Abstract:
A system and method for transmitting multiple radio frequency (RF) channels via an RF coil assembly are provided. An RF coil assembly having a number of coil elements may be configured to transmit a number of RF channels which is less than the number of coil elements thereof. Some implementations may use signal splitters for some or all of the RF channels to produce driving inputs for each coil element. By using more coil elements than RF channels, various embodiments may exhibit increased power efficiency and improved B1 uniformity.

Description:
BACKGROUND OF THE INVENTION 
       [0001]    The present invention relates generally to magnetic resonance (MR) imaging, and more specifically, to a system and method for transmitting multiple radio frequency (RF) channels via a multi-element RF coil assembly. In some embodiments, the number of independent RF channels communicated to the RF coil assembly may be less than the number of coil elements in the assembly. In such a case, one or more of the RF channels may be split and/or phase shifted for application to more than one coil element. 
         [0002]    MR imaging in general is based upon the principle of nuclear magnetic resonance. When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B 0 ), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field, such as a B 1  excitation field, which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, or “longitudinal magnetization”, M Z , may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment M t . A signal is emitted by the excited spins after the excitation signal B 1  is terminated and this signal may be received and processed to form an image. 
         [0003]    When utilizing these signals to produce images, magnetic field gradients (G x , G y , and G Z ) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. 
         [0004]    One goal in MR imaging is to produce a homogenous B 1  excitation field such that a desired magnetization effect caused by the B 1  field will be produced as intended. In classic “birdcage” coils, the energy in an ideal B 1  field, measured in Joules, is determined for the loops of a coil based upon: 
         [0000]        J   N   =J   0 ·cos(ω t+N ΔΦ)  Eqn. 1 
         [0000]    where ω is the characteristic Larmor frequency for the spins of interest, N is the number of loops in the birdcage coil, and ΔΦ is the angular distance between loops. Since all loops are electrically interconnected, the birdcage coil acts like a transmission line with one wavelength existing about the entirety of the structure. In normal 1.5T imaging, the Larmor frequency for protons is about 64 MHz and permittivity of human tissue (or other objects of interest) is generally not a significant factor in producing appreciable B 1  inhomogeneities. 
         [0005]    However, in high field imaging, such as where the composite B fields are on the order of 3T or 7T, the Larmor frequency for protons is higher, due to f=γ·B. For example, in 3T imaging, the Larmor frequency for protons is about 127 MHz. Thus, wavelengths for the RF transmissions become shorter and the permittivity of a tissue to be imaged can become a factor. The relative permittivity of human tissue can have values of ER from about 6 to 70. Significant phase changes as well as signal attenuations can occur in an RF transmission as it passes into and through an imaging tissue under these conditions. Such phenomena can cause both constructive and destructive interference with the RF transmissions from the other loops of an RF coil. Therefore, even when the transmissions from loops of a coil are carefully tuned to produce an ideal homogenous B 1  field, inhomogeneities may still be produced at high Tesla fields and Eqn. 1 may not hold in reality. Inhomogeneous B 1  fields lead to inaccurate flip angle distributions in a field of view (FOV) and dark areas in images. 
         [0006]    One way to help prevent these inhomogeneities is to utilize a technique known as RF shimming. RF shimming involves adjusting the signal inputs for each loop of a coil assembly to account for expected or measured field inhomogeneities. “Passive” RF shimming includes splitting, phase shifting, amplifying, attenuating, or otherwise tuning the same RF waveform to produce varying inputs for each coil. “Dynamic” RF shimming includes producing unique RF waveforms for each coil and accounting for inhomogeneities in the design of the waveforms. 
         [0007]    Classic birdcage coil assemblies are not known to implement RF shimming techniques as effectively as transverse electro-magnetic (TEM) coil assemblies. TEM coil assemblies have individual coil elements which can be driven separately, making them more ideal for multi-channel transmissions such as parallel transmission. However, known TEM coil assemblies generally experience coupling or mutual inductance between neighboring coil elements and non-neighboring coil elements, and between coil elements and the RF shield. This coupling is relied upon to characterize the TEM coil assembly as a single resonator in design techniques. Additionally, known TEM coil assemblies use either one RF waveform input, or a multi-channel RF input which is tailored to the structure of the TEM coils. 
         [0008]    It would therefore be desirable to have a system and method for effectively reducing or eliminating B 1  inhomogeneities in high field imaging. It would be further desirable for such system and method to present an RF coil assembly which is capable of operating without coupling and which is capable of transmitting a variety of multi-channel RF inputs. 
