Inherently de-coupled sandwiched solenoidal array coil

Details Inherently de-coupled sandwiched solenoidal array coil Inherently de-coupled sandwiched solenoidal array coil

2001-08-24

2004-06-15

Shaw, Shawna Jeannine (Department: 3737)

Surgery

Diagnostic testing

Detecting nuclear, electromagnetic, or ultrasonic radiation

C324S318000, C324S322000

Reexamination Certificate

active

06751496

ABSTRACT:

FIELD OF THE INVENTION
This invention relates generally to methods and apparatus for magnetic resonance imaging. In particular, it relates to an electromagnetically de-coupled sandwiched solenoidal array coil for receiving radio frequency magnetic resonance signals in an MRI apparatus.
BACKGROUND AND SUMMARY OF THE INVENTION
Based on principals of nuclear magnetic resonance (NMR), magnetic resonance imaging (MRI) has become a widely accepted medical imaging modality—having evolved tremendously over the last two decades as an important clinical technique for obtaining visual images of tissues and organ structures within the human body. Basically, clinical MRI relies on the detection of NMR signals from abundant hydrogen protons in the human body. These protons are first subjected to a strong radio frequency (RF) electromagnetic wave excitation pulse. If the frequency of the excitation pulse is properly chosen, the protons receive needed RF energy to make a transition to an excited state. Eventually, the excited protons give up their excess energy via a decay process, commonly known as “relaxation”, and return to their original state.
Since the magnetic moment of a proton is a vector quantity, the microscopic behavior of millions of protons considered together is equivalent to the vector sum of the individual magnetic moments of all the protons. For convenience, this sum is typically represented as a single resultant magnetization vector, M
0
, that is aligned with {overscore (B)}
0
(the static main magnetic field). The strong RF excitation pulse used in MRI effectively tips this resultant magnetization vector away from alignment with the static main field {overscore (B)}
0
and causes it to precess before decaying back to an equilibrium alignment with {overscore (B)}
0
. The component of this precessing resultant magnetization vector in a plane perpendicular to {overscore (B)}
0
induces an RF signal, referred to as the nuclear magnetic resonance (NMR) signal, in an RF “pick-up” or “receive” coil(s) placed near the body portion containing the excited protons.
During clinical MRI, the magnetic resonance of protons in different tissues within an anatomical region are made distinguishable through the evocation of a magnetic field gradient along each of three mutually orthogonal spatial directions-the effect of which is to cause protons at different spatial locations to have slightly different nuclear magnetic resonance frequencies. The NMR signals induced in the receive coil can then be processed to reconstruct images of the anatomical structure of interest (i.e., images of the spatial distribution of NMR nuclei which, in many respects, conform to the anatomical structures containing such nuclei).
To obtain the maximum induced signal in a receive coil, the magnetic field of the receive coil—conventionally designated as {overscore (B)}
1
—must be oriented perpendicular to the direction of the static main magnetic field ({overscore (B)}
0
) of the MRI apparatus. For a planar-loop (i.e., a substantially flat loop) type receive coil, that direction is in a direction normal to the plane of the conductive loop(s) of the coil. For a quadrature detection (QD) type coil—which basically consists of two RF receive coils having mutually perpendicularly oriented {overscore (B)}
1
fields—must also have the {overscore (B)}
1
fields of both of its coils oriented perpendicular to the MRI apparatus static field {overscore (B)}
0
to obtain a maximum induced signal.
Due to the unique nature of the clinical MRI environment, there are certain design considerations that are particularly relevant toward obtaining maximum performance from an RF receive coil. For example, the NMR signals induced in an RF receive coil during magnetic resonance imaging are nominally on the order of nanovolts in magnitude while the background ambient electrical noise may be of comparable levels or higher. Therefore, a high performance RF receive coil for clinical MRI needs to be electromagnetically “sensitive” enough to detect the low level NMR signals despite the relatively high levels of background electrical noise. Moreover, other design considerations such as field-of-view, uniformity (i.e., uniformity of the magnetic field generated by the coil) and coil efficiency are also highly relevant to coil performance in the clinical MRI environment X coil uniformity because it can affect image interpretation and coil efficiency because a highly efficient coil allows the same image signal information to be acquired within a shorter time frame.
Theoretical analysis and experimental results have indicated that for many MRI applications using multiple RF receive coils together as a signal receiving array is advantageous for improving coil sensitivity, signal-to-noise ratio and imaging field-of-view. Conventionally, the imaging “field-of-view” (FOV) for an MRI receive coil is defined as the distance indicated between two points on the coil sensitivity profile (i.e., a graph of coil sensitivity vs. distance profile) where the signal drops to 80% of its peak value. In a typical MRI receive coil array arrangement, instead of using a single large FOV but less sensitive coil that covers the entire imaging volume of interest, multiple small FOV but sensitive coils are distributed as an array over the entire imaging volume. Each individual coil of the array covers a small localized volume and the NMR signals received by each coil are simultaneously acquired through corresponding data acquisition channels. Signals from each of the channels are then appropriately combined and processed to construct an image of the complete volume of interest. Due to this ability to simultaneously acquire a signal from multiple sources (i.e., multiple coils) and since each individual signal channel is provided with its own associated detection circuitry, an array type coil can conceivably operate with high efficiency. However, the simultaneous acquisition of a signal from a plurality of individual receive coils necessitates that each coil function independently, free of interaction or coupling.
As two individual coils having the same resonance frequency are brought in close proximity to each other, the common resonance frequency starts to split into two separate frequencies due to the electromagnetic interaction or “coupling” between the coils. Generally, the closer the coils are brought together, the stronger the interaction and the larger the frequency split. Since an MRI receive coil should have its maximum sensitivity optimized for a particular relatively narrow band of frequencies, the resonance frequency splitting can cause sensitivity degradation when two or more receive coils are closely arranged in an array.
Generally, MRI systems are categorized as either a horizontal field type or vertical field type, based on the direction of the static main magnetic field. In a horizontal field system, the static main magnetic field is typically oriented in a superior-inferior direction relative to a patient laying in a prone/supine position. In a vertical field MRI system, the static main magnetic field is oriented in an interior-posterior direction relative to a prone/supine patient. This difference in main field orientation is important in that it affects the ultimate placement and configuration of an RF receive coil(s) used for diagnosis in such systems. More often than not, a receive coil designed specifically for use in a horizontal field system will not be suitable for similar use in a vertical field system and vice versa.
Consequently, horizontal field MRI systems and vertical field MRI systems typically require radically different RF receive coil configurations to obtain the maximum achievable performance from the coil. For example, a planar-loop type receive coil configuration designed for obtaining images of a the human spine works well in a horizontal field MRI system when placed in posterior contact with the back of a patient in supine position. However, the same coil configuration may not work in a vertical field system because, in that case, th

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