using a magnetometer system to analyse the magnetic field of a region of a subject's body, the method comprising: using one or more detectors to detect the time varying magnetic field of a region of a subject's body; using a digitiser to digitise a signal or signals from the one or more detectors, each signal that is digitised including noise and a periodic signal produced by one or more of the one or more detectors due to the time varying magnetic field of the region of the subject's body; and averaging the digitised signal or signals over plural periods; wherein the magnetometer system is configured such that the noise in each signal provided to the digitiser for digitisation is greater than about 25% of the interval between digitisation levels of the digitiser.

2. The method of claim 1, comprising: using one or more arrays of plural detectors to detect the time varying magnetic field of a region of a subject's body.

3.-5. (canceled)

6. The method of claim 1, wherein each detector comprises a planar coil.

7. The method of claim 1, comprising detecting one or more signal features of the periodic signal produced by one or more of the one or more detectors due to the time varying magnetic field of the region of the subject's body that are smaller than the interval between digitisation levels of the digitiser.

8. The method of claim 1, wherein the magnetometer system is configured such that detector noise and system noise in each signal is greater than about 25% of the interval between digitisation levels of the digitiser.

9. The method of claim 1, wherein the magnetometer system is configured such that the noise in each signal comprises or approximates to white noise.

10. The method of claim 1, wherein each detector comprises a coil or plural coils connected in series and has a configuration such that the resistance of the coil or the plural coils connected in series is in the range 5 to 3000 Ohms.

11. The method of claim 1, further comprising using a noise generator to generate at least part of the noise in each signal.

12.-13. (canceled)

14. The method of claim 1, wherein the region of the subject's body whose magnetic field is being analysed comprises one of: the bladder, heart, head, brain, womb or a foetus.

15. A method of analysing the magnetic field of a subject's heart, the method comprising: using the method of claim 1 to analyse the time varying magnetic field of a subject's heart.

16. A magnetometer system for medical use, comprising: one or more detectors for detecting the time varying magnetic field of a region of a subject's body; a digitiser configured to digitise a signal or signals from the one or more detectors, each signal that is digitised including noise and a periodic signal produced by one or more of the one or more detectors due to the time varying magnetic field of the region of the subject's body; and averaging circuitry configured to average the digitised signal or signals over plural periods; wherein the magnetometer system is configured such that the noise in each signal provided to the digitiser for digitisation is greater than about 25% of the interval between digitisation levels of the digitiser.

17. The magnetometer system of claim 16, comprising: one or more arrays of plural detectors for detecting the time varying magnetic field of a region of a subject's body.

18.-20. (canceled)

21. The magnetometer system of claim 16, wherein each detector comprises a planar coil.

22. The magnetometer system of claim 16, wherein the magnetometer system is configured to detect one or more signal features of the periodic signal produced by one or more of the one or more detectors due to the time varying magnetic field of the region of the subject's body that are smaller than the interval between digitisation levels of the digitiser.

23. The magnetometer system of claim 16, wherein the magnetometer system is configured such that detector noise and system noise in each signal is greater than about 25% of the interval between digitisation levels of the digitiser.

24. The magnetometer system of claim 16, wherein the magnetometer system is configured such that the noise in each signal comprises or approximates to white noise.

25. The magnetometer system of claim 16, wherein each detector comprises a coil or plural coils connected in series and has a configuration such that the resistance of the coil or the plural coils connected in series is in the range 5 to 3000 Ohms.

26. The magnetometer system of claim 16, further comprising a noise generator for generating at least part of the noise in each signal.

27. A magnetometer system for medical use, comprising: one or more arrays of planar coils for detecting the time varying magnetic field of a region of a subject's body, each coil having a maximum outer diameter less than 7 cm; and a detection circuit coupled to each coil and configured to convert a current or voltage generated in the coil by a time varying magnetic field to an output signal for use to analyse the time varying magnetic field.

28. A magnetometer system for medical use, comprising: one or more detectors for detecting a time varying magnetic field, each detector comprising a coil or plural coils connected in series, wherein each detector has a maximum outer diameter of 7 cm or less, and a configuration such that the resistance of the coil or the plural coils connected in series is in the range 5 to 3000 Ohms; and a detection circuit coupled to each detector and configured to convert a current or voltage generated in the coil or coils by a time varying magnetic field to an output signal for use to analyse the time varying magnetic field.

