This invention relates to nuclear magnetic resonance (NMR) imaging methods and systems. More specifically, this invention relates to radio frequency (RF) coils used with such apparatus for transmitting and receiving RF signals.
In NMR imaging, a uniform magnetic field B.sub.O is applied to the imaged object along the z axis of a Cartesian coordinate system, the origin of which is centered within the imaged object. The effect of the magnetic field B.sub.0 is to align the object's nuclear spins along the z axis. In response to an RF magnetic signal of the proper frequency, oriented within the x-y plane, the nuclei precess about the z-axis at their Larmor frequencies according to the following equation: EQU .omega.=.gamma. B.sub.0
where .omega. is the Larmor frequency, and .gamma. is the gyromagnetic ratio which is constant and a property of the particular nuclei. Water, because of its relative abundance in biological tissue and the properties of its nuclei, is of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 khz/gauss and therefore in a 1.5 Tesla polarizing magnetic field B.sub.0 the resonant or Larmor frequency of water is approximately 63.9 Mhz.
In a typical imaging sequence, the RF signal centered at the Larmor frequency .omega., is applied to the imaged object by means of a radio frequency (RF) coil. A magnetic field gradient G.sub.z is applied at the time of this RF signal so that only the nuclei in a slice through the object along the x-y plane, which have a resonant frequency .omega., are excited into resonance.
After the excitation of the nuclei in this slice, magnetic field gradients are applied along the x and y axes. The gradient along the x axis, G.sub.x, causes the nuclei to precess at different resonant frequencies depending on their position along the x axis, that is, G.sub.x spatially encodes the precessing nuclei by frequency Similarly, the y axis gradient, G.sub.y, is incremented through a series of values and encodes y position into the rate of change of phase as a function of gradient amplitude, a process typically referred to as phase encoding.
A weak RF signal is generated by the precessing nuclei may be sensed by the RF coil and recorded as an NMR signal. From this NMR signal, a slice image may be derived according to well known reconstruction techniques. An overview NMR image reconstruction is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith.
The quality of the image produced by NMR imaging techniques is dependent, in part, on the strength and uniformity of the RF signal used to excite the nuclei. The strength of the RF magnetic signal directly affects the signal-to-noise ratio of the resultant image. The strength of the RF magnetic field is limited, in practice, by the efficiency of power transfer from the RF generator to the RF coil, the optimum level of NMR excitation, and by the tolerance of the patient to RF power deposition. The uniformity of the magnetic field affects both the slice selectivity of the G.sub.z gradient and the severity of image artifacts that may be produced from phase differences in the excited nuclei introduced by variations in the RF signal.
Referring to FIG. 1, a nucleus 10 has a magnetic moment 12 which may be excited into precession 18 about a static magnetic field B.sub.0 by an RF magnetic signal producing magnetic vector 14 along a plane perpendicular to the static magnetic field B.sub.0.
The excitating RF magnetic field 14 may oscillate along a single axis within the x-y plane. Such an oscillating field may be generated by a "saddle" coil (not shown) comprised of two conductive loops disposed along the axis of oscillation and perpendicular to the static magnetic field B.sub.0 as is known in the art.
A more effective excitation of the nuclear moments 12 may be achieved with a circularly polarized RF magnetic field, i.e. one that produces a rotating magnetic vector 14. Preferably, the magnetic vector 14 rotates within the x-y plane at an angular velocity equal to the Larmor frequency .omega. as shown by arrow 20 in FIG. 1.
It is known that a rotating RF magnetic vector may be generated with certain RF coil structures when the coil structure is excited at its "resonant" frequency. Referring to FIG. 2, one such coil structure 28 for creating a rotating magnetic field is comprised of a pair of conductive hoops 22 spaced along the axis of the static magnetic field B.sub.0. The hoops 22 are joined with a series of conductive segments 24 parallel to axis of the static magnetic field B.sub.0. The hoops 22 and conductive segments 24 have an intrinsic inductance and may be broken along their length with capacitive elements 30 to promote the desired pattern of current flow through the conductive segments 24 when the coil is driven by an external RF generator 26.
When the coil structure 28 is driven a particular frequency, the phase of the current distribution in each axial segment 24 will equal the transverse angle .theta. of the segment 24 measured around the axis of the static magnetic field B.sub.0. This phase distribution is the result of a "delay line" effect of the intrinsic inductance of the hoop elements 22 and the capacitance of the axial segments 24. At the driving frequency the delay line produces a full 360.degree. of phase shift, in the current flowing though the conductive segments 24, for 360.degree. of angular displacement .theta. of the conductive segments 24. As is understood in the art, this current distribution circularly polarizes the RF magnetic field 14 as described above.
Detailed descriptions of several RF coil structures which use the phase shifting properties of various coil geometries at a given frequency, are given in the following U.S. Pat. Nos. assigned to the assignee of the present application and hereby incorporated by reference: 4,680,548, entitled: "Radio Frequency Field Coil for NMR" and issued Jul., 14, 1987; 4,692,705, entitled: "Radio Frequency Field Coil for NMR" and issued Sep. 8, 1987; and, 4,694,255, entitled: "Radio Frequency Field Coil for NMR" and issued Sep. 15, 1987. These designs will be referred to collectively as "resonant RF coils".
Referring still to FIG. 2, the coil structure 28 may be driven by a RF generator directly connected across one of the capacitive elements 30 in an conductive segment 24. Alternatively, U.S. Pat. No. 4,638,253, entitled: "Mutual Inductance NMR RF Coil Matching Device, issued Jan. 20, 1987, teaches a method of inductively coupling an RF source 26 to the coil structure 28. This patent is also assigned to the assignee of the present application and hereby incorporated by reference.
It will be apparent, by application of the law of superposition, that in the resonant coil design, considerable current will flow circumferentially through the conductive hoops 22 which connect the conductive segments 24. This current is the sum of currents flowing through each conductive segments 24 on opposite sides of the coil 28. The circumferential currents produce longitudinal magnetic field components along the B.sub.0 axis (not shown) in distinction from the desired transverse rotating magnetic field 14. These longitudinal field component may adversely affect the axial homogeneity of the generated transverse magnetic field 14.
During an MR imaging sequence, the object to be imaged (also not shown) is placed within the coil volume as defined by the hoops 22 and conductive segments 24. The proximity of the imaged object to the coil structure results in capacitive coupling between the coil 28 and the imaged object and therefore an increased loss of RF power within the imaged object from dielectric heating.
To the extent that the imaged object is not uniform in cross section or is unevenly centered within the RF coil 28, the capacitive coupling to the imaged object will vary among different coil elements as will the dielectric losses coupled to these different coil structures. The effect of this uneven loading on the RF coil 28 will be to "detune" the coil structure upsetting the delay line of the coil structure and hence distorting the phase distribution of the currents in the conductive segments 24. A change in the phase distribution of the axial currents may produce distortion in the reconstructed NMR image and reduce the RF power coupled from the RF generator 26.