- Email: [email protected]
Local head gradient coils: Window(s) of opportunity
NeuroImage 62 (2012) 660–664
Contents lists available at SciVerse ScienceDirect
NeuroImage journal homepage: www.elsevier.com/locate/ynimg
Review
Local head gradient coils: Window(s) of opportunity Eric C. Wong Departments of Radiology and Psychiatry, University of California, San Diego, CA USA
a r t i c l e
i n f o
a b s t r a c t
Article history: Accepted 1 January 2012 Available online 9 January 2012
At the Medical College of Wisconsin (MCW), prior to the 1991 announcement of the discovery of BOLD fMRI, all of the technical pieces that were needed for efficient BOLD fMRI imaging were assembled for other applications, allowing MCW to jump into the fMRI business just days after the announcement. Central among these pieces was single shot EPI, implemented at MCW using a three axis local head gradient coil. This article describes the development of local gradient coil technology at MCW, and a historical perspective on local head gradient coils in general. © 2012 Elsevier Inc. All rights reserved.
Keywords: Local gradient coil Functional MRI (fMRI) Echo planar imaging
Contents Introduction . . . . . . . . The first human head gradient A note on LGC safety . . Historical perspective . . . . A comeback for the LGC? . . Acknowledgments . . . . . References . . . . . . . . .
. . coil . . . . . . . . . .
. . . . . at MCW . . . . . . . . . . . . . . . . . . . . . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
Introduction On August 12, 1991, I sat with Peter Bandettini in the auditorium of the San Francisco Hilton, watching with mouth agape as Tom Brady showed a movie of EPI brain images getting brighter and darker with brain activity—BOLD fMRI. It was stunning, and everybody in the room knew it was a revolutionary moment. Peter was a graduate student at the Medical College of Wisconsin (MCW), and was looking for a thesis project, something combining MRI technology and neuroscience, and there it was, dropped in his lap. I was toward the end of the PhD phase of an MSTP program at MCW, with a focus on local gradient coil (LGC) design and applications. This article describes the history of the LGC program at MCW, and the role it played in the early development of fMRI at that institution. As a starting graduate student in Biophysics at MCW, I had a choice between EPR, for which the department under the direction of Jim Hyde, was world renowned, and MRI, which was an up and coming
E-mail address: [email protected].
1053-8119/$ – see front matter © 2012 Elsevier Inc. All rights reserved. doi:10.1016/j.neuroimage.2012.01.025
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
660 661 663 663 664 664 664
area, but in which MCW had primarily participated in the area of RF coils and the clinical applications thereof. Nevertheless, as an MSTP student, with aspirations of mixing a biophysics background with medicine, MRI was an obvious choice. The training was enriched in a very important way through interactions with GE Medical Systems, which was located only 15 miles to the west in Waukesha, WI. Joint journal club meetings created strong intellectual ties between MCW and GE, which at the time housed Gary Glover, Scott Hinks, Norbert Pelc, Wally Bloch, and Matt Bernstein, among others. The culture of the Biophysics division at MCW was strongly focused on technical innovation. Hypothesis driven research was something that happened elsewhere. In this setting I started in 1988 to look for a thesis project. Anatomical MRI was an established clinical tool, but things that were more functional in nature were very much in the air among MRI research labs: blood flow, diffusion, and metabolism through MRS were hot topics. The IVIM model of LeBihan et al. (1988) was a promising early attempt at measuring tissue perfusion, and sparked interest in this area. One of my early ideas for providing a more specific measure of perfusion was to implement a time of flight measurement on the millimeter scale (Wong and Hyde, 1991). By exciting a comb of
E.C. Wong / NeuroImage 62 (2012) 660–664
very thin (~0.5 mm) slices, and refocusing a second comb between the excited slices, one could in principle image only those spins that moved from one comb to the other. The idea was that one could adjust the spacing and timing of the slices so that diffusion alone could not bridge the gap between excited and refocused slices, and the method would be more specific to perfusion than the IVIM approach. Given the scale of microvascular flow velocities and T2 relaxation times, it was apparent that in order to test the idea, the excited slices would have to be an order of magnitude thinner than was typical for human systems at the time (~ 3 mm), and that the most obvious means of achieving such slice thicknesses was by using much stronger gradients than the 10 mT/m that was available on our scanner. With no small animal systems available in the lab, the only clear route to stronger gradients was to build a local gradient coil. The first human head gradient coil at MCW The body of published gradient coil design methods was fairly limited at the time, and I was inspired by the target field approach of Robert Turner (1986). This method introduced an analytical means of designing a continuous current distribution that would achieve a target field, and was later adapted to include other geometries, and to minimize inductance and power dissipation. I was interested initially in reducing the errors associated with approximating the derived continuous current distribution by a relatively small number of discrete wires, and began to explore numerical methods to directly optimize wire positions. Using a coil defined by straight segments of wire connecting hundreds of nodes, I used a simple gradient descent algorithm to minimize the error in the desired fields over a region of interest (Wong et al., 1991a). For a recent review of gradient coil design methodology, see Hidalgo-Tobon (2010). As an aside, while my approach turned out to be simply gradient descent, it was formulated in my mind as a correlation method. In our lab, Andrzej Jesmanowicz had long been using correlation techniques for things like T1 mapping, pre-calculating model functions at different T1 values, and determining by correlation the best fit T1 for each voxel. I formulated the gradient coil design problem by asking the question—how well does the change in fields associated with an increment of the location of each wire segment correlate with the error fields? These correlation