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Breast Imaging |
1 From the Departments of Radiology (K.H., O.P., D.N.S., P.E.H., N.J.M., J.K., F.A.J.) and Surgery (S.S.), Brigham and Women’s Hospital, Harvard Medical School, 75 Francis St, Boston, MA 02115; and the Department of Radiology, Beth Israel Deaconess Medical Center, Boston, Mass (J.B.). Received April 4, 2000; revision requested May 25; revision received June 23; accepted August 1. Supported in part by NCI program grant 67165, research grant CA 46627, contract 282-97-0080 from the U.S. PHS Office on Women’s Health, and a grant from GE Medical Systems. P.E.H. supported by German Research Council (DFG) grant HU 798/1-1. J.K. supported by the Austrian Science Foundation and the Research and Education Fund of the European Congress of Radiology. Address correspondence to K.H. (e-mail: kullervo@bwh.harvard.edu).
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ABSTRACT |
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 TOP ABSTRACT INTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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MATERIALS AND METHODS: Eleven fibroadenomas in nine patients under local anesthesia were treated with MR imaging-guided FUS. Based on a T2-weighted definition of target volumes, sequential sonications were delivered to treat the entire target. Temperature-sensitive phase-difference–based MR imaging was performed during each sonication to monitor focus localization and tissue temperature changes. After the procedure, T2-weighted and contrast material–enhanced T1-weighted MR imaging were performed to evaluate immediate and long-term effects.
RESULTS: Thermal imaging sequences were improved over the treatment period, with 82% (279 of 342) of the hot spots visible in the last seven treatments. The MR imager was used to measure temperature elevation (12.8°–49.9°C) from these treatments. Eight of the 11 lesions treated demonstrated complete or partial lack of contrast material uptake on posttherapy T1-weighted images. Three lesions showed no marked decrease of contrast material uptake. This lack of effective treatment was most likely due to a lower acoustic power and/or patient movement that caused misregistration. No adverse effects were detected, except for one case of transient edema in the pectoralis muscle 2 days after therapy.
CONCLUSION: MR imaging–guided FUS can be performed to noninvasively coagulate benign breast fibroadenomas.
Index terms: Breast neoplasms, MR, 00.121411, 00.121412, 00.121415, 00.12143 • Breast neoplasms, therapeutic radiology, 00.12986, 00.12989 • Breast neoplasms, therapeutic ultrasound (US), 00.12989, 00.311 • Breast neoplasms, US, 00.12986, 00.12989 • Fibroadenoma, 00.311
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INTRODUCTION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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Lynn and colleagues (1) investigated the potential surgical application of focused ultrasound surgery (FUS) more than 5 decades ago. Since then, FUS has been extensively tested for trackless surgery of the brain in animals (2,3) and humans (4). During the past 1
decades, investigators in clinical trials in which focused ultrasound was used in noninvasive surgery of the prostate gland, kidney, liver, bladder (5–8), and eye (9) have demonstrated the clinical potential of this method.
Although the use of FUS for the ablation of malignant tumors has been repeatedly shown to be promising (10), its widespread acceptance as a soft-tissue ablation method has been limited because of the difficulty of controlling focal spot position, precise target definition, and beam dosimetry. It is fortunate that magnetic resonance (MR) imaging can satisfy these requirements of FUS therapy. MR imaging has excellent anatomic resolution for targeting, high sensitivity for localizing tumors, and temperature sensitivity for detecting temperature elevations. Several MR imaging parameters are temperature sensitive and allow detection of small temperature elevations prior to any induced irreversible tissue damage (11). Thus, the focus can be located at relatively low powers, and the accuracy of targeting can be verified. In addition, by using temperature-sensitive MR imaging sequences, focal temperature elevations and effective thermal doses may be estimated (12,13).
The technical feasibility of performing MR imaging–guided FUS by using MR imaging to guide and monitor therapy has been established (14,15) and demonstrated by several research groups (16–22). The adequate treatment of target volumes can be substantiated not only with intraprocedural temperature-sensitive MR imaging but also with postprocedural T1- and T2-weighted imaging that reveals signal intensity changes in the treated tissue (19,23, 24). Occlusion of the microvasculature within the sonicated tissue can be detected by administering intravenous contrast agent (23,25).
