Abstract:
A magnetic detection coil is provided, which includes a plurality of differential coils. Each differential coil is made of one of superconductors and metallic materials. The differential coils having mutually different loop directions are arranged in parallel at a spatially predetermined distance apart and mutually electrically connected in series. Each differential coil is one of a first-order differential coil and a second-order differential coil.

Description:
CLAIM OF PRIORITY 
     The present application claims priority on Japanese application JP2005-300991 filed on Oct. 14, 2005, the content of which is hereby incorporated by reference into this application. 
     BACKGROUND OF THE INVENTION 
     The present invention relates to an apparatus for measurement of biomagnetism which employs magnetic detection coils and superconducting quantum interference devices. 
     An apparatus for measurement of biomagnetism, which is used for magnetocardiography and magnetoencephalography, has so far employed a technique that a magnetic detection coil made of superconducting wires detects magnetic signals generated by an organism to transmit to a superconducting quantum interference device (hereinafter referred to as SQUID). A SQUID has a superconductor ring with a Josephson junction, in which a voltage between both ends of the Josephson junction cyclically varies with a period of Φ 0 =h/2e (Wb) according to a magnetic flux penetrating the SQUID. 
     It has been a typical technique in magnetocardiography and magnetoencephalography that a magnetic detection coil made of superconducting wires detects magnetic signals generated by a measurement object as a magnetic flux, which is transmitted to a SQUID. The magnetic detection coil helps to eliminate noise due to environmental magnetic fields to increase a Signal/Noise (S/N) ratio. 
       FIG. 17  is a schematic diagram illustrating architecture of a Flux Locked Loop (FLL) circuit in a typical apparatus for measurement of a magnetic field. 
     In an FLL circuit  1700 , current produced by a magnetic flux penetrating a magnetic detection coil  1701  flows through the magnetic detection coil  1701  and an input coil  1702 . Accordingly, the input coil  1702  creates a magnetic flux, which is transmitted to a SQUID  1703 . The SQUID  1703 , which has a superconducting ring with a Joesphson joint, is supplied a bias current by a bias current source  1705 . A voltage between both ends of the SQUID  1703  cyclically varies with a period of Φ 0 =h/2e (Wb) according to a magnetic flux penetrating the SQUID  1703 . In the FLL circuit  1700 , a feedback circuit, which includes a preamp  1706 , an integrator  1707 , a feedback resistor  1708  and a feedback coil  1704 , is provided in a rear stage of the SQUID  1703 . The feedback coil  1704  feeds back a magnetic flux so as to cancel a change in the magnetic flux penetrating the SQUID  1703 . 
     It is possible to obtain current flowing through the feedback coil  1704  by measuring a potential difference between both ends of the feedback resistor  1708 . The magnetic flux penetrating the SQUID  1703  can be calculated based on this current value. 
     A circuit which has the architecture described above is called FLL circuit. The FLL circuit  1700  provides an output voltage proportional to the magnetic field detected by the magnetic detection coil  1701 . 
     Description is given of a typical magnetic detection coil used for an apparatus for measurement of biomagnetism with reference to  FIG. 18 . 
       FIGS. 18A to 18E  are schematic diagrams illustrating magnetic detection coils used for an apparatus for measurement of biomagnetism. 
       FIG. 18A  shows a zero-order differential magnetic detection coil (magnetometer),  FIG. 18B  a first-order differential magnetic detection coil,  FIG. 18C  a second-order differential magnetic detection coil,  FIG. 18D  a zero-order differential magnetic detection coil formed on a thin film substrate and  FIG. 18E  a first-order differential magnetic detection coil formed on a thin film substrate. 
     As shown in these drawings, a magnetic detection coil typically employs architecture in which superconducting wires are wound around a cylindrical bobbin or the other architecture in which a thin film is formed on a substrate. 
     As shown in  FIG. 18A , a zero-order differential magnetic detection coil  181  has a coil  181   a  which is made of a bobbin  1811  wound by one turn of superconducting wire. A magnetic flux Φ M  in the following equation (1) detected by the zero-order differential magnetic detection coil  181  is represented by a magnetic flux Φ 181a  penetrating the coil  181   a  as follows:
 
Φ M =Φ 181a   (1)
 
     Because the magnetic flux Φ M  is equivalent to the magnetic flux Φ 181a  penetrating the coil  181   a , the zero-order differential magnetic detection coil  181  is able to obtain magnetic signals greater than the first-order and second-order differential magnetic detection coils to be described later. However, the zero-order differential magnetic detection coil does not decrease an effect of environmental magnetic fields at all, directly experiencing noise due to the environmental magnetic fields. Accordingly, the zero-order differential magnetic detection coil  181  is usually used in a magnetically shielded room. 
     In this specification, a distance between centers of coils is referred to as “center-to-center distance”. 
     Given the magnetic flux Φ M  shown in  FIG. 18A  is a positive magnetic signal, current flows through the zero-order differential magnetic detection coil  181  in a direction of a fine arrow, which is drawn along the coil. Hereinafter, a signal representative of a magnetic flux pointing upward is defined as a positive magnetic signal. A direction of current, which flows through a coil when a positive magnetic signal is detected, is represented by a fine arrow. 
     As shown in  FIG. 18B , a first-order differential magnetic detection coil  182  has coils  182   a  and  182   b . The coil  182   a  has one turn of superconducting wire, which is wound around a bobbin  1821  in a first direction. The coil  182   b  has one turn of superconducting wires which is wound around the bobbin  1821  in a second direction opposite to the first direction, lying a predetermined distance vertically apart from the coil  182   a . A magnetic flux Φ G1  in the following equation (2) detected by the first-order differential magnetic detection coil  182  is represented by a magnetic flux Φ 182a  penetrating the coil  182   a  and a magnetic flux Φ 182b  penetrating the coil  182   b  as follows:
 
Φ G1 =Φ 182a −Φ 182b   (2)
 