coefficients collectively provide the direction of motion for the coil segments—equivalently, the downhill direction for the error function. Later, Peter Bandettini would come to me shortly after our first BOLD fMRI experiment, asking for suggestions on how to process the data, and pointing out that a simple subtraction between ‘on’ and ‘off’ periods does not take into account the delay and the transition period between states that were apparent in the data. Thinking back to my gradient coil design approach, I suggested that he “construct a model for what he thinks the response should look like, calculate the dot product between that model and the data from each voxel, and make an image of that”. So began the development of correlation based fMRI analysis at MCW. In 1990, I designed and built a three axis gradient coil with a 4inch diameter, and interfaced it to a 1.5 T GE clinical scanner. Using the standard Techron 8607 linear amplifiers, this coil produced 80 mT/m on the two transverse axes, and 210 mT/m on the Z axis at 70 A, with a ramp time of 56 μs. Help with impedance matching of this very low inductance coil to the gradient amplifiers came from Bob Vavrek of GE, and with RF coil expertise from Jim Hyde and Andrzej Jesmanowicz at MCW, we quickly had an array of custom RF coils to use with the gradient coil. Equipped with this hardware and the sophisticated pulse sequence library and programming environment of a clinical scanner, it quickly became obvious that there were many areas of MRI research that could be pushed forward using this technology. I implemented not only my perfusion method using comb excitation and refocusing, but also projects such as mapping diffusion anisotropy in the rabbit lens (Wu et al., 1993), high
661
resolution hand imaging (Wong et al., 1991b), and real time imaging using EPI (Jesmanowicz et al., 1993; Wong, 1992; Wong et al., 1992a). In early August of 1991, as the annual meeting of the SMRM approached, I had data from rabbit brain for a presentation on my perfusion imaging method, but with a few days remaining, I decided that it would be nice to have human data as well, and set out to design and build a human head gradient coil. At the time, the only local gradient coils for human head imaging were single (Z) axis coils (Turner et al., 1990). The first of these was very simple and elegant—two freestanding bundles of loops with counter-rotating currents formed an open and unobtrusive structure and allowed for the use of existing RF coils. The second used distributed current loops which improved the gradient linearity, but required a closed structure and custom RF coils. Dr. Turner himself visited MCW in late 1990, and provided reassurance that it was indeed possible to safely implement human head gradient coils, as well as advice and encouragement for my gradient coil pursuits in general. Having had experience with smaller three axis coils, I was eager to take the next step and determine whether it was possible to achieve the low length to diameter ratio required for a symmetrical three axis human head gradient coil, with sufficient gradient linearity over the brain. While the Z head gradient had already been demonstrated, the return current paths for the X and Y gradients would be the limiting factor in determining the minimum coil length. The coil design that resulted was produced using the same gradient descent algorithm, programmed on a NeXT Cube computer that I bought using a large fraction of my annual graduate student salary. The GUI for the transverse coil design program, along with the design for the X axis of the resulting coil is shown in Fig. 1. The design would generate gradients of 25 mT/m at 100 A, was torque balanced by symmetry, and was as compact as possible, with a cylindrical region of linearity of 20 cm diameter and 15 cm length, just enough to encompass the brain and no more. The Biophysics lab at MCW contained a full machine shop, used primarily for fabrication of many custom devices for EPR, as well as RF coils for MRI. The shop was run very conscientiously by a machinist and RF engineer named Richard (Dick) Johnson, who has sadly since passed away. Dick was a pleasure to work with, and would have been happy to help with the gradient coil project, but if we had enlisted his help, the coil would have been completed, beautifully finished, months later. Therefore, on the weekend of August 3–4, I planned to build the coil from start to finish between the time Dick clocked out on Friday, and when he returned on Monday morning, using ‘accelerated’ machining methods. The crew consisted of my wife Denise, who was already a veteran of several previous gradient and RF coil construction projects and my most trusted co-craftsperson, and Peter Bandettini, who was new to gradient coil construction, but was as now, up for anything. The coil was built using a PVC pipe with 26.5 cm inside diameter which was obtained as a cutoff from a local construction site, 13 Gauge copper magnet wire, fiberglass cloth, and epoxy. 13 Gauge was chosen because a large spool of it, which Jim Hyde had purchased years earlier at an auction for $1, was lying around the lab. That particular diameter turned out to provide an excellent tradeoff between power deposition and physical size, allowing for a compact design with just enough current carrying capacity to enable sustained EPI scanning without water cooling. The Z gradient was the first layer, and was embedded in the PVC pipe. Using a vertical mill with a rotary table, grooves matching the width of the wire were milled into the pipe, and the wire was tapped into the grooves with a hammer (Fig. 2A). The pipe was then wrapped in layers of fiberglass painted with epoxy, built up thick enough to accept the next layer of wires, and cured for about 2 h. The fiberglass/epoxy was milled down to 0.100″ thickness using the vertical mill and the rotary table motorized by energy transported from a lathe across the room using kite string (Fig. 2B). The complex patterns for the transverse layers were printed onto paper, scaled in a copy machine to life size to be transferred to the surface of the coil with pencil (Fig. 2C). The wire patterns for the next two layers, defined by over 3000 points per layer, were milled into the fiberglass by
662
E.C. Wong / NeuroImage 62 (2012) 660–664
Fig. 1. GUI for gradient descent gradient coil design tool. The design shown is one octant of the X gradient coil designed and built in August 1991. The program was written in Objective C and ran on a NeXT Cube computer.