In the present study, we report what are, to our knowledge, the first results of a clinical application of a previously described MR imaging–guided FUS system (26,27). We decided to treat benign fibroadenomas in the breast for several reasons: Fibroadenomas are well-circumscribed benign tumors, easily distinguishable from adjacent normal breast tissue at T2-weighted and contrast material–enhanced T1-weighted MR imaging (28). Therefore, targeting is straightforward, and the response to therapy can be well monitored. Because of its relatively external position and easy immobilization, the breast is ideally suited to this type of therapy. Although a fibroadenoma represents a clinically meaningful target for noninvasive therapy, the final outcome of the treatment is not as critical as in a case of breast cancer. The ultimate goal of this therapeutic development is to establish the feasibility, safety, and effectiveness of noninvasive therapy for the treatment of malignant breast tumors. Given that fibroadenomas manifest in a similar setting and in the same anatomic location, they appear to be an ideal target for evaluation of this treatment approach. The purpose of this study was to test the feasibility of MR imaging–guided FUS for noninvasive treatment of benign fibroadenomas in the breast.
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MATERIALS AND METHODS |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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18 years), having more than a 1.5-cm distance to the skin and rib cage from the sonication plane, and having histologic confirmation after large-core biopsy (14-gauge needle size; biopsy typically performed several weeks before therapy). Women who were pregnant or lactating, had mammograms showing calcifications, or had standard MR imaging contraindications were excluded. The patients undergoing treatment ranged in age from 19 to 38 years, with a mean age of 29 years. All patients gave written informed consent after the nature of the procedure was fully explained. The lesions had a volume of 0.7–6.5 cm3, with a mean volume of 1.9 cm3. Lesion locations among the nine patients were as follows: Six lesions were in the right breast, and five were in the left breast. Four were in the upper outer quadrant; three, in the upper inner quadrant; one, in the lower outer quadrant; one, in the lower inner quadrant; one, in a lower central region; and one, central in the breast.
Equipment
Sonications were performed with a clinical MR imaging–compatible ultrasound surgery system (GE Medical Systems, Milwaukee, Wis) (26,27). A focused piezoelectric transducer with a 100-mm diameter, 80-mm radius of curvature, and 1.5-MHz resonant frequency generated the ultrasound field. The acoustic power output as a function of applied radio-frequency power was measured by using a radiation force technique (29). All results are reported here as applied acoustic power. The half-intensity beam diameter and length were 1.0 and 4.8 mm, respectively. The transducer was found to ablate a tissue volume approximately 4 mm in diameter and 6–7 mm in length during 10-second sonications when animal tissues in vivo were exposed at an acoustic power of about 40 W (27,30). The radio-frequency driving system has previously been described in detail (27).
The transducer was mounted in an acrylic plastic container filled with degassed distilled water and covered with a plastic membrane (0.075-mm-thick polyvinyl chloride) that was secured with an O-ring on top of the container. The acrylic plastic container (length, 83 cm; width, 34 cm; height, 11 cm) was mounted on a standard MR imaging table (Signa; GE Medical Systems). The transducer could be moved 10, 4, and 10 cm, respectively, in the x, y, and z directions by using a computer-controlled positioning device. The first-generation system used a hydraulic mechanism and has previously been described in detail (26,27). The second-generation system used brass lead screws and piezoelectric motors to move the transducer. The motors were placed at the end of the table, and the motion was transferred to the lead screws through telescoping shafts (Fig 1). Linear optic encoders (Computer Optical Products, Chatsworth, Calif) with a revolution of 0.13 mm were used in both systems to provide position feedback to ensure accurate movement.
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A custom-made surface coil (GE Medical Systems) was placed under the patient and around the breast to improve the signal-to-noise ratio. The acoustic coupling between the patient and the water in the positioner was achieved by placing a custom-made plastic bag filled with degassed water on top of the positioner and under the breast. The bag was flexible and contoured around the breast under its weight. The plastic of the bag extended approximately 15 cm beyond the pocket that contained the water and formed a skirt. This bag was attached to the padding on the MR imaging system so that the skirt and bag formed a watertight pocket in the padding. A layer of degassed water was poured on top of the pillow. When the breast was in place, the degassed water in the pocket replaced all of the air around the bottom part of the breast; this procedure ensured good acoustic coupling.