       BRIEF DESCRIPTION OF THE INVENTION 
       [0009]    Embodiments of the present invention provide a system and method of transmitting a multi-channel RF pulse sequence via a multi-element coil array. By decoupling neighboring coil elements, carefully selecting coil widths and spacing, and/or by having more coil elements than transmit channels, increased power efficiency and improved B 1  homogeneity can be achieved. 
         [0010]    In accordance with one aspect of the invention, an MR imaging apparatus includes a main magnet having a bore therethrough and a plurality of gradient coils positioned about the bore of the main magnet. The apparatus further includes an RF coil assembly that is also disposed within the bore of the main magnet. A pulse module of the apparatus is adapted to output a plurality of RF transmit channels to the RF coil assembly for transmission during an imaging sequence. The RF coil assembly has a plurality of individual coil elements which is greater in number than the plurality of RF transmit channels. 
         [0011]    According to another aspect of the invention, a method for configuring an RF transmit system is provided. The method includes the steps of affixing a plurality of coil elements about a frame, connecting an RF pulse input line to a signal splitter, and routing the outputs from the signal splitter to drive less than all of the coil elements. The method also includes the step of connecting at least one additional RF pulse input line to the remainder of the coil elements to drive the coil elements. 
         [0012]    In accordance with a further aspect of the invention, an RF coil assembly is provided. The RF coil assembly includes a volume coil structure which has an opening therethrough and an end ring. The coil assembly also includes a plurality of conductive segments positioned about a surface of the volume coil structure in a transverse electro-magnetic (TEM) arrangement. A driving input array of the coil assembly is configured to receive input signals representing a multi-channel transmission and communicate the signals to drive the conductive segments. The RF coil assembly also has at least one channel splitter connected to receive a single channel of the multi-channel transmission and provide at least two of the input signals. 
         [0013]    Various other features and advantages will be made apparent from the following detailed description and the drawings. 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         [0014]    The drawings illustrate embodiments presently contemplated for carrying out the invention. 
           [0015]    In the drawings: 
           [0016]      FIG. 1  is a schematic block diagram of an exemplary MR imaging system for use with an embodiment of the present invention. 
           [0017]      FIG. 2  is a perspective view of an exemplary RF coil assembly in partial cut-away. 
           [0018]      FIG. 3  is a circuit diagram representing a pair of coil elements of the RF coil assembly of  FIG. 2 . 
           [0019]      FIG. 4  is a block diagram of a signal splitter for a single channel RF transmission. 
           [0020]      FIG. 5  is a block diagram of a pair of signal splitters for a two channel RF transmission 
           [0021]      FIG. 6  is a block diagram of a number of signal splitters for a multi-channel RF transmission 
           [0022]      FIG. 7  is a diagram of one configuration for connecting a signal splitter to coil elements. 
           [0023]      FIG. 8  is a diagram of another configuration for connecting a signal splitter to coil elements. 
       
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
       [0024]    Referring to  FIG. 1 , the major components of a preferred magnetic resonance imaging (MRI) system  10  incorporating an embodiment of the present invention are shown. The operation of the system is controlled from an operator console  12  which includes a keyboard or other input device  13 , a control panel  14 , and a display screen  16 . The console  12  communicates through a link  18  with a separate computer system  20  that enables an operator to control the production and display of images on the display screen  16 . The computer system  20  includes a number of modules which communicate with each other through a backplane  20   a . These include an image processor module  22 , a CPU module  24  and a memory module  26 , known in the art as a frame buffer for storing image data arrays. The computer system  20  is linked to disk storage  28  and tape drive  30  for storage of image data and programs, and communicates with a separate system control  32  through a high speed serial link  34 . The input device  13  can include a mouse, joystick, keyboard, track ball, touch activated screen, light wand, voice control, or any similar or equivalent input device, and may be used for interactive geometry prescription. 
         [0025]    The system control  32  includes a set of modules connected together by a backplane  32   a . These include a CPU module  36  and a pulse generator module  38  which connects to the operator console  12  through a serial link  40 . It is through link  40  that the system control  32  receives commands from the operator to indicate the scan sequence that is to be performed. The pulse generator module  38  operates the system components to carry out the desired scan sequence and produces data which indicates the timing, strength and shape of the RF pulses produced, and the timing and length of the data acquisition window. The pulse generator module  38  connects to a set of gradient amplifiers  42 , to indicate the timing and shape of the gradient pulses that are produced during the scan. The pulse generator module  38  can also receive patient data from a physiological acquisition controller  44  that receives signals from a number of different sensors connected to the patient, such as ECG signals from electrodes attached to the patient. And finally, the pulse generator module  38  connects to a scan room interface circuit  46  which receives signals from various sensors associated with the condition of the patient and the magnet system. It is also through the scan room interface circuit  46  that a patient positioning system  48  receives commands to move the patient to the desired position for the scan. 
