Magnetic resonance imaging and spectroscopy of intraocular tumors

SURVEY OF OPHTHALMOLOGY

CURRENT

VOLUME 33 - NUMBER 6 * MAY-JUNE ism

RESEARCH

EDWARD COTLIER AND ROBERT WEINREB, EDITORS

Magnetic Resonance Imaging and Spectroscopy of Intraocular Tumors NANCY H. KOLODNY, PH.D.,‘-’ EVANGELOS S. GRAGOUDAS, M.D.,’ DONALD J. D’AMICO, M.D.,1 AND DANIEL M. ALBERT, M.D.1 ‘Department of Ophthalmology, Massachusetts Eye and Ear Infirmary and Harward Medical School, Boston, Massachusetts, and Department of Chemistry, Wellesley College, Wellesley, Massachusetts

Abstract. Proton magnetic resonance imaging (‘H MRI) has emerged as a clinically useful tool for the diagnosis of intraocular tumors. During the last four years ‘H MRI characteristics, including spin-lattice relaxation times (T,) and spin-spin relaxation times (Ts), have been established for several types of tumors. The introduction of surface coils to the imaging process has significantly improved the quality of intraocular MR images, leading some clinicians to suggest that ‘H MR images are preferable to CT scans. Another MRI technique, in which sodium-23 (“Na) is imaged rather than protons, is now under development as tool for intraocular diagnosis. The potential of 23Na MRI depends upon the high concentration and “visibility” of sodium in the vitreous body,

and upon the apparent differences in sodium behavior in normal cells vs. tumor cells. The metabolism of normal ocular tissues and intraocular tumors may be probed noninvasively with phosphorus-31 MR spectroscopy (“P MRS). Much progress has been made during the last few years in understanding the appearance of 31P MR spectra of many types of healthy and diseased cells and tissues. Clinical application of this technique to the diagnosis and monitoring of intraocular tumors following conservative treatment will be dependent upon the development of spectroscopy techPiques that collect information from the volume of interest (tumor) only. (Sure Ophthalmol 33:502-514, 1989)

Key words.

magnetic resonance imaging magnetic resonance spectroscopy nuclear magnetic resonance ocular tumors l

l

l

behavior of water molecules in an array of glass capillary tubes placed in a magnet equipped with a magnetic field gradient. 25 At the same time, NMR spectra of biological systems were first obtained.33 NMR equipment was developed to include tnagnets with interior space large enough to accommodate humans, gradients to allow spatial localization, and powerful computers to perform the calculations necessary to produce NMR images and high-resolution spectra. As NMR was adopted by the medical community, its name was changed to MR to eliminate the connotations of the term “nuclear” to the lay public. Today, hospitals throughout the world are acquiring MR systems. For lesions of the brain and spinal cord MR is now the diagnostic procedure of

The phenomenon of nuclear magnetic resonance (NMR) was discovered in 1945 by physicists interested in the behavior in magnetic fields of nuclei in solids and liquids. 3,40At first NMR was the province of physicists who developed the theoretical basis for understanding the interactions of nuclei with magnetic fields and with each other. NMR spectroscopy was soon adopted by chemists who found it a useful tool for the identification of molecules and unraveling of their structures. NMR spectrometers became commercially available during the 1950s and remain an essential tool in the chemistry laboratory. During the early 1970s two parallel developments led to a revolution in the use of NMR. In 1973 Lauturbur demonstrated that cross-sectional images could be produced by observing the NMR 502

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choice. MR images of the brain have been added to medical school curricula since they demonstrate its anatomy so well. Papers using MR imaging appear frequently and in large numbers in radiology journals. For the ophthalmologist, MR images and spectra are a relatively new tool. Proton MR imaging has been added to the techniques of computed tomography (CT) and ultrasonography as a method for diagnosing intraocular lesions. This article reviews the current use of MR for the study of intraocular tumors, and presents new MR techniques that may be useful in the future.

I. Physical Principles of Magnetic Resonance Magnetic resonance images and spectra are produced when the atomic nuclei present in a molecule, such as the protons in water, interact with radiofrequency energy in the presence of a magnetic field. The physical basis for this interaction is described in Section I. In Section II we describe the techniques used to produce images and spectra. These sections serve as background for, but are not required for an understanding of, the survey of the literature on applications of magnetic resonance to intraocular tumors presented in Section III. A. PROPERTIES OF NUCLEI Our ability to produce magnetic resonance images or spectra of biological tissues depends on the properties of their nuclei, as well as their concentration in the tissue of interest. These nuclear properties include spin, magnetogyric ratio, and natural abundance. As a result of their internal structures, nuclei possess a fundamental property known as spin, give the symbol I. Nuclear spin values are either integral, half-integral or zero. Only nuclei with values of I not equal to zero are “MR active.” Thus, since many biologically important nuclei, such as the common isotopes of oxygen and carbon, have spins with values of zero, they may not be used to produce MR images or spectra. Most nuclei, however, have nonzero spins. The nuclei of interest for MR studies of ocular tumors are protons (‘H, I = l/2), sodium-23 (23Na, I = 3/2) and phosphorus-31 (3’P, I = l/2). Each of these nuclei is 100% naturally abundant in biological tissue. All nuclei are positively charged. We picture the nucleus with I = l/2 as a “spinning sphere” of positive charge. A nucleus with I = 312 is called quadrupolar, and behaves as a “spinning football” of positive charge. The combination of non-zero spin and charge confers upon the nucleus a magnetic moment pN (Fig. 1). The magnitude of kN is proportional to the magnitude of the spin and to another

Fig. 1. Schematic representation of the behavior of a nucleus in a magnetic field. Due to its inherent spin, the positively charged nucleus possesses kN. In the presence of magnetic field the magnetic moment rotates about tary movement (known as precession) teristic frequency, called the Larmor

magnetic moment B, along the z-axis, the z-axis. This rooccurs at a characfrequency.

inherent property of the nucleus, its magnetogyric ratio, yN, according to the equation, pN = yNhI. Nuclei with larger magnetogyric ratios, such as protons, are more “MR sensitive.” B. INTERACTIONS OF NUCLEI WITH MAGNETIC AND RADIOFREQUENCY FIELDS 1. Excitation If a homogeneous external magnetic field is applied to a collection of nuclei, such as the protons present in the water in the human body, the nuclei interact with the magnetic field in a particular manner. In the classical model of magnetic resonance, the spinning nucleus acts like a gyroscope, precessing about the external magnetic field with a characteristic frequency, known as the Larmor, or resonance frequency, oN (Fig. 1). The value of the resonance frequency is proportional to the magnetogyric ratio of a given nucleus and to the strength of the applied magnetic field B,: wN = yNBo. Thus, the higher the magnetic field, the higher the resonance frequency. For example, in a clinical MR imager at 0.15 tesla (T) the proton resonance frequency is about 6 MHz, while at 1.5 T it is about 60 MHz. A higher resonance frequency produces MR images and spectra with better signal/noise ratios, and thus many contemporary clinical systems are being manufactured with 1.5 T magnetic fields. Unless an additional stimulus is applied, the nu-