29.-34. (canceled)

Description

[0213] A number of preferred embodiments of the present invention will now be described by way of example only and with reference to the accompanying drawings, in which:

[0214] FIG. 1 shows schematically the use of an embodiment of the present invention for detecting the magnetic field of a subject's heart;

[0215] FIGS. 2-5 show further exemplary arrangements of the use of an embodiment of the present invention when detecting the magnetic field of a subject's heart;

[0216] FIG. 6 shows schematically a coil arrangement in accordance with an embodiment of the present invention;

[0217] FIG. 7 shows a further exemplary arrangement of the use of an embodiment of the present invention when detecting the magnetic field of a subject's heart;

[0218] FIG. 8 shows a probability distribution;

[0219] FIG. 9 shows Fisher information as a function of noise;

[0220] FIGS. 10 and 11 show schematically magnetometer systems in accordance with embodiments of the present invention;

[0221] FIG. 12 shows schematically the operation of a reverse biased Zener diode in accordance with an embodiment of the present invention;

[0222] FIG. 13 shows the output of various white noise generators;

[0223] FIGS. 14-19 show coil arrangements in accordance with embodiments of the present invention; and

[0224] FIG. 20 shows a plot of coil resistance versus noise voltage output of a magnetometer system configured in accordance with an embodiment of the present invention.

[0225] Like reference numerals are used for like components where appropriate in the Figures.

[0226] FIG. 1 shows schematically the basic arrangement of a preferred embodiment of a magnetometer system that may be operated in accordance with the present invention. This magnetometer system is specifically intended for use as a cardiac magnetometer (for use to detect the magnetic field of a subject's heart). However, the same magnetometer design can be used to detect the magnetic field produced by other body regions, for example for detecting and diagnosing bladder conditions, pre-term labour, foetal abnormalities and for magnetoencephalography. Thus, although the present embodiment is described with particular reference to cardio-magnetometry, it should be noted that the present embodiment (and the present invention) extends to other medical uses as well.

[0227] The magnetometer system comprises a detector 40 coupled to a detection circuit 41 that may contain a number of components. The detector 40 may be an induction coil 40, e.g., as described further below.

[0228] The detection circuit 41 may comprise a low impedance pre amplifier, such as a microphone amplifier, that is connected to the coil 40, a low pass filter, e.g. with a frequency cut off of 250 Hz, and a notch filter to remove line noise (e.g. 50 Hz).

[0229] The current output from the coil 40 is processed and converted to a voltage by the detection circuit 41 and provided to an analogue to digital converter (ADC) 42 which digitises the analogue signal from the coil 40 and provides it to a data acquisition system 43.

[0230] A biological signal that is correlated to the heartbeat, e.g. an ECG or Pulse-Ox trigger from the test subject is used as a detection trigger for the digital signal acquisition, and the digitised signal over a number of trigger pulses is then binned into appropriate signal bins, and the signal bins overlaid or averaged, by the data acquisition unit 43.

[0231] The coil 40 and detection circuit 41 may be arranged such that the coil 40 and the preamplifier of the detection circuit 41 are arranged together in a sensor head or probe which is then joined by a wire to a processing circuit that comprises the remaining components of the detection circuit 41. Connecting the sensor head (probe) and the processing circuit by wire allows the processing circuit to be spaced from the sensor head (probe) in use.

[0232] With this magnetometer, the sensor head (probe) will be used as a magnetic probe by placing it in the vicinity of the magnetic fields of interest.

[0233] FIG. 2 shows an improvement over the FIG. 1 arrangement, which uses in particular the technique of gradient subtraction to try to compensate for background noise. In this case, an inverse coil 44 is used to attempt to subtract the effect of the background noise magnetic field from the signal detected by the probe coil 40. The inverse coil 44 will, as is known in the art, be equally sensitive to any background magnetic field, but only weakly sensitive to the subject's magnetic field. The inverse coil 44 can be accurately matched to the pick-up coil 40 by, for example, using a movable laminated core to tune the performance of the inverse coil to that of the pick-up coil 40.

[0234] FIG. 3 shows an alternative gradient subtraction arrangement. In this case, both coils 40, 44 have the same orientation, but their respective signals are subtracted using a differential amplifier 45. Again, the best operation is achieved by accurately matching the coils and the performance of the detection circuits 41. Again, a movable laminated core can be used to tune the performance of one coil to match the performance of the other.

[0235] FIG. 4 shows a further preferred arrangement. This circuit operates on the same principle as the arrangement of FIG. 3, but uses a more sophisticated method of field cancellation, and passive coil matching. In particular, a known global magnetic field 44 is introduced to both coils 40, 44 to try to remove background magnetic field interference.

[0236] In this circuit, the outputs from the detection circuits 41 are passed through respective amplifiers 47, 48, respectively, before being provided to the differential amplifier 45. At least one of the amplifiers 47, 48 is tuneable. In use, a known global field 46, such as 50 Hz line noise, or a signal, such as a 1 kHz signal, applied by a signal generator 49, is introduced to both coils 40, 44. The presence of a signal on this frequency on the output of the differential amplifier 45, which can be observed, for example, using an oscilloscope 50, will then indicate that the coils 40, 44 are not matched. An amplifier control 51 can then be used to tune the tuneable voltage controlled amplifier 48 to eliminate the global noise on the output of the differential amplifier 45 thereby matching the outputs from the two coils appropriately.