         [0026]    The gradient waveforms produced by the pulse generator module  38  are applied to the gradient amplifier system  42  having Gx, Gy, and Gz amplifiers. Each gradient amplifier excites a corresponding physical gradient coil in a gradient coil assembly generally designated  50  to produce the magnetic field gradients used for spatially encoding acquired signals. The gradient coil assembly  50  forms part of a magnet assembly  52  which includes a polarizing magnet  54  and an RF coil assembly  56 . RF coil assembly  56  includes a number of coil elements  57 . A transceiver module  58  in the system control  32  produces pulses which are amplified by an RF amplifier  60  and coupled to the RF coil  56  by a transmit/receive switch  62 . The resulting signals emitted by the excited nuclei in the patient may be sensed by the same RF coil  56  or a separate RF coil (not shown). The sensed signals are coupled through the transmit/receive switch  62  to a preamplifier  64 . The amplified MR signals are demodulated, filtered, and digitized in the receiver section of the transceiver  58 . The transmit/receive switch  62  is controlled by a signal from the pulse generator module  38  to electrically connect the RF amplifier  60  to the coil  56  during the transmit mode and to connect the preamplifier  64  to the coil  56  during the receive mode. The transmit/receive switch  62  can also enable a separate RF coil (for example, a head coil or surface coil) to be used in either the transmit or receive mode. 
         [0027]    The MR signals picked up by the RF coil  56  are digitized by the transceiver module  58  and transferred to a memory module  66  in the system control  32 . A scan is complete when an array of raw k-space data has been acquired in the memory module  66 . This raw k-space data is rearranged into separate k-space data arrays for each image to be reconstructed, and each of these is input to a data processor  68  which operates to Fourier transform the data into an array of image data. This image data is conveyed through the serial link  34  to the computer system  20  where it is stored in memory, such as disk storage  28 . In response to commands received from the operator console  12 , this image data may be archived in long term storage, such as on the optical disk drive  30 , or it may be further processed by the image processor  22  and conveyed to the operator console  12  and presented on the display  16 . 
         [0028]      FIG. 2  depicts one exemplary configuration for an RF coil assembly for use in an MR imaging system, such as that shown in  FIG. 1 . The RF coil assembly  80  of  FIG. 2  is a generally cylindrical-shaped volume coil assembly. It is also contemplated that embodiments of the invention may be implemented with other coil types, such as head coils. RF coil assembly  80  is configured to transmit one or more RF waveforms to generate a B1 field in an interior volume  82  thereof. In the embodiment shown, RF coil assembly  80  is arranged as a transverse electro-magnetic (TEM) coil assembly. In this regard, RF coil assembly  80  includes an outer RF shield  84 , surrounding a hollow cylindrical structure or frame  86 . RF shield  84  may be formed of a copper mesh or other conductive materials suitable for shielding RF transmissions. 
         [0029]    The RF shield  84  of RF coil assembly  80  is shown in partial cut-away to reveal three coil elements  88   a ,  88   b ,  88   c . As arranged, RF coil assembly  80  has a total of sixteen coil elements  88  spaced evenly about the circumference of frame  86 . However, it is to be understood that various configurations of RF coil assembly  80  may have any number of coil elements. As can be seen from the depiction of coil elements  88   a ,  88   b , and  88   c , the coil elements  88  are laid out lengthwise as conductive segments along the primary axis of frame  86 . In one embodiment, the coil elements  88  are copper strips having widths of approximately 1.25 inches and lengths of approximately 460 mm. However, it is appreciated that similar widths and lengths may equivalently achieve the advantages discussed herein when the coil elements  88  are evenly spaced about frame  86 . 
         [0030]    Coil elements  88  are each in electrical communication with the common RF shield  84  via connectors  90  at the ends of the coil elements  88 . Thus, RF shield  84  may act as a current return path when the coil elements  88  are being used for RF transmission. On one end ring  94  of coil assembly  80 , a decoupling element  92  is attached between each coil element  88 . As will be explained below, decoupling neighboring coil elements  88  provides for improved control over transmissions from each coil element. 