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clei will precess at the same orientation with respect to the static magnetic field indefinitely. This state is called the equilibrium state. In order to produce MR images or spectra, a second magnetic field, B,, is applied. This field is not static, as is B,, but can be thought of as rotating with a particular frequency, in the radiofrequency range. If the frequency of B, equals the resonance frequency, the nuclei will be excited. M, the sum, per unit volume, of all the individual nuclear magnetic moments in an assembly of identical nuclei, will begin to rotate about the direction of B,. B, is applied perpendicular to B,, for a period of time (on the order of microseconds) such that the magnetic moment tilts either 90 or 180 degrees (“) from its equilibrium position. Thus we have what are referred to as 90” (r/2) pulses, and 180” (n) pulses common to most MR procedures. (See Fig. 2.) The resonance frequency of a nucleus also depends upon how it is bonded to other nulei in the molecule in which it is located. Both protons in any water molecule (H,O) have the same resonance frequency, but protons bonded to carbons in fat molecules have a different resonance frequency. Phosphorus-31 nuclei have a range of resonance frequencies, known as chemical shifts, that vary among the biologically important phosphates, making possible the use of 3’P MR spectroscopy for the identification and quantitation of these molecules in normal and diseased tissues. Since sodium-23 nuclei have a spin of 312 rather than l/2, their behavior differs from that of protons and phosphorus in that they may experience a range of resonance frequencies due to the motional behavior of their environment. 2. Relaxation

a. Tl, Spin-lattice Relaxation Time Once a collection of magnetic moments, indicated by M in Fig. 3, has been tilted away from its equilibrium position by the application of a radiofrequency pulse (B,), it will return to equilibrium, tilting back to its original position through processes called relaxation. As they relax, the nuclei emit signals at their resonance frequencies in the radiofrequency range. The time constant for the return to equilibrium is called T,, the spin-lattice relaxation time. “Spin lattice” refers to the mechanisms by which the relaxation occurs, and T, is characteristic of the environment of the excited nuclei. T, values also depend on the strength of the static magnetic field, B,, and thus caution must be exercised in assuming that T, values are constant from one MR system to another. The range of T, values that have been observed for aqueous protons in ocular tissues is between 0.3 and 2.5 s at 1.5 T.

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Fig. 2. The sum, per unit volume, ofall the individual nuclear magnetic moments in an assembly of identical nuclei is called M. M is tilted away from its equilibrium position along the x-axis by the application of a second magnetic field, B,, shown here along the x-axis. B, may be applied for a time (on the order of microseconds) long enough to tilt M 90” from its equilibrium position (a ?r/2 pulse), or for twice that length of time, so that M is tilted 180” (a ITpulse). The 7r/2 and P pulses are the most common pulses used in MR imaging and spectroscopy.

Although exact descriptions of the processes that cause particular T, values are not yet available, a number of insights have been gained into proton T, values of several ocular tissues. Phosphorus-31 nuclei, with I = l/2 as for protons, have similar relaxation mechanisms. T, values are much more difficult to measure for 31P nuclei, since these nuclei are much more dilute than aqueous protons, but they appear to fall in the same range as for protons. The situation is significantly different for 23Na, however. With a spin of 312, 23Na possess a quadrupole moment. Relaxation mechanisms for 23Na are different from those for spin-l/2 nuclei and 23Na T, values are 60 ms or less in biological tissues. This, and the correspondingly short T, values for 23Na, complicate the acquisition of 23Na MR images and spectra. However, as will be described below, techniques have been developed that overcome these complications and make 23Na a promising MR imaging nucleus, particularly for the eye.

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b. T2, Spin-spin Relaxation Time A second relaxation process occurs following the radiofrequency pulse, also shown in Fig. 3. Rather than simply “tilting” together back toward the static magnetic field, the magnetic moments begin to behave independently of one another, and become “dephased.” This process, characterized by a time T,, is called spin-spin relaxation. Spin-spin relaxation is due to the interactions between the excited nucleus and neighboring nuclei. Generally, in dilute solutions, the values of T, and T, for a particular spin are approximately equal. However, in biological tissues the values may diverge, with T, being shorter than T,. T, is also less strongly dependent on the strength of B, than is T,.

II. Techniques of Magnetic Resonance

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A. MR IMAGING 1. Equipment Required As shown in Fig. 4, a typical magnetic resonance imaging system consists of several components. The main magnet, large enough to enclose the patient, generates the static magnetic field, B,. A set of magnetic field gradient coils (and their associated power supplies) imparts spatial coding by generating AB. One or more probes (and their associated radiofrequency hardware) excite the nuclei by generating B,, and then detect the signals emitted by the nuclei as they relax. A computer controls the collection of data and their processing into images. For clinical studies, the magnet must be large enough to comfortably accommodate a patient. Most contemporary MR magnets, built of a shielded solenoid of superconducting wire cooled to liquid helium temperature (4K), surrounded by a liquid nitrogen bath at 77K, have a clear bore of 70-90 cm. Superconducting magnets, once charged, operate with no electrical losses. The field strengths of these magnets in clinical use range from 0.15T to 1.5 T. Some clinical MR systems use electromagnets, requiring a constant input of electrical power and cooling with running water. While electromagnets can operate at up to 1.5 T, most are designed to operate at lower field strengths. In electromagnets, as shown in Fig. 4, imaging is performed in a gap between the poles rather than within a bore as in the case of superconducting magnets. The magnet provides the primary magnetic field, B,, and thus determines the frequency at which the images are acquired. To be useful for MR imaging, the static magnetic field (B,) must be stable and highly homogeneous within a sufficiently large volume to contain the region of the body to be studied. Imaging, however, requires disruption of the homogeneous field in a

Fig. 3. M relaxes back to its equilibrium position along the x-axis via spin-lattice relaxation processes, characterized by a time constant T,. At the same time, the individual magnetic moments comprising M dephase in the xy plane via spin-spin relaxation processes, characterized by a time constant T,.

(pz:“,,I Fig. 4. A block diagram of the major components of magnetic resonance imaging and spectroscopy systems.