[0237] Most preferably in this arrangement, a known global field of 1 kHz or so is applied to both coils, so as to achieve the appropriate coil matching for the gradient subtraction, but also a filter to remove 50 Hz noise is applied to the output signal.

[0238] FIG. 5 shows a further variation on the FIG. 4 arrangement, but in this case using active coil matching. Thus, in this arrangement, the outputs of the coils 40, 44 are again channelled to appropriate detection circuits 41, and then to respective amplifiers 47, 48, at least one of which is tuneable. However, the tuneable amplifier 48 is tuned in this arrangement to remove the common mode noise using a lock-in amplifier 52 or similar voltage controller that is appropriately coupled to the output from the differential amplifier 45 and the signal generator 49.

[0239] The above embodiments of the present invention show arrangements in which there is a single pickup coil that may be used to detect the magnetic field of the subject's heart. In these arrangements, in order then to make a diagnostic scan of the magnetic fields generated by a subject's heart, the single pickup coil can be moved appropriately over the subject's chest to take readings at appropriate spatial positions over the subject's chest. The readings can then be collected and used to compile appropriate magnetic field scans of the subject's heart.

[0240] It would also be possible to arrange a plurality of coil and detection circuit arrangements, e.g. of the form shown in FIG. 1, in an array, and to then use such an array to take measurements of the magnetic field generated by a subject's heart. In this case, the array of coils could be used to take readings from plural positions over a subject's chest simultaneously, thereby, e.g., avoiding or reducing the need to take readings using the same coil at different positions over the subject's chest.

[0241] FIG. 6 shows a suitable coil array arrangement that has an array 60 of 16 detection coils 61, which may be then placed over a subject's chest to measure the magnetic field of a subject's heart at 16 sampling positions over the subject's chest. In this case, each coil 61 of the array 60 should be configured as described above and connected to its own respective detection circuit (i.e. each individual coil 61 will be arranged and have a detection circuit connected to it as shown in FIG. 1). The output signals from the respective coils 61 can then be combined and used appropriately to generate a magnetic scan of the subject's heart.

[0242] Other array arrangements could be used, if desired, such as circular arrays, irregular arrays, etc.

[0243] More (or less) coils could be provided in the array, e.g. up to 50 coils, or more than 50 coils. For example, where it is desired to measure the magnetic field of a different region of a subject's body (i.e. other than the heart), then an increased number of coils may be provided so as to provide an appropriate number of sampling points and an appropriate spatial coverage for the region of the subject's body in question.

[0244] It would also be possible in this arrangement to use some of the coils 61 to detect the background magnetic field for the purposes of background noise subtraction, rather than for detecting the wanted field of the subject's heart. For example, the outer coils 62 of the array could be used as background field detectors, with the signals detected by those coils then being subtracted appropriately from the signals detected by the remaining coils of the array. Other arrangements for background noise subtraction would, of course, be possible.

[0245] It would also be possible to have multiple layers of arrays of the form shown in FIG. 6, if desired. In this case, there could, for example, be two such arrays, one on top of each other, with the array that is closer to the subject's chest being used to detect the magnetic field generated by the subject's heart, and the array that is further away being used for the purposes of background noise detection.

[0246] To measure the magnetic fields generated by the heart, the above arrangements can be used to compile magnetic field scans of a subject's heart by collecting magnetic field measurements at intervals over the subject's chest. False colour images, for example, can then be compiled for any section of the heartbeat, and the scans then used, for example by comparison with known reference images, to diagnose various cardiac conditions. Moreover this can be done for significantly lower costs both in terms of installation and on-going running costs, than existing cardiac magnetometry devices.

[0247] FIG. 7 shows an exemplary arrangement of the magnetometer as it is envisaged it may be used in a hospital, for example. The magnetometer 30 is a portable device that may be wheeled to a patient's bedside 31 where it is then used to take a scan of the patient's heart (e.g.). There is no need for any magnetic shielding, cryogenic cooling, etc. The magnetometer 30 can be used in the normal ward environment.

[0248] It should be noted here that the signal generated by the pick-up coil in the present embodiments (and invention) will be the derivative of the useful signal, so the output signal can be (and preferably is) integrated over time to generate the wanted, useful signal. Such integration will also have the effect of tending to remove the effect of noise from the signal (provided the noise amplitude is not too big). Furthermore, the noise will remain in the integrated signal and so can be recovered if desired or needed, by taking the derivative of the integrated signal. The magnetometer system can be used in an analogous manner to detect and analyse other medically useful magnetic fields produced by other regions of the body, such as the bladder, head, brain, a foetus, etc.

