45Ultra-High-Field MRI

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L. L. Wald, PhD; E. Adalsteinsson, PhD

MGH/MIT/HMS Athinoula A. Martinos Center for Functional and Structural Biomedical Imaging, Charlestown Navy Yard, Building 149, 13th Street (2301), Charlestown, MA 02129-2060, USA

C O N T E N T S

45.1 Introduction 511

45.2 Ultra-High-Field MRI Systems 511 45.3 Spatially Tailored RF Excitation 512 45.4 Acceleration of Transmit Pulses with Parallel Excitation 514

45.5 Future System Requirements for Parallel-Excitation Techniques 515 45.5.1 Multiple, Fast Transmit Channels 515 45.5.2 SAR Considerations 516

45.5.3 Pulse Calculation 517 45.5.4 Large-Flip-Angle Pulses 517

45.6 RF Transmit Arrays for Parallel Excitation 517 45.6.1 General Considerations 517

45.6.2 Strip-Line Arrays 518 45.6.3 Loop-Coil Transmit Arrays 518 45.7 Conclusion 520

References 520

Parallel-Excitation Techniques for 45

Ultra-High-Field MRI

Lawrence L . Wald and Elfar Adalsteinsson

controlled spatially varying fl ip angle and/or excitation phase. While the benefi ts of spatially tailored excitation pulses have been known for some time, the potential to decrease the pulse length of spatially tailored excitations to practical durations for clinical imaging using parallel-transmission methods raises the exciting possibility of their routine use. If the technical issues are addressed, this technology will be a timely addition to ultra-high fi eld strength (7 T and above) studies. At these fi elds, the wavelength effects of the excitation fi elds in the head or body cause considerable transmit B₁ inhomogeneity. This issue could be addressed with spatially tailored excitation pulses with accelerated gradient trajectories and parallel excitation channels. This chapter will look into the potential needs of this nascence fi eld.

45.2

Ultra-High-Field MRI Systems

High-fi eld MRI offers many potential advantages to clinical and scientifi c studies, including increased sensitivity and in many cases improved image contrast. In synergy, these effects promise improved characterization of brain function and anatomy in health and disease. However, spatially inhomogene- ous B₁ transmit fi elds intrinsic to the shortened RF wavelength in biological tissue present one of the outstanding methodological challenges to bringing ultra-high-fi eld systems from the research to the clinical arena. The spatial homogeneity of the transmit fi eld is especially problematic since tissue contrast is a function of the excitation fl ip angle for many clinical and research imaging applications. Thus, unlike detection inhomogeneity that manifests primarily as image intensity shading, a non-uniformly transmit- ted B₁ fi eld results in spatially dependent tissue contrast and therefore reduced diagnostic power.

45.1

Introduction

The success of parallel-imaging methods and their impact on image encoding have sparked a great deal of interest in parallel-excitation arrays and the potential to utilize the spatial information in an array during RF transmission. One of the principal applications of parallel-excitation technology will likely be spatially tailored RF pulses, excitation pulses with a carefully

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The problem is quite severe for the human head at 7 T. In this case, the wavelength of the RF in the tissue is approximately 14 cm, comparable to the size of the head. The “dielectric center brightening” is the well-known result of this wavelength effect for most coil types. At 7 T, the transmit B₁ fi eld is peaked at the center of the head, undergoes minima at approximately the location of the insular cortex and then partially recovers toward the scalp. Since the head is not symmetric, the response to the head-coil interac- tion is also not symmetric, but presents a complex three-dimensional spatial profi le in amplitude and phase. While it is tempting to pass this dramatic fi eld variation off as a “coil problem,” there are no current coil designs for 7-T brain imaging that produce uniform transmit-fi eld profi les. Figure 45.1 illustrates the center brightening in a 7-T proton-density-weighted image acquired with a “uniform” birdcage coil.