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known, software controllable and reproducible fashion via the magnetic field gradients (AB). The gradients cause the magnetic field to vary linearly in the x, y and z directions. The z gradient is shown schematically as AB, in Fig. 4. Because of this variation, superimposed on the primary magnetic field, each region of the body in this complex magnetic field is labeled with a unique set of frequencies that depend on its position in the magnetic field. Since frequency is directly related to position, information collected during the radiofrequency pulse stage can be converted by the computer into a recognizable image. In a magnetic resonance imaging procedure, nuclei must be excited in order to produce the signals that can be processed to yield spatial information and image intensities that reflect the nuclei’s inherent relaxation times, T, and Ts. Radiofrequency pulses are used to excite the nuclei within the selected region of the body. These pulses are delivered by a probe. Radiofrequency probes can be constructed in many ways, but for the purposes of this review, can be divided into two types: probes consisting of volume coils and their associated tuned circuits and probes consisting of surface coils and their associated tuned circuits (Fig. 5). Volume coils, often called head coils or body coils depending upon how much of the body pulses

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pulse has been terminated. They allow the creation of MR images of large regions of the body and are the most widely used type of probe. Surface coils, as their name implies, do not enclose, but rather rest upon the surface of the body. Depending upon the size and design of a surface coil, it will generally excite (or detect signals from) a region within the body of hemispherical shape to a depth equal to the radius of the coil. However, the excitation is not uniform, being strongest close to the coil, and weakening with distance. Surface coils are very useful for small, localized regions of the body, such as the eye and orbit. Often, a volume coil is used to excite the nuclei and a surface coil to detect the resulting signals. It is the computer and its associated pulse programmer that control that MR study. The computer runs software that directs the pulse programmer to turn the magnetic field gradients and radiofrequency pulses on and off at specified times. The computer also allows the operator to set the strengths of the gradients and radiofrequency pulses. Following the delivery of the radiofrequency pulses, the computer also receives the sig-

Fig. 5. Top: A volume coil, used for MR imaging or spectroscopy, uniformly excites all nuclei of a particular type enclosed within it. This coil is a “head coil”, suitable for SIP MR spectroscopy at 1.5T. Bottom: A surface coil, used for MR imaging or spectroscopy, excites all nuclei of a particular type in an adjacent hemispherical region.

nals from the probe (converted by an analog/digital converter), and performs the necessary calculations (Fourier transforms) to produce images. The computer is also connected to a display device that allows viewing of the images. Schematic representations of two typical pulse programs used in our studies of ocular tissues are shown in Fig. 6. Of particular note are the timing parameters known as TE and TR. As is true for virtually all MR imaging sequences, these sequences produce radiofrequency signals in the form of “spin echoes.” The time until the appearance of the echo is controlled by the pulse sequence, and is called TE. TE is shown explicitly in Fig. 6B. In Fig. 6A the echo occurs after the second radiofrequency pulse at a time equal to the spacing between it and the first radiofrequency pulse, with TE equalling the sum of these time intervals. In order for high quality images to be obtained, pulse sequences must be applied repeatedly and the collected data added

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Fig. 6. A. Pulse program used to produce a ‘H MR image on an imaging system constructed at the M.I.T. MRI Facility under a joint program between M.I.T. and IBM. The lines entitled “Readout”, “Phase Encode” and “Slice Select” display the timing of the signals transmitted to the power supplies for the x, y and z gradients. These signals determine the times at which the gradients are turned on and off. The times may be varied, as indicated by the parameters Dl-D7. The line entitled “RF” displays the timing and shape of the radiofrequency pulses transmitted to the sample. The first pulse is a “soft” ?r/2 pulse, and the second a “soft” rr pulse, producing a spin echo detected during the “sample” period. The lengths of the pulses are determined by the parameters P6 and P7. “Soft” rf pulses excite nuclei with a narrow range of resonance frequencies.

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MR images are two-dimensional representations of the intensities of the radiofrequency signals received by the probe from a region of the body. The intensities are displayed in an array of pixels (usually 256 x 256). The dimensions of the pixels vary, depending upon the imaging system and its software, but are generally in the square millimeter range. The parameters that determine the intensities of the signals from the pixels include nuclear (‘H or 23Na) density, T, and/or T,, and flow rate (intrinsic tissue characteristics) and slice thickness (computer controlled). To emphasize one or another of these parameters one chooses a particular imaging sequence. Throughout the MR literature images are known by the parameter they emphasize, for example, T, or T,.

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The ‘H MR image shown in Fig. 7, for example, is called a T,-weighted image. To produce this image, the pulse sequence shown in Fig. 6A was used with TE = 17 ms and TR = 3 17 ms. In such an image, protons with short T, values are bright (hyperintense), while those with long T, values are dim (hypointense). T,-weighted images generally result from pulse sequences with short TEs. In T,-weighted images, protons with long Tss are hyperintense, while those with short T,s are hypointense. These images are produced by sequences with longer TE and TR values. Often a series of images with different values of TE are obtained for T,-weighted studies. Factors that influence the lengths of T, and T, in intraocular tumors are discussed below. Some MR systems are equipped with software that allows the calculation of T, and T, values from selected regions of an image. Most MR images are used directly, however, as a visual tool. 23Na imaging req uires much shorter echo times (TEs) than ‘H MR imaging, since T, and T, for 23Na

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Fig. 7. A T,-weighted ‘H MR image of an enucleated human eye with uveal melanoma. The tumor, with superimposed hemorrhage, appears as a hyperintense region.*]

are less than 60 ms. Unless TE is short enough, all of the signals produced by the relaxing sodium nuclei will be gone before data collection occurs. The pulse sequence in Fig. 6B allows the collection of 8 echoes within 27 ms following the radiofrequency pulse. This technique is still in its infancy, and is not available on commercial systems at this time.

ET AL

Fig. 8. A 31P MR spectrum of perfused human uveal melanoma cells. Peaks are identified as: SP, sugar phosphates; PE, phosphorylethanolamine; PC, phosphorylcholine; Pi, inorganic phosphate; GPE, glycerol 3phosphorylethanolamine; GPC, glycerol J-phosphorylcholine; PCr, phosphocreatine; ATP (y, cx, and 8), adenosine 5’-triphosphate (y, (Y, and 8); NAD, nicotinamide adenine dinucleotide; UDPG, uridine diphosphoglucose. The spectrum was obtained on an 8.45 T magnet at the High Field NMR Resource, Francis Bitter National Magnet Laboratory, Massachusetts Institute of Technology, Cambridge, MA.