[0249] In these embodiments, the bio-physical magnetic fields of interest are typically very small, and can be so small that they are significantly below spurious background noise signals. The conventional approach to detect these signals would be to attempt to remove the background using passive shielding and/or active cancellation.

[0250] However, the Applicants have recognised that at least some noise is actually desirable, and in particular that a noisy signal can increase the sensitivity of the apparatus to sub-threshold signals when attempting to recover very small periodic signals.

[0251] In the present embodiments, noise is used to enable sub-threshold signals to be detected. Sub-threshold in this sense refers to signals with a voltage amplitude that is smaller than the smallest voltage separation in the analogue to digital conversion (ADC) system 42. Such signals appear as a single structureless output with zero information content. However, noise can be used to increase the effective peak-to-peak voltage of the signal, triggering transitions in the digital conversion. Hence the noise allows a signal to be detected with positive information content.

[0252] To detect the repeating signal, a trigger function is used to inform the detection apparatus of the presence of the signal. The underlying target signal is acquired repeatedly, and signal averaging techniques are applied to the data.

[0253] The detected voltage will comprise two basic elements: a detected field that includes both the background and target fields, and an additional noise element. The noise element has a number of potential sources, including thermal noise, Johnson noise and intensity noise. Such noise sources are unavoidable and will therefore be present on any detection system. This noise will be random and incoherent (and therefore uncorrelated) both spatially (between detectors) and as a function of time (for a single detector). Any noise component that does not conform to this model (such as environmental noise) can be dealt with in post processing of the data.

[0254] The processing of the signal involves the detection of the signal as a voltage that is digitised into a set of discrete levels. These levels are indexed by an integer value, denoted by x, referring to the strength, size or level of the signal with maximum level, X. If the width of the digitized levels is v, which represents the minimum detectable voltage movement, the maximum signal that can be detected is V=Xv.

[0255] Digitisation of the signal results in the loss of information. If the magnitude of the total signal S for any is less than v, then digitisation will result in a complete loss of information. The Fisher information is a way of measuring the amount of information that an observable random variable x carries about an unknown parameter Q upon which x depends.

[0256] The total signal to be detected is a function of two elements, the biomagnetic signal Q and environmental noise a, giving a signal S(Q,). The extent to which x reveals information about Q, is determined by the degree to which the distributions of Q and a can be distinguished.

[0257] To determine the extent to which x can be used to measure Q, the Fisher information, F(x,Q), is used, which is a measurement of the amount of information that an observable x carries about the otherwise unknown parameter Q upon which the probability of x depends.

[0258] In other words, a change in the signal S can produce a commensurate change in x, thereby changing the information. The extent to which this occurs (how much of a change in the Q is required to change x) determines the information that x contains. If v is larger than a then x is unlikely to change and therefore carries very little information, if x changes frequently then a single sample of x contains more information.

[0259] If f(x,Q) is the likelihood of x with respect to Q, then the Fisher information F(x,Q) is given by:

[00002]$F\left(x,Q\right)={\left(\frac{d}{\mathrm{dQ}}.Math.\log \left(f\left(x,Q\right)\right)\right)}^2.Math..Math.f\left(x,Q\right).Math.\mathrm{dx},$

where the likelihood function, f(x,Q)=P(x|Q,), is given by the distribution of likely values that x can have for a given Q and noise level a.

[0260] FIG. 8 shows the probability distribution for x, which is determined by Q and the noise level.

[0261] The final component of the signal that must be considered is the impact of noise. For this example, the noise is considered to take the form of a small displacement:

0<Q<<v,

which is accompanied by random (incoherent) Gaussian white noise displacements. Under these conditions the conditional probability distribution for x takes the form:

P(x|Q,)=A Exp[(xQ).sup.2/(2.sup.2)],

where A is a normalisation amplitude and the noise amplitude a is the width of a Gaussian distribution centered at Q.

[0262] In this case the Fisher information becomes:

[00003]$F\left(x,Q\right)=\frac{\mathrm{Exp}\left(-\frac{{\left(x-Q\right)}^2}{2.Math..Math.{}^2}\right).Math.{\left(x-Q\right)}^2}{{}^4}.Math..Math.\mathrm{dx}.$

[0263] FIG. 9 shows the Fisher information as a function of the noise level a. As can be seen from FIG. 9, below a threshold of noise amplitude, a, the information drops to zero and it is impossible to extract information about Q from S. Hence, noise is in fact essential to extract the signal. Moreover, and counter-intuitively, there is no maximum noise level beyond which it is impossible to extract Q.

[0264] The Fisher information peaks when v/2 and then falls monotonically, however it always remains above zero. Hence, the amount of information that can be extracted per iteration reduces, and more averaging is required to calculate an accurate value for Q. However, importantly it does not return to zero and therefore it is possible to determine Q for even large values of noise.