The B₁-ﬁ eld inhomogeneity leads to unwanted spatial variations in the tissue contrast for most pulse sequences. For example, the contrast in a typi- cal T1-weighted pulse sequence will vary within the head from proton-density-weighted to heavily T1- weighted in a simple FLASH image using a “uniform”

mode birdcage (BC) or transverse-electromagnetic (TEM) type structure at 7 T. The severity of the effect depends on the contrast’s dependence on B₁, and since the problem arises during excitation, it is not easily dealt with in post-processing. Where the

intrinsic contrast information is not present locally, no amount of image manipulation can substitute for the missing information. In addition to mitigating the non-uniform transmit B₁ proﬁ le at high ﬁ elds, spatially tailored RF excitation offers the potential to minimize local susceptibility dephasing dropouts in T2*-weighted sequences by pre-phasing the RF excitation to cancel the expected off-resonance effects (Stenger et al. 2003; Stenger et al. 2000; Stenger et al. 2002; Glover and Lai 1998).

45.3

Spatially Tailored RF Excitation

RF excitations appropriately modulated in amplitude and phase during time-varying gradients offers the potential of spatially tailored RF phase and amplitude in the excitation (Pauly et al. 1989). Although these types of pulses have been demonstrated for many years, there are serious engineering challenges to their routine and practical use. First and foremost is the lengthy encoding period needed for these pulses (as long as 50 ms). One of the most promising ways to address this limitation is the extension of parallel reconstruction methods to parallel excitation to accelerate the excitation encoding. A practical goal is

Fig. 45.1. Coronal and axial 7-T proton-density-weighted brain images acquired with a “uniform”

band-pass birdcage coil. For a low-fl ip-angle gradient-echo exam, the image intensity scales as the square of B₁ (one factor of B₁ for the transmit effi ciency and one for the receive effi ciency).

Variations in the transmit efﬁ ciency are more problematic than the receive inhomogeneities since they lead to contrast alterations

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to achieve a 5-ms duration 3D excitation pulse with a spatial proﬁ le that can mitigate the observed B₁ pattern in the head at 7 T in birdcage-like coils. This short duration is needed to be useful in common anatomical imaging sequences such as MPRAGE and FLASH.

Shaping the 2D spatial fl ip-angle distribution of an RF excitation requires the use of modulated RF amplitude and phase, while the gradients trace a particular k-space trajectory, typically a spiral or echo-planar path. In practice, we fi rst chose a target magnetization map (which is proportional to the fl ip-angle map for small fl ip angles). For example, the target magnetization map might be the inverse of the measured B₁ profi le for a given coil in the head. The calculation of the needed RF waveform is greatly simplifi ed in the low-fl ip angle case where it can be reduced to a k-space analysis (Pauly et al. 1989). The RF excitation during such a gradient traversal is viewed as a series of small fl ip angle excitations short enough so the gradient can be thought of as constant during these sub-excitations. The phase and amplitude of these small RF pulses is altered so that the deposition of RF energy in the “excitation k-space” matrix is the Fourier transform of the desired spatial fl ip-angle map. This design method relies on the linear nature of the Bloch equations for small fl ip angles. Then the Fourier transform of the k-space excitation matrix is a good approximation of the desired excitation profi le (Fig. 45.2).

Extending this concept to 3D proﬁ les requires covering k-space in three directions and is there-

fore considerably more time consuming in addition to requiring additional gradient trajectory designs.

Full 3D excitation is required for many applications including a slice-selective excitation with an in-plane tailored ﬂ ip-angle pattern. In both the 2D and 3D cases, the gradient trajectory is chosen to cover the needed area or volume in k-space, using similar considerations as for image encoding. For our current applications, we constrain the excitation and 3D gradient designs to single-shot applications. Single-shot 3D image encoding is rarely (if ever) performed, and it is plausible for excitation trajectories only because relatively low spatial resolutions are sufﬁ cient (1 cm or less) to mitigate dielectric-center-brightening patterns. As in image encoding, there is likely to be one or more spatial direction that is considerably more slowly encoded than the other(s), leaving this direction open to off-resonance artefacts.