         [0031]    Coupling or mutual inductance experienced between non-neighboring coil elements, such as between coil element  88   a  and coil element  88   c , is diminished due to the relatively narrow width of the coil elements  88  and the distance therebetween. That is, compared to known TEM-type coil assemblies, coil elements  88  are relatively small and spaced further apart. Mutual inductance experienced between non-neighboring coil elements is, therefore, not significant in comparison to the strength of the signals applied to the coil elements  88  for transmission. 
         [0032]    Additionally, the sixteen element RF coil array  80  achieves an improved power efficiency over an eight element RF coil array (not shown). In whole body volume coils, the RF shield  84  is often relatively close to the coil elements  88 . As a result, the efficiency of the coil assembly  80  is determined by losses in the structure as opposed to losses caused by the imaging subject. With more elements, the current per coil element  88  used to drive the coil assembly  80  can be reduced. Compared to an eight element coil assembly, a sixteen element coil assembly can have a power efficiency improved by a factor of two. 
         [0033]    Referring now to  FIG. 3 , a circuit model of a pair of neighboring coil elements  96 ,  98  is shown, illustrating a minimization of the coupling or mutual inductance therebetween. Each coil element  96 ,  98  may be modeled as a current loop having a first capacitance  100 ,  102 , then an inductance  104 ,  106 , and another capacitance  108 ,  110 . The current return paths  120 ,  122  are provided via the RF shield shown in  FIG. 2 . The circuit model of  FIG. 3  illustrates that, when input signals  112 ,  114  are applied to the coil elements  96 ,  98 , a coupling or mutual inductance  116  is experienced between the two coil elements. That is, a portion of the energy from the signal  112  applied at coil  96  is transferred to coil  98 , and vice versa. However, when a decoupling element, such as a capacitor  118 , is connected between the coils  96 ,  98  at one end, this mutual inductance is substantially eliminated. Each coil element  96 ,  98  can thus be controlled more independently of other coils. Due to such a decoupling, an RF coil assembly such as that shown in  FIG. 2  need not be treated as a single resonator. 
         [0034]    Referring now to  FIG. 4 , a block diagram of a signal splitting configuration for providing driving inputs to sixteen coil elements is shown. The configuration of  FIG. 4  is a single mode configuration in which one RF waveform or transmit channel  130  is split to provide a number of driving inputs. As shown, RF waveform  130  is applied first to a sixteen-way power splitter  132 . The outputs of power splitter  132  are then appropriately amplitude adjusted by attenuators  134  and phase shifted by phase shifters  136 , then amplified by sixteen amplifiers  138 . The degrees of individual amplitude and phase adjustment of each output signal  140  may be determined according to a passive RF shimming design technique. Each output signal  140  is therefore appropriately scaled and phase shifted to produce a substantially homogenous B 1  field when transmitted by coil elements evenly spaced about an RF coil assembly. 
         [0035]      FIG. 5  depicts a block diagram of an alternative signal splitting configuration for providing driving inputs to sixteen coil elements. Two separate transmit channels or RF waveforms  150 ,  152  may be designed and used as inputs. The RF waveforms  150 ,  152  may be designed individually according to a dynamic RF shimming design technique. That is, the RF waveform inputs  150 ,  152  themselves are designed to account for and/or correct expected or measured B1 inhomogeneities. Each RF waveform  150 ,  152  of  FIG. 5  is applied to a respective eight-way power splitter  154 , a series of attenuators  156 , a series of phase shifters  158 , and a series of amplifiers  160 , to produce eight distinct driving input signals for eight separate, independently driven coil elements (not shown). In other words, a passive RF shimming technique may be used to divide the signals from each RF waveform for improved B1 homogeneity in addition to the dynamic RF shimming technique used to design the waveforms  150 ,  152 . 
         [0036]      FIG. 6  shows another implementation of signal splitting for an RF coil assembly having sixteen coil elements. Eight independent transmit channels or RF waveforms  170 - 184  are used as inputs. In one embodiment, the eight RF waveforms  170 - 184  in the aggregate comprise a multi-channel parallel transmission. For example, RF waveforms  170 - 184  may be designed according to a SENSE or GRAPPA parallel transmission technique. The waveforms may further be designed in addition, or as an alternative, to parallel transmission techniques, according to a dynamic RF shimming technique. Each RF waveform  170 - 184  is applied to a two-way power splitter  186 , a pair of attenuators  188 , a pair of phase shifters  190 , and a pair of amplifiers  192  to produce a pair of driving input signals  194 - 208 , respectively. 
