INV ITEDP A P E R

Design, Performance, andApplications of a HybridX-Ray/MR System forInterventional GuidanceDigital flat-panel detectors make it possible to view combined x-ray and magnetic

resonance images to more-accurately guide medical diagnosis and treatment.

By Rebecca Fahrig, Arundhuti Ganguly, Prasheel Lillaney, John Bracken,

John A. Rowlands, Zhifei Wen, Huanzhou Yu, Viola Rieke, Juan M. Santos,

Kim Butts Pauly, Daniel Y. Sze, Joan K. Frisoli, Bruce L. Daniel, and Norbert J. Pelc

ABSTRACT | Image-guided minimally invasive procedures

have made a substantial impact in improving patient manage-

ment, reducing the cost, morbidity, and mortality of treatments

and making therapies available to patients who would other-

wise have no option. X-ray fluoroscopy and magnetic reso-

nance imaging (MRI) are two powerful tools for guiding

interventional procedures but with very different strengths

and weaknesses. X-ray fluoroscopy offers very high spatial and

temporal resolution and is excellent for guiding and deploying

devices. MRI offers tomographic imaging with complete

freedom of plane orientation, outstanding soft tissue discrim-

ination, and the ability to portray physiological responses

during treatment. We have shown that it is feasible to fully

integrate an X-ray fluoroscopy system into the bore of an

interventional MR scanner to provide a single congruent field

of view, with integration requiring minor modifications to the

flat-panel digital detector, and using a static-anode X-ray tube.

Given the limited availability of the MR scanner platform (0.5T

GE Signa SP magnet), and the X-ray fluence limitations of the

static-anode X-ray tube, we are now investigating the technol-

ogy developments required to place a rotating-anode digital

flat-panel X-ray system immediately adjacent to a closed-bore

MRI system. These types of hybrid systems could have

enormous impact in the diagnosis and treatment of oncologic,

cardiovascular, and other disorders.

KEYWORDS | Hybrid imaging; interventional image guidance;

magnetic resonance imaging; X-ray fluoroscopy

I . INTRODUCTION

Combining X-ray fluoroscopy and magnetic resonance

imaging (MRI) systems for guidance of interventional

procedures has recently becomemore available, with several

vendors offering dual-modality suites [1]–[8]. In most cases,

these suites require a long travel distance between the two

imaging modalities, on the order of several tens of feet. Thisdesign strategy was driven by the use of relatively standard

X-ray components (e.g., image intensifiers and rotating

anode X-ray tubes) and by the need to ensure minimal

impact of one system on the other. As a result, switching

back and forth between the two modalities is time

consuming and awkward, leading to potential risk and loss

of image registration. It would also invariably lead to less

Manuscript received July 17, 2007; revised xxxx. This work was supported in part by the

National Institutes of Health under Grants NIBIB R01 EB000198 and RR09784, GE

Healthcare, NSERC, and the Lucas Foundation.

R. Fahrig, A. Ganguly, V. Rieke, K. B. Pauly, D. Y. Sze, J. K. Frisoli, B. L. Daniel, andN. J. Pelc are with the Department of Radiology, School of Medicine, Stanford

University, Stanford, CA 94301 USA (e-mail: fahrig@stanford.edu).

P. Lillaney is with the Departments of Radiology and Bioengineering, Stanford

University, Stanford, CA 94301 USA.

J. Bracken and J. A. Rowlands are with the Department of Medical Biophysics,

University of Toronto and Sunnybrook Health Sciences Center, Toronto,

ON M4N 3M5, Canada.

Z. Wen was with the Departments of Radiology and Physics, Stanford University,

Stanford, CA 94301 USA. He is now with the Department of Medical Physics,

University of Wisconsin-Madison, Madison, WI 53706 USA.

H. Yu was with the Departments of Radiology and Electrical Engineering, Stanford

University, Stanford, CA 94301 USA. He is now with GE Healthcare, Menlo Park,

CA 94086 USA.

J. M. Santos is with the Department of Electrical Engineering, Stanford University,

Stanford, CA 94301 USA.

Digital Object Identifier: 10.1109/JPROC.2007.913506

468 Proceedings of the IEEE | Vol. 96, No. 3, March 2008 0018-9219/$25.00 �2008 IEEE

than optimal use of the modalities since one may bereluctant to move the patient repeatedly. In a recent paper

by Dick [9] it was emphasized that during interventional

procedures in such suites Bmultiple external devices and

connections required careful supervision and several minutesof preparation when moving in either direction.[

The advent of digital flat-panel detectors that are more

resistant than X-ray image intensifiers to stray magnetic

fields opened the door to alternative hybrid imaging systemsthat place both modalities in close proximity. This allows

faster, easier switching between modalities, maintaining

multimodality registration and providing a truly hybrid

imaging system. This technology enabled the development

of our static-anode X-ray fluoroscopy system integrated

into the bore of a vertical gap interventional MRI scanner

(Signa SP, GE Milwaukee, WI)Vthe SP-XMR. Here we

describe the technical challenges that were overcome toallow close integration of the two imaging systems

including the following critical issues: 1) X-ray tube and

detector operation within the magnetic field, 2) MR field

homogeneity and noise properties with the X-ray system in

place, and 3) registration and integration approaches.

Our experience with the SP-XMR system is guiding

development of our next generation of truly hybrid XMR

systems. Since the vertical gap MRI system is no longercommercially available and was limited in field to 0.5 T,

attention is turning to short closed-bore systems, such as

the Espree (Siemens Medical Solutions, Erlangen,

Germany; 1.2 m long, 75 cm inner bore diameter). This

and/or similar systems from other vendors may become

the most prevalent interventional MRI platform, and we

therefore propose to widen the availability of truly hybrid

XMR systems by developing the technology for placementof an X-ray system directly adjacent to a high-field closed-

bore MRI system while maintaining the shortest possible

distance between the two systems.