Tailored 3D pulses have been pursued to mitigate susceptibility dephasing in fMRI (Stenger et al. 2003;

Stenger et al. 2000; Stenger et al. 2002; Glover and Lai 1998). In this case, relatively high spatial resolution is needed in all three spatial directions, requiring very long pulse durations. For example, Stenger and colleagues (2003) utilized four excitations of 20 ms each to achieve a 5-mm slice-selective pulse and 3.7- mm in-plane spatial resolution and 0.1-mm transi- tions in the slice direction. Recently, Stenger and col- leagues have pursued tailored 3D pulses to address B₁ inhomogeneity (Saekho et al. 2005). For this application, they applied the “stack of spirals” trajectory approach, which allowed the relatively slowly vary-

Fig. 45.2. Standard demonstration of a tailored excitation pulse. Target fl ip-angle profi le (left) is used to compute the RF pulse phase and amplitude as a function of position in excitation k-space. The amplitude k-space matrix of the pulse along a spiral trajectory is shown (middle). The right-hand image shows a low fl ip angle proton-density-weighted image acquired with this pulse at 7 T. A grey-scale target fl ip-angle map is thus “burned into” an oil phantom at 7 T using a uniform birdcage coil. Single 16-ms spiral readout, 2-mm resolution. Pulse is not slice-selective or encoded in the through-plane direction and is therefore of limited practical value. Pulse design courtesy of K. Setsompop, MIT

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ing B₁ proﬁ le to be addressed in all three directions.

The excitation resolution was 2×2×1.25 cm³_, and the pulse required 22 ms, which is still unacceptably long for many applications. In addition, this approach is not suitable to 2D imaging since it does not provide a sharp slice proﬁ le.

A slice-selective spatial excitation pulse with in- plane B₁ mitigation but retaining sharp slice-select pro- ﬁ les is also a desirable goal for 2D imaging sequences.

A method of 3D single-channel excitation to select a 2D slice with in-plane B₁ compensation was demonstrated by Ulloa et al. (2005) and Saekho et al. (2005).

In this case, high spatial resolution is needed in the slice-select direction (defi ned here as the z direction) to achieve thin, well-defi ned slice profi les, but only 1- cm resolution is needed in the in-plane (x, y) direction to follow the slowly varying B₁ maps. Therefore, the gradient trajectory usually considered consists of lines or “spokes” in the z direction extending far in the k_z direction (for the needed sharpness in the slice select profi le). A sinc-like pulse is played out for each of these k_z lines to insure a slice selective pulse. The k_z spokes with differing excitation amplitudes and phases are repeated at a multitude of locations in the k_x, k_y plane.

Figure 45.3 shows an image acquired in a low-dielectric oil phantom at 7 T with a uniform mode birdcage coil.

This low-dielectric phantom produces uniform images

with conventional slice-selective pulses. In this image a

“spokes” trajectory pulse was used that imposed a 30%

reduction in B₁ in the center of the phantom with a Gaussian profi le to reduce dielectric center brighten- ing. Only 13 “spokes” (lines in the k_z directions) were needed spaced to give a 1-cm resolution to the in- plane fl ip-angle profi le. Each spoke used a traditional sinc-shaped excitation to produce a clean slice select in z. The duration of the pulse was only 4.1 ms using the fast head gradients on this 7-T system. Thus, useful pulse durations are potentially feasible for this type of excitation given a modest acceleration with parallel- excitation methods. For example, two-fold acceleration will produce a 2-ms pulse, a short enough duration to substitute for the slice-selective pulse in most 2D sequences without excessive implications for the pulse sequence strategy.

45.4

Acceleration of Transmit Pulses with Parallel Excitation

The traversal of transmission k-space can be accelerated in analogy to the under-sampling of k-space in accelerated image encoding (Katscher et al. 2002;

Zhu 2002; Katscher et al. 2003; Zhu 2004). Here, the under-sampling of transmission k-space is com- pensated for by simultaneous transmission from multiple transmit coils. Acceleration of the transmit pulse through parallel excitation has the potential to provide the considerable acceleration needed to bring the duration of the 3D tailored RF pulses to a length useable for common imaging applications. A plausible goal for a slice-selective pulse is a duration of less than 5 ms and accommodation of in-plane spatial B₁ variations on a scale of 1 cm of resolution. Preliminary results suggest that these goals can be met on a state-of-the-art 7-T head-gradient system with a two-fold acceleration from transmit SENSE. With four-fold acceleration, the pulse duration could be potentially reduced to less than 2 ms, making these tailored excitations truly interchange- able with conventional slice-select pulses in most pulse sequences.