A treatment plan was developed from a contiguous set of coronal MR images (Fig 2). On one to three contiguous images, the number of which depended on the maximum lesion diameter in the sagittal direction, the radiologist (O.P., D.N.S.) outlined the individual tumor by using the system software, on the basis of the T2-weighted images described previously. A 5-mm distance between the sonication planes was chosen to match the length of the ablation zone. The ablations were closely packed, with a lateral distance of about 2.8 mm that resulted in overlapping treatment zones, thus ensuring that all intratumoral tissue was treated.
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MR Imaging
Pretreatment MR imaging confirmed that the biopsy-proved fibroadenoma was in an accessible location, was of a suitable size, and demonstrated enhancement after intravenous contrast material administration. Pretreatment and follow-up MR images of the breast were routinely obtained by using a dedicated breast coil (Medrad, Indianola, Pa). MR imaging examinations were performed with a 1.5-T standard whole-body system (Signa; GE Medical Systems) by using a field of view of 160–250 mm and a matrix size of 256 x 256. A T1-weighted fast spin-echo (SE) sequence (400–700/9–12 [repetition time msec/echo time msec], three data acquisitions, 3-mm interleaved section thickness, and 16-Hz bandwidth) with fat suppression and a T2-weighted fast SE sequence (2,000–4,000/100; echo train length, eight; two data acquisitions; and 3-mm section thickness) with fat saturation were performed in most cases. Images were obtained in transverse, coronal, and sagittal planes.
For contrast-enhanced images, T1-weighted imaging started about 4 minutes after intravenous injection of the gadolinium-based contrast agent (0.1 mmol per kilogram of body weight) (gadopentetate dimeglumine [Magnevist; Berlex, Wayne, NJ]). Fibroadenoma volumes were calculated by assuming the volume v of an ellipsoid v = 4/3 
abc, with a, b, and c being half the diameter in three orthogonal directions, as measured on the T2-weighted images.
Temperature Monitoring with MR Imaging
The temperature elevations during the 10-second sonications were monitored across the focal plane by obtaining temperature-sensitive MR images before, during, and after sonication. The temperature mapping evolved during the clinical trial. The first treatment was monitored by performing T1-weighted gradient-echo imaging with no visible hot spots. The next four treatments used phase-difference imaging, but only one image was obtained during the sonications with manual triggering. In the last seven treatments, phase images obtained before, during, and after the sonications with computer triggering to ensure accurate timing were used. Phase imaging was performed to estimate the temperature-dependent proton-resonant frequency shift (31). Phase maps were acquired by performing the following fast spoiled gradient-echo sequence: 27.3/13.5, 30° flip angle, 7.2-kHz bandwidth, 256 x 128 resolution, 16-cm field of view, and 3-mm section thickness.
Five to 10 images were obtained in a series, with a total acquisition time of 18–36 seconds. The first image acquisition was triggered 4 seconds prior to the start of sonication. The imager was programmed to reconstruct the magnitude and the real and imaginary images for each of these times. The real and imaginary images were used to calculate the phase difference between the two times, as described in references 32 and 33. The temperature dependence of the proton-resonant frequency has been shown to be linear above the coagulation threshold (27,34,35).
To judge the adequacy of the treatment, the MR imaging–derived temperature information was analyzed, and the temperature-time history for each image voxel was calculated. The peak temperature derived from the MR images was used as a guide to ensure that adequate power was delivered to coagulate the target tissue. The power level was set after the initial localization pulse by increasing the power and repeating the sonications in the middle of the tumor. The power was kept constant at each sonication level, unless several sonications indicated that the temperatures were too low. In such a case, the power was increased for all remaining sonications.