B. MR SPECTROSCOPY 1. Equipment Required MR spectroscopy requires the same equipment as MR imaging. However, whereas all imaging techniques depend upon magnetic field gradients, only “volume selective” spectroscopy techniques do. Spectroscopy requires higher magnetic fields than imaging, in order to provide sufficient sensitivity and resolution. Thus, only 1.5 T systems are very useful for spectroscopy. 2. Information Obtained In MR spectroscopy, the information obtained is generally in the form of a spectrum, as shown Fig. 8. The spectrum displays a series of peaks along the xaxis, according to their values of resonance frequency, or “chemical shift” with respect to an accepted reference compound. The y-axis displays the peaks’ intensities. The peaks are not narrow spikes; their shapes and widths provide additional information about the physical and chemical environment of the nuclei producing the peaks. Wellresolved spectra are those in which closely spaced peaks are sufficiently separated to be assigned chemical shift values. Under certain circumstances, in which sufficient time is allowed to elapse between data accumulations, the intensity of a peak is a direct reflection of the concentration of the nucleus producing the peak. Since most spectra are produced with less time between data accumulations,

peak intensities must be corrected using the T, values of the nuclei in order to obtain reliable concentration information.

III. Applications of MR to Intraocular Tumors A. PROTON MR IMAGING MRI is an attractive technique for the physician because of its excellent tissue specificity, its threedimensional capability and its lack of ionizing radiation. Its major disadvantages are the high cost of the necessary equipment, and the contraindications of the technique for patients who are claustrophobic or who have particular types of metallic prostheses or pacemakers. Typical MR imaging sessions last about one hour, with each image requiring less than five minutes of data acquisition. During the last four years proton MRI has been developed for the diagnosis of intraocular lesions. Some of the work presented in. 1984 and 1985 has been reviewed by Sobel et al. 48 The studies that have appeared in the literature since then have demonstrated improvements in sensitivity and resolution with the use of surface, rather than volume coils. In addition, they have established an acceptable set of parameters for the appearance in MR images of intraocular tumors such as uveal melanomas. The studies may be divided into those that have focused on enucleated eyes or eyes in vim, and those that

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have used surface or volume coils to produce the proton MR images. The earliest study of proton MRI of enucleated eyes is that of Gonzales et al.14 Using an 8 cm bore 1.4 T magnet and a volume coil, the feasibility of this technique for ocular imaging was established, based on the contrast obtainable between the bovine lens, vitreous body, and simulated vitreous hemorrhage and liquefaction. Enucleated human eyes bearing intraocular tumors, and, often, accompanying retinal detachments, have since been examined at 1.4 T”s51 or 1.5 T.” The images obtained in these studies have produced several interesting observations. Gomori et all5 demonstrated that the values of T, and T, of choroidal melanomas are related to the melanin content of the tumors. This important observation was explained by the nature of melanin as a free-radical containing species. Since the melanin free radicals should be accessible to the water protons visualized in proton MR images, and since the free radicals produce paramagnetic proton relaxation enhancement, both T, and T, will be shorter in the presence of melanin than in its absence. Thus, tumors with high melanin content demonstrated T, and T, values 10% and 20%, respectively, of those of amelanotic tumors. On T,weighted images, therefore, the pigmented tumors were markedly bright (hyperintense) with respect to the vitreous body and the optic nerve. On Trweighted images the pigmented tumors were hypointense. Amelantic tumors, on the other hand, were hyperintense with respect to the vitreous but isointense with respect to the optic nerve on T,weighted images. On T,-weighted images amelanotic tumors were hypointense with respect to the vitreous but isointense with respect to the optic nerve. Another important result of this study was the hyperintense appearance in T,-weighted proton MR images of subretinal effusions. Similar intensity results were obtained by terPenning et a15’ and by Kolodny et al.*’ We found that degree of pigmentation of choroidal melanomas was correlated with the brightness of the tumors in T,-weighted proton MR images. In addition, the brightness of subretinal efhrsions appeared to depend upon their age; chronic subretinal hemorrhage was hyperintense with respect to the vitreous body and a mildly pigmented tumor. On the other hand, acute hemorrhage was hypointense with respect to all of these tissues. This change in the appearance of hemorrhages was explained as resulting from the conversion of hemoglobin to methemoglobin, the latter being a paramagnetic species that, like melanin, causes paramagnetic proton relaxation enhancement. An additional important result reported in this paper was the production of

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high quality proton MR images of eyes with tantalum rings sutured to the sclera. These rings are used for localization of choroidal melanomas before proton beam irradiation. The safety and effectiveness of imaging such eyes has been questioned, and we have found that as well as being an effective technique, there are no dangers connected with MR imaging in the presence of tantalum rings.46 In viva proton MR imaging of intraocular tumors has evolved during the last four years into a technique that, some feel, rivals or surpasses CT. The advances in MRI that account for this assessment include higher sensitivity, resulting in shorter imaging times and thinner image slices, and better in-plane resolution, resulting in clearer images and thus smaller discernable lesions. Increases in sensitivity of MR imaging are due to the more widespread use of higher field strength magnets and to the use of surface, rather than volume coils, for signal reception. Higher magnetic held strengths produce higher signal/noise ratios, despite the concomitant increase in T, that can lower the signal/ noise unless longer TR’s are used. Sassani and Osbakken41 presented proton MR images obtained at 0.15 T using a volume radiofrequency coil. In these cross-sectional images of the heads of normal volunteers and patients with treated and untreated choroidal and ciliary body melanoma, “reasonable anatomic delineation” of the eye (vitreous and lens) and orbital tissues was obtained. They noted that orbital fat was the tissue with greatest intensity in all T,-weighted images. The treated melanoma was not visible, while the untreated ciliary body melanoma was hyperintense with respect to the brain and hypointense with respect to orbital fat. Slice thicknesses were estimated to be 2.6 mm in their best images, obtained using a saturation-recovery pulse sequence; no value was given for inplane resolution. The size of the detected melanoma was not given. In another study at 0.15 T published two years later, Worthington et al54 used both a volume coil (slice thickness 10 mm) and a surface coil (slice thickness 5 mm, in-plane resolution 0.7 mm). The superior resolution of the images obtained with the surface coil was obvious. Five cases of intraocular melanoma were studied, and it was stated that T, and T, values were short. Sobel et al were able to detect ocular lesions 5 mm and larger, with slice thicknesses of 7 mm.*’ They used a 0.5 T magnet and a volume coil, and obtained both T,- and T,-weighted images. Six choroida1 melanomas were observed most effectively with a pulse sequence in which TE = 28 ms and TR = 500 ms. The tumors were hyperintense with respect to the vitreous. In T,-weighted images the tumors were hypointense. Values were calculated for T,