II . THE SP-XMR SYSTEM

A. System ConfigurationThe current implementation of the SP-XMR consists of

1) the GE BInova[ fluoroscopy capable flat-panel detector

with a 20� 20 cm2 area and 0.2� 0.2 mm2 pixels slightly

modified for operation at high magnetic fields, placed at a

field of 0.5 T; 2) the OEC 9800 workstation, which

controls X-ray exposure, image acquisition, processing and

display, data communication, and archiving; and 3) the

fixed-anode X-ray generator and mainframe from an OEC8800 mobile C-arm system modified to allow separation of

the high-voltage generator from the tube (distance 10 m)

and with the X-ray tube mounted in a small housing. Both

the X-ray tube and housing were modified to remove all

magnetic components. The X-ray tube is in a field of

strength�0.2 T when placed in the overhead bracket, with

the external magnetic field aligned with the internally

applied electric field between the anode and cathode. Thehigh-voltage cables are passed through waveguides in the

penetration panel, with radio-frequency (RF) shielding

between the penetration panel and generator to eliminate

RF leakage into the room. The X-ray tube is installed in the

bracket between the two donut-shaped superconducting

magnets, allowing a true anterior-posterior (AP) view of a

patient supported axially in the magnet. A homemade

collimator with remote controls has been built and installed,allowing intraprocedural modification of the X-ray field of

view (FOV). Construction of an integrated cradle-handling

platform for side-dock procedures allows positioning flexi-

bility of the detector. The overall system is shown in Fig. 1.

We have found that automatic exposure control, online

image subtraction [e.g., for digital subtraction angiography

(DSA) studies], continuous image storage, and image trans-

mission via Digital Imaging and Communications inMedicine (DICOM) format to other systems (e.g., the MR

real-time imaging workstation) are critical during clinical

interventions. Amore detailed description of systemupgrades

and modifications has been provided previously [10].

B. Performance of the X-Ray System in theMagnetic Field

Conventional fluoroscopic X-ray systems present, inprinciple, several problems that prevent operation in a

magnetic field. X-ray image intensifiers (XRIIs) and X-ray

tubes both contain electrons moving in vacuum through

relatively large distances (20–60 cm and 1–2 cm, respec-

tively). The deflection of these moving electrons by a

magnetic field can be considerable, making both devices

difficult to operate in the stray field of a powerful

electromagnet. Fortunately, the invention and developmentof the solid state flat-panel fluoroscopic imager to replace the

XRII has reduced the problem of using a detector in a

magnetic field to manageable proportions. There is,

unfortunately, no such simple replacement possible for the

X-ray tube, for which a vacuum approach is essential. For

this case, a more incremental approach is necessary and is

discussed below. In addition, we need to understand the

operation of the induction motor used, ingeniously to permitthe rotation of the anode inside the vacuum of the tube,

without the need for a rotating vacuum seal. Finding a

replacement for this mechanism is challenging, and so

means for extending the utility of a conventional tube in a

magnetic field are carefully considered.

DetectorVInvestigation of Modulation Transfer Function(MTF) and Detective Quantum Efficiency (DQE): Flat-paneldetectors can be either direct conversion (X-ray transduced

directly into a charge image by a photoconductor such as

amorphous selenium and stored on individual pixel

capacitors) or indirect conversion (X-ray transduced to light

in a phosphor, often cesium iodide, and this light is absorbed

and converted into a charge image by individual photodiodes

at each pixel). In both conversion methods, a very large

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Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 469

integrated circuit built on a monolithic glass substrate

coordinates the acquisition, temporary storage, and readoutof the latent charge image. This large integrated circuit is

called an active matrix and uses control (gate) lines and data

(readout) lines formed lithographically to sequence out the

latent image charge stored at the pixels (see Fig. 2). The core

technology enabling such large integrated circuits is

hydrogenated amorphous silicon (a-Si:H), which is the

same technology used in liquid crystal displays for laptop and

desktop monitors [11].There is no need for any ferromagnetic components on

the active matrix array itself, and in practice, the

manufacturing conditions exclude them. This is fortunate,

as it would be very difficult and expensive to change these

processes. In contrast, however, the auxiliary electronics

needed to drive the active matrix (gate drivers, data line

amplifiers and multiplexers, power supplies, transformers,

and timing circuits) built on a printed circuit board as wellas the physical housing are generally made without any

consideration of their magnetic compatibility, and there-

fore it is to be expected that multiple magnetically

sensitive and ferromagnetic components will be present.

However, since these are built individually from compo-

nents and not mass-produced in a clean room as is the

active matrix, there is ample opportunity to modify and

eliminate troublesome parts.The only active components on an active matrix array are

thin-film transistors [(TFTs), metal–oxide–semiconductor

field-effect transistors], which are used as switches to

control the flow of charge from the pixels to the data lines.When the TFT is off, there is no current passing through the

a-Si:H channel and therefore no opportunity for the

magnetic field to affect it in any way; when it is on, a very

Fig. 2. Schematic diagram of themain components of an activematrix

array. The only active components are the thin-film transistors,

which have currents that are too small to be affected by external

magnetic fields.

Fig. 1. Installation and components of the SP-XMR hybrid system. (a) X-ray system installed in the bore of the SP magnet with patient shown

in ‘‘side dock’’ geometry. (b) High-voltage generator and mainframe in the equipment room of the magnet. (c) Digital flat-panel detector.