Follow-up
Patient response to FUS therapy at MR imaging, as compared with patient status at the time of treatment, was judged by one of the authors (P.E.H.) on the basis of the decrease in size of the contrast-enhancing region on T1-weighted images after intravenous contrast material administration. Partial response was defined as a 50%–90% decrease in the contrast-enhancing region, and complete response was defined as a greater than 90% decrease in contrast material uptake. Minor response was indicated when there was a 10%–49% decrease in the contrast-enhancing region. In addition, tumor volume was measured on the basis of T2-weighted images and then compared with the initial tumor volume status. This was necessary because T2-weighted images were used as the treatment-planning platform. The imaging parameters used for the pretreatment images were used in follow-up imaging.
To evaluate potential acute effects such as edema and skin damage, the first follow-up imaging examination was performed by one of the authors immediately after therapy. Further follow-up imaging was performed at 2 and 10 days, at 6 months, and then at individually prescribed times. The median follow-up time was 6 months (range, 1.5 months to 4 years). Routine follow-up included clinical examination and MR imaging. The patient graded acute pain and discomfort at each follow-up visit as none, slight, moderate, or severe. Diagnostic US examinations and mammography were performed if individually indicated. Tumor volumes before and after therapy were compared by performing a two-tailed paired t test.
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RESULTS |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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Five of the 12 treatments resulted in only limited tumor devascularization after therapy, as demonstrated on postcontrast T1-weighted images (Fig 6). Two of the five patients who underwent these treatments had these results because of insufficient power administration, which resulted in subthreshold temperature levels that developed in the treatment zone. The first patient was deliberately given low-power exposures (about 30 W). In the second patient, the power level was low (32 W) because the patient experienced substantial pain at levels over 35 W. This pain can be explained by the administration of only 5 mL of local anesthetic. After 48 hours, the tumor was still enhancing on T1-weighted images. As a consequence, the tumor was treated again 4 weeks later by using 10 mL of a local anesthetic, which led to complete response in the follow-up images (Table 2).
A third patient experienced treatment failure due to the placement of excess local anesthetic anterior to the fibroadenoma. Unavoidable microscopic bubbles in the local anesthetic injected in front of the fibroadenoma probably caused scattering of the ultrasound beam and thus limited the power delivered to the tumor. In support of this hypothesis, the patient developed a bruise after therapy; this was the only clinical complication witnessed in any of our patients.
In the fourth patient, failure was due to patient motion. A follow-up examination in which T2-weighted imaging was performed 2 days after treatment demonstrated some edematous changes in the pectoralis major muscle that was adjacent to the fibroadenoma. On these images, however, the fibroadenoma itself was unchanged when compared with that on the pretreatment images. This adverse effect (tissue edema), visible only on the images, did not result in identifiable clinical complications or complaints, and no posttreatment therapy was necessary.
In the fifth patient, treatment produced devascularization in part of the tumor; the tumor was not completely treated owing to patient motion that, in turn, caused tumor motion and necessitated recommencement of treatment. These factors, together with equipment problems, prolonged treatment, and the sonications had to be terminated prior to the completion of therapy. An attempt was made 6 weeks later to complete treatment, but the tissue changes induced by the prior treatment made tumor targeting difficult. Although the posttherapy contrast-enhanced images showed a small nonenhancing tissue volume close to the tumor, no additional tumor was coagulated.
In these and other cases, the acute pain and discomfort were well tolerated. In four treatments, pain was graded by the patient as none; in four treatments, as slight; in two treatments, as moderate; and in one treatment with insufficient local anesthesia, as severe. Aside from local pain and moderate swelling, no other complications were seen in any of the patients. Other potential complications such as skin burns, substantial swelling, and sterile abscess were not observed. Typically, the most prominent findings after treatment were the changes that resulted from core biopsy prior to FUS treatment. One patient developed a bruise on the skin of the breast. Some tenderness around the treatment zone was common for up to 10 days after the procedure, but most patients described it as minor.
Correlated to the mild clinically adverse findings immediately and up to 10 days after therapy, transient edema surrounding the target volume was visible on T2-weighted images. On occasion, the edema extended to the subcutaneous region. By 6 months after treatment, no visible signs or side effects persisted, and no late or delayed adverse effects were noted. In three patients, additional bilateral mammograms were obtained and showed no treatment-related changes up to 4 years after therapy.