     The reason why the magnetic flux Φ 182b  has a minus value is that the coil  182   b  is wound in the opposite direction. 
     It should be noted that taking a difference is referred to as “differentiating”, taking a first-order difference as “first-order differentiating” and taking a second-order difference as “second-order differentiating” in this specification. 
     The coil  182   a  is located proximity to a detection object and the coil  182   b  is located relatively away from it. Because environmental magnetic fields of spatial uniformity are cancelled, it is possible to detect only a magnetic flux deriving from the detection object. 
     It should be noted that a vertical direction is meant to represent a direction perpendicular to a plane including a coil and a horizontal direction is meant to represent a direction in parallel with this plane. In this connection, it may be possible to allow the vertical direction to coincide with a direction of measurement of a magnetic field. 
     As shown in  FIG. 18C , a second-order differential magnetic detection coil  183  has coils  183   a ,  183   b  and  183   c . The coil  183   a  has one turn of superconducting wire wound around a bobbin  1831  in a first direction. The coil  183   b  has two turns of superconducting wire wound around the bobbin  1831  in a second direction opposite to the first direction, lying a predetermined distance vertically apart from the coil  183   a . The coil  183   c  has one turn of superconducting wire wound around the bobbin  1831  in the first direction, lying a predetermined distance vertically apart from the coil  183   b . A magnetic flux Φ G2  in the following equation (3) detected by the second-order differential magnetic detection coil  183  is represented by a magnetic flux Φ 183a  penetrating the coil  183   a , a magnetic flux Φ 183b  penetrating the coil  183   b  and a magnetic flux Φ 183c  penetrating the coil  183   c  as follows:
 
Φ G2 =Φ 183a −2Φ 183b +Φ 183c   (3)
 
     As described above, the second-order differential magnetic detection coil  183  differentiates magnetic fluxes in two steps in a vertical direction, thereby decreasing an effect due to both environmental magnetic fields of spatial uniformity and environmental magnetic fields having a first-order gradient. As a result, the second-order differential magnetic detection coil  183  is able to decrease the effect of the environmental magnetic fields more than the first-order differential magnetic detection coil  182  which is only able to decrease the effect of the environmental magnetic fields of spatial uniformity. When a magnetic signal is included in the magnetic flux Φ 183b  penetrating the coil  183   b  and the magnetic flux Φ 183c  penetrating the coil  183   c , a magnetic signal detected by the second-order differential magnetic detection coil  183  will decrease. It is understood that the higher order a differential magnetic detection coil possesses, the less effect of environmental magnetic fields will exist. However, a detected magnetic signal will decrease accordingly. In this way, a tradeoff study is necessary to solve the irreconcilable situations described above. A magnetometer, a first-order differential magnetic detection coil, or a second-order differential magnetic detection coil has been so far typically used for biomagnetism in conjunction with a magnetically shielded room according to magnitude of environmental magnetic fields. 
     There is a technique to use a superconducting thin film instead of superconducting wires for a magnetic detection coil.  FIG. 18D  shows a zero-order differential magnetic detection coil  184 , in which a superconducting thin film is formed on a substrate  1841 . A magnetic flux Φ 184a  detected by a coil  184   a  is transmitted to a SQUID  1842  formed on the same substrate  1841 .  FIG. 18E  shows a first-order differential magnetic detection coil  185 , in which coils  185   a  and  185   b  having directions opposite to each other are formed on a substrate  1851 . A difference Φ 185a −Φ 185b  between a magnetic flux Φ 185a  detected by the coil  185   a  and a magnetic flux Φ 185b  detected by the coil  185   b  is transmitted to a SQUID  1852  formed on the same substrate  1851 . An advantage of using the superconducting thin film is that it is possible to determine and materialize an accurate area of a magnetic detection coil. 
     Several types of arrangements for magnetic detection coils have been proposed taking into account characteristic features of the differential coils described above. As an example, a technique has been proposed, in which plural types of magnetic detection coils having different differential orders are placed at a measurement point so as to calculate and estimate magnetic field sources or a distribution of the magnetic field sources in an organism. Patent document No. 1: Japanese Published Patent Application 09-084777 (claim 1, paragraph 0015, FIG. 1). 
     However, as shown in  FIGS. 18A to 18C , there have been limited arrangements in which only a magnetic field differentiated in a certain direction is detected. These arrangements have a problem that when environmental magnetic fields are strong because a magnetically shielded room is not available, for example, it is not possible to adequately reduce an effect of them. One possible technique is to increase order for a differential magnetic detection coil so as to decrease this effect. Although the effect can be decreased by this technique, it is inevitably accompanied by a reduction in a magnetic signal to be detected. 
     A magnetic detection coil using a superconducting thin film has a problem that it is intrinsically difficult to form a coil three-dimensionally. 
     In addition, it appears to be unproductive to combine these two types of coils, because magnetic detection coils with three-dimensional structure as shown in  FIGS. 18A to 18C  and magnetic detection coils formed on a superconducting thin film as shown in  FIGS. 18D and 18E  are different from each other in terms of usage, structure and manufacturing processes. 
     SUMMARY OF THE INVENTION 
     In view of the problems described above, the present invention has an object to provide a magnetic detection coil and an apparatus for measurement of a magnetic field, which acquire increased S/N ratios as a result of not only avoiding a decrease in detection sensitivity for magnetic signals but also decreasing an effect due to environmental magnetic fields. 
     It is an aspect of the present invention to provide a magnetic detection coil including a plurality of differential coils. Each differential coil is made of one of a superconductor and metallic materials. The differential coils having mutually different loop directions are arranged in parallel at spatially predetermined intervals and mutually electrically connected. Each differential coil is one of a first-order differential coil and a second-order differential coil. 
     The present invention provides the magnetic detection coil and the apparatus for measurement of a magnetic field, which has the advantage of increased S/N ratios as a result of not only avoiding a decrease in detection sensitivity for magnetic signals but also decreasing an effect due to environmental magnetic fields. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a perspective view showing a magnetic detection coil according to an embodiment (case 1 of the present invention). 
         FIG. 2  is a perspective view showing a magnetic detection coil according to an embodiment (case 2 of the present invention). 
         FIG. 3A  is a graph showing simulation results for a distribution of signal strength detected by a zero-order differential magnetic detection coil shown in  FIG. 18A .  FIG. 3B  is a graph showing a projection of  FIG. 3A  on an x-B z  plane. 
         FIG. 4A  is a graph showing simulation results for a distribution of signal strength detected by a first-order differential magnetic detection coil shown in  FIG. 18B .  FIG. 4B  is a graph showing a projection of  FIG. 4A  on an x-B z  plane. 
         FIG. 5A  is a graph showing simulation results for a distribution of signal strength detected by a second-order differential magnetic detection coil shown in  FIG. 18C .  FIG. 5B  is a graph showing a projection of  FIG. 5A  on an x-B z  plane. 
         FIG. 6A  is a graph showing simulation results for a distribution of signal strength detected by a magnetic detection coil shown in  FIG. 1 .  FIG. 6B  is a graph showing a projection of  FIG. 6A  on an x-B z  plane. 
         FIG. 7A  is a graph showing simulation results for a distribution of signal strength detected by a magnetic detection coil shown in  FIG. 2 .  FIG. 7B  is a graph showing a projection of  FIG. 7A  on an x-B z  plane. 
         FIG. 8  is a graph showing experimental results which demonstrate advantages provided by a magnetic detection coil according to an embodiment of the present invention. 
         FIG. 9  is a schematic diagram illustrating an apparatus for measurement of a magnetic field used for an experiment of a magnetic detection coil according to an embodiment of the present invention. 
         FIG. 10  is a schematic diagram illustrating an apparatus for measurement of a magnetic field according to an embodiment of the present invention. 
         FIG. 11A ,  FIG. 11B  and  FIG. 11C  are graphs showing histories of output  1 , output  2  and output  3  shown in  FIG. 9 , respectively. 
         FIG. 12  is a perspective view showing an arrangement of magnetic detection coils according to an embodiment of the present invention. 
         FIG. 13A  is a top view schematically illustrating a magnetic detection coil pair shown in  FIG. 12 .  FIG. 13B  illustrates relationship among a current vector of a magnetic field source and magnetic flux densities in a z-direction.  FIG. 13C  is a schematic diagram illustrating an arrangement of 64 magnetic detection coil pairs, each pair the same as that shown in  FIG. 12 , in a form of 8 by 8 lattices. 
         FIG. 14  is a perspective view illustrating overall architecture for an apparatus for measurement of a magnetic field according to an embodiment of the present invention. 
         FIG. 15  is a perspective view illustrating an apparatus for magnetocardiography of an unborn child according to an embodiment of the present invention. 
         FIG. 16  is a perspective view illustrating an apparatus for magnetoencephalography according to an embodiment. 
         FIG. 17  is a schematic diagram illustrating architecture of a Flux Locked Loop (FLL) circuit in a typical apparatus for measurement of a magnetic field. 
         FIG. 18A  shows a zero-order differential magnetic detection coil (magnetometer).  FIG. 18B  shows a first-order differential magnetic detection coil.  FIG. 18C  shows a second-order differential magnetic detection coil.  FIG. 18D  shows a zero-order differential magnetic detection coil formed on a thin film substrate.  FIG. 18E  shows a first-order differential magnetic detection coil formed on a thin film substrate. 
     
