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    • 专利摘要:
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    公开号:KR19990033805A
    申请号:KR1019970055236
    申请日:1997-10-27
    公开日:1999-05-15
    发明作者:김정회；이수열；이정한；이완
    申请人:윤종용；삼성전자 주식회사；
    IPC主号:
    专利说明:

    Design method of planar gradient magnetic coil for magnetic resonance imaging device
    The present invention relates to a design method of a planar gradient magnetic coil used in an open magnetic resonance imaging apparatus, and more particularly, to a planar gradient magnetic coil designed by applying a current density optimization technique to minimize inductance induced in a coil. It is about.
    In general, the magnetic resonance imaging apparatus is an imaging apparatus widely used to diagnose various diseases. For such magnetic resonance imaging apparatus, "Insertable bipolar gradient coils for magnetic resonance imaging," M. A. Petropoulos et al., Rev. Sci. Instrum ,. pp. 2639-2645, 62 (11), Nov. It was introduced in 1991.
    Magnetic resonance imaging with an open structure to provide more comfort to a patient with recent phobia or for interventional studies such as a biopsy needle or simple surgery during magnetic resonance imaging. The device is spreading. In order to make an open magnetic resonance imaging apparatus, the structure of a main magnet forming a main magnetic field must first be open. The open main magnet which is used a lot now is mainly a permanent magnet. The structure having the largest open area among the shapes of the open main magnet introduced until recently is the C-type permanent magnet 1 as illustrated in FIG. 1. The permanent magnet 1 creates a shooting space between the opposite pole faces and makes a magnetic path between the two poles C-shaped so that the shooting space can be approached in three directions among four directions. Doing. In order to construct an open magnetic resonance imaging apparatus, not only the main magnet 1, but also the inclined magnetic field coil 2, which makes a gradient magnetic field, must have an open structure. The most widely used structure of the open gradient magnetic coil 2 is a planar structure as shown in FIG. The planar gradient magnetic coil 2 has its shape maintained by two parallel planar structures. That is, the windings 2: gradient magnetic coils which form gradient magnetic fields are disposed on the two planar structures, respectively. The planar structure is made of an insulator with a suitable mechanical strength to support the electromagnetic forces generated when injecting current into the coil. And reference numeral 3 is a high frequency coil.
    The shape of the coil disposed on the planar structure is determined to form a desired gradient magnetic field in the imaging space. When the output capacity of the gradient magnetic coil (2) driver is small, in order to obtain a gradient magnetic field having a desired strength, a method of increasing the number of turns after determining the shape of the winding is conventionally used. In other words, the strength of the gradient magnetic field is increased in proportion to the number of turns. It is possible to easily increase the gradient magnetic field strength by increasing the number of turns, but there is a problem that the inductance increases in proportion to the square of the number of turns. This is because when the inductance of the gradient magnetic coil is increased, the high speed switching of the gradient magnetic coil is difficult. Recently, in order to minimize the inductance of gradient magnetic coils, a method of spatially distributing windings of gradient magnetic coils has been proposed. In this method, the winding shape is designed by minimizing the magnetic field energy formed by the gradient magnetic coil in the spatial frequency domain. As the designed winding shape is similar to the fingerprint, as shown in FIG. Also called gradient magnetic coil. This fingerprint type gradient magnetic coil is known to have an advantage that the inductance is greatly reduced as compared to the gradient magnetic coil of the conventional winding concentrated type. However, the fingerprint gradient magnetic coil has a problem in that the linearity of the gradient magnetic field is poor. In particular, the planar gradient magnetic coil is known to have a very bad linearity of the gradient magnetic field.
    The present invention was devised to improve the above problems, and the design of the planar gradient magnetic coil to increase the linearity of the gradient magnetic field formed by converting the current density of the planar gradient magnetic coil having the smallest inductance in the spatial frequency region and converting it The purpose is to provide a method.
    1 is a cross-sectional view showing a schematic configuration of a general open magnetic resonance imaging apparatus;
    2 is an explanatory diagram for showing the basic structure of the planar gradient magnetic coil of FIG.
    Figure 3 is a plan view showing the winding form of the planar X-direction gradient magnetic coil designed in a way to minimize the inductance by a conventional design method,
    4 is a plan view showing the winding form of the planar X-direction gradient magnetic coil designed by the design method according to the present invention.