[0265] Thus, in the present embodiment, the magnetometer system is configured such that the noise in the signal (i.e. the noise amplitude of the signal or, where the noise comprises white noise or other Gaussian noise, the standard deviation of the noise amplitude of the signal) provided (input) to the digitiser is greater than about 25% of the interval between digitisation levels of the ADC, e.g. about 50% of the interval between digitisation levels of the ADC.

[0266] This solves the problem of maintaining a significantly large dynamic range in the presence of a large noise. It is possible to retain information at the small signal end while also capturing the full deviation of the noise. This shows that in fact, in the presence of noise it is possible to do away with low level sensitivity as long as the noise is sufficiently high.

[0267] Hence a smaller induction coil that is insufficiently sensitive to detect bio-magnetic fields in a low noise environment can be used in a higher noise environment to detect the same small bio-magnetic signals.

[0268] In order to extract data from the signal, a trigger is used to gate the coil output, and the signal is split into a number of gate cycles. Power-line and other noise background noise sources can be removed by filtering if required, but Johnson shot noise and other noise sources are retained to increase the dynamic range. The sub-signals are averaged to produce the output.

[0269] As each trigger instance delivers a small amount of information, this final step of averaging is used to deliver the final information content of the signal. Fine detail and small elements of the signal can be future extracted with more averaging. Accordingly, in the present embodiment a signal that is smaller than the interval between digitisation levels of the digitiser, i.e. a signal that would otherwise be smaller than the minimum signal detectable by the digitiser can be detected. Similarly, a signal that is smaller than the noise, i.e. a signal with a noise ratio less than one, can be detected.

[0270] In particular, the P wave, QRS wave and/or T wave of the time varying magnetic field of a subject's heart or other signal features of interest, that is or are smaller than the interval between digitisation levels, can be detected.

[0271] The signal to noise ratio for the magnetometer system of the present embodiment for a typical heart scan conducted in a typical noise environment may be around 120 dB (although it should be noted that the noise environment may be subject to significant variation). The signal to noise ratio for the system when the electronics system is terminated with a representative impedance (i.e. so as to exclude external environmental noise) may be around 55 dB. The signal level amplitude of signal features of interest is significantly lower than the quantisation level of the ADC.

[0272] As shown in FIG. 10, potential sources of noise in the magnetometer system according to the present embodiment include environmental noise, detector noise, and other system noise such as electronics noise, etc.

[0273] In the present embodiment, the combination of detector noise and system noise (i.e. the local noise) is controlled to be greater than about 25% of the interval between digitisation levels of the digitiser after the amplifier 41 and prior to the digitiser 42 (e.g. about 50% of the interval between digitisation levels of the digitiser), and the environmental noise is not controlled in this manner. In the present embodiment, most of the environmental noise is removed by using a notch filter to remove the mains (line) frequency (50 Hz in the UK) from the signal prior to digitisation. A low pass filter with an appropriate cut-off frequency (e.g. 40 Hz, where the line frequency is 50 Hz) may additionally be used to remove any remaining high frequency environmental noise.

[0274] This has a number of advantages. Firstly, detector noise and system noise (local noise) can be controlled relatively easily and relatively more accurately, e.g. by choice of detector, system components, etc. In comparison, control of environmental noise typically requires complex and expensive arrangements such as shielding, etc., and unexpected sources of environmental noise may exist which cannot be accounted for and controlled.

[0275] Secondly, detector noise and system noise (local noise) will typically have a relatively cleaner spectrum when compared with environmental noise, e.g. may comprise or may approximate to white noise. This simplifies the signal processing required to obtain the final output signal, and ensures that the noise present in the system contains relatively little or no structure that could otherwise interfere with the periodic signal of interest.

[0276] The desired amount of noise in the system can be provided in a number of different ways.

[0277] For example, the detector system can be designed in order to provide sufficient noise. In this case, at least some of the desired noise may be provided by detector noise such as shot noise, Johnson noise etc., and/or at least some of the desired noise may be provided by system noise such as electronics noise, etc. System noise in this context includes sources of noise within the magnetometer system (e.g. the magnetometer electronics) that introduce noise to the signal before the signal is digitised by the digitiser.

[0278] This advantageously relaxes the design constraints on the coils, and allows, e.g. the detectors to be designed with less emphasis on the signal to noise characteristics of the coils, and with relatively more emphasis on size, shape, and weight considerations in order to provide a magnetometer system that is more suitable for medical magnetometry (and in particular for magneto cardiography).

[0279] Additionally or alternatively, noise can be added to the system. FIG. 11 shows one such embodiment, in which a white noise generator 70 is used to add noise to the signal. This allows for a greater degree of control over the level of noise present in the system, and may allow the noise to more closely approximate to white noise.