    
    
     DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     An embodiment of the present invention is now described in detail with reference to the drawings. In the drawings referenced hereinafter, items having a same function are given a same reference number. 
     Low-temperature superconducting materials or high-temperature superconducting materials can be used as superconducting materials for a magnetic detection coil used in an apparatus described in the following embodiment. A low-temperature superconducting material, which has a low superconductor transition temperature, works as a superconductor in a low-temperature (liquid helium temperature) environment. In contrast, a high-temperature superconducting material, which has a high superconductor transition temperature, works as a superconductor in a high-temperature (liquid nitrogen temperature) environment. It may be possible to alternatively adopt a superconducting material having a superconductor transition temperature between the liquid helium temperature and the liquid nitrogen temperature or a superconducting material having a higher superconductor transition temperature than the liquid nitrogen temperature. It is understood that materials having high electric conductivity such as copper can be used for a member of a magnetic detection coil. 
     a. Magnetic Detection Coil—1 
       FIG. 1  is a perspective view showing a magnetic detection coil according to an embodiment. 
     A magnetic detection coil  1  includes coils  1   a ,  1   b ,  1   c  and  1   d , and bobbins  11   a  and  11   b . The coil la is formed by one turn of superconducting wire wound around the bobbin  11   a  in a first direction. The coil  1   b  is formed by one turn of superconducting wire wound around the bobbin  11   a  in a second direction opposite to the first direction, lying vertically a predetermined distance apart from the coil  1   a . The coil  1   c  is formed by one turn of superconducting wire wound around the bobbin  11   b  in the second direction, which is positioned horizontally a predetermined distance from the coil  1   a . The coil  1   d  is formed by one turn of superconducting wire wound around the bobbin  11   b  in the first direction, lying vertically a predetermined distance apart from the coil  1   c . To summarize, the magnetic detection coil  1  is made of one continuous wire. The coils  1   a  and  1   c  are located in a same plane and the coils  1   b  and  1   d  are located in a same plane. In other words, first-order differential coils are arranged in parallel spaced a predetermined distance apart. A magnetic flux Φ P1  in the following equation (4) detected by the magnetic detection coil  1  is represented by a magnetic flux Φ 1a  penetrating the coil  1   a , a magnetic flux Φ 1b  penetrating the coil  1   b , a magnetic flux Φ 1c  penetrating the coil  1   c  and a magnetic flux Φ 1d  penetrating the coil  1   d  as follows:
 
Φ P1 =(Φ 1a −Φ 1b )−(Φ 1c −Φ 1d )  (4)
 
     The magnetic detection coil  1  according to this embodiment is a magnetic detection coil which has a first-order differential function not only in an axial direction (vertical direction) with respect to the bobbin  11   a  (first term) and the bobbin  11   b  (second term) but also in a horizontal direction. Because the magnetic detection coil  1  detects magnetic signals which are first-order differentiated in both vertical and horizontal directions, it is possible to reduce an effect due to environmental magnetic fields more than the first-order differential magnetic detection coil  182  shown in  FIG. 18B . 
     b. Magnetic Detection Coil—2 
       FIG. 2  is a perspective view showing a magnetic detection coil according to an embodiment. 
     A magnetic detection coil  2  includes coils  2   a ,  2   b ,  2   c ,  2   d ,  2   e  and  2   f , and bobbins  21   a  and  21   b . The coil  2   a  is formed by one turn of superconducting wire wound around the bobbin  21   a  in a first direction. The coil  2   b  is formed by two turns of superconducting wire wound around the bobbin  21   a  in a second direction opposite to the first direction, lying vertically a predetermined distance apart from the coil  2   a . The coil  2   c  is formed by one turn of superconducting wire wound around the bobbin  21   a  in the first direction, lying vertically a predetermined distance apart from the coil  2   b . The coil  2   d  is formed by one turn of superconducting wire wound around the bobbin  21   b  in the second direction, which is positioned horizontally a predetermined distance from the coil  2   a . The coil  2   e  is formed by two turns of superconducting wire wound around the bobbin  21   b  in the first direction, lying vertically a predetermined distance apart from the coil  2   d . The coil  2   f  is formed by one turn of superconducting wire wound around the bobbin  21   b  in the second direction, lying vertically a predetermined distance apart from the coil  2   e . To summarize, the magnetic detection coil  2  is made of one continuous wire. Both pair of the coils  2   a  and  2   d , the coils  2   b  and  2   e , and the coils  2   c  and  2   f  are located on a same plane, respectively. In other words, second-order differential coils are arranged in parallel spaced a predetermined distance apart. A magnetic flux Φ P2  in the following equation (5) detected by the magnetic detection coil  2  is represented by a magnetic flux Φ 2a  penetrating the coil  2   a , a magnetic flux Φ 2b  penetrating the coil  2   b , a magnetic flux Φ 2c  penetrating the coil  2   c , a magnetic flux Φ 2d  penetrating the coil  2   d , a magnetic flux Φ 2e  penetrating the coil  2   e  and a magnetic flux Φ 2f  penetrating the coil  2   f  as follows:
 