For spatially selective RF pulses, the excitation consists of a regular deposition of RF energy as a function of position in the excitation k-space deﬁ ned by the gradient trajectory being played out during the excitation. Since the concept of k-space in spatially

Fig. 45.3. Gaussian dip proﬁ le shown in an oil phantom with a uniform BC coil at 7 T. The 4-ms pulse is slice-selective and utilizes the “spokes” trajectory. Pulse design courtesy of K. Setsompop, MIT

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selective RF pulses is dual to that of image encoding, the principles of parallel acceleration of the excitation gradient trajectory follow by analogy. An array of multiple transmit coils is needed, each of which exhibits a spatially different excitation pattern (fl ip- angle map) and is driven by an independent RF waveform. The gradient waveform is played out during the excitation and modulates the transverse magnetization created. The amplitude and phases of the excitations as a function of position in the excitation k- space are determined by the Fourier transform of the target fl ip-angle map for the low-fl ip-angle case.

Since multiple, spatially differing transmit channels are used, there is the opportunity to under- sample in the excitation k-space. This shortens the path to be traversed without sacrifi cing spatial defi - nition. Since its introduction, transmit SENSE has been further explored by a number of groups, mainly using simulations on single-transmit-channel systems, but also on a three-channel small-bore scanner (Ullmann et al. 2005) and a clinical imager outfi t- ted with eight channels (Zhu et al. 2005). Figure 45.4 shows a spatially tailored 2D excitation, accelerated using a prototype 3-T whole-body system with eight independent transmit channels (Siemens Medical Solutions, Erlangen Germany). The 2D spiral exci- tation trajectory was accelerated from R=1 to R=8, giving pulse durations ranging from 9.4 ms to 1.2 ms.

The data in Fig. 45.4 suggest a potential for transmit SENSE to tolerate considerably more acceleration without degradation than seen with image encoding.

This might be possible since a major source of degradation is noise ampliﬁ cation during the inversion of the under-sampled image. In the image-encoding case, we seldom have sensitivity to spare, while the signal-to-noise ratio of the transmit pulses is high since there is relatively little noise in the transmit system (Fig. 45.5).

45.5

Future System Requirements for Parallel-Excitation Techniques

45.5.1

Multiple, Fast Transmit Channels

A key component of parallel-excitation techniques is that the MR system is capable of simultaneously exciting with different RF waveforms on multiple excitation channels. The waveform is carefully calculated with a time-varying phase and amplitude modulation that is synchronized to the gradient traversal. Simply splitting a single RF waveform and exciting multiple

Fig. 45.4. Acceleration of a 2D spiral trajectory spatial-excitation pulse (target proﬁ le is the “MIT” logo). Accelerated from R=1 (top left) to 8 (bottom right) using the eight-channel transmit array of Fig. 45.7. The R=1 pulse has a duration of 9.4 ms, the R=8 pulse 1.2 ms. Some of the noise in the high-acceleration images is due to the pulse-amplitude normalization; the high-rate pulses have a lower ﬂ ip angle. Pulse design courtesy of K. Setsompop, MIT

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coils with a global phase shift and amplitude altera- tion is a form of B₁ shimming, but does not provide the required spatial control for accelerating spatially tailored RF pulses. Since the spatial resolution of the excitation pattern needed is low compared to imag- ing (at least for mitigating B₁ inhomogeneity), the gradient trajectory is expected to be slew-rate limited.

This makes fast, low-inductance gradient sets such as a dedicated head-gradient system advantageous.

It also requires a ﬁ ner gradient and RF raster than found on most clinical systems. Ideally, the gradient and RF waveforms should be deﬁ ned on a 1-µs raster.