[0280] In this regard, the Applicants have recognised that in practice it may be possible for any inherent system noise and any environmental noise to fall below about 25% of the interval between digitisation levels in use, for example where environmental noise is particularly small and/or is otherwise removed from the detector signal(s) and/or depending on the design of the detector(s), etc., and that this would result in a poor output signal (for the reasons given above). Accordingly, in order to ensure that there is sufficient noise in the signal(s) that is digitised, the magnetometer system may be appropriately designed, and/or a noise generator may be used to increase the noise to greater than about 25% of the digitisation interval.

[0281] As shown in FIG. 12, white noise can be generated by reverse biasing a Zener diode and amplifying the output to achieve the required amplitude.

[0282] Alternative means of generating white noise that may be used include digital techniques such as a shift register arrangement using an arrangement of shift registers with feedback. Such designs may be implemented using field programmable gate arrays. FIG. 13 shows the output of three different white noise generators in accordance with various embodiments.

[0283] The amount of noise that is added to the signal by the noise generator 70 may be manually controlled and/or optimised, e.g. by means of a dial or other device, or automatically controlled and/or optimised, e.g. by means of software or otherwise. For example, the amount of added noise may be optimised by varying the amount of added noise, and monitoring the effects of the added noise on the recovered signal until an optimised signal is provided.

[0284] Additionally or alternatively, the amount of added noise may be optimised by measuring the natural noise in the system, and then controlling the added noise in response to the measurement.

[0285] In this case, the system noise (or natural noise) can be measured by switching the detector at the input of the amplifier out of circuit, and terminating the input with an equivalent impedance, and then either directly determining the number of noise bits on the ADC or adding noise until the measured system noise is two times the initial measurement, whereupon the added noise will be equal to the system noise.

[0286] The amount of noise may be controlled or optimised as part of the initialisation of the magnetometer system, e.g. automatically as part of an initialisation routine for the magnetometer system.

[0287] The design of the induction coil for use in the preferred embodiments of the present invention will now be described. The coil should have an overall size that permits spatial resolution suitable for magneto cardiography.

[0288] Johnson noise is normally considered to be the limit of sensitivity of a magnetic detection coil, and therefore conventional coil designs aim to optimise signal over Johnson noise. Designs optimising signal to noise that limit Johnson noise are beneficial when other environmental noise sources are limited. However, in any practical circumstances, most sources of environmental background noise are much larger than either the target signal or the Johnson noise. Hence, the design criteria limiting Johnson noise can be substantially relaxed, indeed, as described above, Poissonian distributed noise, such as Johnson noise actually plays a positive role in signal extraction.

[0289] The frequency of the relevant magnetic signals of the heart is between 1 Hz and 60 Hz. Thus, the coil of the present embodiments is designed to be sensitive to magnetic fields at these frequencies.

[0290] As shown in FIG. 14, in the present embodiment, each coil comprises a planar coil 80, i.e., a coil with plural turns arranged in a single plane. Each planar coil 80 comprises a conductor track 81 arranged in a spiral. Any suitable conductor may be used for the planar coil, such as copper, gold, silver, carbon nanotubes, graphene, or similar.

[0291] Each planar coil 80 may have a maximum outer diameter D between 2 and 6 cm. The conductor track has width W (a thickness in the direction parallel to the plane in which the planar coil's plural turns are arranged) of around 1 mm or less. For example, the width may be between around 0.1 and 1 mm, or between around 0.3 and 0.5 mm.

[0292] The turns of the planar coil are spaced apart (in the direction parallel to the plane in which the planar coil's plural turns are arranged) with a gap G of around 1 mm or less. For example, the spacing may be around 0.1 mm or less or 0.01 mm or less.

[0293] The conductor track is relatively flat in the direction orthogonal to the plane in which the coil's plural turns are arranged, and has a thickness Z (in the direction orthogonal to the plane in which the coil's plural turns are arranged) of around 0.2 mm or less. For example, the thickness Z may be around 0.1 mm or less, or around 0.05 mm or less. The thickness Z should not be less than around 0.035 mm.

[0294] The planar (spiral) coil may comprise desired number of turns, such as 2 or more turns.

[0295] Each planar coil 80 comprises a spiral conductor arranged on or within an electrically insulating substrate 82. The insulating substrate supports the conductor, and is accordingly substantially rigid. Any suitable insulator can be used for the insulating substrate, such as glass, plastic, reinforced plastic, etc. The substrate may also comprise a printed circuit board (PCB) material, such as FR4 and the like.

[0296] The Applicants have recognised that it can be beneficial to increase the width W of the conductor track (its thickness in the direction parallel to the plane in which the planar coil's plural turns are arranged), and correspondingly to reduce the size of the gap G between each of the turns of the planar coil. This has the effect of reducing the resistance of the coil, while increasing the useful signal collecting area of the coil.