Φ P2 =(Φ 2a −2Φ 2b +Φ 2c )−(Φ 2d −2Φ 2e +Φ 2f )  (5)
 
     The magnetic detection coil  2  according to this embodiment is one which has not only a second-order differential function in an axial direction (vertical direction) with respect to the bobbin  21   a  (first term) and the bobbin  21   b  (second term), but also a first-order differential function in a horizontal direction. Because the magnetic detection coil  2  detects magnetic signals which are second-order differentiated in the vertical direction and first-order differentiated in the horizontal direction, it is possible to reduce an effect due to environmental magnetic fields more than the second-order differential magnetic detection coil  183  shown in  FIG. 18C . 
     It should be noted that a circular shape of the coil used in the magnetic detection coils shown in  FIG. 1  and  FIG. 2  is an example and it may be possible to alternatively select other shapes such as a polygonal shape. 
     c. Simulation Results 
     Description is given of advantages in terms of signal strength of magnetic signals (hereinafter referred to as “signal strength”) detected by the magnetic detection coils according to the embodiments with reference to  FIGS. 3 to 7 . 
     A current dipole of a cardiac muscle of a typical thirty-week unborn child is assumed as a magnetic signal source in a simulation. (See A. Kandori, T. Miyashita, and K. Tsukada, “A vector fetal magnetocardiogram system with high sensitivity” Review of Scientific Instruments USA, December 1999, Volume 70, P.4702). Assuming that current concentrates at one point r 0 =(0,0, −z 0 ), a current dipole Q is defined by the following equation (6), where J(r) is a current density at r=(x, y, z) (See “Basic mathematical and electromagnetic concepts of the biomagnetic inverse problem” J. Sarvas, Physics in Medicine and Biology, January 1987, Volume 32 P.11).
 
 J ( r )=δ( r−r   0 ) Q   (6)
 
δ(r−r 0 ) represents a delta function. A magnetic flux density B(r) due to the current dipole Q is represented by the following equation (7).
 
 B ( r )=(μ 0 /4π) Q ×( r−r   0 )/| r−r   0 | 3   (7)
 
     Values are set as follows: Q=(0, 250, 0) (nA·m), z 0 =80 (mm) and a vertical distance between neighboring coils is equal to 50 (mm). z 0  represents a distance from the magnetic detection coil to the current dipole Q, which is hereinafter meant to denote a distance between a plane made by a coil of the magnetic detection coil and the current dipole Q. μ 0 =4π/10 7  is a magnetic permeability in a vacuum condition. 
     A description is given of signal strength detected by a zero-order differential magnetic detection coil with reference to  FIG. 3A  and  FIG. 3B  along with  FIG. 18A . 
       FIG. 3A  is a graph showing a distribution of signal strength detected by a zero-order differential magnetic detection coil shown in  FIG. 18A .  FIG. 3B  is a graph showing a projection of  FIG. 3A  on an x-B z  plane. 
     In  FIGS. 3A and 3B , a central position of the coil  181   a  of the zero-order differential magnetic detection coil  181  (see  FIG. 18A ) is set equal to (x, y, 0) (m). This means that the signal strength is equivalent to a distribution B z (x, y, 0), which is a distribution of a z-component B z (r) of the magnetic flux density B(r) of the equation (7) on an xy-plane. Partially differentiating B z (x, y, 0) with respect to x, which has maximum and minimum values, a distribution of magnetic field generated by the current dipole Q is represented by the following equation (8).
 
∂ B   z   /∂x =(μ Q /4π)(√{square root over (2 x )}+√{square root over (( y   2   +z   0   2 ))}) (√{square root over (2 x )}−√{square root over (( y   2   +z   0   2 ))})/( x   2   +y   2   +z   0   2 ) 5/2   (8)
 
     According to the equation (8), B z  takes a maximum value at x=(y 2 +z 0   2 )/2 and a minimum value at x=√{square root over ((y 2 +z 0   2 )/2)}, respectively. 
     Substituting x=−√{square root over ((y 2 +z 0   2 )/2)} and x=√{square root over ((y 2 +z 0   2 )/2)} for B z (x, y, 0), equations (9) and (10) are respectively obtained.
 
 B   z (−√{square root over (( y   2   +z   0   2 )/2)},  y , 0)=(μ Q /4π)(2/3√{square root over (3)})/( y   2   +z   0   2 )  (9)
 
 B   z (√{square root over (( y   2   +z   0   2 )/2)},  y , 0)=−(μ Q /4π)(2/3√{square root over (3)})/( y   2   +z   0   2 )  (10)
 
     According to the equations (9) and (10), B z (x, y, 0) takes a maximum value B z   max  at P max =(−z 0 /√{square root over (2)}, 0, 0) and a minimum value B z   min  at P min =(z 0 /√{square root over (2)}, 0, 0). Signal strengths are represented by equations (11) and (12), respectively.
 
 B   z   max =(μ Q/ 4π)(2/3√{square root over (3)})/ z   0   2   (11)
 
 B   z   min =−(μ 0   Q/ 4π)(2/3√{square root over (3)})/ z   0   2   (12)
 
     Reading out signal strengths at P max  and P min  from  FIG. 3B , it is known that B z   max  and B z   min  are equal to 1.5 pT and −1.5 pT, respectively. 
     A center-to-center distance d of the points P max  and P min  is represented by the following equation.
 
d=√{square root over (2z 0 )}  (13)
 