hand using a Dremel tool (Fig. 2D). The grooves had to be milled to a depth of 0.072″ to match the diameter of the wire, leaving only 0.028″ of clearance to the layer below, which turned out to be too little. Nicks in the layer below were patched with 5-minute epoxy and Teflon tape, wire was pounded into the grooves (Fig. 2E), and the process was repeated for the final layer. Over the course of the weekend, nearly every machine in the shop was used in some ‘off label’ capacity, but by Monday morning the coil was done, and the shop was back to normal, save one drop of dried epoxy on the floor, which Dick dutifully noted. He knew that we had built a gradient coil, but did not ask how. Over the next few days, cables and an RF shield were installed, and a crude linear transmit/receive RF coil was built into the coil. On August 7, 1991, the gradient coil was ready for test. After clinical scanning hours, Peter, Denise and I brought the coil to our 1.5 T scanner and went through an installation process that would become nearly a daily procedure at MCW for the next several years: gradient amplifiers powered down; daughter boards on the gradient amplifiers swapped to custom boards with RC values adjusted to provide stable feedback under a low inductance load; whole body gradient coils disconnected and LGC connected at terminal blocks at the back of the scanner—using a
titanium screwdriver; LGC simply set on the patient table and strapped down with a Velcro patient restraint; and gradient power back on. After some practice, this took about 10 min. The gradient and RF coils worked immediately, first on a phantom, and later the same night on Denise's head. EPI imaging also worked immediately, but we were not able to get useable human data using my comb excitation perfusion imaging method, as the flow encoding process proved too sensitive to patient motion. Nevertheless, we had in place all the necessary hardware and pulse sequence tools for BOLD fMRI just days before we would learn of its invention: a homemade single shot gradient echo EPI pulse sequence and recon; a local head gradient coil to support EPI; and a custom head RF coil. A comprehensive account of the birth and development of fMRI at MCW is given in an article by Peter Bandettini in this issue. A second generation gradient coil was built the next year (Wong et al., 1992a), and would serve the lab at both 1.5 T and 3 T for the next 7 years. This coil was 2 cm larger in diameter, with similar performance specifications, but was built using precision computer controlled machining techniques. A final key hardware element was an elliptical endcapped birdcage coil (Wong et al., 1992b), which was
E.C. Wong / NeuroImage 62 (2012) 660–664
663
Fig. 2. Construction of our first head gradient coil in August 1991. A: Z gradient embedded in PVC pipe. B: Peter Bandettini milling first fiberglass/epoxy layer round for X gradient coil. A kite string (yellow arrow) brought power from a lathe across the room to power the rotary table that turns the coil. C: Denise Wong and Peter tracing X gradient pattern onto coil. D: Eric hand milling grooves for X gradient. E: Denise and Eric tapping wires into place. F: Finished coil, just after the first test. Peter is holding a hardcopy of the first scan in his right hand (note sketches of a design for a surface gradient coil in the background).
an amalgamation of design elements introduced by Cecil Hayes (1986; Hayes et al., 1985), and proved to deliver state of the art whole brain SNR for the remainder of the 1990s. A note on LGC safety Safety issues for use of LGCs in humans include two primary considerations: peripheral nerve stimulation (PNS), and magnetic forces. One of the compelling reasons for using LGCs is that PNS is fundamentally reduced (see below), but because slew rates can be very high, and the current carrying wires can be very close to the body, local PNS is still a concern. The highest dB/dT typically occurs at the shoulders where the body closely approaches a region of the coil that contains return currents for the X gradient. While it was found to be possible to induce PNS when the shoulders are in contact with our coil, a spacing of just 1–2 cm was found to mitigate these effects. Magnetic forces generated within the coil were a more serious concern. While the coil was entirely torque balanced by symmetry, two scenarios could in principle generate net torques or even net forces. While situated in the uniform portion of the magnetic field, no closed loop can generate net forces. However, if the coil were to be pulsed while in the fringe field at the end of the magnet bore, net torques and/or forces could be generated. For this reason, it was critical that the coil was never operated outside of the homogeneous region of the magnet. On our scanner, the control software
would only allow our LGC to be operated when the head portion of the patient table was at isocenter, providing a measure of safety. Nevertheless, it was the job of the operator to insure that the coil was installed at the proper location, and awareness of this potential concern was a critical aspect of operator training. A second scenario under which the coil could produce net torques is a failure that breaks the symmetry of the current pattern. While none of the Z gradient windings produce torques at all, each half of the X and Y coils produce on the order of 100 N m of torque, balanced by the other half. By design, each coil is driven in series, such that any open circuit failure results in zero current, and hence zero torque. A failure in which a portion of one coil shorts out but leaves an intact current path, or shorts to another coil is conceivable, but extremely unlikely, as it would require that an unintended high current connection be spontaneously established. A major goal of the coil construction process is to prevent this from happening. The mechanical connection of the coil to the table, combined with the mass of the coil, is the final safety features. Thankfully, after 20 years and millions of scans, I have heard of no gradient coil failure other than a simple open circuit. Historical perspective Local head gradient coils at MCW and other labs occupied a unique window of opportunity for approximately the decade of the 1990s. In
664
E.C. Wong / NeuroImage 62 (2012) 660–664
