Sensor configuration in magnetometer for medical use

Abstract

A magnetometer system for medical use comprises one or more induction coils for detecting a time varying magnetic field. Each coil has a maximum outer diameter of 10 cm or less, and a configuration such that the ratio of the coil's length to its outer diameter is 0.9 or more, and the ratio of the coil's inner diameter to its outer diameter is 0.6 or more. Each induction coil comprises a magnetic core. The magnetometer system further comprises 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.

Claims

1. A magnetometer system for medical use, comprising: one or more induction coils for detecting a time varying magnetic field, each induction coil of the one or more induction coils having: a maximum outer diameter of 10 cm or less; a configuration such that a ratio of its length to its outer diameter is in a range 0.9 to 3, and a ratio of its inner diameter to its outer diameter is in a range 0.6 to 1; and a magnetic core; the magnetometer system further comprising a detection circuit coupled to each induction coil of the one or more induction coils and configured to convert a current or voltage generated in each induction coil of the one or more induction coils by the time varying magnetic field to an output signal for use to analyse the time varying magnetic field.

2. The magnetometer system of claim 1, wherein the one or more induction coils comprise plural induction coils arranged in one or more two or three dimensional arrays.

3. The magnetometer system of claim 1, wherein each induction coil of the one or more induction coils comprises plural layers of turns.

4. The magnetometer system of claim 1, wherein each induction coil of the one or more induction coils has a winding length of 10 cm or less.

5. The magnetometer system of claim 1, wherein each induction coil of the one or more induction coils has a configuration such that the ratio of its length to its outer diameter is in the range 1 to 1.5.

6. The magnetometer system of claim 1, wherein each induction coil of the one or more induction coils comprises wire having a radius less than 0.2 mm.

7. The magnetometer system of claim 1, wherein each magnetic core comprises a material with a relative permeability, μ.sub.r, of at least 1000.

8. The magnetometer system of claim 1, wherein each magnetic core comprises a magnetic amorphous metal alloy, a nano-crystalline material, a nickel-iron alloy or a cobalt-iron alloy.

9. The magnetometer system of claim 1, wherein for each coil of the one or more induction coils, the ratio of the coil's magnetic core's outer diameter to the coil's inner diameter, D.sub.c:D.sub.i, is 0.8 or more.

10. The magnetometer system of claim 9, wherein for each coil of the one or more induction coils, the ratio of the coil's magnetic core's outer diameter to the coil's inner diameter, D.sub.c:D.sub.i, is 0.9 or more.

11. The magnetometer system of claim 1, wherein the ratio of each magnetic core's length to its outer diameter, l.sub.c:D.sub.c, is at least 1.

12. The magnetometer system of claim 1, wherein each magnetic core is hollow.

13. A cardiac magnetometer system for analysing the magnetic field of a region of a subject's body, comprising the magnetometer system of claim 1.

14. The use of the magnetometer system of claim 1 for analysing the time varying magnetic field generated by a region of a subject's body.

15. A coil for use to detect a time varying magnetic field of a region of a subject's body, the coil comprising: an induction coil having a maximum outer diameter of 10 cm or less, and a configuration such that a ratio of the induction coil's length to its outer diameter is in a range 0.9 to 3, and a ratio of the induction coil's inner diameter to its outer diameter is 0.6 to 1; and a magnetic core.

16. A method of analysing a magnetic field of a region of a subject's body, the method comprising: using one or more induction coils to detect a time varying magnetic field of a region of the subject's body, each induction coil of the one or more induction coils having: a maximum outer diameter of 10 cm or less, and a configuration such that a ratio of its length to its outer diameter is in a range 0.9 to 3, and a ratio of the its inner diameter to its outer diameter is in a range 0.6 to 1; and a magnetic core; the method further comprising converting a current or voltage generated in each induction coil of the one or more induction coils by the time varying magnetic field of the region of the subject's body to an output signal; and using the output signal or signals from the one or more induction coils to analyse the time varying magnetic field generated by the region of the subject's body.

17. The method of claim 16, comprising using the one or more induction coils to detect the time varying magnetic field of the region of the subject's body in a non-magnetically shielded environment.

18. The method of claim 16, wherein the region of the subject's body comprises one of: the abdomen, bladder, heart, head, brain, chest, womb, one or more foetuses, or a muscle.

19. A method of analysing the magnetic field of a subject's heart, the method comprising: using the method of claim 16 to analyse the time varying magnetic field of the subject's heart.