     A description is given of a signal strength detected by a first-order differential magnetic detection coil with reference to  FIG. 4A  and  FIG. 4B  along with  FIG. 18B . 
       FIG. 4A  is a graph showing simulation results for a distribution of signal strength detected by a first-order differential magnetic detection coil shown in  FIG. 18B .  FIG. 4B  is a graph showing a projection of  FIG. 4A  on an x-B z  plane. 
     In  FIGS. 4A and 4B , a current dipole of a cardiac muscle of a typical thirty-week unborn child is assumed as a magnetic signal source in a simulation and a central position of the coil  182   a  of the first-order differential magnetic detection coil  182  is set equal to (x, y, 0) (m), similarly with  FIGS. 3A and 3B . 
     It is known from  FIG. 4B  that a maximum signal strength detected by the first-order differential magnetic detection coil  182  is substantially equal to 1 pT. 
     Description is given of a signal strength detected by a second-order differential magnetic detection coil with reference to  FIG. 5A  and  FIG. 5B  along with  FIG. 18C . 
       FIG. 5A  is a graph showing simulation results for a distribution of signal strength detected by a second-order differential magnetic detection coil shown in  FIG. 18C .  FIG. 5B  is a graph showing a projection of  FIG. 5A  on an x-B z  plane. 
     Similarly with  FIG. 3A  and  FIG. 3B , a current dipole of a cardiac muscle of a typical thirty-week unborn child is assumed as a magnetic signal source in a simulation and a central position of the coil  183   a  of the second-order differential magnetic detection coil  183  is set equal to (x, y, 0) (m). It is known from  FIG. 5B  that a maximum signal strength detected by the second-order differential magnetic detection coil  183  is substantially equal to 0.75 pT. 
     According to  FIG. 3B ,  FIG. 4B  and  FIG. 5B , it is known that the signal strength detected by the first-order differential magnetic detection coil  182  is lower than the zero-order differential magnetic detection coil  181 , and the signal strength detected by the second-order differential magnetic detection coil  183  is lower than the first-order differential magnetic detection coil  182 . 
     A description is given of a signal strength detected by a magnetic detection coil  1  shown in  FIG. 1  with reference to  FIG. 6A  and  FIG. 6B . 
       FIG. 6A  is a graph showing simulation results for a distribution of signal strength detected by a magnetic detection coil  1  shown in  FIG. 1 .  FIG. 6B  is a graph showing a projection of  FIG. 6A  on an x-B z  plane. 
     The center-to-center distance d=√{square root over (2z 0 )} (see equation (13)) between centers P 1a  and P 1c  of the coils  1   a  and  1   c  is set substantially equal to 113 (mm) so that an output of the magnetic detection coil  1  according to this embodiment becomes maximal. Also, a middle point of a line connecting the points P 1a  and P 1c  is set equal to (x, y, 0) (m). It is observed that the signal strength detected by the magnetic detection coil  1  takes a maximum value at x=0 and y=0, right above the dipole Q. The maximum value obtained from  FIG. 6B  is substantially equal to 2 pT. 
     A description is given of a signal strength detected by a magnetic detection coil  2  shown in  FIG. 2  with reference to  FIG. 7A  and  FIG. 7B . 
       FIG. 7A  is a graph showing simulation results for a distribution of signal strength detected by a magnetic detection coil shown in  FIG. 2 .  FIG. 7B  is a graph showing a projection of  FIG. 7A  on an x-B z  plane. 
     The center-to-center distance d=√{square root over (2z 0 )} (see equation (13)) between centers P 2a  and P 2d  of the coils  2   a  and  2   d  is set substantially equal to 113 (mm) so that an output of the magnetic detection coil  2  becomes maximal. Also, a middle point of a line connecting the points P 2a  and P 2d  is set equal to (x, y, 0) (m). It is observed that the signal strength detected by the magnetic detection coil  2  takes a maximum value at x=0 and y=0, right above the dipole Q. The maximum value obtained from  FIG. 7B  is substantially equal to 1.5 pT. 
     As described above, the magnetic detection coil  1  shown in  FIG. 1  is able to detect a magnetic signal twice as large as that detected by the first-order differential magnetic detection coil  182  shown in  FIG. 18B . 
     Because the magnetic detection coil  1  is expected to decrease an effect of environmental magnetic fields more efficiently than the first-order differential magnetic detection coil  182 , it is understood that the magnetic detection coil  1  is able to provide a higher S/N ratio than the first-order differential magnetic detection coil  182 . 
     A magnetic detection coil  1  or  2  is practically applied to an FLL circuit as a magnetic detection coil, as shown in  FIGS. 9 and 10 . 
     A description is given of advantages of the magnetic detection coil according to this embodiment with reference to  FIG. 8  along with  FIGS. 2  and  18 . 
       FIG. 8  is a graph showing experimental results which demonstrate advantages given by the magnetic detection coil according to this embodiment. 
     In  FIG. 8 , a vertical axis represents strength (T/Hz 1/2 ) of field noise detected by a magnetic detection coil and a horizontal axis a frequency (Hz) of the field noise. 
     A curve (a) in  FIG. 8  represents a frequency characteristic of output from a fluxgate magnetometer outside a magnetic shield, namely a frequency characteristic of environmental magnetic fields. A curve (b) in  FIG. 8  represents a frequency characteristic of output from a second-order differential magnetic detection coil  183  of  FIG. 18C  outside a magnetic shield. Diameters of coils  183   a ,  183   b  and  183   c  of the second-order differential magnetic detection coil  183  were set equal to 18 mm, respectively. A vertical distance between the two neighboring coils was set equal to 50 mm. A curve (c) in  FIG. 8  represents a frequency characteristic of output from a magnetic detection coil  2  shown in  FIG. 2  outside a magnetic shield. Diameters of coils  2   a  to  2   f  of the magnetic detection coil  2  used for an experiment were set equal to 18 mm, respectively. A vertical distance between the two neighboring coils was set equal to 50 mm. In addition, a center-to-center distance between bobbins  21   a  and  21   b  was set equal to 25√{square root over (2)} mm. A curve (d) in  FIG. 8  represents a frequency characteristic of output from a second-order differential magnetic detection coil  183  inside a magnetic shield, which is the same as the magnetic detection coil used for calculating the curve (b) of  FIG. 8 . 
     A reduction rate Sp (dB) is defined by an equation (14), where a magnetic flux density calculated from an output of a magnetic detection coil is denoted as B p  and a magnetic flux density of environmental fields as B a . The reduction rate acquired for a frequency band 0.5-49 Hz by the second-order differential magnetic detection coil  183  was 32-40 dB as shown by curve (b) of  FIG. 8 . On the other hand, the reduction rate acquired by the magnetic detection coil  2  was 41-58 dB as shown by curve (c) of  FIG. 8 . The reduction rate acquired by the second-order differential magnetic detection coil  183  with a magnetic shield was 54-83 dB as shown by curve (d) of  FIG. 8 .
 