1991 the gradient performance of standard commercial whole body scanners was around 10 mT/m and 20 mT/m/ms, while our first head LGC produced 25 mT/m and 400 mT/m/ms, using the same high fidelity linear amplifiers. At this time, the utility of fast gradients for EPI and related methods, and of strong gradients for diffusion and high resolution imaging was already very clear, and of course, the field of fMRI was exploding. The coil itself fit into a very narrow technical niche. Because it was small, short, and symmetrical, it could be made unshielded, without generating significant eddy current effects. In fact, our second head LGC was used to explore spiral imaging, and produced measured spiral trajectories that were accurate to well within one Δk unit with no eddy current correction (Luh and Wong, 1995). The small size and single layer allowed it to be installed quickly and easily, and also allowed it to operate without water cooling. This size would not be practical in a clinical setting because the ROI covered only the brain, and the ergonomics were not ideal. It was also used before parallel acquisition became common, and a simple transmit/receive head coil, which could be made very compact, was still the standard. Even small increases in the size requirements of the coil would significantly increase the complexity of an LGC project. An increase in the diameter would quickly result in increased eddy current effects, inability to mount the coil at isocenter without a custom patient table, and require increased length, which in turn would interfere with the shoulders in a symmetrical design. If a shield layer was required, then the weight would increase from about 25 lbs to at least three times that, and water cooling would be obligatory, further complicating the installation. Over the next 10 years, the major manufacturers gradually pushed whole body gradient performance to approximately 30 mT/m and 120 mT/m/ms, using increasingly large and expensive gradient amplifiers, pushing through the engineering challenges associated with higher power, voltage, cooling requirements, main fields, and vibration. Thus around 2000, EPI, diffusion imaging, and other gradient intensive applications could be performed reasonably efficiently on state of the art whole body systems. Of course, LGCs could further improve performance, but the motivation from the application point of view was limited, and because of the considerations described above, an LGC project would at that point be much more complicated. Indeed, most of the LGC efforts both in academia and in industry that continued through this period were gravitating towards shielded and water cooled gradient sets that were in the hundreds of pounds, requiring special rigging and customized tables to install. In addition, the advances in gradient amplifier technology, including special acceleration circuits and switched mode amplifiers, made interfacing LGCs much more difficult, as these were either tuned to the characteristics of the gradient coil, or depended on the high inductance of the gradient coil to smooth voltage fluctuations from amplifier switching. Thus, for several reasons, interest in LGCs waned. A comeback for the LGC? In the past few years, the industry has come up firmly against a different limitation on gradient performance, that of peripheral nerve
stimulation (PNS). For whole body gradient systems, this limits the useable gradient slew rate to approximately 150–200 mT/m/ms. Because of their smaller size, the field slew rate for a given gradient slew rate is lower for the LGC, and they can be operated at least 400 mT/m/ms, and probably significantly higher, without PNS. This would greatly accelerate high resolution whole brain BOLD fMRI, which is of particular current interest for the study of functional connectivity. Gradient amplitudes in excess of 100 mT/m can also be produced now using head LGCs, which is also about twice the amplitude of current whole body systems, and would be very useful for diffusion imaging. Indeed, the current Human Connectome Project (Connectome, 2010), has as a significant component the development of advanced gradient hardware specifically for these applications. In addition, advances in electronics have produced gradient amplifiers that can again drive low inductance loads, so the technical feasibility of LGCs for use with clinical scanners is also improving. Together, all these factors suggest that there is now a second window of opportunity for LGC. Acknowledgments I would like to thank Peter Bandettini for helpful suggestions and review of this document. Eric Wong is supported by NIH R01 EB002096. References Hayes, C.E., 1986. An End-cap Birdcage Resonator for Quadrature Head Imaging. Society for Magnetic Resonance in Medicine, Montreal, p. 39. Hayes, C.E., Edelstein, W.A., Schenk, J.F., Mueller, O.M., 1985. An efficient, highly homogeneous radiofrequency coil for whole-body NMR imaging at 1.5 T. J. Magn. Reson. 63, 622–628. Hidalgo-Tobon, S.S., 2010. Theory of gradient coil design methods for magnetic resonance imaging. Concepts Magn. Reson. A 36A, 223–242. Jesmanowicz, A., Wong, E.C., Hyde, J.S., 1993. Phase correction for EPI using internal reference lines. Proc., SMRM, 12th Annual Meeting, New York, p. 1239. LeBihan, D., Breton, E., Lallemand, D., Aubin, M.-L., Vignaud, J., Laval-Jeantet, M., 1988. Separation of diffusion and perfusion in intravoxel incoherent motion MR imaging. Radiology 168, 497–505. Luh, W.-M., Wong, E.C., 1995. Single-shot spiral imaging of the human head using a local head gradient coil. Third Meeting, Society of Magnetic Resonance, Nice, France, p. 621. The NIH Human Connectome Project, 2010. http://humanconnectome.org/2010. Turner, R., 1986. A target field approach to optimal coil design. J. Appl. Phys. 19, L147. Turner, R., Le Bihan, D., Maier, J., Vavrek, R., Hedges, L.K., Pekar, J., 1990. Echo-planar imaging of intravoxel incoherent motion. Radiology 177, 407–414. Wong, E.C., 1992. Shim insensitive phase correction for EPI using a two echo reference scan. Proc., SMRM, 11th Annual Meeting, Berlin, p. 4514. Wong, E.C., Hyde, J.S., 1991. Perfusion imaging by interleaved excitation. Tenth Annual Meeting, Society of Magnetic REsonance in Medicine, San Francisco, p. 791. Wong, E.C., Jesmanowicz, A., Hyde, J.S., 1991a. Coil optimization for MRI by conjugate gradient descent. Magn. Reson. Med. 21, 39–48. Wong, E.C., Jesmanowicz, A., Hyde, J.S., 1991b. High resolution short TE MR imaging of the fingers and wrist with a local gradient coil. Radiology 181, 393–397. Wong, E.C., Bandettini, P.A., Hyde, J.S., 1992a. Echo-planar imaging of the human brain using a three axis local gradient coil. Eleventh Annual Meeting, Society of Magnetic Resonance in Medicine, Berlin, p. 105. Wong, E.C., Boskamp, E., Hyde, J.S., 1992b. A volume optimized quadrature elliptical endcap birdcage brain coil. Eleventh Annual Meeting, Society of Magnetic Resonance in Medicine, Berlin, p. 4015. Wu, J.C., Wong, E.C., Arrindell, E.L., Simons, K.B., Jesmanowicz, A., Hyde, J.S., 1993. In-vivo determination of anisotropic diffusion of water, T1, and T2 in the rabbit lens by high resolution MRI. Invest. Ophthalmol. Vis. Res. 34, 2151–2158.