Description

(1) A number of embodiments of the technology described herein will now be described by way of example only and with reference to the accompanying drawings, in which:

(2) FIG. 1 shows schematically the use of an embodiment of the technology described herein for detecting the magnetic field of a subject's heart;

(3) FIGS. 2-5 show further exemplary arrangements of the use of an embodiment of the technology described herein when detecting the magnetic field of a subject's heart;

(4) FIG. 6A shows schematically a coil arrangement in accordance with an embodiment of the technology described herein, and FIG. 6B shows schematically another coil arrangement in accordance with an embodiment of the technology described herein;

(5) FIG. 7 shows a further exemplary arrangement of the use of an embodiment of the technology described herein when detecting the magnetic field of a subject's heart;

(6) FIG. 8 shows schematically the coil configuration for a Brooks coil;

(7) FIG. 9 illustrates the effect of a coil's outer diameter on the coil's inductance;

(8) FIG. 10 illustrates the effect of a core's aspect ratio on its effective permeability;

(9) FIG. 11 illustrates the effect of a core's aspect ratio on its effective permeability;

(10) FIG. 12 shows schematically a single layer coil and a multi-layer coil;

(11) FIG. 13A shows schematically a coil in accordance with an embodiment of the technology described herein, and FIG. 13B shows schematically another coil in accordance with another embodiment of the technology described herein; and

(12) FIG. 14 shows schematically a coil in accordance with an embodiment of the technology described herein.

(13) Like reference numerals are used for like components where appropriate in the Figures.

(14) FIG. 1 shows schematically the basic arrangement of various embodiments of a magnetometer system that may be operated in accordance with the technology described herein. 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 technology described herein) extends to other medical uses as well.

(15) The magnetometer system comprises an induction coil 40 coupled to a detection circuit 41 that may contain a number of components.

(16) The detection circuit 41 may comprise a low impedance pre amplifier, such as a microphone amplifier, that is connected to the coil 40, and one or more filters, e.g. one or more a low pass filters, one or more high pass filters, one or more band pass filters, and/or one or more notch filters e.g. to remove line noise (e.g. 50 or 60 Hz and harmonics).

(17) 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.

(18) A biological signal that is correlated to the heartbeat, e.g. an ECG or Pulse-Ox trigger from the test subject may be 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. Other arrangements would, however, be possible.

(19) 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.

(20) 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.

(21) 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. (Other techniques could, however, be used). 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 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 pickup coil 40 by, for example, using a movable laminated core to tune the performance of the inverse coil to that of the pickup coil 40.

(22) 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.

(23) FIG. 4 shows a further embodiment. 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.

(24) 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 or 60 Hz (and harmonics) line noise, or a signal, such as a 1 kHz signal, applied by a signal generator 49, is introduced to both coils 40, 44.

(25) 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.

(26) In particular embodiments 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 or 60 Hz (and harmonics) noise is applied to the output signal.

(27) 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.

(28) The above embodiments of the technology described herein 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.

(29) 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.

(30) FIGS. 6A and 6B show suitable coil array arrangements that have 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. FIG. 6A shows a regular rectangular array and FIG. 6B shows a regular hexagonal array. In these cases, 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.

(31) Other array arrangements could be used, if desired, such as circular arrays, irregular arrays, etc.

(32) 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.

(33) 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.

(34) 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.

(35) 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.

(36) 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. (Magnetic shielding and/or cooling could, however, be provided if desired.)

(37) In the technology described herein, each coil's 61 length, l, its outer diameter, D, and its inner diameter, D.sub.i, are carefully selected in order to improve the coil's 61 sensitivity to bio-magnetic fields.

(38) In its simplest form, an induction coil is an electronic component that responds to changes in a magnetic field by producing an electromotive force (EMF, or voltage difference) in opposition (due to Lenz's law) to the field that produced this force. From this induced potential difference (voltage), a current will flow through the coil.

(39) It has been shown mathematically that the maximum possible inductance of a coil with an air core for a given length of wire is the Brooks coil.