(d)Static-anodeX-ray tube insertwithmagnetic components removed. (e)Closeupofcollimatorcontrols,hiddenbehindX-ray imagedisplay in (a).

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470 Proceedings of the IEEE | Vol. 96, No. 3, March 2008

tiny current flows in the channel of the TFT. However, thedrift velocity of the carriers is so low that the deflection of

the electrons is negligible; thus the TFT is also immune to

magnetic fields when on. In addition, the very small size of

the current also ensures that there is no magnetic field to

speak of generated by the TFT. Flat-panel detectors also

contain photodiodes: X-ray photodiodes in direct conversion

systems and optical photodiodes in indirect conversion

systems. Here again the intrinsically low drift velocities ofcarriers and the tiny quantities of generated charge cannot

either be deflected significantly by a magnetic field or

generate any external field themselves.

Thus in conclusion, flat-panel detectors (both direct and

indirect conversion) contain no essential intrinsic compo-

nents that are magnetic or cause magnetic disturbances, but

do typically contain extrinsic components or housings that

must be modified or replaced. We found that simplemodifications to the onboard power conditioning circuit

(replacement of inductive filters and extension of cables so

that power supplies experience fields below 0.015 T) of the

GE Inova flat panel was sufficient to ensure robust operation

of the detector when the panel was placed in fields as high as

0.5 T. To verify this, we carried out two experiments to

compare the resolution and noise properties of the detector

at fields of 0 and 0.5 T using standard techniques to measureMTF using a slanted edge and DQE [12]. We found that

inherent detector resolution was not affected by the

magnetic field, although overall system resolution measured

at the location of the patient showed a decrease of�10% for

spatial frequencies of 0.5 cycles/mm and above due to a

change in the focal spot size (see below). The noise in the

X-ray imaging system was identical inside and out of the

magnetic field.

X-Ray TubeVInvestigation of Focal Spot Size andDistribution With E and B Parallel: In our SP-XMR system,

the strength of the magnetic field at the location of the

tube is �0.2 T. Interaction between this strong magnetic

field and the electron beam can lead to deflection and/or

defocusing of the beam. The relatively simple design of the

system allowed us to minimize the deflection of theaccelerating electrons by aligning the anode-cathode

direction with the main magnetic field that, by symmetry,

is perpendicular to the plane midway between the two

donuts (see Fig. 1). Each component of the X-ray tube was

investigated in detail, including the effects of the ac

filament current causing mechanical oscillation of the

filament (shown to be negligible) [10] and the impact of

field strength on focal spot distribution, described here.A finite-element (FE) program (OPERA-3d; Vector

Fields, U.K.) was used to simulate the electron trajectories;

care was taken to match the modeled geometry to our fixed-

anode X-ray tube [13]. Thermionic emission was modeled

using Child’s law current limit model (temperature: 2500 K;

work function: 4.5 eV; total tube current: 4 mA). The

electrostatic Poisson’s equation was solved numerically to

obtain the electrostatic field, including the effects caused byspace charge in the electron beam. The space-charge density

was found by calculating the trajectories of electrons emitted

from the filament under the influence of the electrostatic

field and an external uniform magnetic field. Focal spot

Bimages[ were obtained by plotting and projecting the

electron current density map on the target through a

modeled pinhole to an image plane.

Experimental verification of themodel was carried out byobtaining pinhole images of the focal spot (30 �m tantalum

pinhole 12.4 cm from the X-ray tube, with an overall

magnification of 8.9) with tube and detector mounted on an

MR-compatible stand. A Tesla meter (F.W. Bell 5080; Sypris

Test & Measurement, Orlando, FL) was used to define the

point midway between the two donuts where the magnetic

field strength was 0.208 TVa field strength equivalent to

that at the focal spot location of the installed system. Imageswere acquired at 65 kVp and 4 mA at 0.208 T, and for

comparison in the anteroom of the MR suite at a residual

field of less than 0.001 T using the same stand and geometry.

Results for fields of 0 and 0.208 T are shown in Fig. 3.

The increase in focal spot size shown here accounts for the

change in system resolution (MTF) when the imaging

system operates within the magnetic field. These results

show that even low field strengths affect the size and currentdensity distribution of the focal spot of the X-ray tube.

C. Performance of the MR System in the Presence ofthe X-Ray System

Three main concerns arise when placing X-ray

components close to theMR FOV: 1) magnetic components

Fig. 3.Magnetic field dependent focal spot size at two field strengths:

top ¼ 0 mT, bottom ¼ 208 mT showing FE simulations of the electron

trajectories, the resulting simulated focal spots, and measured focal

spots obtained by placing the nonmagnetic X-ray tube at the fields

strengths indicated.

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Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 471

could cause B0 field inhomogeneity leading to imagedistortion and signal loss; 2) eddy currents could cause

artifacts; and 3) RF noise could cause both loss of signal-to-

noise ratio (SNR) and zipper artifact. As a first step, we

reduced, as far as possible, magnetic components in both

tube and detector. In the X-ray tube, the nickel focusing

cup and all screws, etc., were replaced with nonmagnetic

stainless steel, and the tube housing was constructed out of

brass. For the detector, the housing was redesigned inaluminumwith a carbon-fiber front face, all housing screws

were replaced with nonmagnetic stainless steel, and nickel-

plated connectors were replaced with plastic equivalents.

The packaging of some small components in the detector

was ferromagnetic, and nonmagnetic alternatives were not

available. Even though these components were very small,

they still caused some distortion in B. Below we describe a

solution for this problem as well as our investigations ofeddy currents and RF noise.