 S   p =20 log| B   a   /B   p |  (14)
 
     The reason why reduction rates differ from one another according to frequencies is that a distance between a source of an environmental magnetic field and a magnetic detection coil depends on a frequency. Generally speaking, the more apart from a magnetic detection coil a magnetic field source is located, the less steep magnetic gradient the magnetic field source has. A reduction rate of the magnetic detection coil tends to be higher, accordingly. Introducing a value, which is obtained by integration of field noise over a bandwidth of 0.5-49 Hz, a noise reduction rate of 40 dB was calculated for the second-order differential magnetic detection coil  183  as shown by the curve (b) in  FIG. 8 . This value corresponds to a magnetic signal value, which results from an output of a magnetic detection coil filtered by a bandpass filter having a passing band of 0.5-49 Hz. In contrast, a noise reduction rate of the magnetic detection coil  2  was 54 dB as shown by curve (c) of  FIG. 8 . Also, a noise reduction rate for a combination of a magnetic shield and the second-order differential magnetic detection coil  183  was 73 dB as shown by curve (d) of  FIG. 8 . The results described above lead to an observation that the noise reduction rate of the magnetic detection coil  2  is higher by 14 dB than that of the second-order differential magnetic detection coil  183 . 
     Description is given of advantages given by a magnetic detection coil  2  shown in  FIG. 2  with reference to  FIGS. 9 to 11  along with  FIGS. 2 and 18 . 
       FIG. 9  is a schematic diagram illustrating an apparatus for measurement of a magnetic field, which is used for an experiment demonstrating advantages provided by the magnetic detection coil according to this embodiment. 
     The objective of this experiment is to compare wave shapes of magnetic signals detected by magnetic detection coils in magnetocardiography, which is carried out for adults using a second-order differential magnetic detection coil  183  and a magnetic detection coil  900  of this embodiment without a magnetic shield. 
     Magnetic fluxes detected by two pieces of second-order differential magnetic detection coils  901  and  902  are transmitted to SQUID substrates  903  and  904 , respectively. It should be noted that each of the SQUID substrates  903  and  904  includes an input coil  1702 , a SQUID  1703  and a feedback coil  1704  shown in  FIG. 17 . The second-order differential magnetic detection coils  901  and  902  each have a diameter 18 mm and a vertical distance 50 mm between neighboring coil loops. A center-to-center distance between the second-order differential magnetic detection coils  901  and  902  is set equal to 25√{square root over (2)} mm. This means that two pieces of second-order differential magnetic detection coils are arranged in parallel. The second-order differential magnetic detection coils  901  and  902  and the SQUID substrate  903  and  904  are cooled in a cryostat  905 . Liquid helium is charged in the cryostat  905 , which is thermally insulated by a vacuum insulation layer. The SQUID substrates  903  and  904  are controlled by FLL circuits  906  and  907 , respectively. An output from the FLL circuit  906  and an output from the FLL circuit  907 , which are transformed by analogue to digital (AD) converters  908  and  909 , respectively, are sent to a digital signal processor (DSP)  910 , where the output undergoes real-time digital signal processing. In digital signal processing conducted by the DSP  910 , a differential element  911  generates a difference between two input signals. Filters  912 ,  913  and  914  each include a notch filter for eliminating noise in frequencies of a commercial power supply and a bandpass filter having a passing band of 1-50 Hz. In this connection, output  1  and output  2  are magnetic signals, which are detected by the second-order differential magnetic detection coils  901  and  902  respectively, and subjected to a filtering process. Output  3  is a differential signal between the magnetic signals detected by the second-order differential magnetic detection coils  901  and  902 , and it is subjected to a filtering process. In this way, the output  3  in  FIG. 9  corresponds to the magnetic signal detected by the magnetic detection coil  2  in  FIG. 2  which is subjected to a filtering process. 
       FIG. 10  is a schematic diagram illustrating an apparatus for measurement of a magnetic field according to an embodiment. 
     Magnetic fluxes detected by a magnetic detection coil  1001  having the similar architecture to the magnetic detection coil  2  shown in  FIG. 2  is transmitted to a SQUID substrate  1002 . The SQUID substrate  1002  includes an input coil  1702 , a SQUID  1703  and a feedback coil  1704  shown in  FIG. 17 . The magnetic detection coil  1001  and the SQUID substrate  1002  are cooled in a cryostat  1003 . Liquid helium is charged in the cryostat  1003 , which is thermally insulated by a vacuum insulation layer. The SQUID substrate  1002  is controlled by an FLL circuit  1004 . An output from the FLL circuit  1004 , which is transformed by an AD converter  1005 , is sent to a DSP  1006 , where the output undergoes real-time digital signal processing. Magnetic fluxes detected by the magnetic detection coil  1001  correspond to a difference between fluxes detected by the second-order differential magnetic detection coils  901  and  902  shown in  FIG. 9 . It is known from the simulation results shown in  FIGS. 4 to 6  described above that the magnetic detection coil  1001  is able to detect larger magnetic signals than the second-order differential magnetic detection coils  901  and  902 . The magnetic coil  1001  is able not only to decrease noise more but also to detect larger magnetic signals than the second-order differential magnetic detection coils  901  and  902 . In this way, the magnetic detection coil  1001  is able to provide a higher S/N ratio for detection of magnetic signals compared with the second-order differential magnetic detection coils  901  and  902 . 
     It should be noted that the apparatus for measurement of a magnetic field shown in  FIGS. 9 and 10  can be applied to practical use shown in  FIGS. 14 to 16  to be described later. 
       FIG. 11A  to  FIG. 11C  are graphs showing results of magnetocardiography conducted for an adult by an apparatus for measurement of a magnetic field shown in  FIG. 9 . In each graph, a vertical axis represents magnetic flux density (pT) and a horizontal axis represents time (sec). 
       FIG. 11A ,  FIG. 11B  and  FIG. 11C  are graphs showing histories of the output  1 , the output  2  and the output  3  shown in  FIG. 9 , respectively. 
     Although a peak of QRS wave is observable in cardiomagnetic wave shapes shown in  FIG. 11A  and  FIG. 11B , noise of about some tens of hertz and fluctuations with respect to a baseline curve are prominently observed. This is due to the fact that an effect of environmental magnetic fields is not sufficiently decreased. In contrast, cardiomagnetic wave shapes are clearly detected in  FIG. 11C . The noise of some tens of hertz and the fluctuations with respect to the baseline curve, which appear in  FIGS. 11A and 11B , are decreased by differentiation of in plane direction of the magnetic detection coil. In this way, T waves in addition to QRS waves are clearly observed. The experimental results described above demonstrate that the magnetic detection coil  2  of  FIG. 2  possesses a higher S/N ratio than the second-order differential magnetic detection coil  183  of  FIG. 18C  and it is able to obtain clear cardiomagnetic wave shapes even if measurement is conducted without a magnetic shield. 
     Because an effect due to environmental magnetic fields is decreased more for magnetic signals (output  3  of  FIG. 9 ) than for those (output  1  and output  2  of  FIG. 9 ) obtained by the individual second-order differential magnetic detection coils  901  and  902 , the magnetic detection coil  900  is able to transmit magnetic signals with reduced field noise. 
     Wave shapes obtained from output of the apparatus for measurement of a magnetic field shown in  FIG. 10  are substantially the same as those obtained from the output  3  of the apparatus for measurement of a magnetic field shown in  FIG. 9 . Because the apparatus for measurement of a magnetic field shown in  FIG. 10  reduces the number of components such as a SQUID, an FLL circuit and an AD converter as well as signal processing to half compared with its counterpart shown in  FIG. 9 , it is possible to obtain signals with high S/N ratios at a lower cost. 
     In  FIG. 9 , it may be possible to alternatively adopt a first-order differential magnetic detection coil in place of the second-order differential magnetic detection coil. In  FIG. 10 , it may also be possible to alternatively adopt a magnetic detection coil  1  of  FIG. 1  in place of a magnetic detection coil  2  of  FIG. 2 . 
     d. Arrangement of Magnetic Detection Coils 
     A description is given of an example for arrangement of magnetic detection coils according to an embodiment with reference to  FIG. 12  and  FIG. 13  along with  FIG. 2 . 
       FIG. 12  is a perspective view showing an arrangement of magnetic detection coils according to this embodiment. 
     Magnetic detection coils  12   a  and  12   b  each have the similar architecture as a magnetic detection coil  2  shown in  FIG. 2 . In other words, the magnetic detection coils  12   a  and  12   b  each have a pair of differential magnetic detection coils. The magnetic detection coil  12   a  includes a coil  1201  of second-order differential type and a coil  1202  of the same type having a winding direction opposite to the coil  1201 . Similarly, the magnetic detection coil  12   b  includes a coil  1203  of second-order differential type similar to the architecture of  FIG. 18C  and a coil  1204  of the same type having a winding direction opposite to the coil  1201 . A pair of the magnetic detection coils  12   a  and  12   b  is referred to as a magnetic detection coil pair  12 . It should be noted that the magnetic detection coils  12   a  and  12   b  are arranged so that their directions of first-order differential in a horizontal direction are perpendicular to each other. 
       FIG. 13A  is a top view schematically illustrating a magnetic detection coil pair of  FIG. 12 .  FIG. 13B  illustrates a relationship among a current vector I of a magnetic field source and magnetic flux densities B 1  and B 2  in a z-direction. The magnetic flux density B 1  detected by the magnetic detection coil  12   a  is differentiated in an x-axis direction. The magnetic flux density B 2  detected by the magnetic detection coil  12   b  is differentiated in a y-axis direction.  FIG. 13C  is a schematic diagram illustrating an arrangement of 64 magnetic detection coil pairs, each pair the same as that shown in  FIG. 12 , in a form of 8 by 8 lattices. 
     Generally speaking, when current such as a cardio muscle current flows in an x-axis direction, a magnetic field generated by the current is obtained through magnetic signals detected by a magnetic detection coil, which is differentiated in a y-axis direction. If a magnetic detection coil differentiated in an x-axis direction is used on the other hand, almost no signals are detected. This means that when a magnetic detection coil horizontally differentiated is used, it is preferable to select a magnetic detection coil which is differentiated in a direction perpendicular to that of a current of a magnetic field source. Because a direction of current to be measured, such as a cardio muscle current, is not known in advance, it is preferable to arrange two pieces of magnetic detection coils  2  according to this embodiment so that they perpendicularly intersect each other, similarly with the magnetic detection coil pair  12  shown in  FIG. 12 . 
     It is possible to calculate a vector sum of the magnetic flux density B 1  detected by the magnetic detection coil  12   a  and the magnetic flux density B 2  detected by the magnetic detection coil  12   b  by an equation (15).
 