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(534625 ms) and T, (54-67 ms), all shorter than that of the vitreous, as expected. A single retinoblastoma was also imaged, and behaved similarly to the melanomas. However, the small size of the retinoblastoma prevented calculation of T, and T,. It was pointed out that calcifications in this lesion observed by CT are not visible with MR. Again, inplane resolution was not stated, but lesions greater than 5 mm in size were accurately detected, while smaller lesions were not. It was not clear whether this was due to resolution or sensitivity. Also at 0.5 T, Haik et al” produced MR images with a volume coil whose clarity they assessed as approaching that of CT. This study of choroidal tumors confirmed the results that T, values for melanomas are shorter than those of other tumors, which, in turn, are longer than those of the corresponding normal tissues. Of seven melanomas studied, five were clearly demonstrated. A tumor of height less than 1.5 mm was not detectable. They also showed that the presence of a 15 mm round cobalt-60 implant obliterated all ocular and orbital detail. Thus, the feasibility of using MR imaging on patients with cobalt implants was denied. T, and T, values were found to behave as expected, using the white matter of the brain as the standard. In contrast, intraocular metastatic carcinoma displayed varying T, and T, values, which the authors speculated might be due to amount of associated hemorrhage or exudate. In a similar study by Edwards et al” at 0.6 T using a volume coil, slice thicknesses were 8 mm and pixel size (or in-plane resolution) 0.9 x 1.8 mm. Except for calcification, MR was able to detect all lesions detectable by parallel CT studies. MR’s advantages were stated by Edwards et al to include the ability to perform multiplanar techniques without additional positioning of the patient. As in all other studies, melanomas were found to have short T, values, but these authors found long T, values. Since this is different from the results of all other studies reported to date, it is possible that the images were displaying hemorrhage rather than, or in addition to, tumor. In fact, Edwards et al speculate that the characteristics of the MR image of the tumors was due to intratumor hemorrhage. At 1.5 T, Mafee et al were able to produce MR images with slice thicknesses of 3-5 mm, using a volume coil for most cases, and a 3-inch surface coil for one case.z8,2gThey pointed out the better spatial resolution and greatly improved image quality of the latter image. They were able to demonstrate all intraocular lesions greater than 3 mm in height with the volume coil; with the surface coil they could demonstrate an elongated discoid melanoma less than 3 mm in height. Once again, in T,-weighted images uveal melanomas were unique among

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malignant tumors in that they were hyperintense, while in T,-weighted images they were hypointense with respect to the vitreous. They were thus able to differentiate uveal melanomas from associated subretinal effusion (short T, and long T._,), choroidal hemangiomas (long T, and long T2) and choroidal metastases (short T, and long T2). Mafee et al” also noted that calcification was much more difficult to recognize in MR images of retinoblastomas than in CT scans, but found that dense calcification was seen as an area of hypointensity in both T,- and T,weighted images obtained at 1.5 T. Several papers have appeared that focus on the advantages of using surface coils for obtaining MR images of the eye and orbit. These advantages hold at 0.5 T, 0.6 T and 1.5 T, the field strengths at which these studies were done. Schenck et a142 demonstrated that whereas they were able to obtain images using volume coils that had pixel sizes of about 1 x 1 mm and slice thicknesses of 4 mm, with surface coils of lo-14 cm used as receivers, pixel sizes could be reduced to 0.5 x 0.5 mm and slice thicknesses to 3-5 mm. These reductions were obtained without increasing scan time. Bilaniuk et al’ then used these surface coils to obtain MR images of ocular and orbital lesions. The smallest lesion they detected was 3.9 mm high, with a 9 mm base. They observed the usual relative T, and T, values that are characteristic of melanomas. De Keizer et al” also used a surface coil, placed like a diving mask over the face of the patient, to study eleven cases at 0.5 T. They obtained MR images with 5 mm slice thicknesses, and the same relative intensities based on T, and T, values as above. A tumor of height less than 2 mm (as detected by ultrasonography) was not observed by MR, but all tumors in this study greater than 3 mm were observed. Similarly, Sullivan and Harms50 used a 13 cm surface coil centered directly over the orbit of interest to produce proton MR images at 0.6 T. They used the vitreous as the internal standard for signal intensity. Patients were requested to keep their eyes closed during imaging, while in all other studies patients were requested to keep their eyes open and fixate on a particular point. Relative intensities suggested that values of T, and T, in three melanomas were short with respect to the vitreous body. Areas of calcification apparent in CT scans of two retinoblastomas were not identified in MR images. Otherwise, MR was found to be superior to CT in the detection of orbital neoplasms. A 7.6 cm diameter surface coil was used by Chambers et al6 to produce images at 1.5 T with 3 mm slice thicknesses. Twenty-four of thirty-one uveal lesions were well-visualized using T,-weighted images. These lesions were all greater than 3 mm in elevation. Lesion intensi-

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ties were consistent with the literature, hyperintense in T,-weighted images and hypointense in T2weighted images. Using surface coils placed directly over the eye of interest, high quality proton MR images may be obtained in clinically acceptable times. The relaxation behavior of uveal melanomas appears to be unique within the eye, with both T, and T2 appearing to have short values. Actual measurements ofT, and T, do not appear to be necessary to differentiate melanomas from other intraocular lesions. Rather, observation of a lesion that is hyperintense on a T,-weighted image and hypointense on a T,weighted image strongly suggests that the lesion is a pigmented melanoma. Amelanotic melanomas may be more difficult to diagnose using proton MRI, since the short T, and T2 values are probably due to the presence of melanin.

Fig. 9. A 23Na MR image of a bovine eye with induced hemorrhage. The image was obtained on a 1.9T magnet at the Brigham and Women’s Hospital. The data from eight echoes acquired in 25.9 milliseconds using the pulse sequence shown in Figure 6B were used to produce the image.