Contents lists available at SciVerse ScienceDirect
NeuroImage journal homepage: www.elsevier.com/locate/ynimg
Review
Local head gradient coils: Window(s) of opportunity Eric C. Wong Departments of Radiology and Psychiatry, University of California, San Diego, CA USA
a r t i c l e
i n f o
a b s t r a c t
Article history: Accepted 1 January 2012 Available online 9 January 2012
At the Medical College of Wisconsin (MCW), prior to the 1991 announcement of the discovery of BOLD fMRI, all of the technical pieces that were needed for efficient BOLD fMRI imaging were assembled for other applications, allowing MCW to jump into the fMRI business just days after the announcement. Central among these pieces was single shot EPI, implemented at MCW using a three axis local head gradient coil. This article describes the development of local gradient coil technology at MCW, and a historical perspective on local head gradient coils in general. © 2012 Elsevier Inc. All rights reserved.
Keywords: Local gradient coil Functional MRI (fMRI) Echo planar imaging
Contents Introduction . . . . . . . . The first human head gradient A note on LGC safety . . Historical perspective . . . . A comeback for the LGC? . . Acknowledgments . . . . . References . . . . . . . . .
. . coil . . . . . . . . . .
. . . . . at MCW . . . . . . . . . . . . . . . . . . . . . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
Introduction On August 12, 1991, I sat with Peter Bandettini in the auditorium of the San Francisco Hilton, watching with mouth agape as Tom Brady showed a movie of EPI brain images getting brighter and darker with brain activity—BOLD fMRI. It was stunning, and everybody in the room knew it was a revolutionary moment. Peter was a graduate student at the Medical College of Wisconsin (MCW), and was looking for a thesis project, something combining MRI technology and neuroscience, and there it was, dropped in his lap. I was toward the end of the PhD phase of an MSTP program at MCW, with a focus on local gradient coil (LGC) design and applications. This article describes the history of the LGC program at MCW, and the role it played in the early development of fMRI at that institution. As a starting graduate student in Biophysics at MCW, I had a choice between EPR, for which the department under the direction of Jim Hyde, was world renowned, and MRI, which was an up and coming
E-mail address: [email protected].
1053-8119/$ – see front matter © 2012 Elsevier Inc. All rights reserved. doi:10.1016/j.neuroimage.2012.01.025
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
. . . . . . .
660 661 663 663 664 664 664
area, but in which MCW had primarily participated in the area of RF coils and the clinical applications thereof. Nevertheless, as an MSTP student, with aspirations of mixing a biophysics background with medicine, MRI was an obvious choice. The training was enriched in a very important way through interactions with GE Medical Systems, which was located only 15 miles to the west in Waukesha, WI. Joint journal club meetings created strong intellectual ties between MCW and GE, which at the time housed Gary Glover, Scott Hinks, Norbert Pelc, Wally Bloch, and Matt Bernstein, among others. The culture of the Biophysics division at MCW was strongly focused on technical innovation. Hypothesis driven research was something that happened elsewhere. In this setting I started in 1988 to look for a thesis project. Anatomical MRI was an established clinical tool, but things that were more functional in nature were very much in the air among MRI research labs: blood flow, diffusion, and metabolism through MRS were hot topics. The IVIM model of LeBihan et al. (1988) was a promising early attempt at measuring tissue perfusion, and sparked interest in this area. One of my early ideas for providing a more specific measure of perfusion was to implement a time of flight measurement on the millimeter scale (Wong and Hyde, 1991). By exciting a comb of
E.C. Wong / NeuroImage 62 (2012) 660–664
very thin (~0.5 mm) slices, and refocusing a second comb between the excited slices, one could in principle image only those spins that moved from one comb to the other. The idea was that one could adjust the spacing and timing of the slices so that diffusion alone could not bridge the gap between excited and refocused slices, and the method would be more specific to perfusion than the IVIM approach. Given the scale of microvascular flow velocities and T2 relaxation times, it was apparent that in order to test the idea, the excited slices would have to be an order of magnitude thinner than was typical for human systems at the time (~ 3 mm), and that the most obvious means of achieving such slice thicknesses was by using much stronger gradients than the 10 mT/m that was available on our scanner. With no small animal systems available in the lab, the only clear route to stronger gradients was to build a local gradient coil. The first human head gradient coil at MCW The body of published gradient coil design methods was fairly limited at the time, and I was inspired by the target field approach of Robert Turner (1986). This method introduced an analytical means of designing a continuous current distribution that would achieve a target field, and was later adapted to include other geometries, and to minimize inductance and power dissipation. I was interested initially in reducing the errors associated with approximating the derived continuous current distribution by a relatively small number of discrete wires, and began to explore numerical methods to directly optimize wire positions. Using a coil defined by straight segments of wire connecting hundreds of nodes, I used a simple gradient descent algorithm to minimize the error in the desired fields over a region of interest (Wong et al., 1991a). For a recent review of gradient coil design methodology, see Hidalgo-Tobon (2010). As an aside, while my approach turned out to be simply gradient descent, it was formulated in my mind as a correlation method. In our lab, Andrzej Jesmanowicz had long been using correlation techniques for things like T1 mapping, pre-calculating model functions at different T1 values, and determining by correlation the best fit T1 for each voxel. I formulated the gradient coil design problem by asking the question—how well does the change in fields associated with an increment of the location of each wire segment correlate with the error fields? These correlation coefficients collectively provide the direction