(40) FIG. 8 illustrates the design of the conventional Brooks coil. Here the winding cross-section is square and the overall diameter of the coil has a width of 4 times one of the sides of the square. The inductance L for the Brooks coils is given by the equation:
L≅0.02591hN.sup.2μ.sub.0H
where h is the height or length of one side of the square winding cross section, N is the total number of turns, μ.sub.0 is the permeability of free space, and H is the magnetic field strength. This can also be expressed in terms of the mean winding radius (r.sub.mean) of the coil as follows:
L≅0.016994r.sub.meanN.sup.2μ.sub.0H

(41) H can also be expressed as BA which represents the magnetic flux density B multiplied by the cross-sectional area A of the coil:
H=BA=Bπr.sub.mean.sup.2

(42) It can be seen from these equations that to increase the inductance L of an air-core coil, either the radius of the coil r.sub.mean or the number of turns N must be increased. However, both of these add to the electrical resistance of the coil.

(43) FIG. 9 illustrates the effect of a coil's outer diameter, D, on the coil's inductance, L. In FIG. 9, the solid (lower) line shows the change in measured inductance L with diameter D of a coil that does not have a magnetic core, while the dashed (upper) line shows the change in measured inductance L with diameter D of a coil that has a soft magnetic core. Each coil measured in FIG. 9 had a fixed number of (30) turns. FIG. 9 shows that the presence of a soft magnetic core improves the inductance of the coil. Moreover, FIG. 9 shows that, for a fixed number of turns, the inductance L of a (single-layer) coil increases with its diameter (and cross-sectional area), allowing it to cut more magnetic flux lines.

(44) Most conventional coil designs are based upon the Brooks coil and winding cross-sections. Air core coils are commonly used because they do not saturate easily and experience low losses, particularly at higher frequencies. Often air-gaps are deliberately introduced to certain inductors to reduce core saturation. Inner-to-outer diameter ratios are typically small. Large gauges of wire are chosen to reduce resistance/noise which results in physically large and heavy coils.

(45) Another way to increase the sensitivity of an induction coil without increasing resistance is to introduce a soft-magnetic (ferrous) material (core) into the centre of the coil.

(46) Ferrous cores are materials possessing high magnetic permeability and can be used to guide and confine magnetic fields. When introduced to an induction coil, they can greatly enhance the magnetic field strength. Ferrous cores act as flux concentrators within the coil which draw magnetic field lines to themselves, greatly increasing the inductance of the coil.

(47) The inductance of a single-layer coil with a ferrous core is given by the following formula:
L=μ.sub.eμ.sub.0N.sup.2BA
Here, μ.sub.e refers to the effective permeability of the ferrous material in the centre (which is equal to 1 in the case of air).

(48) FIG. 9 shows the increase in inductance for each coil with the introduction of a soft magnetic material core. It can be seen that a 10 mm diameter coil with a core has an equivalent induction to a 50 mm diameter coil without a core despite having 25× less cross-sectional area (and significantly, lower resistance).

(49) However magnetic cores are not without their downsides as they can introduce losses primarily through hysteresis and eddy currents. High permeability alone is not a sufficient for a material to be selected as a magnetic core. Generally speaking materials with low coercivity are preferred as it allows them to respond to changing (AC) fields with lower losses (materials with high coercivity can be considered permanent magnets).

(50) A number of coils that use a magnetic core and that are configured in accordance with embodiments will now be discussed.

(51) Increasing the inductance of a coil has a number of positive effects, including increasing the coil's sensitivity to magnetic fields, and increasing the time constant of the voltage rise time, thereby shifting the frequency response of the coil to lower frequencies (which are more typical of biological signals) and acting as a choke for higher-frequency sources of noise.

(52) According to various embodiments, amorphous metallic alloys (sometimes referred to as metallic glasses or glassy metals) cores are used, e.g. in place of conventional pressed iron powder cores. These materials 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.

(53) By alloying with certain magnetic materials such as iron, cobalt, and nickel, very high magnetic permeability and susceptibility materials are possible, such as Metglas 2714a or FINEMET. Their high resistance reduces eddy current losses when subjected to alternating magnetic fields; their low coercivity also reduces losses.

(54) As such, a core of, for example, Metglas 2714a, nano-crystalline materials (i.e. polycrystalline materials with very small grain sizes, the space between which are filled with amorphous material), or MuMetal, may be used.