B. SODIUM MR IMAGING

Sodium MRI is a sensitive indicator of tissue damage and necrosis. Its sensitivity is due in large part to the considerable variations in both the absolute intracellular sodium concentration and the ten-fold difference between intracellular and extracellular regions. Furthermore, the z9Na nucleus has a quadrupole moment and this renders it extremely sensitive to its local environment. This sensitivity manifests itself in changes in T2 and in the “visibility factor,” or percent of the theoretically possible intensity which is actually observed. Variations in intensity in 23Na MR images are thus due to a combination of concentration, T, and “visibility.” The primary factors that have limited the use of 23Na MRI as a clinical tool are its 1OOO-fold lower concentration than protons and its very short T2, necessitating special pulse sequences for data acquisition. Nevertheless, several studies applying 23Na MR to ocular imaging have appeared. A 23Na MRI study of bovine eyes demonstrated that at 2.7 T the anterior aqueous and posterior vitreous may be clearly distinguished from the lens, sclera, cornea, zonular fibers and ciliary body.13 The pulse sequence used to produce these images included an echo time of 30 ms, much too long to capture sufficient signal from nuclei with expected T, values of 26.5 and 2.9 ms, as in the lens, whereas T2 of vitreous z3Na has been found to be about 60 ms.3g To visualize the lens, therefore, a “limited angular view planar-integral projection reconstruction” pulse sequence with TE = 1 ms was used. A solenoidal volume coil that enclosed the eye was used for all image acquisition. With TE = 12 ms, we have been able to produce 23Na MR images at 1.5 T of enucleated human eyes bearing uveal melanomas.20~z’ These images, ob-

tained

with a surface

echo sequence

coil and a gradient

recalled

in 10 to 20 minutes of imaging time,

demonstrate hyperintensity for vitreous and aqueous as well as varying intensity for subretinal hemorrhage and melanoma. Work currently underway in our laboratory has yielded the pulse sequence shown in Fig. 6B. 23 With this pulse sequence we have produced three-dimensional data sets that can be processed to produce “‘Na MR images of 16 or 32 slices in each dimension of bovine eyes with induced hemorrhage and liquified vitreous. As shown in Fig. 9, these images display excellent contrast and resolution. Furthermore, the data may be processed to yield localized values of T,. 29Na MR spectra of sequential regions of the eye have also been obtained using a technique called rotating frame zeugmatography,g but the length of time required for the technique and the uncertain shape of the region from which data are acquired make this technique less useful than 23Na MRI. The concentration of *‘Na in cellular tissue is l/l 0 that in the vitreous. Thus, as we have found, intraocular tumors are expected to be extremely hypointense with respect to the vitreous in z3Na MR images. However, Turski et a15* have shown a marked elevation of sodium signal related to canine cerebral gliosarcoma tissue. They suggested that there may be an increase in intracellular sodium in these cells as they are undergoing rapid proliferation. Human brainstem tumors38 and astrocytomas (grade 34) also exhibited increased sodium signal compared with that of adjacent brain tissue, while tumors treated with radiation therapy demonstrated a marked reduction in sodium signal. Thus, in-

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Fig. 10. A. Pulse program (FDRESS) used to produce spectra from a selected disk-shaped region. The region is selected based upon its distance from the surface coil. The gradient and RF lines in the program are similar to those described in Figure 6.

KOLODNY ET AL

1989 P6

P6

I

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Fig. 10. B. Pulse program (ISIS) used to produce spectra from a selected cubic region. The magnetic field gradients in

Gradients

the x, y and z directions are turned on and off sequentially. Three “soft” a radiofrequency pulses are applied, followed by a “hard” 7~12pulse.

‘

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formation about the proliferative state of intraocular tumors may, in the future, be obtainable from 23Na MR images. MR SPECTROSCOPY

The clinical potential of3’P MR spectroscopy presents an exciting but difficult challenge to the investigator. 31P MR spectra contain information about the identity and amounts of phosphorus-containing metabolites present in the tissue studied. The challenges of the technique lie in 1) understanding the role of these metabolites, and, in particular in the case of ocular tumors, how they are related to the state of the tumors pre- and post-therapy, and 2) producing spectra that include information from the tumor only, i.e., producing “localized” spectra that exclude signals from surrounding tissue. During the past four years several reviews have appeared on the use of 31P MR spectroscopy in understanding phosphorus metabolites’ roles in ocular tissues and in many types of tumor.‘2*‘6~4gThese reviews concentrate on correlating each of the peaks in 3’P MR spectra with a single metabolite. Extensive use of perchloric acid extracts of freezeclamped tissues enabled such correlations to be made. Typical data for the isolated porcine and human crystalline lens and cornea appear in Greiner

, I I

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et al.r6 Isolated intact bovine retinas,’ cultured human retinal pigmented epithelial cells,30 and intact bovine eyes 43 have also been studied using “P MR spectroscopy. Although many ocular tissues produce 3rP MR spectra since they contain metabolically active cells (the vitreous being an important exception), most of them are not present in sufficient amounts to be detectable in clinically acceptable times. In the normal intact eye, the lens produces the strongest “P MR signal.43 In the melanoma-bearing eye, tumors of elevations greater than 3 mm also produce strong “P MR signals. We have determined 3’P MR spectroscopic characteristics of human uveal melanoma by studying cultured cell pellets,rg perfused cultural cells,‘* tumors in athymic nude mice’ and enucleated tumor-bearing eyes. 22 Fig. 8 shows the 31P MR spectrum typical of all of these systems, demonstrating that the phosphorus metabolite levels in choroidal melanoma differ from those in normal human lens in the following ways. All of the melanoma tumors and cell lines from which spectra were obtained exhibited elevated levels of the membrane phospholipid metabolite monoester phosphorylcholine (PC) and phosphodiesters glycerol S-phosphorylethanolamine (GPE) and glycerol S-phosphorylcholine (GPC). Most also had elevated levels of

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phosphorylethanolamine (PE). There was, however, significant variation in the relative levels of these four compounds. We are currently investigating the biochemical causes of the variations of the levels of phosphomonoand -diesters in perfused cultured human ocular melanoma cells.32 In similar studies of retinoblastoma cells (Y-79), Miceli et al found correlations between the nutritional status of the cells and the levels of these compounds.g’ Since the only ocular structure demonstrating elevated levels of these peaks is the tumor, the peaks serve an important purpose for the possible future clinical application of 3’P MR spectroscopy to the diagnosis of ocular melanomas. Other 31P MR peaks, common to all metabolically active tissue, should allow follow-up after conservative treatment of ocular melanomas. These peaks, caused by the “high energy” compounds adenosine 5’-triphosphate (ATP) and phosphocreatine (PCr), have been shown to reflect the overall metabolic status of many types of human tumor cells as they grow, s6 recover from hypoxia53 or respond to therIn general, as the cells’ metabolic status apy* 26,34,35,45 is compromised, the levels of the peaks due to these high energy compounds decrease. Model studies using Greene melanoma implanted in rabbit eyes have been used to determine the capability of3’P MR spectroscopy for the assessment of tumor viability pre- and post-treatment.8*is*24 These studies showed significant decreases in the high energy compounds, but highlighted the most important remaining challenge to be met if 31P MR spectroscopy is to be clinically useful. The challenge of obtaining spectra from localized regions of the eye, so that tissues other than the tumor would not contribute “P MR signals, was met by Kindy-Degnan et al in their model study by applying a 3’P MR surface coil to the sclera adjacent to the tumor.18 This produced useful results, but it required surgery and therefore obviated one of MR’s most attractive features, its noninvasiveness. Another model study used an external surface coil, but spectra obtained included peaks probably due to muscle.24 A similar problem was encountered in the study of human ocular melanoma implanted in athymic nude mice.’ Several techniques have become available that allow the acquisition of 3’P MR spectra from a localized region using either a surface or volume coil external to the subject. Bottomley4 introduced the depth-resolved surface coil spectroscopy (DRESS) technique for the noninvasive study of high-energy phosphate metabolism in human heart, and has used it to observe myocardial infarction.5 Our version of the pulse sequence for this technique is shown in Fig. 10A. DRESS uses the inhomogeneous radiofrequency field of the surface coil combined