 B   0 =√{square root over ( B   1   2   +B   2   2 )}  (15)
 
     The equation (15) enables reliable detection of a magnetic field generated by a current source irrespective of a direction of the current source to be measured. 
     When a current vector of a magnetic field source is denoted as I=(I x , I y ), an x-component I x  and a y-component I y  of the current are approximated by an equation (16) using changes ΔB z /Δx and ΔB z /Δy in magnetic flux densities. ΔB z /Δx is a change in a magnetic flux density in a z-axis direction, which is first-order differentiated in an x-axis direction. ΔB z /Δy is a change in a magnetic flux density in a z-axis direction, which is first-order differentiated in a y-axis direction (see H. Hosaka and D. Cohen, “Visual determination of generators of the magnetocardiogram” Journal of Electrocardiology USA, 1976, Volume 9, pp. 426-432).
 
(I x , I y )∝(−ΔB z /Δy, ΔB z /Δx)  (16)
 
     An x-component I x  and y-component I y  of the current of a magnetic field source are approximated by an equation (17) using the magnetic flux density B 1  differentiated in an x-axis direction detected by the magnetic detection coil  12   a  and the magnetic flux density B 2  differentiated in a y-axis direction detected by the magnetic detection coil  12   b.  
 
(I x , I y )∝(−B 2 , B 1 )  (17)
 