(55) The Applicants have recognised that the effective permeability (μ.sub.e) of the magnetic core will depend both of the relative permeability (μ.sub.r) of the magnetic material, and also on the geometry of the core. In particular, the effective permeability (μ.sub.e) depends on the core's geometry-dependent demagnetizing factor N.sub.demag:

(56) ${\mu }_e=\frac{{\mu }_r}{1+N_{\mathrm{demag}}.Math.\left({\mu }_r-1\right)},\mathrm{where}\text{:}$$N_{\mathrm{demag}}\cong \frac{D_C^2}{l_C^2}.Math.\left(\ln \frac{2l_C}{D_C}-1\right).$
Here, D.sub.c and l.sub.c are the diameter and length of the core.

(57) For sufficiently large relative permeabilities, the effective core permeability is almost independent of material properties because this formula simplifies to:

(58) ${\mu }_e\approx \frac{1}{N_{\mathrm{demag}}}$

(59) FIG. 10 illustrates the effect of a core's aspect ratio on its effective permeability, μ.sub.e. As can be seen from FIG. 10, a material with a relative permeability of 10,000 may have an effective permeability of 4 when the length and diameter of the core are equal, or an effective permeability of >1000 when the core is 100 times longer than it is wide.

(60) This can be seen more readily by re-plotting the data from air-core coils (depicted in FIG. 9) and scaling the inductance to be relative to that of the coil without a core present. This is shown in FIG. 11. Here, with the addition of a fixed core length (50.8 mm), but variable core diameter (x-axis), drastically different values of inductance are seen for the same electrical resistance coil. Coils with a core diameter of 5 mm (10:1 aspect ratio) exhibit a ˜17 times increase in measured inductance compared to only a ˜2.5 times increase for coils with a 50 mm diameter core (˜1:1 aspect ratio).

(61) Taking this insight to its logical extreme, the most sensitive coils have a high permeability core, many turns, present a large cross-sectional area and maintain a high aspect ratio. Unfortunately physical constraints mean that it is not practical to produce a coil larger than a certain length and so a compromise must be struck.

(62) In this regard, the outer diameter of the coil, D, should be limited to around 10 cm or less, in order to provide a coil having an overall size that can achieve a spatial resolution that is suitable for medical magnetometry (and in particular for magneto cardiography).

(63) The ratio of the coil's length to its outer diameter should be relatively large (i.e. 0.9 or more), so that the coil is relatively long (along its axis) for its width. This means that the coil can comprise a magnetic core that has a relatively large length to diameter ratio, l.sub.c:D.sub.c, (and accordingly a high effective permeability, μ.sub.e), and accordingly that the coil will have a relatively high inductance, L.

(64) However, the magnetic field strength falls off proportionally with 1/r.sup.3, so turns that are twice as far away from the source of the magnetic field experience a field strength reduced by a factor of 8. Thus, for example, turns 10 cm from the magnetic field source (e.g. the top of heart, or middle, etc.) will experience a magnetic field strength of 12.5% of the strength experienced at 5 cm from the source; turns 15 cm from the source will experience a magnetic field strength of 3.7% the strength experienced at 5 cm, and 29% of the strength experienced at 10 cm; and turns 20 cm from the source will experience a magnetic field strength of 1.56% of the strength experienced at 5 cm, 12.5% of the strength experienced at 10 cm, and 42% of the strength experienced at 15 cm.

(65) This means that turns at the top and bottom of coil experience very different field strengths. This in turn means that there is no benefit in designing a very long coil despite improvements due to the aspect ratio. From these consideration, it was determined that the optimum coil length, l, is ˜50 mm. Beyond this length, the field of the heart weakens and diverges significantly, and the magnetometer device becomes less practical and unwieldy.

(66) Furthermore, the Applicants have recognised that the ratio of the coil's inner diameter to its outer diameter (i.e. the ratio of the inner diameter of the winding(s) to the outer diameter of the winding(s)), D.sub.i:D, should be relatively large, i.e. 0.6 or more. This means that the coil's winding(s) are packed relatively tightly in the direction orthogonal to the core's axis (i.e. have a relatively narrow spread of radial distances from the coil's axis in the direction orthogonal to the coil's axis). This in turns means that the turns of each layer of the coil will be relatively close to the core.

(67) In addition, arranging the outer surface of the core to be in contact with the coil winding(s), means that the winding(s) are as close as possible to the core. Turns of the coil in direct contact or in close proximity with the core receive a large boost to their measured inductance values, but the effect can be almost negligible for the outer turns.

(68) On the other hand, each coil should comprise plural layers of turns, since increasing the number of layers of turns has the effect of increasing the coil's inductance (e.g. without increasing the coil's length, l). However, increasing the number of layers of turns will decrease the ratio of the coil's inner diameter to its outer diameter, D.sub.i:D.