with a single magnetic field gradient to produce spectra from disk-shaped regions at varying distances from the coil. Thus, signals from a tumor may be obtained while those from the lens are excluded. However, depending upon the location of the tumor, unwanted signals from muscle or orbital tissue could also be obtained. Combining DRESS with specially designed Faraday shields may help eliminate such unwanted signals and work is in progress in our laboratory in this area. Another technique for localized MR spectroscopy, image-selected in viva spectroscopy (ISIS), was developed by Ordidge et a1.37 Beginning with a proton MR image to localize the region of interest, the technique then obtains an MR spectrum from that region using three magnetic field gradients and an appropriate combination of radiofrequency pulses, as shown in Fig. 10B. ISIS has been implemented at i .5T with a “P MR volume coil inserted into a ‘H volume coil, and has been used to evaluate response of brain tumors to treatment.44 The smallest volume of interest examined in this study was 40 cm3; useful data were acquired from this region in 30 minutes. The “P MR spectra showed the expected changes in high energy phosphates following pharmacotherapy and radiation therapy. Thus, although this method has promise, it has not been used for regions small enough to be useful for evaluating ocular tumors. Ways to increase the sensitivity of ISIS will have to be found, or new techniques developed, before volume-selectivity spectroscopy can be applied to ophthalmology. Continuing advances in the field of localized MR ,spectroscopy make this seem possible in the next few years.

References 1.

Albert DM, Svitra PP, Kolodny NH et al: 3’P MR Spectroscopy of human

ocular tumors in athymic mice. Invest C@1986 Bilaniuk LT, Schenck JF, Zimmerman RA et al: Ocular and orbital lesions: Surface coil MR imaging. Radiology 156: 669-674, 1985 Bloch F, Hansen WW, Packard M: Nuclear induction. Phy Rev 69~127, 1946 Bottomley PA: Noninvasive study of high-energy phosphate metabolism in human heart by depth-resolved “P NMR spectroscopy. Science 229.769-772, 1985 Bottomley PA, Herfkens RJ, Smith LS, Bashore TM: Altered phosphate metabolism in myocardial infarction: P-31 MR spectroscopy. Radiology I65:703-707, 1987 Chambers RB, Davidorf FH, McAdoo JF, Chakeres DW: Magnetic resonance imaging of uveal melanomas. Arch O,bhthalmol10_5:917-921, 1987 Chapman D, Kemp CM, Morris PC, Pons M: Studies of cellular metabolism in isolated intact bovine retinas by 3’P nmr. FEBS Letters I43:293-295, 1982 D’Amico DJ, Kolodny NH, Gragoudas ES et al: “P nuclear magnetic resonance spectra of ocular melanoma in rabbits: Metabolic correlates of proton-beam induced tumor regression. Invest O$thalmol Vis Sci 26 (suppl):209, 1985 D’Amico DJ, Kolodny NH, Talley AR et al: Sodium-23 MR techniques for study of diabetic retinopathy. InvestOphthalmol Vis Sci 27 (supp1):145, 1986

thulmnl Vis Sci 27(suppl):86,

2.

3. 4.

5.

6.

514

Surv Ophthalmol 33 (6) May-June 1989

10. de Keizer RJW, Vielvoye GJ, de Wolff-Rouendaal D: Nuclear magnetic resonance imaging of intraocular tumors. Am] Ophthulmol 102:4388-4441, 1986 11. Edwards JH, Hyman RA, Vacirca SJ et al: 0.6T magnetic resonance imaging of the orbit. AJNR 6~253-258, 1985 12. Evanochko WT, Ng TC, Glickson JD: Application of in viva NMR spectroscopy to cancer. Magn Reson Med 1:508-534, 1984 13. Garner WH, Hilal SK, Lee S-W et al: Sodium-23 magnetic resonance imaging of the eye and lens. Proc Nat1 Acad Sci USA 83:1901-1905, 1986 14. Gonzales RG, Cheng H-M, Barnett P et al: Nuclear magnetic resonance imaging of the vitreous body. Science 2231399 400, 1984 15. Gomori JM, Grossman RI, Shields JA et al: Choroidal melanomas: Correlation of NMR spectroscopy and MR imaging. Radiology 158: 443-445, 1986 16. Greiner JV, Kopp SJ, Glonek T: Phosphorus nuclear magnetic resonance and ocular metabolism. Sun, Ophthulmol 30:189-202, 1985 17. Haik BG, Saint Louis L, Smith ME et al: Magnetic resonance imaging in choroidal tumors. Ann Ophthulmol 19:218-238, 1987 18. Haik BG, Saint Louis L, Smith ME et al: Magnetic resonance imaging in choroidal tumors. Ann OphU&zol 19:218-238, 1987 18. Kindy-Degnan N, MUJJay T, Char DH et al: Assessment of intraocular tumor viability: A pre- and post-treatment comparison of fine needle biopsy and 31 P magnetic resonance spectroscopy. Invest Ophthalmol Vis Sci 29 (supFl):210, 1988 19. Kolodny NH, Albert DM, Epstein J et al: Characterization of human uveal melanoma cells by Phosphorus-31 nuclear magnetic resonance spectroscopy. Am J Ophthalmol 100: 38-44, 1985 20. Kolodny NH, Gragoudas ES, D’Amico DJ et al: Multinuclear MR imaging and spectroscopy of human ocular tumors. Invest Ophthulmol Vis Sci 28 (suppl): 116, 1987 21. Kolodny NH, Gragoudas ES, D’Amico DJ et al: Proton and sodium-23 magnetic resonance-imaging of human ocular tissues: A model study. Arch Ophthalmnl 105:1532-1536, 1987 22. Kolodny NH, Gragoudas ES, D’Amico DJ et al: Preliminary results on phosphorus-31 NMR evaluation of human uveal melanoma in enucleated eyes. Ophthalmology 95:66&673, 1988 23. Kolodny NH, Kohler SJ, D’Amico DA et al: High-resolution sodium-23 MRI of intravitreal pathology. Invest Ophthalmol Vis Sci 29 (suppl):338, 1988 24. Kolodny NH, Svitra PP, Gragoudas E et al: In vivo “P MR spectroscopy of rabbit choroidal melanoma. Invest Ophthalmol Vis Sci 27 (suf$d):259, 1986 25. Lauturbur PC: Image formation by induced local interactions: Examples employing nuclear magnetic resonance. Nature 242: 190, 1973 26. Marincek B, Friedburg H: Calculated Tl and T2 MR images and s’P MR spectra in rabbit thigh V2 carcinoma to monitor the effect of steroid administration. Znvest Radio1 20:701707, 1985 27. Mafee MF, Goldberg MF, Greenwald MJ et al: Retinoblastom and simulatine lesions: Role of CT and MR imaging. ” I Radio1 Clin North AL 25:667- 682, 1987 28. Mafee MF, Peyman GA, Grisolano JE et al: Malignant uveal melanoma and simulating lesions: MR imaging and evaluation. Radiology 160: 773- 780, 1986 29. Mafee MF, Peyman GA, Peace JH et al: Magnetic resonance imaging in the evaluation and differentiation of uveal melanoma. Ophthalmology 94:341-348, I987 30. Miceli MV, Kan L-S, Newsome DA: Phosphorus-31 NMR spectroscopy of cultured human retinal pigmented epithelial cells. Invest Ophthulmol Vis Sci 28;70-75, 1987 3 1. Miceli MV, Kan L-S, Newsome DA: Phosphorus-3 1 NMR of cultured retinoblastoma cells: Correlations with metabolic indices. Invest Ophthalmol V~.TSci 28 (supPl):25, 1987 32. Minichiello M, Albert DM, Kolodny NH et al: Phosphorus31 MRS monitoring ofhuman ocular melanoma metabolism