of motion for the coil segments—equivalently, the downhill direction for the error function. Later, Peter Bandettini would come to me shortly after our first BOLD fMRI experiment, asking for suggestions on how to process the data, and pointing out that a simple subtraction between ‘on’ and ‘off’ periods does not take into account the delay and the transition period between states that were apparent in the data. Thinking back to my gradient coil design approach, I suggested that he “construct a model for what he thinks the response should look like, calculate the dot product between that model and the data from each voxel, and make an image of that”. So began the development of correlation based fMRI analysis at MCW. In 1990, I designed and built a three axis gradient coil with a 4inch diameter, and interfaced it to a 1.5 T GE clinical scanner. Using the standard Techron 8607 linear amplifiers, this coil produced 80 mT/m on the two transverse axes, and 210 mT/m on the Z axis at 70 A, with a ramp time of 56 μs. Help with impedance matching of this very low inductance coil to the gradient amplifiers came from Bob Vavrek of GE, and with RF coil expertise from Jim Hyde and Andrzej Jesmanowicz at MCW, we quickly had an array of custom RF coils to use with the gradient coil. Equipped with this hardware and the sophisticated pulse sequence library and programming environment of a clinical scanner, it quickly became obvious that there were many areas of MRI research that could be pushed forward using this technology. I implemented not only my perfusion method using comb excitation and refocusing, but also projects such as mapping diffusion anisotropy in the rabbit lens (Wu et al., 1993), high
661
resolution hand imaging (Wong et al., 1991b), and real time imaging using EPI (Jesmanowicz et al., 1993; Wong, 1992; Wong et al., 1992a). In early August of 1991, as the annual meeting of the SMRM approached, I had data from rabbit brain for a presentation on my perfusion imaging method, but with a few days remaining, I decided that it would be nice to have human data as well, and set out to design and build a human head gradient coil. At the time, the only local gradient coils for human head imaging were single (Z) axis coils (Turner et al., 1990). The first of these was very simple and elegant—two freestanding bundles of loops with counter-rotating currents formed an open and unobtrusive structure and allowed for the use of existing RF coils. The second used distributed current loops which improved the gradient linearity, but required a closed structure and custom RF coils. Dr. Turner himself visited MCW in late 1990, and provided reassurance that it was indeed possible to safely implement human head gradient coils, as well as advice and encouragement for my gradient coil pursuits in general. Having had experience with smaller three axis coils, I was eager to take the next step and determine whether it was possible to achieve the low length to diameter ratio required for a symmetrical three axis human head gradient coil, with sufficient gradient linearity over the brain. While the Z head gradient had already been demonstrated, the return current paths for the X and Y gradients would be the limiting factor in determining the minimum coil length. The coil design that resulted was produced using the same gradient descent algorithm, programmed on a NeXT Cube computer that I bought using a large fraction of my annual graduate student salary. The GUI for the transverse coil design program, along with the design for the X axis of the resulting coil is shown in Fig. 1. The design would generate gradients of 25 mT/m at 100 A, was torque balanced by symmetry, and was as compact as possible, with a cylindrical region of linearity of 20 cm diameter and 15 cm length, just enough to encompass the brain and no more. The Biophysics lab at MCW contained a full machine shop, used primarily for fabrication of many custom devices for EPR, as well as RF coils for MRI. The shop was run very conscientiously by a machinist and RF engineer named Richard (Dick) Johnson, who has sadly since passed away. Dick was a pleasure to work with, and would have been happy to help with the gradient coil project, but if we had enlisted his help, the coil would have been completed, beautifully finished, months later. Therefore, on the weekend of August 3–4, I planned to build the coil from start to finish between the time Dick clocked out on Friday, and when he returned on Monday morning, using ‘accelerated’ machining methods. The crew consisted of my wife Denise, who was already a veteran of several previous gradient and RF coil construction projects and my most trusted co-craftsperson, and Peter Bandettini, who was new to gradient coil construction, but was as now, up for anything. The coil was built using a PVC pipe with 26.5 cm inside diameter which was obtained as a cutoff from a local construction site, 13 Gauge copper magnet wire, fiberglass cloth, and epoxy. 13 Gauge was chosen because a large spool of it, which Jim Hyde had purchased years earlier at an auction for $1, was lying around the lab. That particular diameter turned out to provide an excellent tradeoff between power deposition and physical size, allowing for a compact design with just enough current carrying capacity to enable sustained EPI scanning without water cooling. The Z gradient was the first layer, and was embedded in the PVC pipe. Using a vertical mill with a rotary table, grooves matching the width of the wire were milled into the pipe, and the wire was tapped into the grooves with a hammer (Fig. 2A). The pipe was then wrapped in layers of fiberglass painted with epoxy, built up thick enough to accept the next layer of wires, and cured for about 2 h. The fiberglass/epoxy was milled down to 0.100″ thickness using the vertical mill and the rotary table motorized by energy transported from a lathe across the room using kite string (Fig. 2B). The complex patterns for the transverse layers were printed onto paper, scaled in a copy machine to life size to be transferred to the surface of the coil with pencil (Fig. 2C). The wire patterns for the next two layers, defined by over 3000 points per layer, were milled into the fiberglass by
662
E.C. Wong / NeuroImage 62 (2012) 660–664
Fig. 1. GUI for gradient descent gradient coil design tool. The design shown is one octant of the X gradient coil designed and built in August 1991. The program was written in Objective C and ran on a NeXT Cube computer.