Although currently only a few experimental systems have more than two independent excitation channels, there are few fundamental barriers to making this technology widespread. However, the beneﬁ ts of parallel transmission will not be widely explored until more systems are available.

45.5.2

SAR Considerations

Although there are many similarities between parallel-excitaiton techniques such as transmit SENSE and the acceleration of image encoding using parallel-receive coils (regular parallel imaging), there are also some important differences. Firstly, in the receive

case, the data acquired from the multiple receive channels can be digitally combined with great fl exibility in post-processing. For the parallel-excitation case, the only fl exibility is in the pre-calculation of the RF waveforms. These waveforms are based on the target profi le, knowledge of the transmit profi les of the individual coils, and knowledge of the gradient trajectory. The B₁ fi elds from the multiple transmit coils simply superimpose as magnetic fi elds within the object with no opportunity for adjustment after the excitation. The spatial distribution of the transverse magnetization (and thus the fl ip-angle map) is modulated by the gradient waveform in concert with the RF waveform allowing cancellation or build- up of the magnetization at certain spatial locations as desired. The electric fi elds (which govern the SAR distribution) are not affected by the gradients and simply superimpose among the coil elements.

Therefore, the SAR distribution, although affected by the transmit pulses and array geometry, is not an explicit part of current RF design. Nevertheless, it is an important consideration for high-fi eld imaging, and future work is needed to include SAR explicitly in the design of the RF pulses to enable a fl exible trade off between RF excitation properties (due to the magnetic RF fi eld, B₁) and SAR distribution (due to the electric RF fi eld, E₁). There is potential for the E₁ fi elds from the coils to superimpose constructively,

Fig. 45.5. Accelerated ﬁ ve-spoke trajectory at 3 T with eight-channel parallel system (pulse duration: 5.4 ms). In this case, the RF waveforms were calculated to allow a very inhomogeneous array (that of Figs. 45.7 and 45.8) to produce a uniform in-plane excitation (left image) with a sharp slice-selective proﬁ le (right image)

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creating a local SAR hot spot. Therefore, in addition to monitoring the average power from each channel, the system must make some estimate of how the local E₁ ﬁ elds will superimpose so that the local SAR limits are not exceeded.

45.5.3

Pulse Calculation

The development of an automated system for cal- culating tailored RF pulses based on the B₁ proﬁ le measured in an individual subject and utilizable by an MR technologist will be required to utilize paral- lel-excitation techniques fully for B₁-inhomogeneity mitigation. In this scenario, a 3D B₁-ﬁ eld-mapping sequence would be incorporated into an automated pre-scan, and the image reconstruction program would calculate the 3D B₁ map and the RF phase and amplitude waveforms for accelerated RF pulses needed for subsequent studies. Currently, the calculation of accelerated 3D pulses can take several min- utes, but substantial speed-ups are anticipated with coding optimization and faster processors such as the new generation of 64-bit machines now available on clinical scanners for image reconstruction.

45.5.4

Large-Flip-Angle Pulses

All work performed to-date has assumed the small- fl ip-angle approximation. While this approximation provides for elegant and computationally tractable RF designs, large fl ip-angle pulses are central to many clinical pulse sequences, and the low-fl ip-angle con- straint needs to be addressed. Thus, parallel-excitation methods stand where slice-select pulse design stood 25 years ago when a sinc function was used for slice selection. The non-linearity of the Bloch equations for high fl ip angles causes the sinc-shape RF pulses to result in an imperfect slice profi le for larger fl ip angles, but it is a good initial guess. More sophisti- cated calculation methods such as the Shinnar-LeRoux algorithm (Shinnar and Leigh 1989; Shinnar et al.

1989a; Shinnar et al. 1989b; Shinnar et al. 1989c;

Pauly et al. 1991), optimal control theory (Mao et al. 1986; Conolly et al. 1989) or simulated annealing (Geen et al. 1989) have emerged to produce pulses that appear sinc-like to fi rst approximation, but provide improved slice-select profi les for large-fl ip-angle excitations. It remains an open question how to best

proceed along these lines for accelerated tailored RF pulses, but the analogous nature of the problem gives hope that similar innovations will be made.