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DISCUSSION |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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In the current study, we have shown that noninvasive treatment of fibroadenomas with MR imaging–guided FUS is both feasible and safe, without marked adverse effects. Our results also demonstrate the advantages of combining a noninvasive focused ultrasound technique with MR imaging in a clinical setting. Furthermore, we have shown that MR imaging is suitable for both FUS treatment planning and delineation of FUS therapy–induced changes in the female breast. The clear visibility of FUS-induced changes in the breast tissue, such as edema and devascularization, further enhances the usefulness of MR imaging guidance. With respect to FUS, the most important feature of MR imaging, however, is probably its distinctive use in localizing the hot spot and measuring temperature changes during treatment. This use of MR imaging allows the monitoring of both the efficacy and safety of focused ultrasound therapy.
Most of the hot spots were visible on the serial phase-difference images; this finding ensured that the right target was sonicated. Some of the hot spots, mainly those at the edge of the tumor, were not visible; we presume that this was because the sonication location was primarily fat. The proton-resonant frequency shift is not temperature-sensitive in fat. The temperature history during the sonications was obtained in about 172 (48%) of the 356 sonications in the last seven treatments (Table 1) that used the triggered series of phase images to monitor temperature elevation. Relatively small patient motion during the sonications was the reason that many of the temperature measurements were unreliable. When the motion was large, the focus was not seen at all. The temperature was calculated by subtracting the phase before the sonications from the phase during the sonications to obtain the temperature elevation. Any motion can be easily detected by monitoring the difference signal in voxels outside the sonicated volume, so the operator can avoid using motion-distorted temperature values. It is clear that the therapy could benefit from thermal mapping sequences that are less motion sensitive. Several promising methods may decrease motion sensitivity (40,41). The other factor influencing temperature measurements is the influence of fat on the temperature coefficients in voxels that have both fat and water. Fat suppression may help to decrease this problem (42).
In the 11 lesions treated, eight therapies were partially or nearly completely successful. Success was judged on the basis of postcontrast T1-weighted images showing a partial or complete lack of contrast material uptake, which implies devascularization, and tissue necrosis similar to that in the findings of various histologic studies of focused US effects in animals in vivo (17,23,43–46). These contrast changes were present immediately after treatment and lasted the entire follow-up period. Thus, only one follow-up imaging session may be needed to determine the treatment result. The best time to perform this may be after approximately 1 week, when treatment-related tissue swelling has resolved. Moreover, after 6 months, the mean volume based on T2-weighted imaging of the treated fibroadenomas decreased, whereas that of nontreated fibroadenomas in the same patients increased or remained stable. Furthermore, at clinical examination, the treated fibroadenomas were not only smaller but also softer. In three patients, posttherapy mammograms were obtained and demonstrated tumor size decrease without detectable scar formation.
Four treatments resulted in no reduction in contrast material uptake in the tumors. One treatment resulted in reduced contrast material uptake in part of the tumor. As described in Results, this was caused by insufficient power in two cases, injection of local anesthetic in front of the lesion in one case, and patient movement that caused the therapy to miss the tumor in one case. The power problem was solved in one case by retreating at a higher power level. This case demonstrates how posttherapeutic MR imaging can be performed to ensure that the desired end result has been achieved. The other cases show that if local anesthetic is used, it should be placed beyond the lesion and that tumor location should be monitored throughout therapy. These guidelines were followed in the remaining treatments, which resulted in success. One case was terminated because of the extensive length of the treatment, which was caused by patient motion and system failure. The attempt to continue treatment 6 weeks later was unsuccessful because of difficulties in identifying the tumor on T2-weighted images; this difficulty was partially due to the changes induced by the previous treatment. However, the remaining tumor was readily visible in contrast-enhanced images obtained after treatment. This indicates that there may be cases in which tumors are not adequately identified on T2-weighted images alone.
A potential source of the different imaging responses and patient discomfort during treatment might have been the relatively long treatment time (45–240 minutes). The transducer used in the present study, which had a frequency of 1.5 MHz, produces thermal lesions approximately 4 mm wide and 6–7 mm long. These lesion sizes were relatively similar to the ones induced with other single FUS systems for thermal ablation (47,48). Therefore, when using single-focus transducers, larger tumors must be treated with a series of overlapping lesions and with a delay between sonications to avoid heating the tissue on the path to the target zone (36,49,50). Considerably shorter treatment times can be achieved by using continuous-wave ultrasonic phased-array technology (49,51). For example, in a recent study in which a phased-array applicator was used (52), it was demonstrated that a single 20-second sonication produced thermal lesions greater than 1 cm3 in pig liver in vivo.