Magnetic Field Homogeneity: To obtain a map of the

main magnetic field within the imaging volume of the

scanner, complex images of a large doped-water phantom

were collected at two echo times using a two-dimensional

gradient echo (GRE) sequence (34 � 34 cm2 FOV, 66

3 mm slices, (�TE of 2 or 3 ms). From the unwrappedphase difference map of two echo time images, the raw

field map was calculated. The zeroth-order offset and the

first-order linear component were removed since they can

be handled by the scanner center frequency and linear

shims. We assume that the ferromagnetic components

(FCs) inside the detector are located at a few discrete

locations and are small diameter (� a few centimeters)

much smaller than the isocenter-to-detector distance (�30cm) (see Fig. 1). Each FC may therefore be approximated

by a magnetic dipole pointing in the direction of the

external field and located near the FC. This suggests that

appropriate placement of permanent magnets may be used

to eliminate the effect of the magnetic components. For

this to be practical, the intrinsic coercivity of the

permanent magnet must be greater than the external

magnetic field, which ensures that the permanent magnetwill not be demagnetized by the external field. Therefore

the magnetic field disruption caused by the detector can be

canceled by the permanent magnets oriented in a direction

opposite to B.

This passive shimming approach was investigated

using NdFeB magnets (Dexter Magnetic Technologies,

Elk Grove Village, IL) of various sizes and strengths, with

optimum placement determined using a fitting procedure.The field degradation from the detector and the

effectiveness of the passive shimming were quantified

by measuring the peak-to-peak field inhomogeneity (p-p)

and the standard deviation of the field ð�fÞ in a

cylindrical volume ð�28 � 20 cm2Þ. It was also evaluatedqualitatively with a steady-state free precession (SSFP)

sequence (TR/TE¼ 11:64=5:56 ms). The SSFP sequence

was selected for these tests since it is important forinterventional guidance and is very sensitive to magnetic

field inhomogeneity.

The p-p and �f were 5.4 and 0.55 �T, respectively,without the detector present. The inhomogeneity param-

eters were much larger with the detector in place, 12.6

and 1.68 �T, respectively. Nonlinear least squares

fitting, optimizing both strengths and locations of the

dipoles, predicted that two ideal dipoles would reducethe field inhomogeneity with the detector to 5.5 (p-p)

and 0.52 �T ð�fÞ. These two dipoles were physically

realized using two sets of permanent magnets, each

set consisting of a number of identical NdFeB magnets

(0.2 � 0.2 � 0.5 in3Þ matching the ideal dipole

strength. The two sets of magnets were placed just

below the detector in custom-made holders at locations

prescribed by the fitting (relative to the center of thedetector, with the power and other cables pointing in

the �y direction, one at x ¼ �8:5, y ¼ 1:5 cm and the

second at x ¼ 9:9, y ¼ 1:8 cm). The field map of the

detector with the shimming magnets in place showed a

deviation of 5.6 (p-p) and 0.57 �T ð�fÞ, compared to 5.4

and 0.55 �T for the baseline field with no detector in

the magnet. Fig. 4 clearly shows the severe impact of the

inhomogeneity due the presence of the detector and theimprovement achieved from the passive shimming with

high-coercivity magnets.

Dynamic Effects: Wrapping the detector in multiple

layers of aluminum foil (�100 �m in total) provides

sufficient reduction of RF noise such that the X-ray

detector can be left on during MR imaging. This allows the

detector to reach temperature equilibrium, which, bystabilizing the gain, reduces the fixed pattern noise of the

X-ray images. To verify the performance of the shield, MR

images of a water phantom were taken using a GRE

sequence (TR/TE 150/minfull) and a transmit/receive

Fig. 4. SSFP images showing the effects of field homogeneity with

(a) baseline (i.e., no X-ray detector present), (b) with the detector in

the bore of the magnet, and (c) with the detector in the bore of the

magnetic but with permanent magnet shimming to reduce the

effects of main B-field inhomogeneity.

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472 Proceedings of the IEEE | Vol. 96, No. 3, March 2008

flexible body coil. Signal-to-noise ratios for baseline,detector in and on with no shielding, and detector in

and on with Al foil shielding were 7.48, 5.97, and 7.18,

respectively. The small decrease of 4% from baseline was

considered acceptable for operation of the system.

Presence of the Al foil attenuates the X-ray beam by only

1% at 50 keV, the mean energy of our X-ray beam when

imaging at 80 kVp.

We carried out a study of eddy current effects using astandard image-based system-level test. Eddy current

measurements were made using the manufacturer’s tools

(BGrafidy[) and with the assistance of the service

engineer. The X-ray tube had no effect, at least in part

because is it outside the volume in which the gradient coil

has its main effect. When the X-ray detector was in

position below the patient cradle (and well within the

volume affected by the gradient coil), changes in the time-dependent field produced by the gradient coil could be

observed with Grafidy. The effects were of a magnitude

and temporal dependence that was felt could be compen-

sated using the preemphasis filter. We considered having

two eddy-current compensation coefficients, to be used

when the detector is inside or outside of the bore,

respectively. However, before recalibrating the eddy-

current compensation, we examined whether any effectscould be seen in images. Seeing no significant effects, we

decided this complication was not necessary at this time.

In other configurationsVfor example, with stronger

gradients or larger field-of-view detectors, the situation

could be different and dedicated eddy-current corrections

may be necessary.

D. System IntegrationFor maximum utility, the X-ray and MR units in a

hybrid system must be integrated such that switching

between modalities is fast, easy, and efficient; registration

between systems is maintained; and each system can be

used to prescribe the acquisition geometry of the other.