[0297] Increasing the height Z of the conductor track (its thickness in the direction orthogonal to the plane in which the coil's plural turns are arranged) also has the effect of reducing the resistance of the coil. However, this can affect the minimum spacing between turns that it is feasible to achieve in practice. For example, where chemical etching techniques are used to form the spiral track 81, increasing the conductor height can increase the minimum gap between each of the turns of the spiral that it is feasible to achieve in practice.

[0298] In one exemplary arrangement, the thickness Z is 0.035 mm, and the gap width G is 0.127 mm. Where the thickness Z is increased above around 0.05 mm, the track gap has to be increased. In another exemplary arrangement (e.g. using a 0.14 mm thick FR4 PCB), the thickness Z is 0.14 mm, and the gap width G is 0.25 mm. The Applicants have found that these value ranges are achievable in practice and provide planar coils that are capable of producing a useful output signal.

[0299] Laser etching may be used to achieve even smaller gap widths G between the turns of the planar coil, e.g. as small as 3-5 m.

[0300] Each planar coil may optional further comprise a magnetic core. This has been found to improve performance of the planar coils. A magnetic core may be located within (at the axial centre of) each spiral planar coil, or otherwise adjacent to each spiral planar coil. Each planar spiral coil may have an inner diameter in the range 4-35 mm, or 15-22 mm, and correspondingly each magnetic core may have a diameter in the range 4-35 mm, or 15-22 mm.

[0301] The magnetic core should be made from a material with a high relative permeability such as a ferrite or other magnetic material. For example, the magnetic core may be made from a magnetic amorphous metal alloy and/or a nano-crystalline material. These materials can exhibit very high magnetic permeabilities, but can be lighter than other magnetic materials such as iron powder.

[0302] Nanocrystalline materials are poly-crystalline materials with very small grain sizes, the space between which is filled with amorphous materials.

[0303] Amorphous metals (sometimes referred to as metallic glasses or glassy metals) differ from traditional metallic materials and alloys in that they have highly disordered atomic structures instead of conventional crystalline or poly-crystalline lattices, and as such have a number or unique properties. They are typically produced from a mixture of differently sized metallic atoms which are quench-cooled at millions of degrees per second, removing the thermal energy for atoms to move and form ordered domains or grains. By alloying with certain magnetic materials such as iron, cobalt, and nickel, very high magnetic permeability and susceptibility materials are possible. Their high(er) resistance (similar to that of their molten components) reduces eddy current losses when subjected to alternating magnetic fields. Their low coercivity also reduces losses.

[0304] One exemplary such material is known as MetGlas 2714a (Metallic Glass Alloy).

[0305] Utilising planar coils is beneficial, particular where as discussed above the magnetometer system comprises plural detectors arranged in plural layers, since this can result in a magnetometer system having a size, shape, and weight that is more suitable for medical magnetometry (and in particular for magneto cardiography). For example, FIG. 15 shows a suitable coil array 60 arrangement that has an array of planar detection coils 80, which may be then placed over a subject's chest to measure the magnetic field of a subject's heart at 7 sampling positions over the subject's chest.

[0306] The or each array may comprise 30-45 planar coils, e.g. 37 planar coils, arranged in a two dimensional array. This number of planar coils has been found to provide particularly good spatial coverage for the human heart.

[0307] As shown in FIG. 16, a single layer of planar coils may be utilised for measuring the magnetic field of a subject's heart. Alternatively, as shown in FIG. 17, since the planar coils are relatively flat in the out-of-plane direction, plural layers of planar coils may be conveniently utilised for measuring the magnetic field of a subject's heart.

[0308] The one or more arrays of planar coils may comprise 20-120 layers of planar coils, e.g. 40-90 layers, one above the other. This has been found to result in a magnetometer system that produces a useful signal, while having a weight that is suitable for medical magnetometry (and in particular for magneto cardiography).

[0309] Each layer may be formed on its own PCB layer, or plural layers of planar coils may be arranged on or within a single multi-layer PCB, e.g. so as to reduce the overall weight of the magnetometer system.

[0310] Where the magnetometer system comprises plural detectors, some or all of the detectors may be connected, e.g. in parallel and/or in series. Connecting plural detectors in series will have the effect of increasing the induced voltage for a given magnetic field strength. Connecting plural detectors in parallel will have the effect of reducing the thermal noise (Johnson noise) in the detectors. As shown in FIG. 18, a combination of series and parallel connections can be used to optimise the balance of voltage and noise performance of the detectors.

[0311] As shown in FIG. 19, one or more or each detector in the magnetometer system can be arranged in a gradiometer configuration, i.e. where two detectors are co-axially aligned (in the direction orthogonal to the plane in which each coil's windings are arranged), and where the signal from each of the coils is summed, e.g. to provide a measure of a change in the magnetic field in space.