This initial series of patient treatments provided information that has been used to improve the treatments. The following changes will be made for future clinical treatments: First, a new, more reliable positioning system that will allow faster treatments has been designed and constructed by TxSonics (Haifa, Israel). Second, a more optimal surface coil that will improve the signal-to-noise ratio of targeting and thermal imaging has been designed by GE CRD (Schenectady, NY) for breast treatments. Third, online temperature and thermal dose calculation programs have been developed to provide a display of thermal exposures during treatment. This workstation-based system also allows patient motion to be monitored. Fourth, a phased-array system has been approved by the Food and Drug Administration to produce larger focal spots and thus increase the size of the tumors that can be treated in a reasonable time. Fifth, double-echo sequences have been tested with short echo times to evaluate the T1 changes that indicate temperature elevation in fat and long echo times for phase imaging in the tumor tissue. Finally, the use of contrast-enhanced images for targeting will allow tumors that are not visible on T2-weighted images to be treated.
Two additional factors should be considered for elimination of patient pain, which may contribute to motion during treatment. First, our approach of injecting the local anesthetic behind the lesion may be suboptimal, and more work needs to be done to eliminate pain in the treated volume. Second, the closest allowed distance of the US focus from the skin and ribs was 15 mm. In agreement with our animal test results, no skin damage was observed. Therefore, the focal depth distance of 15 mm could probably be reduced in future trials. The distance from the rib cage was also 15 mm. This is probably close to the limit with this transducer size and may have been a source of patient pain. The focus-to-bone distance could have been reduced by using a larger-diameter transducer or transducer array.
One of the most important potential applications of FUS ablation is the treatment of breast cancer. The noninvasive treatment of fibroadenomas with FUS has limited promise in its own right for some indications; for example, in patients who are uncomfortable with palpable lumps and prefer noninvasive therapy to avoid surgical scars or who are not candidates for open surgery for medical reasons. However, the present study should be regarded more as a feasibility study, the first step toward a therapy for breast cancer. A noninvasive ablative method of tumor debulking has great appeal to patients who want to preserve the integrity of the breast. Whether MR imaging–guided FUS can replace the current practice of lumpectomy plus radiation therapy in select patients who have breast cancer with small tumors and the appropriate histologic background remains an open question. It is clear that tumor definition with MR imaging is better than that with direct surgical visualization or any other imaging modality. At the same time, it is more difficult to detect the true extent of invasive breast cancer than to detect that of fibroadenomas, and important prognostic information relating to margins and complete histologic information may be lost by using the noninvasive technique.
In summary, in this study we have successfully performed clinical MR imaging–guided FUS. The therapy proved safe and effective in breast fibroadenomas. MR imaging–guided FUS is recommended for further testing as an alternative to invasive surgery for patients who have fibroadenomas. Moreover, and probably more important, MR imaging–guided FUS has the potential to become an important modality for the local treatment of malignant breast tumors in combination with radiation therapy.
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ACKNOWLEDGMENTS |
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FOOTNOTES |
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Author contributions: Guarantors of integrity of entire study, K.H., F.A.J., P.E.H.; study concepts and design, K.H., F.A.J., O.P., S.S.; definition of intellectual content, O.P., P.E.H., K.H., F.A.J.; literature research, K.H., O.P., P.E.H.; clinical studies, all authors; data acquisition, O.P., D.N.S., P.E.H., N.J.M., J.K., K.H.; data analysis, O.P., P.E.H., N.J.M., K.H.; statistical analysis, P.E.H.; manuscript preparation, O.P., P.E.H., N.J.M., K.H., F.A.J.; manuscript editing, review, and final version approval, all authors.
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REFERENCES |
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TOPABSTRACTINTRODUCTIONMATERIALS AND METHODS RESULTSDISCUSSIONREFERENCES |
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