Patient motion between X-ray and MR imaging should also

be minimized, and thus all objects should be MR and X-ray

compatible so that they do not need to be moved in and outof the FOV as the imaging modality is switched. Our

registration approach, and construction and design of an

X-ray compatible RF coil, are described here.

System Co-Registration: The key step in coregistration is

to accurately determine the positions of the X-ray

components in MR coordinatesVthe geometric param-

eters described by the projection matrix. Sixteen fiducialmarkers, provided by the crossing points of lead-cored

fishing line visible in both MR and X-ray images, are

placed in the FOV; detection is carried out semiautomat-

ically in X-ray and MR images [14]. By fitting these MR

and X-ray measurements to the nonlinear function

describing the projection process, the values of the

geometric parameters are determined. Two possible

sources of registration error, fiducial marker measurementnoise and MR gradient nonlinearity, were studied. To

illustrate the use of the calibrated geometric parameters

and to validate their accuracy, we used X-ray images as

Bscouts[ to prescribe MR slices. The calibration process

was implemented as a GUI-based package (Fig. 5) on the

same computer platform as RTHawk, a robust real-time

interactive imaging platform for MR scanner control

[15]–[20], developed at Stanford.

X-Ray Compatible RF Coils: For the design of an X-ray

compatible RF-coil, all coil components that potentially

lie in the FOV of the X-ray image, i.e., in the path of the

X-ray beam, should have minimal X-ray attenuation. To

minimize the attenuation by the loop conductor, we used

aluminum (Al), which has a significantly lower atomic

number ðZ ¼ 13Þ than the commonly used copper (Cu,Z ¼ 29) [21]. Using aluminum foil, we built a four-

element abdominal phased array (PA). The capacitors and

copper lead wires, detuning circuits, and baluns were

placed outside the X-ray FOV, so that in normal use they

would not appear in the X-ray image.

MR image quality of the X-ray compatible PA was

compared to a single-channel receive-only coil (SPGR, TR/

TE/flip/BW¼ 40 ms=5:5 ms=50=31 kHz) and showed ap-proximately a 60% improvement in SNR. Measurements

of the quality factor of the coils resulted in an unloaded/

loaded Q ratio of �270/40 for the individual elements of

the array. For comparison, a conventional abdominal PA

coil of similar design had an unloaded/loaded Q ratio of

185/50. This showed that replacing copper conductors

with aluminum did not noticeably influence the noise

performance of the coils. An X-ray/MR image pair,acquired during a transjugular intrahepatic portosystemic

shunt (TIPS) placement using an X-ray compatible coil, is

shown in Fig. 6. This image can be compared with those

shown in Fig. 7, early X-ray images acquired without the

X-ray compatible coil.

E. Clinical Operation of the SP-XMR SystemThe X-ray tube and generator are permanently installed

in the MR suite and the equipment room, respectively. Prior

to a clinical procedure, the X-ray detector is placed on a

modified mount under the patient table, and an X-ray

compatible cradle is installed. Switching from X-ray to MR

takes only 60 s and requires the following: a) X-ray system

software is placed in a Bstandby[ state and b) X-ray detectorand display are powered off to prevent RF noise from the

display. Switching fromMR to X-ray takes the same amountof time. Note that during the first group or procedures

carried using SP-XMR guidance, the detector was lowered

by 5 ft during MR scanning, but this is no longer necessary

with the passive shimming described above.

To date, we have completed 56 patient studies under

SP-XMR guidance: 16 TIPS [22], 11 hysterosalpingo-

grams, a brain biopsy [23], a vaginal reconstruction, a cyst

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Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 473

drainage, a facial vascular malformation sclerotherapy, 11other arteriovenous malformations, a prostate seed

implantation, 3 arthrograms, and 10 cystography studies

looking for renal reflux in infants and young children. For

all studies, several switches back and forth between imagingmodalities were typically required, with X-ray typically

providing device guidance or visualization of small,

contrast-filled structures, while MR was used to visualize

critical soft-tissue structures or to provide three-dimen-sional physiological information.

One particularly illustrative example of the utility of

the XMR system was provided during the creation of a

neovagina for an 18-year-old patient with a severe

congenital anomaly (cloacal atresia) where the caudal

spine was unformed and she was born without a urinary

bladder, urethra, rectum, anus, and vagina. During a dozen

previous surgeries, a loop of small bowel was used to createa neobladder, with a urostomy for drainage, and a

colostomy was constructed to allow evacuation of the

bowel. The patient also had separate halves of a uterus,

neither of which was connected to a cervix or vagina.

Menstrual cycles had started in one half of the uterus,

Fig. 5. (a) System schematic showing the relationship among magnet control, MR console, real-time imaging module, and X-ray/MR calibration

module. The GUI interface for the calibration module is shown in (b), with an MR image of one plane of the X-ray/MR compatible lead-core

fishing line phantom and the projection matrix that defines the relationship between the X-ray system FOV and the MR system FOV.

The RTHawk real-time interface is shown in (c). The software XScout permits definition of an MR plane of interest by drawing a line on

the X-ray projection image and combining the known location of the line in the detector plane with the calibrated projection matrix

values from XCalib. The parameters of the plane of interest are then passed to RTHawk.

Fig. 6. Composite imagegenerated fromimagesacquiredduringaTIPS

procedure. Thewhite arrowhead indicates the trocar and needle being

directed towards the hepatic vein (arrow) during MR guidance of

the puncture. The X-ray image was acquired later in the procedure

during balloon expansion of the stent along the tract of the needle.

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474 Proceedings of the IEEE | Vol. 96, No. 3, March 2008

causing accumulation of blood in the contained uterine

cavity and attached fallopian tube. The other half

contained no menstruating tissue but developed a large

retained stone. A kidney had also migrated incorrectlyduring development, resting in the bottom of the pelvis.