hand using a Dremel tool (Fig. 2D). The grooves had to be milled to a depth of 0.072″ to match the diameter of the wire, leaving only 0.028″ of clearance to the layer below, which turned out to be too little. Nicks in the layer below were patched with 5-minute epoxy and Teflon tape, wire was pounded into the grooves (Fig. 2E), and the process was repeated for the final layer. Over the course of the weekend, nearly every machine in the shop was used in some ‘off label’ capacity, but by Monday morning the coil was done, and the shop was back to normal, save one drop of dried epoxy on the floor, which Dick dutifully noted. He knew that we had built a gradient coil, but did not ask how. Over the next few days, cables and an RF shield were installed, and a crude linear transmit/receive RF coil was built into the coil. On August 7, 1991, the gradient coil was ready for test. After clinical scanning hours, Peter, Denise and I brought the coil to our 1.5 T scanner and went through an installation process that would become nearly a daily procedure at MCW for the next several years: gradient amplifiers powered down; daughter boards on the gradient amplifiers swapped to custom boards with RC values adjusted to provide stable feedback under a low inductance load; whole body gradient coils disconnected and LGC connected at terminal blocks at the back of the scanner—using a
titanium screwdriver; LGC simply set on the patient table and strapped down with a Velcro patient restraint; and gradient power back on. After some practice, this took about 10 min. The gradient and RF coils worked immediately, first on a phantom, and later the same night on Denise's head. EPI imaging also worked immediately, but we were not able to get useable human data using my comb excitation perfusion imaging method, as the flow encoding process proved too sensitive to patient motion. Nevertheless, we had in place all the necessary hardware and pulse sequence tools for BOLD fMRI just days before we would learn of its invention: a homemade single shot gradient echo EPI pulse sequence and recon; a local head gradient coil to support EPI; and a custom head RF coil. A comprehensive account of the birth and development of fMRI at MCW is given in an article by Peter Bandettini in this issue. A second generation gradient coil was built the next year (Wong et al., 1992a), and would serve the lab at both 1.5 T and 3 T for the next 7 years. This coil was 2 cm larger in diameter, with similar performance specifications, but was built using precision computer controlled machining techniques. A final key hardware element was an elliptical endcapped birdcage coil (Wong et al., 1992b), which was
E.C. Wong / NeuroImage 62 (2012) 660–664
663
Fig. 2. Construction of our first head gradient coil in August 1991. A: Z gradient embedded in PVC pipe. B: Peter Bandettini milling first fiberglass/epoxy layer round for X gradient coil. A kite string (yellow arrow) brought power from a lathe across the room to power the rotary table that turns the coil. C: Denise Wong and Peter tracing X gradient pattern onto coil. D: Eric hand milling grooves for X gradient. E: Denise and Eric tapping wires into place. F: Finished coil, just after the first test. Peter is holding a hardcopy of the first scan in his right hand (note sketches of a design for a surface gradient coil in the background).
an amalgamation of design elements introduced by Cecil Hayes (1986; Hayes et al., 1985), and proved to deliver state of the art whole brain SNR for the remainder of the 1990s. A note on LGC safety Safety issues for use of LGCs in humans include two primary considerations: peripheral nerve stimulation (PNS), and magnetic forces. One of the compelling reasons for using LGCs is that PNS is fundamentally reduced (see below), but because slew rates can be very high, and the current carrying wires can be very close to the body, local PNS is still a concern. The highest dB/dT typically occurs at the shoulders where the body closely approaches a region of the coil that contains return currents for the X gradient. While it was found to be possible to induce PNS when the shoulders are in contact with our coil, a spacing of just 1–2 cm was found to mitigate these effects. Magnetic forces generated within the coil were a more serious concern. While the coil was entirely torque balanced by symmetry, two scenarios could in principle generate net torques or even net forces. While situated in the uniform portion of the magnetic field, no closed loop can generate net forces. However, if the coil were to be pulsed while in the fringe field at the end of the magnet bore, net torques and/or forces could be generated. For this reason, it was critical that the coil was never operated outside of the homogeneous region of the magnet. On our scanner, the control software
would only allow our LGC to be operated when the head portion of the patient table was at isocenter, providing a measure of safety. Nevertheless, it was the job of the operator to insure that the coil was installed at the proper location, and awareness of this potential concern was a critical aspect of operator training. A second scenario under which the coil could produce net torques is a failure that breaks the symmetry of the current pattern. While none of the Z gradient windings produce torques at all, each half of the X and Y coils produce on the order of 100 N m of torque, balanced by the other half. By design, each coil is driven in series, such that any open circuit failure results in zero current, and hence zero torque. A failure in which a portion of one coil shorts out but leaves an intact current path, or shorts to another coil is conceivable, but extremely unlikely, as it would require that an unintended high current connection be spontaneously established. A major goal of the coil construction process is to prevent this from happening. The mechanical connection of the coil to the table, combined with the mass of the coil, is the final safety features. Thankfully, after 20 years and millions of scans, I have heard of no gradient coil failure other than a simple open circuit. Historical perspective Local head gradient coils at MCW and other labs occupied a unique window of opportunity for approximately the decade of the 1990s. In
664
E.C. Wong / NeuroImage 62 (2012) 660–664