    * Explanation of symbols for main parts of the drawings
    1: main magnet 2: gradient magnetic coil
    3: coarse-grain coil 4: patient
    5 patient carrier 6 flat structure
    7 winding 8 winding
    In order to achieve the above object, the method of designing a planar gradient magnetic coil according to the present invention comprises the steps of: (a) setting a condition of the gradient magnetic field; (b) a current density satisfying the condition of the gradient magnetic field but having a minimum inductance; Obtaining a function in the spatial frequency domain, (c) obtaining a current density function in the spatial frequency domain in which the spatial frequency domain is scaled by a predetermined magnification, and (d) obtaining a scaled current density function in the spatial frequency domain. And (e) extracting the shape of the gradient magnetic field coil from the current density function converted into the spatial position.
    In the present invention, the step (b), wherein α and β are the spatial frequencies in the X and Z directions, respectively, and i is a current density function converted to the spatial frequencies α and β, i,
    In step (c), α and β are the spatial frequencies in the X and Z directions, respectively, and the current density functions are obtained when the magnifications of the spatial frequencies α and β are respectively k k and k β. i,
    It is preferable to obtain | require by.
    Hereinafter, a design method of a planar gradient magnetic coil according to the present invention will be described in detail with reference to the accompanying drawings.
    FIG. 2 illustrates a basic planar structure of a planar gradient magnetic coil of the magnetic resonance imaging apparatus of FIG. 1. In Fig. 2 the direction of the main magnetic field is in the y direction and the distance between the two planar structures 6 is 2a. Since the current forming the gradient magnetic field is only on two planes, this current density function can be expressed by Equation 1 below.
    [Equation 1]
    In Equation 1 above Wow Is the current density component in the X and Z directions on a plane where y = a, and Wow Denotes current density components in the X and Z directions on a plane where y = -a, respectively. Green * s function to find the magnetic field produced by this current density in the spatial frequency domain Is expressed as a Fourier transform equation.
    [Equation 2]
    In Equation 2, α and β represent spatial frequencies in the X and Z directions, respectively. Is a two-dimensional Fourier transform Let's do it. The same symbol is used for the current density function of the z component. Since the current density function on the plane must satisfy the continuous equation, the following equation (3) is established.
    [Equation 3]
    Therefore, the following equation (4) holds.
    [Equation 4]
    Magnetic field vector potential to find the magnetic field formed by the current density function The components of Equation 5, Equation 6 and Equation 7 are respectively obtained.
    [Equation 5]
    [Equation 6]
    [Equation 7]
    If the y component of the magnetic flux density in the same direction as the main magnetic field is obtained by applying a curl to the magnetic field vector potential above, Equation 8 is obtained.
    [Equation 8]
    To minimize the inductance of gradient magnetic coils, the magnetic field energy generated by gradient magnetic coils should be minimized. The magnetic field energy W formed by the gradient magnetic coil in the X direction or the -X direction is obtained from the following equation (9).
    [Equation 9]
    If the current density function that minimizes the magnetic field energy W is imposed under the constraint that the magnetic flux density of the desired size is made at a predetermined position, the inductance of the gradient magnetic coil composed of this current density function will be minimized. The constraints to be used are shown below.
    i = 1.2, ........ For N
    Where N is the number of spatial positions @ at which the desired magnetic flux densities B y , i are to be produced. The spatial position should be appropriately determined at the position where the magnetic resonance image is to be obtained, and the magnitude of the magnetic flux density should be determined to form the gradient magnetic field in the desired direction. W is In order to minimize this under the constraints, Lagrange equation is established and solved to obtain the current density function as shown in Equation 10 below.
    [Equation 10]
    In Equation 10 above, λ i represents a Lagrange multiplier. When the equation for λ i is established, the following equations (11) and (12) are obtained.
    [Equation 11]
    [Equation 12]
    The current density function for the gradient magnetic coil in the Y direction can also be obtained in a small process. As a result, the equation is equivalent to replacing all cosh functions of Equations 10 and 12 with the sinh function.
    In Equation 10, which represents the current density function forming the gradient magnetic field in the X-direction or Z-direction in the spatial frequency domain, the current density function i is obtained by scaling the spatial frequency axes α and β by kα and kβ times, respectively. Get
    [Equation 13]
    Lagrange multiplier λ ' i in Equation 13 is obtained by Equation 14 and Equation 15 below.