The challenge, therefore, was to create a connection

between the perineal dimple (vaginal remnant on the skin

surface) and the menstruating uterus in order to allow

normal menstrual flow while avoiding the neobladder,

kidney, bowel, and surrounding blood vessels. Using our

hybrid system, we performed percutaneous transperineal

puncture of the imperforate uterus under MR guidanceusing numerous oblique planes of imaging, verified the

path of the needle relative to sensitive structures by

injecting contrast under X-ray imaging, and then created a

neovagina by progressively balloon-dilating the channel

along the needle tract. Images acquired during the

procedure are shown in Fig. 7. The ability to obtain

three-dimensional MRI images, combined with the speed,

sensitivity, projectional imaging, and spatial resolution ofreal-time X-ray fluoroscopy, allowed successful tract

creation without injury to surrounding structures in thispatient with no surgical option.

III . THE FUTURE OF TRULYHYBRID XMR

The SP-XMR system, based on a vertically open magnet,

allows the X-ray and MR systems to have overlapping fields

of view with no need to move the patient or equipment

when switching between modalities. This is ideal from

many respects, including integration of the information

from the two techniques. On the other hand, the image

quality performance of both the X-ray and MR systems islimited. The MR performance is limited by the low field

strength and modest gradient performance of the Signa-SP

while the X-ray performance is limited by X-ray tube

output (static anode X-ray tube) and the inability to easily

obtain views other than the AP projection. It is possible to

increase the performance of the X-ray system within this

architecture. Nonetheless, it is doubtful that the X-ray

subsystem could match the capabilities of a high perfor-mance X-ray angiography system. At the opposite extreme,

if the X-ray and MR systems are widely separated so as to

prevent any interaction between them, one would be able

to obtain the best performance from each system at the

expense of integration of the two modalities. The wide

separation would also preclude frequent switching be-

tween modalities, which most likely would lead to limited

use of one or the other.Because of the limitations of commercially available MR

systems based on open magnets, many sites investigating

interventional MRI are turning to short, closed-bore

systems, such as the Siemens Espree (1.2 m long, 75 cm

inner bore diameter) for interventional MRI guidance.

While these systems do not match the full capabilities of the

highest performanceMR systems, their capabilities are quite

good (e.g., allowing functionalMRI, diffusion, and perfusionstudies) while the wider and shorter bore provides much

better patient access than in conventional systems (though

not as good as that in open systems). Placement of an X-ray

fluoroscopy system directly adjacent to a closed-bore MRI

system, while maintaining the shortest possible distance

between the two systems, offers an interesting compromise

between the two extreme architectures described above.

With the Siemens Espree, for example, a separation of only1.1 m is realizable between the imaging volumes of the two

systems, and the mechanical transport system of the MR

scanner can be used to maintain easy registration between

the fluoroscopy and MR systems. This patient/table travel is

on the same order as is currently seen in the cath-angio suite

when a clinician inserts a catheter-based device through

groin access and then follows the catheter into a structure in

the brain.As stated above, one important limitation of our

current X-ray/MR system is X-ray tube output. In fact,

in the minimally invasive abdominal interventions for

Fig. 7. (a)–(c) MR images acquired during passage of the needle from

the perineal dimple into the isolated left-sided uterus. (d) Guidewire

coiled in the lumen of the uterus (in place to maintain tract) and the

injection of radiopaque contrast during retraction of the needle

along the tract in order to verify that the needle did not injure the

neobladder or other structure. (e) Injection of contrast directly into

the neobladder through the stoma; a faint blush from the previous

injection [i.e., shown in (c) during withdrawal of the needle] can be

seen; there is no evidence of connection between bladder and tract.

In this early study, the white arrowheads point to artifacts

originating from the non-X-ray compatible patient cradle; black

arrowheads point to artifacts from the non-X-ray compatible RF coil.

Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 475

which the XMR system was used, we could not treat

heavy individuals because the heat loading capacity,

and thus X-ray flux, of the static-anode X-ray tube is

limited. Meeting current X-ray image quality standards

requires a rotating-anode X-ray tube to enable veryshort, high exposure, low-noise imaging. When com-

bined with a small separation between X-ray and MR

systems, the X-ray tube experiences fringe fields of

significant magnitude with a direction that is highly

dependent on position.

Schematics of two possible implementations, for

guidance of neuro-angiographic interventions and for

guidance of cardiac interventions, are shown in Fig. 8with geometry of closest approach given the somewhat

different imaging requirements. Coronary artery X-ray

imaging requires high output with very short pulses for

optimum visualization of cardiac vessels (typically 9 10 kW

pulses, with a 5 ms pulse width, at 80 kVp). The need for

oblique views of the cardiac vessels places the X-ray tube

farther from the magnet shroud since angulation is

required. The neuroimaging application is less exactingwith respect to X-ray tube output; longer X-ray pulse

widths can be tolerated since motion is small. However,

since most device navigation can be achieved using

standard anteroposterior and lateral views, the X-ray

system can be placed closer to the magnet shroud. Both

applications require high resolution, i.e., good focal spot

distribution, for visualization of small vessels, with a focal

spot size on the order of 0.8 mm � 0.8 mm2. Note that thedistance from the patient table to the X-ray tube is

determined based on government regulations that fix the

minimum source-to-skin distance at greater than 38 cm,

and therefore typical values for the focal spot to center of

rotation of clinical C-arm systems is �66 cm (i.e., for

patient total thickness on the order of 40 cm).