45.6

RF Transmit Arrays for Parallel Excitation

45.6.1

General Considerations

While a great deal of effort has been placed in design- ing receive-only arrays, until recently the design of arrays for transmitting has been limited to cases where the use of a uniform body RF excitation was not available (Kim et al. 2003; Peterson et al. 2003).

More recent work has focused on developing a transmit array based on surface-coil designs, whose excitation ﬁ eld fall-off would partially compensate for the dielectric-center-brightening effect (Pinkerton et al. 2005) or allow ﬂ exible excitation and recep- tion in a single array ultimately providing either B₁ shimming approaches or parallel-excitaiton applications (Lee et al. 2001; Lee et al. 2004; Adriany et al.

2005; Adriany et al. 2003). The work toward transmit manipulation has almost exclusively focused on the decoupled lumped-element strip-line array.

A separate receive coil is required to map the B₁ profi les since all phase information in the B₁ fi eld is lost if the same coil is used for both transmit and receive (the complex transmit fi eld scales as B₁ and the receive profi le as the complex conjugate, B₁*;

therefore, the detected image from the transmit/

receive coil, which is the product, loses all phase information). Since a separate receive coil is needed to measure the complex B₁ maps, the body coil is useful for this purpose due to its uniformity in the oil phantoms we used at 3 T.

A second over-all issue with transmit arrays is the lack of an equivalent of preamplifi er decoupling (Roemer et al. 1990) to help reduce inductive coupling between array elements. In preamplifi er coupling, the coil is matched to present the impedance needed to optimize the noise fi gure of the preamplifi er (usually 50 Ohm), but the preamplifi er itself does not present a 50-Ohm impedance to the coil. Instead, it is confi g- ured to present a high impedance in series with the loop of the coil. In practice, this is achieved by having the preamplifi er impedance transformed to a low impedance across a PIN-diode trap circuit, which in

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turn presents a high impedance in series with the coil.

With a high impedance effectively inserted in the coil loop, inductive coupling is reduced since the currents and therefore magnetic ﬁ elds that mediate the inductive coupling are reduced. It is not possible to extend this concept to the transmit array in a straightforward way since power transmission along a 50-Ohm coax necessitates a 50-Ohm RF source. However, it might be possible to engineer an RF source and transmission line with the needed properties. Some preliminary designs have been discussed (Nam and Wright 2005; Kurpad et al. 2005).

45.6.2

Strip-Line Arrays

In the strip-line approach, each element consists of a linear conductor possibly broken with series capaci- tors and connected to the ground plane at either end via a capacitor. Thus, the circuit resembles the standard “pi” model of a transmission line. The elements resemble the transmission line element of a TEM coil (Vaughan et al. 1994). The strip-line elements showed a favorable lack of coupling to each-other when used in their original (no-lumped-element) conﬁ guration and closely shielded and trimmed to a O/4 length (Lee et al. 2001). But most work has found it simpler to use the lumped-element approach and either have the elements far from one-another (Lee et al. 2004) or decoupled with capacitive networks (Adriany et al.

2005) or both (Lee et al. 2004) (Fig. 45.6).

The ability of the ﬁ elds of the strip-line to penetrate into the subject critically depends on the spacing of the line from its ground plane. The closer the line is to the ground plane, the more “bottled up” the ﬁ elds are (as is desired for a transmission line), and the less inductive coupling to neighboring coils exists. But,

this also keeps the ﬁ elds away from the subject. Relax- ing this distance allows more penetration, but also allows the otherwise tight design to resemble more a square surface coil on its side, with the associated element-to-element coupling of this geometry. In fact, it is the inductive coupling between the strip-line elements of the TEM coil that gives the TEM volume res- onator its intrinsic mode structure. In the TEM coil, the thinner the coil annulus, the lower the element- to-element coupling (and narrower the mode spacing), and a higher the fraction of the stored energy is held outside of the subject (between the rung and shield). So, at least in their practical implementation, the strip-line approach like the loop approach is not free from coupling issues.