1991 the gradient performance of standard commercial whole body scanners was around 10 mT/m and 20 mT/m/ms, while our first head LGC produced 25 mT/m and 400 mT/m/ms, using the same high fidelity linear amplifiers. At this time, the utility of fast gradients for EPI and related methods, and of strong gradients for diffusion and high resolution imaging was already very clear, and of course, the field of fMRI was exploding. The coil itself fit into a very narrow technical niche. Because it was small, short, and symmetrical, it could be made unshielded, without generating significant eddy current effects. In fact, our second head LGC was used to explore spiral imaging, and produced measured spiral trajectories that were accurate to well within one Δk unit with no eddy current correction (Luh and Wong, 1995). The small size and single layer allowed it to be installed quickly and easily, and also allowed it to operate without water cooling. This size would not be practical in a clinical setting because the ROI covered only the brain, and the ergonomics were not ideal. It was also used before parallel acquisition became common, and a simple transmit/receive head coil, which could be made very compact, was still the standard. Even small increases in the size requirements of the coil would significantly increase the complexity of an LGC project. An increase in the diameter would quickly result in increased eddy current effects, inability to mount the coil at isocenter without a custom patient table, and require increased length, which in turn would interfere with the shoulders in a symmetrical design. If a shield layer was required, then the weight would increase from about 25 lbs to at least three times that, and water cooling would be obligatory, further complicating the installation. Over the next 10 years, the major manufacturers gradually pushed whole body gradient performance to approximately 30 mT/m and 120 mT/m/ms, using increasingly large and expensive gradient amplifiers, pushing through the engineering challenges associated with higher power, voltage, cooling requirements, main fields, and vibration. Thus around 2000, EPI, diffusion imaging, and other gradient intensive applications could be performed reasonably efficiently on state of the art whole body systems. Of course, LGCs could further improve performance, but the motivation from the application point of view was limited, and because of the considerations described above, an LGC project would at that point be much more complicated. Indeed, most of the LGC efforts both in academia and in industry that continued through this period were gravitating towards shielded and water cooled gradient sets that were in the hundreds of pounds, requiring special rigging and customized tables to install. In addition, the advances in gradient amplifier technology, including special acceleration circuits and switched mode amplifiers, made interfacing LGCs much more difficult, as these were either tuned to the characteristics of the gradient coil, or depended on the high inductance of the gradient coil to smooth voltage fluctuations from amplifier switching. Thus, for several reasons, interest in LGCs waned. A comeback for the LGC? In the past few years, the industry has come up firmly against a different limitation on gradient performance, that of peripheral nerve
stimulation (PNS). For whole body gradient systems, this limits the useable gradient slew rate to approximately 150–200 mT/m/ms. Because of their smaller size, the field slew rate for a given gradient slew rate is lower for the LGC, and they can be operated at least 400 mT/m/ms, and probably significantly higher, without PNS. This would greatly accelerate high resolution whole brain BOLD fMRI, which is of particular current interest for the study of functional connectivity. Gradient amplitudes in excess of 100 mT/m can also be produced now using head LGCs, which is also about twice the amplitude of current whole body systems, and would be very useful for diffusion imaging. Indeed, the current Human Connectome Project (Connectome, 2010), has as a significant component the development of advanced gradient hardware specifically for these applications. In addition, advances in electronics have produced gradient amplifiers that can again drive low inductance loads, so the technical feasibility of LGCs for use with clinical scanners is also improving. Together, all these factors suggest that there is now a second window of opportunity for LGC. Acknowledgments I would like to thank Peter Bandettini for helpful suggestions and review of this document. Eric Wong is supported by NIH R01 EB002096. References Hayes, C.E., 1986. An End-cap Birdcage Resonator for Quadrature Head Imaging. Society for Magnetic Resonance in Medicine, Montreal, p. 39. Hayes, C.E., Edelstein, W.A., Schenk, J.F., Mueller, O.M., 1985. An efficient, highly homogeneous radiofrequency coil for whole-body NMR imaging at 1.5 T. J. Magn. Reson. 63, 622–628. Hidalgo-Tobon, S.S., 2010. Theory of gradient coil design methods for magnetic resonance imaging. Concepts Magn. Reson. A 36A, 223–242. Jesmanowicz, A., Wong, E.C., Hyde, J.S., 1993. Phase correction for EPI using internal reference lines. Proc., SMRM, 12th Annual Meeting, New York, p. 1239. LeBihan, D., Breton, E., Lallemand, D., Aubin, M.-L., Vignaud, J., Laval-Jeantet, M., 1988. Separation of diffusion and perfusion in intravoxel incoherent motion MR imaging. Radiology 168, 497–505. Luh, W.-M., Wong, E.C., 1995. Single-shot spiral imaging of the human head using a local head gradient coil. Third Meeting, Society of Magnetic Resonance, Nice, France, p. 621. The NIH Human Connectome Project, 2010. http://humanconnectome.org/2010. Turner, R., 1986. A target field approach to optimal coil design. J. Appl. Phys. 19, L147. Turner, R., Le Bihan, D., Maier, J., Vavrek, R., Hedges, L.K., Pekar, J., 1990. Echo-planar imaging of intravoxel incoherent motion. Radiology 177, 407–414. Wong, E.C., 1992. Shim insensitive phase correction for EPI using a two echo reference scan. Proc., SMRM, 11th Annual Meeting, Berlin, p. 4514. Wong, E.C., Hyde, J.S., 1991. Perfusion imaging by interleaved excitation. Tenth Annual Meeting, Society of Magnetic REsonance in Medicine, San Francisco, p. 791. Wong, E.C., Jesmanowicz, A., Hyde, J.S., 1991a. Coil optimization for MRI by conjugate gradient descent. Magn. Reson. Med. 21, 39–48. Wong, E.C., Jesmanowicz, A., Hyde, J.S., 1991b. High resolution short TE MR imaging of the fingers and wrist with a local gradient coil. Radiology 181, 393–397. Wong, E.C., Bandettini, P.A., Hyde, J.S., 1992a. Echo-planar imaging of the human brain using a three axis local gradient coil. Eleventh Annual Meeting, Society of Magnetic Resonance in Medicine, Berlin, p. 105. Wong, E.C., Boskamp, E., Hyde, J.S., 1992b. A volume optimized quadrature elliptical endcap birdcage brain coil. Eleventh Annual Meeting, Society of Magnetic Resonance in Medicine, Berlin, p. 4015. Wu, J.C., Wong, E.C., Arrindell, E.L., Simons, K.B., Jesmanowicz, A., Hyde, J.S., 1993. In-vivo determination of anisotropic diffusion of water, T1, and T2 in the rabbit lens by high resolution MRI. Invest. Ophthalmol. Vis. Res. 34, 2151–2158.