(69) This can be addressed to some degree by using a wire with a smaller gauge and hence cross-sectional area, and so more turns (and thus increased wire length) can be added in the same volume, or an identical number of turns can be placed in close proximity to the core. This is illustrated by FIG. 12.

(70) This change comes at the expense of increased resistance as it is proportional to cross-sectional area (A):

(71) $R=\frac{\rho l}{A}$
However, if the resistance becomes too high, then the Johnson-Nyquist noise can becomes a problem (this can increase the number of cycle averages needed and prolong scan times, or increase the size of the smallest detectable feature), and the amplification electronics need to be suitably modified to ensure sufficient current flow. Though the temperature can be reduced to minimize noise levels, cryogenic refrigerants (such as liquid nitrogen or helium) are not practical, e.g. in terms of cost and safe containment.

(72) The above factors must be carefully balanced when designing coils for use in medical magnetometry. In this regard, the Applicants have found that a particularly balance between the above described competing factors can be found by providing a coil or coils with the following configuration:

(73) $D\approx 4.7\mathrm{cm};$$l\approx 5\mathrm{cm};\mathrm{and}$$\frac{Di}{D}\approx 0.745$
where D is the outer diameter of the coil, l is length of the coil, and D.sub.i is the inner diameter of the coil. This arrangement is depicted in FIG. 13B. Coils having these proportions have been found to have a relatively high inductance, L.

(74) In various other particular embodiments, the or each coil has the following configuration:

(75) $D\approx 4\mathrm{cm};$$l\approx 5\mathrm{cm};\mathrm{and}$$\frac{Di}{D}\approx 0.625$
where D is the outer diameter of the coil, l is length of the coil, and D.sub.i is the inner diameter of the coil. This arrangement is depicted in FIG. 13A. Coils having these proportions have been found to have an even higher inductance, L.

(76) The Applicants have also found that it is not necessary to have a completely solid core, so long as the overall dimensions of the core are maintained. Indeed, a thin strip or ribbon (e.g. <35 μm thick) of Metglas 2714a foil rolled into a hollow cylinder (or e.g. a laminated stack of layers formed into a hollow cylinder) and placed into a coil can yield a similar (or even greater) increase in coil inductance to a ferrite rod of the same overall dimensions because of its high relative permeability (μ.sub.r). This results in significant reductions in both material costs and coil weight. Similar benefits can be obtained using plural laminated layers of foil.

(77) The Applicants have also found that, for these high aspect ratio core shapes, it is important to place the core material in direct contact with (or close to) the windings. This minimizes the potential for leakage inductance and partially serves as a choke to filter out undesired high-frequency noise. As such, the coils may comprise self-supporting bonded coils (i.e. instead of winding onto a bobbin and introducing an “air” gap between core and wire).

(78) The wire used is 0.25 mm copper or copper clad aluminium. By reducing the wire gauge and increasing the length of the coil, many more turns are able to be wound, significantly increasing the inductance of the coil. By using copper clad aluminium, the weight of the coil is significantly reduced (e.g. compared to copper). If the weight of the coil grows too large, then the cost and engineering challenge of safely fixing them above the patient increases. Copper clad aluminium can offer significant (>50%) weight reductions at the price of increased resistance.

(79) The coils according to various embodiments are around 10 times more resistive, so exhibit around 3 times more thermal noise than the coils described in WO2014/006387. The inductance however is around 11 times higher, so the signal to noise ratio is improved by a factor of more than 3.

(80) Although as shown in FIG. 13, the coil's 61 core 70 may have the same length as the winding 71, as illustrated by in FIG. 14, it would also be possible for the core 70 to be longer than the length of the winding 71. This can increase the aspect ratio of the core 70, and so increase its effective permeability. (It should be noted that FIG. 14 is for illustrative purposes only, and is not to scale.)

(81) The presence of a magnetic core significantly increases the inductance of the coils. The use of air cored coils having the configuration described herein to detect biological magnetic fields of interest would necessitate significantly increased scan times in order to obtain the same signal to noise ratio.

(82) It can be seen from the above that the technology described herein, in its 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. With appropriate data treatment, slight patient motion will not significantly degrade the image.

(83) This is achieved, in embodiments of the technology described herein at least, by using an improved design of detection coil that has a particular configuration and that is configured to detect the time varying magnetic of the (e.g.) heart.