    [Equation 14]
    [Equation 15]
    The same calculation can be made for the gradient coil in the Y direction, and the resulting expression is equivalent to substituting the sinh function instead of the cosh function of all equations.
    An example of designing a planar X-direction gradient magnetic coil using the above design method is as shown in FIG. 4. The designed gradient magnetic coil was composed of two sides, and the distance between the two sides was 50 cm and the size of each side was limited to a circle having a diameter of 100 cm so that the gradient magnetic coil could be easily attached to the pole surface of the permanent magnet. And the scaling factor was k (alpha) = k (beta) = 0.6. That is, since FIG. 3 shows the shape of the winding 7 of the coil designed by the method of minimizing the inductance using Equation 10, FIG. 4 illustrates the current density function of scaling the spatial frequency axis in Equation 10. It shows the form of the winding 8 of the coil designed by. The number of constraints used when designing this gradient coil was 4, which is shown in Table 1.
    In the constraints of Table 1, points 1 and 3 are intended to produce a gradient magnetic field of 4 Gauss / cm in the X direction, and points 1, 2, and 4 have uniform magnetic flux density values at positions with the same x value. It is limited. The winding 8 is obtained by integrating the current density function in the X direction in the Z direction and dividing the integrated value into contour lines at equal intervals. In other words, the winding type can be said to approximate the discrete current winding current density function having a continuous distribution. In order to compare the magnetic field linearity of the gradient magnetic coil designed in this way with the magnetic field linearity of the gradient magnetic coil designed by the prior art, the magnetic field linearity error was calculated. In the ellipsoidal cylinder 40cm long, 30cm short, and 30cm long, similar to the human body, 100 points are equally spaced, and the magnetic flux density values at these points are calculated. Error was calculated. The linearity error represents the difference between the ideal magnetic flux density value and the actual magnetic flux density value. The magnetic flux density was calculated by applying the Bio-Savart law after dividing the curved winding into short segments of about 3,000 to 5,000. As a result, the gradient magnetic coil designed by the method illustrated in FIG. 4 has a linearity of 200% improvement over the gradient magnetic coil of the prior art illustrated in FIG. 3. In addition, the winding of FIG. 4 has a simpler form than the winding of FIG.
    As described above, the planar gradient magnetic coil according to the present invention has the following effects.
    1. It improves gradient magnetic field linearity of planar gradient magnetic coil designed by inductance minimization technique.
    2. It is easier to manufacture the gradient magnetic coil by simplifying the coil shape of the planar gradient magnetic coil designed by the inductance minimization technique.
    权利要求:
    Claims (3)
    [1" claim-type="Currently amended] (A) setting the conditions of the gradient magnetic field,
    (B) obtaining a current density function having a minimum inductance in the spatial frequency domain satisfying the condition of the gradient magnetic field;
    (C) obtaining a current density function in the spatial frequency domain in which the spatial frequency domain is scaled by a predetermined magnification;
    (D) converting the scaled current density function into a spatial position in the spatial frequency domain; and
    (E) extracting the shape of the gradient magnetic coil from the current density function converted into the spatial position.
    [2" claim-type="Currently amended] The method of claim 1, wherein the step (b) is where α and β are spatial frequencies in the X and Z directions, respectively, and i is a current density function converted into the spatial frequency α and β.

    A method of designing a planar gradient magnetic field coil for a magnetic resonance imaging apparatus, characterized in that obtained by
    [3" claim-type="Currently amended] The current density function according to claim 1, wherein in step (c), α and β are spatial frequencies in the X and Z directions, respectively, and the magnifications of the spatial frequencies α and β are respectively k k and k β. Is i ',

    A method of designing a planar gradient magnetic field coil for a magnetic resonance imaging apparatus, characterized in that obtained by
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    同族专利:
    公开号 | 公开日
    引用文献:
    公开号 | 申请日 | 公开日 | 申请人 | 专利标题
    法律状态:
    1997-10-27|Application filed by 윤종용, 삼성전자 주식회사
    1997-10-27|Priority to KR1019970055236A
    1999-05-15|Publication of KR19990033805A
    优先权:
    申请号 | 申请日 | 专利标题
    KR1019970055236A|KR19990033805A|1997-10-27|1997-10-27|Design method of planar gradient magnetic coil for magnetic resonance imaging device|
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