We have shown that a thin-film transistor-based large-

area digital X-ray detector is essentially immune to

external magnetic fields and have verified operation in

magnetic fields as high as 0.5 T. We have also shown that

the electron optics of the X-ray tube are very sensitive toexternal fields, and therefore we orient the motor of the

X-ray tube towards the shroud of the magnet so that it

experiences a somewhat higher field than is present at the

focal spot (separation between motor and electron beam is

�10 cm). An estimate of the magnetic fields that the X-ray

tube components would experience, for the geometries in

Fig. 8, is summarized in Table 1 for a 1.5 T Signa magnet

(GE Medical Systems).

Fig. 8. Two possible implementations of the closed-bore XMR system. The installation shown in (a) permits anteroposterior and lateral

views, while the geometry of (b) with the X-ray system placed slightly farther away from the shroud of the magnet allows angulation

for cardiac applications.

Table 1 Strength of the Magnetic Field at the Focal Spot Location (F.S.)

and at the Motor for the Two Geometries Illustrated in Fig. 8. Br is the Field

Strength in the Radial Direction and Bz is the Field Strength in the Axial

Direction. For the Neurointerventional Applications (Neuro), the Distance

From the Center of the MR Field of View to the X-Ray Tube Components Is

Determined Only by the Length of the Magnet and the Radial Location of

the X-Ray Tube. For the Cardiac Applications (Cardiac and Car. Oblique),

the Need for Angulation Initially Places the X-Ray Tube Farther From the

Magnet Shroud, But Angling the Gantry Then Places the Tube at Smaller R

and Therefore in Higher Magnetic Fields

Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

476 Proceedings of the IEEE | Vol. 96, No. 3, March 2008

These geometries have guided our initial investigationsof the operation of a rotating-anode X-ray tube when

placed in external fields of the magnitudes described

above, and the behavior of both a standard induction motor

and the electron optics were characterized [24].

X-Ray Tube Motor in Fringe Field of Closed-Bore Magnet:When an induction motor is operating in the presence of

an external magnetic field, the stator soft iron ringbecomes partially magnetized by the external field, in

addition to the magnetization produced by the stator coils.

These effects might impact the performance of the motor

and hence decrease the frequency of anode rotation (f). Itis important to maintain high f to ensure proper

distribution and dissipation of heat from the focal track

to the body of the anode.

A Machlett DX 69B rotating anode X-ray tube insertwith a 4-in 400 000 heat units anode was used to test the

effect of external magnetic field on the induction motor.

The tube insert was removed from the housing to permit

visual contact with the anode, which allowed measure-

ment of rotational frequency using a calibrated Strobotac

Type 1531 strobe light (GenRad, Condord, MA). The high

voltage and filament cables were disconnected, and the

tube was mounted on a precision-machined aluminumstage with a rotation axis about the central point of the

induction motor soft iron ring in order to provide accurate

angular position control (see Fig. 9 inset). A transformer

circuit was used to drive the motor. Mimicking the usual

startup and run procedure, a 220 V root mean square

(VRMS) ac current was applied to the coils for 2 s to

initiate the rotation. The voltage was then switched to50 VRMS for steady-state operation.

The mounted tube was advanced axially into the fringe

field of a full body, unshielded 1.5T GE Signa CV/I MRI

system (S3 magnet), providing a relatively uniform field

of known direction throughout the insert volume,

depending on the distance of the insert from the magnet.

We assumed that B is parallel to the MRI central axis and

uniform throughout the induction motor volume andverified that the gradient in Bz throughout the motor

volume did not exceed 4% at any point and that Br always

remained below 4% of Bz along the center line of the

magnet. We evaluated the performance of the induction

motor for fields up to 0.24 T and angles between B and

the axis of rotation of the motor ranging from 0 to 90� inincrements of 15�.

The results are summarized in Fig. 9 and show that atransverse field (a field at 90� to the anode axis of rotation)has a significant effect on motor operation with rotation

speed falling below 3000 rpm for fields greater than 0.03 T.

Electron Optics in Fringe Field of Closed-Bore Magnet:When an X-ray tube is placed in a magnetic field of

arbitrary direction, the electron trajectories are the joint

result of the electric field (from the electrodes and thefocusing cup) and the magnetic field. One direct method to

study these trajectories is to solve the equations of motion

of the electrons. The problem may be simplified to a

classical one of an electron, with an effective initial velocity

v0, moving in a combination of electric and magnetic fields

E and B, both being uniform and static. The regime of

interest for this investigation was the presence of a smallBwith arbitrary direction relative to E. The motion of theelectrons may be viewed as a combination of 1) linear

uniformly accelerated motion in the direction of the

magnetic field with an initial velocity due to the electric

field component in that direction, 2) a drift at a constant

velocity, and 3) a rotation about B at a constant speed.

We verified our analytical solution using both the FE

model described above and for fields up to 150 Gauss (8 A)

using a Helmholtz coil pair (inner radius 21 cm, 380 turnsof magnet wire, dc power supply: 20 A maximum current,

50 V, PowerTen, Elgar Electronics Corp., San Diego, CA)

to generate a known field at the location of the focal spot of

our static anode X-ray tube. The resulting deflection has a

measured slope of �0.25 mm/A, which agrees well with

the perturbation theory calculation of 0.28 mm/A and can

be converted to deflection per tesla using the Helmholtz-

coil field-current coefficient of 1.870� 103 T/A to get aslope of 150 mm/T or�2 mm per 0.0150 T [25]. Using this

value, we calculated the expected deflection of the e-beam

in the rotating-anode X-ray tube due to a Br in the range of

0.015 to 0.1000 T. For the largest Br of 0.1022 T, the total

deflection could be as high as 15 mm, which could easily

drive the focal spot off the anode track in a rotating anode

tube with similar electron optic geometry.