Used as a receive coil to make a magnitude image formed in a normal sum of squares fashion, the strip line array shows highest reception sensitivity near the elements similar to other arrays. The transmit ﬁ elds for a single element are also peaked near the element.

But depending on the relative phases, much different spatial patterns can be obtained in the transmit B₁ map due to constructive and destructive interference between the elements (van de Moortele et al. 2005).

While the use of these cancellations and super-posi- tions has been theoretically examined for B₁ shimming (Collins et al. 2005; Ibrahim et al. 2001), they are also the basis for the ability of the parallel-excitaiton method to accelerate tailored RF pulses.

45.6.3

Loop-Coil Transmit Arrays

In this confi guration, many of the considerations of receive arrays apply to the transmit array. Although the tight design of the stripline is attractive, one potential benefi t of a transmit array of loop elements is the ability to minimize nearest neighbor interac- tions with coil overlap. A second potential benefi t is the effi cient nature of the loop coil. Figure 45.7 show a prototype eight-channel 3-T transmit array used for the transmit-SENSE experiments shown in Figs. 45.4 and 45.5. Eight circular loops with diameter of 15.6 cm were placed on an acrylic former with outer diameter of 27.6 cm. Figure 45.8 shows B₁ amplitude and phase maps acquired with each coil excited independently. Considerable next-nearest- neighbor coupling and coupling between opposite elements is visible. Nevertheless, as long as the B₁ fi eld maps are accurate, they will be correctly taken into account in the pulse calculation.

Fig. 45.6. Strip-line geometry for lumped-element conﬁ gura- tion (ground plane not shown). After Lee et al. (2004)

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Fig. 45.8. Magnitude (top) and phase (bottom) maps of the eight-channel transmit overlapping-loop array. Data were acquired with the array coil for RF transmission and the body RF coil for reception. Measurements courtesy of A. Alagappan, MGH Fig. 45.7. Eight-channel transmit array of overlapping

circular loop coils for 3 T. Each element has a transmit/

receive switch, cable trap, and PIN-diode coil detun- ing to allow measurement of the B₁ map with a separate receive-only coil. S12 measures between nearest neighbors were –22 dB and –17 dB between next-nearest neighbors. Coupling between opposite coils on the cylinder was about –15 dB. Coil design and construc- tion courtesy of A. Alagappan, MGH

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45.7 Conclusion

Theoretical work on parallel RF transmission and recent experimental implementations on prototype systems indicate that parallel excitation has the potential to overcome critical obstacles to robust and routine human scanning at high fi eld strength. As these developments are extended to 16- and 32-channel neuroimaging arrays, it seems likely that very- high-fi eld human brain imaging will be possible with an essentially constant fl ip angle across the human head with RF pulse durations comparable to current slice-selective pulses. While most work has been con- centrated on head-sized transmitters, the problem of B₁ homogeneity in the torso becomes signifi cant at lower B₀ fi elds. Thus, the fi rst clinical applications might be for body transmit coils at 3 T.

Of course, intriguing research questions remain open in several areas, including optimal coil array designs that minimize element couplings and maxi- mize spatial orthogonality of individual channels, the estimation and real-time monitoring of local SAR during simultaneous application of RF in excitation coil arrays, and the development of rapid and robust RF pulse designs that extend the current low-ﬂ ip- angle domain to an arbitrary excitation angle, such as spin echoes, saturation, and inversions pulses. How- ever, with continued active research in these areas, progress is likely to accelerate, and one can envision logical additions to the architecture of a current clinical scanner that readily accommodates the requirements of a general parallel RF excitation system.

These include coil arrays optimized for parallel transmit and receive, modular blocks of RF synthesis and amplifi cation with individual characterization of amplifi er nonlinearities, subject-specifi c models of local SAR deposition for monitoring, and a rapid B₁- estimation pre-scan that feeds fast RF-pulse-design software capable of incorporating E₁ constraints.

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