     In this way, the magnetic detection coil pair  12  is able to detect the current of the magnetic field source approximately as a current vector. 
     As shown in  FIG. 13B , the current can be represented as a vector by the magnetic flux densities B 1  and B 2  detected by the magnetic detection coils  12   a  and  12   b , respectively. 
     As described above, it is possible to detect a distribution of a magnetic field by arranging a plurality of magnetic detection coil pairs  12 . In addition, the equation (17) allows detection of a distribution of current vectors of a magnetic field source (current vector field). In this way, it is possible to estimate a location where a current of cardio muscle flows without worrying about a direction of the current for magnetocardiography. It is also possible to estimate a location where a neural current flows without worrying about a direction of the current for magnetoencephalography. 
     An apparatus  1400  for measurement of a magnetic field with magnetic detection coils according to an embodiment is now described with reference to  FIG. 14  as well as  FIGS. 1 and 2 . 
       FIG. 14  is a perspective view illustrating overall architecture for the apparatus for measurement of a magnetic field according to this embodiment. 
     In the apparatus  1400 , magnetic detection coils  1  shown in  FIG. 1  or magnetic detection coils shown in  FIG. 2  and a SQUID are kept in low temperatures inside a cryostat  1401 . With regard to arrangement of the magnetic detection coils, pairs of magnetic detection coils, each pair having two magnetic detection coils as shown in  FIG. 12 , are arranged in a configuration shown in  FIG. 13 . It should be noted that each magnetic detection coil is so arranged that a plane of a coil loop of a magnetic detection coil is parallel with a bottom plane of the cryostat  1401 . Liquid helium is charged in the cryostat  1401 , which is thermally insulated by a vacuum insulation layer. The cryostat  1401  is supported by a gantry  1402 . A testee for measurement of biomagnetism lies on a bed  1403 , and a height and horizontal position of the bed  1403  are adjusted so that a measurement area (a chest or back in case of magnetocardiography, for example) is positioned near the bottom plane of the cryostat  1401 . A measurement and control circuit  1404  controls a SQUID magnetometer to transform detected magnetic signals into voltage signals, which are transmitted to a signal process and display device  1405 . The device  1405  is able to eliminate an effect of environmental magnetic fields by a DSP so as to obtain magnetic signals generated by an organism of the testee and to display in real time a wave shape of magnetocardiography or magnetoencephalography, a diagram showing isomagnetic lines, a diagram showing current distribution and the like. 
     e. Example for Application of Magnetic Detection Coils: Apparatus for Magnetocardiography of Unborn Child 
     A description is given of an apparatus  1500  for magnetocardiography of an unborn child with reference to  FIG. 15  along with  FIGS. 1 and 2 . 
       FIG. 15  is a perspective view illustrating the apparatus for magnetocardiography of an unborn child according to an embodiment. 
     In the apparatus  1500 , magnetic detection coils  1  shown in  FIG. 1  or magnetic detection coils shown in  FIG. 2  and a SQUID are kept in low temperatures inside a cryostat  1501 . With regard to arrangement of the magnetic detection coils, pairs of magnetic detection coils, each pair having two magnetic detection coils as shown in  FIG. 12 , are arranged in a configuration shown in  FIG. 13 . It should be noted that each magnetic detection coil is so arranged that a plane of a coil loop of a magnetic detection coil is in parallel with a bottom plane of the cryostat  1501 . Liquid helium is charged in the cryostat  1501 , which is thermally insulated by a vacuum insulation layer. The cryostat  1501  is supported by a gantry  1502 . The cryostat  1501  is not only movable in horizontal and vertical directions, but also adjustable in diagonal direction. A position of the cryostat  1501  is so adjusted that magnetic detection coils are positioned near an abdominal region of a mother  1503 . A measurement and control circuit  1505  controls a SQUID magnetometer to transform detected magnetic signals into voltage signals, which are transmitted to a signal process and display device  1506 . The device  1506  eliminates an effect of environmental magnetic fields and cardiomagnetic signals deriving from the mother  1503  by a DSP so as to detect cardiomagnetic signals of an unborn child  1504 . The device  1506  displays in real time not only a cardiomagnetic wave shape  1507 , but also a heart rate  1508 , which is calculated based on the cardiomagnetic signals from the unborn child  1504 . The apparatus for magnetocardiography of an unborn child described above is able to monitor in real time the cardiomagnetic wave shape  1507  and the heart rate  1508  of the unborn chilled  1504 . 
     f. Example for Application of Magnetic Detection Coils: Apparatus for Magnetoencephalography 
     Description is given of an apparatus  1600  for magnetoencephalography with reference to  FIG. 16  along with  FIGS. 1 and 2 . 
       FIG. 16  is a perspective view illustrating the apparatus for magnetoencephalography according to an embodiment. 
     In the apparatus  1600 , magnetic detection coils  1  shown in  FIG. 1  or magnetic detection coils  2  shown in  FIG. 2  and SQUID&#39;s are kept in low temperatures inside cryostats  1601   a  and  1601   b . With regard to arrangement of the magnetic detection coils, pairs of magnetic detection coils, each pair having two magnetic detection coils as shown in  FIG. 12 , are arranged in a configuration shown in  FIG. 13 . It should be noted that each magnetic detection coil is so arranged that a plane of a coil loop of the magnetic detection coil is in parallel with a side surface of the cryostat  1601   a  or  1601   b . Liquid helium is charged in the cryostats  1601   a  and  1601   b , which are thermally insulated by vacuum insulation layers. The cryostats  1601   a  and  1601   b  are supported by a gantry  1602 . The cryostats  1601   a  and  1601   b  are not only movable in horizontal and vertical directions, but also adjustable in diagonal direction. Positions of the cryostats  1601   a  and  1601   b  are so adjusted that magnetic detection coils are positioned near a head of a testee  1603 . Measurement and control circuits  1605   a  and  1605   b  control SQUID magnetometers to transform detected magnetic signals into voltage signals, which are transmitted to a signal process and display device  1606 . The device  1606  eliminates an effect of environmental magnetic fields by a DSP so as to detect encephalomagnetic signals of the testee  1603 , displaying real-time encephalomagnetic wave shapes  1607   a  and  1607   b . The apparatus  1600  includes a device  1604  for stimulating auditory sense, which is used for giving vocal stimuli to ears of the testee  1603 . The signal process and display device  1606  monitors a real-time reaction of the testee  1603 . The device  1606  is able to calculate a time difference between peaks of the encephalomagnetic wave shapes  1607   a  and  1607   b , displaying in real time a transmission time  1608 . The apparatus  1600  is able to measure a spontaneous brain magnetic field and a phenomenon-related brain magnetic field in addition to a brain magnetic field evoked by a sensory stimulus, which is a response to an auditory stimulus, visual stimulus and somatic sensation stimulus. 
     A magnetic detection coil, which is a first-order or second-order differential magnetic detection coil, has been described as an example for the embodiment. The magnetic detection coil according to this embodiment detects signals differentiated in two different directions. It may be possible to alternatively select a magnetic detection coil which is third-order or more differentiated in a vertical direction, for example. 
     In the embodiment described above, magnetocardiography has been picked up as an example. The apparatus for measurement of biomagnetism according to this embodiment can be applied to measurement of a magnetic field generated through neural activities by a brain of a testee and to measurement of a heart magnetic field of an unborn child inside a mother. 
     In the embodiment described above, a SQUID magnetometer has been selected as an example to transform magnetic fluxes detected by the magnetic detection coils into voltage values. Other than this example, it may be possible to alternatively adopt other magnetometers such as a magnetoresistance element, a giant magnetoresistance element, a fluxgate magnetometer, an optical pumping magnetometer and the like. It may also be possible to alternatively adopt a cryogenic cooler and a SQUID cooled by liquid nitrogen if it is made of high-temperature superconducting material. 
     This embodiment provides an apparatus for measurement of biomagnetism having a high S/N ratio which is able to carry out more sensitive and accurate measurement of biomagnetism, which allows measurement of biomagnetism under an environment without a magnetic shield.