         [0037]    From the implementations shown in  FIGS. 4-6 , it will be appreciated that the features and advantages described herein may be extended to equivalently apply for an arbitrary number of RF waveform inputs which is less than or equal to the number of coil elements to be driven. Moreover, it is contemplated that the division of each independent RF transmit channel need not be equal. For example, in a four-channel transmission via an eight element coil assembly, one RF waveform could be divided into four driving inputs, another RF waveform could be divided into two driving inputs, and the remaining two RF waveforms could be applied to coil elements without splitting. In short, embodiments of the invention allow for M independent transmit channels to be applied to an N-element coil array, where M is less than or equal to N. 
         [0038]    Referring now to  FIG. 7 , one pair of driving input signals  210 ,  212  from the signal splitting configuration of  FIG. 6  is shown connected to two coil elements  214 ,  216  of an RF coil assembly  218 . According to passive RF shimming techniques, for a homogenous B 1  field, driving input signals that are split from the same RF waveform or transmit channel are generally phase shifted according to the angular position at which they will be applied with respect to one another. In the embodiment shown, a first driving input  210  is applied at coil element  214 . The second driving input is applied at a coil element  216  opposite the first coil element  214  about the circumference of coil assembly  218 . Therefore, the second driving input is phase shifted by 180 degrees from the first driving input signal. 
         [0039]    In contrast,  FIG. 8  shows an alternative configuration in which driving input signals are connected more sequentially. As shown, one RF waveform  220  is divided by a signal splitting arrangement  222  into five driving input signals  224 - 232 . The driving input signals  224 - 232  are connected to five neighboring coil elements  234 - 242  on a sixteen element RF coil assembly. Therefore, each driving input signal has a phase shift of 22.5 degrees from its neighbors. The first driving input signal  224  is not phase shifted and is applied at a first coil element  234 . The second driving input signal  226  is phase shifted by 22.5 degrees from the first driving input signal  224  and is applied to a second coil element  236  next to the first coil element  234 . The third driving input signal  228  is phase shifted by 22.5 degrees from the second driving input signal and 45 degrees from the first input signal. Third driving input signal  228  is connected to a third coil element  238  next to the second coil element  236 . Similarly, fourth driving input signal  230  and fifth driving input signal  232  are phase shifted another 22.5 degrees each, and are applied at coil elements  240  and  242 , respectively. Thus, it can be seen from the configurations shown in  FIGS. 7 and 8  that RF waveforms may be divided and applied at any coil elements at any positions about an RF coil assembly. 
         [0040]    Therefore, a number of embodiments of the present invention have been described. It has been demonstrated that an M-channel RF pulse sequence may be transmitted via an N-element RF coil assembly to achieve a variety of features and advantages. 
         [0041]    In one embodiment of the present invention, an MR apparatus includes a main magnet, a plurality of gradient coils, an RF coil assembly, and a pulse module. The main magnet has a bore therethrough, about which the plurality of gradient coils are positioned. The RF coil assembly is disposed within the bore, and has a plurality of individual coil elements. The pulse module is adapted to output a plurality of RF transmit channels to the RF coil assembly for transmission during an imaging sequence. The number of individual coil elements of the RF coil assembly is greater than the number of RF transmit channels. 
         [0042]    Another embodiment includes a method for configuring an MR transmit system. The method includes affixing a plurality of coil elements about a frame and connecting a first RF pulse input line to a first signal splitter. The outputs of the signal splitter are routed to drive less than all of the plurality of coil elements. The method further includes connecting at least one additional RF pulse input line to the remaining coil elements to drive the coil elements. 
         [0043]    In a further embodiment of the invention, an RF coil assembly includes a volume coil structure, a driving input array, and at least one channel splitter. The volume coil array has an opening therethrough, a plurality of conductive segments positioned about its surface in a TEM arrangement, and an end ring. The driving input array receives a plurality of input signals representing a multi-channel transmission communicates the signals in order to drive the plurality of conductive segments. The at least one channel splitter is connected to receive a single channel of the multi-channel transmission and to provide at least two of the plurality of input signals. 
         [0044]    Accordingly, the present invention has been described in terms of the preferred embodiment. It is recognized that equivalents, alternatives, and modifications, aside from those expressly stated, are possible and within the scope of the appending claims.