[0312] As described above, according to the Fisher information, the presence of some noise allows small signals to be detected that would otherwise be undetectable. The Applicants have furthermore recognised that there is an optimum noise range that enables signal detection in a reasonable time.

[0313] As described above, the Fisher information is always positive, so that more noise always enhances small signals. However, the Fisher information reduces exponentially above a threshold of about 50% of the interval between digitisation levels of the digitiser. Higher levels of noise accordingly mean that more periods of the signal must be averaged in order to obtain an output signal with a useful information content. As such, an upper noise limit is set by the maximum practical scan time.

[0314] The Applicants have furthermore recognised that controlling the resistance of the coil or group of plural coils connected in series is equivalent to controlling the local noise. There are three main noise sources in the apparatus, namely (i) Johnson noise from the coil windings, (ii) current noise generated in the system, and (iii) circuit noise in the amplifier circuit. Noise sources (i) and (ii) are due to the induction coil and can be controlled for optimum noise performance.

[0315] Since adding more coils to the system increases the resistance, this can lead to the result that adding more coils to the sensor, in some cases, does not improve the system.

[0316] FIG. 20 shows a plot of coil resistance against noise voltage output. As the number of turns on a coil increases so does the resistance of the coil. The resistance impacts Johnson noise, as well as the noise produced by the amplifier, i.e. in the form of voltage noise and current noise (this varies to some extent dependent on the circuit selected). Johnson noise remains the dominant source of noise.

[0317] There is an upper limit for the scan time that a patient can tolerate of around 10-15 minutes. Preferably this is less than 5 minutes and ideally less than 3 minutes. This time limit limits the maximum noise that the system can practically be configured to have, and in turn limits the maximum resistance that the coil or coils can practically be configured to have. In particular, the system should operate with resistance values that are to the left of the 15 minute point in FIG. 20.

[0318] Accordingly, the optimum resistance range for the coil or coils is around 5-3000 Ohms, preferably 10-2000 Ohms, more preferably 10 to 200 Ohms.

[0319] The magnetometer system of the present embodiment can be configured to have a particular desired coil resistance by appropriate selection of the type and/or design of detector (e.g. induction coil and/or planar coil), the number of turns on the or each coil, the number of coils (e.g. connected in series), and/or the resistance per unit length (e.g. the cross-sectional area) of the wire, etc.

[0320] In one exemplary arrangement, each planar coil comprises a layer of a copper clad printed circuit board FR4 material. Each planar coil has a track width W of 0.1 mm and gap width G of 0.1 mm. The planar coil is in a spiral form with an inner diameter of 4 mm and an outer diameter 48 mm.

[0321] The PCB comprises a multi-layer PCB having ten layers of planar coils.

[0322] Each of the individual layers is connected in series. The combined resistance for this arrangement is 400 Ohms. For an array of 37 planar coils with ten layers, the device can detect the magnetic field of a human heart with a scan time of around 1 hour.

[0323] Where a second ten layer PCB is added in series (a total of 20 layers), the device can detect the magnetic field of a human heart in 40 minutes. Where a third ten layer PCB is added in series (a total of 30 layers) the device can detect the magnetic field of a human heart in 15 minutes.

[0324] Where, however, additional layers are added in series, then the signal deteriorates. For example, a device having five ten layer PCBs (a total of 50 layers) requires >30 minutes to detect the magnetic field of a human heart. The resistance of this system is around 2000 Ohms.

[0325] In contrast, where thirty layers are placed in parallel with thirty layers (resistance 600 Ohms), the magnetic field of a human heart can be detected in 10 minutes. Where thirty layers are placed in parallel with thirty layers and again in parallel with thirty layers (resistance 400 Ohms) the magnetic field of a human heart can be detected in 5 minutes.

[0326] It will accordingly be appreciated that the amount of the noise in the system can be controlled by controlling the resistance of the system, which in turn can be controlled by selecting an appropriate combination of series and parallel connections between plural planar coils.

[0327] It can be seen from the above that the present invention, in its preferred embodiments at least, provides a magnetic imaging device that can be deployed effectively from both a medical and cost perspective in a wide range of clinical environments, e.g. for use when detecting magnetic fields generated by the heart. The magnetometer is, in particular, advantageous in terms of its cost, its practicality for use in clinical environments, and its ability to be rapidly deployed for near patient diagnosis and for a wide range of applications. It is non-contact, works through clothing, fast, compact and portable and affordable. An image can be recovered with high resolution after a minute of signal recording and absolute single beat sensitivity is potentially possible. Patient motion of up to 1-2 cm will not significantly degrade the image.

[0328] This is achieved, in the preferred embodiments of the present invention at least, by detecting, digitising and averaging plural repeating periods of the time varying magnetic field of a region of a subject's body overall plural periods, where the system is arranged such that the noise present in the signal to be digitised is greater than about 25% of the interval between digitisation levels of the digitiser.