Fig. 9. The effect of an externally applied magnetic field on the

induction motor of a rotating anode X-ray tube. The speed of rotation

was measured for a range of field magnitudes and orientations. The

measured data in the inset show the results for low magnetic fields at

large angles. The inset photograph shows the precision-machined

aluminum stand used tomaintain a known angle between the external

field and the axis of rotation of the motor.

Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 477

We have built a nonmagnetic rotating-anode X-ray tubeinsert with minimal magnetic components (Rytech Inc.,

Toronto, ON, Canada) including removal of the iron core

between the stator windings, and replacement of all Ni and

magnetic stainless steel with nonmagnetic stainless steel

(see Fig. 10). The main goal of this effort is to allow

verification of our FE model of e-beam deflection for a

geometry where the field at the focal spot is determined onlyby the external field and not disturbed by other magneticmaterial. Preliminary results are shown in Fig. 11, where

good agreement between the FE model and measured

deflection is seen for a small field of 195 G in a directionequivalent to Br. These preliminary investigations point to

the need for active deflection coils with feedback to control

the magnetic field at the location of the focal spot.

SummaryVClosed-Bore XMR System: These preliminary

results illustrate the challenges to the tube design yet to be

met before full integration between a closed-bore MR

system and an X-ray fluoroscopy system can be achieved.The two main challenges to be addressed are X-ray tube

motor operation and efficiency and focal spot deflection.

The first challenge will possibly require redesign of the X-ray

tube motor to provide alternative mechanical approaches to

maintain anode rotation at the desired speed. A recent

publication [26] outlines several approaches that could be

considered, including springs and piezoceramic motors. The

second challenge can likely be met using a combination ofmagnetic shielding and active deflection coils. We measured

the field inside and outside of a 3-mm-thick box of mu metal

and found that a field reduction of �0.05 T can be achieved

in fields up to �0.1 T; new formulations of such metals are

quoted as having screening factors of five to seven for static

fields, with a saturation field of 0.8 T (MagnoShield

COMPOUND panels, Aaronia AG, DE-54597 Euscheid,

Germany). The impact of motor options, shielding, andactive coil solutions on the field homogeneity of the magnet

must, of course, be considered, and some compromise in

X-ray tube operation may be necessary.

IV. CONCLUSIONS

The hybrid X-ray/MR system developed using the Signa SP

open bore magnet has proven the utility of closeintegration between the two systems, facilitating fast and

Fig. 10. (a) An X-ray tube built to our specifications (Rytech Inc.) with minimal magnetic material including (b) an iron-free stator

and (c) a nickel-free focusing cup. The tube is shown during heating and outgassing in (d).

Fig. 11. Our first experiment with the nonmagnetic X-ray tube to

determine deflection of the focal spot when placed in a magnetic field

that is oriented perpendicular to the anode-cathode direction or in

the Br direction (see Fig. 8). Our finite-element simulation (same scale

as experimental data) shows good agreement with the simulations,

with both indicating a deflection of approximately 5 mm.

Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

478 Proceedings of the IEEE | Vol. 96, No. 3, March 2008

repeated switching between the two imaging modalitiesduring an interventional procedure. We have shown that,

in principle, TFT-based flat-panel digital detectors are

immune to magnetic field effects and that static-anode

X-ray tubes operate well in magnetic fields if alignment

between the magnetic and electric field is maintained.

The homogeneity of the magnetic field can be maintained

in spite of the proximity of the X-ray detector to the MR

field of view, and RF noise and eddy currents do notsignificantly degrade MR image quality. We have also

carried out preliminary experiments to evaluate the

performance of a rotating-anode X-ray tube in the fringe

fields of a closed-bore 1.5 T magnet. Some modification

of X-ray tube design will likely be required for closest

integration of X-ray fluoroscopy with a closed-bore MR

system, although relatively simple solutions may be

possible. Our experience has shown that interventionalphysicians need, and are demanding, better image

guidance than is currently available. XMR systems are,

we believe, practical systems that can achieve the goal of

better guidance. While some technical challenges remain,

there are several possible approaches to overcoming these

challenges and making the dream of a practical, dual-

modality image guidance system a reality. h

Acknowledgment

The authors would like to thank S. Williams,

P. Neysmith, H. Easton, N. R. Bennett, and G. DeCrescenzo

for technical assistance. They also greatly appreciate the

technical and surgical assistance provided by C. Cooper and

G. Johnson during patient procedures.

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Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

Vol. 96, No. 3, March 2008 | Proceedings of the IEEE 479

ABOUT THE AUTHORS

Rebecca Fahrig, photograph and biography not available at the time of

publication.

Arundhuti Ganguly, photograph and biography not available at the time

of publication.

Prasheel Lillaney, photograph and biography not available at the time

of publication.

John Bracken, photograph and biography not available at the time of

publication.

John A. Rowlands, photograph and biography not available at the time

of publication.

Zhifei Wen, photograph and biography not available at the time of

publication.

Huanzhou Yu, photograph and biography not available at the time of

publication.

Viola Rieke, photograph and biography not available at the time of

publication.

Juan M. Santos, photograph and biography not available at the time of

publication.

Kim Butts Pauly, photograph and biography not available at the time of

publication.

Daniel Y. Sze, photograph and biography not available at the time of

publication.

Joan K. Frisoli, photograph and biography not available at the time of

publication.

Bruce L. Daniel, photograph and biography not available at the time of

publication.

Norbert J. Pelc, photograph and biography not available at the time of

publication.

Fahrig et al. : Hybrid X-Ray/MR System for Interventional Guidance

480 Proceedings of the IEEE | Vol. 96, No. 3, March 2008