KOLODNY ET AL

33. 34.

35.

36.

37.

38.

39.

40.

41.

42.

43.

44.

45.

46.

47. 48.

49.

50. 51.

52.

53.

54.

employing a perfusion culture system. Invest Ophthulmol Vis Sci 28 (su#d):25 (1987) Moon RB, Richards HJ: Determination of intracellular pH by “P magnetic resonance. J Biol Chem 248:7276-7277,1973 Naruse S, Hirakawa K, Horikawa Yet al: Measurements of in vivo 3’P nuclear magnetic resonance spectra in neuroectodermal tumors for the evaluation of the effects of chemotherapy. Cancer Research 45:2429-2443, 1985 Naruse S, Horikawa Y, Tanaka C et al: Evaluation of the effects of photoradiation therapy on brain tumors with in vivo “P MR spectroscopy. Radiology 160:827-830, 1986 Okunieff P, Koutcher JA, Gerweck L et al: Tumor size dependent changes in a murine Iibrosarcoma: use ofinvivo s’P NMR for non-invasive evaluation of tumor metabolic status. Int J Radiation Oncology Biol Phys 12:793-799, 1986 Ordidge RJ, Connelly A, Lohman JAB: Image-selected in vivo spectroscopy (ISIS). A new technique for spatially selective NMR spectroscopy. J Magn Reson 66:283-294, 1986 Perman WH, Turski PA, Houston LW et al: Methodology of in vivo human sodium MRI imaging at 1.5T. Radiology 160:811-820, 1986 Pettegrew JW, Glonek T, Minshew NJ et al: Sodium-23 NMRofintact bovine lens and vitreous humor.] Map Reson 63:439-444, 1985 Purcell EM, Torrey HC, Pound RV: Resonance absorption by nuclear magnetic moments in a solid. Phys Res 69:37-38, 1946 Sassani JW, Osbakken MD: Anatomic features of the eye disclosed with nuclear magnetic resonance imaging. Arch Ophthalmol 102:541-546, 1984 Schenck JF, Hart HR Jr, Foster TH et al: Improved MR imaging of the orbit at 1.5 T with surface coils. AJNR 6:193-196, 1985 Schleich T, Matson GB, Willis JAet al: Surface coil phosphorus-31 nuclear magnetic resonance studies of the intact eye. Exfi Eye Res 40.343-55, 1985 Segebarth CM, Baleriaux DF, Arnold DL et al: MR Imageguided P-31 MR spectroscopy in the evaluation of brain tumor treatment. Radiology 165:215-219, 1987 Semmler W, Gademann G, Bachert-Baumann Pet al: Monitoring human tumor response to therapy by means of P-3 1 MR spectroscopy. Radiology 166:533-539, 1988 Smith EV, Kolodny NH, Gragoudas ES, et al: Safety and effectiveness of magnetic resonance imaging of choroidal melanoma patients with episcleral tantalum rings after proton beam irradiation. Am J Ophthalmol 105:695-696, 1988 Sobel DF, Kelly W, Kjos BO et al: MR imaging oforbital and ocular disease. AJNR 6:259-264, 1985 Sobel DF, Moseley IF, Brand-Zawadzki M: Magnetic resonance imaging (MRI) of the eye and orbit, in Gonzales C, Becker M (Eds): Diagnostic Imaging in Ophthalmology. Springer, New York, 1986, pp 99-113 Sostman JD, Armitage IM, Fisher JJ: NMR in cancer. I. High resolution spectroscopy of tumors. Magn Reson Imaging 2:265-278, 1984 Sullivan JA, Harms SE: Surface-coil MR imaging of orbital neoplasms. AJNR 7:29-34, 1986 terpenning BJ, Cheng H-M, Barnett P et al: MR imaging of enucleated human eyes at 1.4 Tesla. J Comput Assist Tomagr 10:551-559, 1986 Turski PA, Houston LW, Perman WH et al: Experimental and human brain neoplasms: Detection with in vivo sodium MR imaging. Radiology 163:245-249, 1987 West-Jordan JA, Smith A, Myint S et al: 3 1 P NMR studies on recovery from hypoxia of human tumor cells. Map Reson Med 5:182-185, 1987 Worthihgton BS, Wright JE, Curati WL, et al: The role of magnetic resonance imaging techniques in the evaluation of orbital and ocular disease. Clin Radio1 37:219-226, 1986

Reprint requests should be addressed to Evangelos S. Gragoudas, M.D., Retina Service, Massachusetts Eye and Ear Infirmary, 243 Charles Street, Boston, MA 02114.