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Abstract Refers to the similar tower crane, this design is composed by the system design and the lazy arm design to the QTZ500 tower crane. In the lazy arm design progress, it has carried Finite Element method on the analysis computation, and used ANSYS10.0 software. According to the entire machine main performance parameter, various organizations type and the steel structure pattern has been determined. The design parameter of operating modes which are composed of nose increase, the cross center increase and the root increase. Through the suitable simplification to the lazy arm, the lazy arm finite element model is establishment applied ANSYS10.0 software, and then exerted various operating modes load, carried on the solution. Then ANSYS10.0 software can calculate various pitch points stress situation, various units receive the axial stress size, and the lazy arm distortion size under various operating modes. Also it can demonstrate the animation in the process of the lazy arm increase. It has clearly displayed the lazy arm stress performance under various operating modes. Through the revision for model parameter, the analysis comparison is carried on the different model. Because the stress condition and rigidity condition of different model is compared under the same operating mode, and the generalized analysis intensity and the rigidity condition is carried on, a most reasonable model parameter can be obtained, though the intensity and the rigidity examination regarding this model, then the final parameter result of the lazy arm can be obtained. Key words: QTZ500 tower crane Lazy arm Finite element analysis ANSYS10.0 河北建筑工程學(xué)院 畢業(yè)設(shè)計（論文）外文資料翻譯 系別： 機械工程系 專業(yè)： 機械設(shè)計制造及其自動化 班級： 機053班 姓名： 張勇杰 學(xué)號： 22號 外文出處： European Journal of Radiology 附 件：1、外文原文；2、外文資料翻譯譯文。 指導(dǎo)教師評語： 簽字： 年 月 日 MR imaging at high magnetic fields Masaya Takahashi a, *, Hidemasa Uematsu b , Hiroto Hatabu a a Department of Radiology, Beth Israel Deaconess Medical Center, Boston, MA 02115, USA b Department of Radiology, University of Pennsylvania Medical Center, Philadelphia, PA, USA Received 12 November 2002; received in revised form 13 November 2002; accepted 14 November 2002 Abstract Recently, more investigators have been applying higher magnetic field strengths (3C1/4 Tesla) in research and clinical settings. Higher magnetic field strength is expected to afford higher spatial resolution and/or a decrease in the length of total scan time due to its higher signal intensity. Besides MR signal intensity, however, there are several factors which are magnetic field dependent, thus the same set of imaging parameters at lower magnetic field strengths would provide differences in signal or contrast to noise ratios at 3 T or higher. Therefore, an outcome of the combined effect of all these factors should be considered to estimate the change in usefulness at different magnetic fields. The objective of this article is to illustrate the practical scientific applications, focusing on MR imaging, of higher magnetic field strength. First, we will discuss previous literature and our experiments to demonstrate several changes that lead to a number of practical applications in MR imaging, e.g. in relaxation times, effects of contrast agent, design of RF coils, maintaining a safety profile and in switching magnetic field strength. Second, we discuss what will be required to gain the maximum benefit of high magnetic field when the current magnetic field (5/1.5 T) is switched to 3 or 4 T. In addition, we discuss MR microscopy, which is one of the anticipated applications of high magnetic field strength to understand the quantitative estimation of the gain benefit and other considerations to help establish a practically available imaging protocol. # 2002 Elsevier Science Ireland Ltd. All rights reserved. Keywords: Magnetic resonance imaging; Higher magnetic field strength; Contrast agent 1. Introduction Thanks to recent technological development, whole- body magnetic resonance (MR) scanners at higher magnetic field strengths (]/3T)have been introduced into research and clinical settings. In the beginning, one of the main reasons to install higher fields was its higher sensitivity to the blood oxygenation level-dependent effect for functional MR imaging of the brain [1]. Recently, more investigators applied these higher mag- netic field strengths to both research and conventional clinical settings. The expectation for higher magnetic fields in MRI is the improvement in signal-to-noise ratio (SNR) due to higher signal intensity (SI), where the most significant benefit is to decrease the length of time required to obtain images. Then, higher spatial resolu- tion may be achievable. One question is how it improves or practically how beneficial it is when we switch the current magnetic field (5/1.5 T) to 3 or 4 T. Several studies have reported and discussed the advantages of higher magnetic field in, for example, delineation of various brain lesions [1] or cardiac structures [2,3]. Dougherty et al. [2] reported that the SNR of the anterior myocardium at 4 T was 2.9 times higher than that of the same region at 1.5 T. Bernstein et al. demonstrated contrast enhanced imaging at 3 T and concluded that higher spatial resolution at 3 T could improve diagnostic accuracy [4]. In addition, if higher magnetic field can provide better image quality, it may be reasonable to expect a reduction in total injection of contrast agent, for example, in MR angiography which needs to cover a larger area of the peripheral artery [5] or the lung [6,7]. However, such speculation would be difficult to prove as higher magnetic fields change other imaging aspects besides SNR. Many theoretical and experimental studies havebeen employed to demonstrate the magnetic field dependen- cies. Besides SNR, the magnetic field-dependence is * Corresponding author. Tel.: C27/1-617-667-0198; fax: C27/1-617-667- 7021. E-mail address: mtakahas@caregroup.harvard.edu (M. Takahashi). European Journal of Radiology 46 (2003) 45C1/52 0720-048X/02/$ - see front matter # 2002 Elsevier Science Ireland Ltd. All rights reserved. PII: S 0 7 2 0 - 0 4 8 X ( 0 2 ) 0 0 3 3 1 - 5 well-documented in tissue relaxation times [8C1/10],as well as in MR contrast agent effects (e.g. R1, R2 or R2* relaxivities) [11,12]. SNR depends upon imaging para- meters, RF coil sensitivity and machine adjustments, such as magnetic field homogeneity, accuracy in excita- tion/refocusing pulse settings, etc. These theoretical and experimentally proven properties suggest that imaging parameters must be reconfigured for different magnetic fields. Unlike relaxation time and MR contrast agent effects, the benefit to signal intensity at higher magnetic field should be compared under nearly identical experi- mental conditions. Therefore, it is imperative to quan- tify the practical differences in terms of SNR and contrast-to-noise ratios (CNR) between higher and lower (B/1.5 T) magnetic fields. However, the studies of direct comparisons between SNRs and CNRs as an outcome of the combined effect of several magnetic field-dependent parameters at different fields compared with the theoretical values are substantially sparse. Hence, it is still unclear how much benefit we can gain in SNR or what we can/should do in switching a current magnetic field strength (5/1.5 T in most cases) to a higher magnetic field. In this article, we consider the magnetic field dependent alterations, e.g. MR signal on the image, relaxation times, effects of contrast agent, design of RF coil and safety profile. Then, we evaluate the scientific expectations for MR imaging on a higher magnetic field to quantify the scientific and technical issues relative to safe human experimentation. Further, the feasibility of MR microscopy, which is one of the expectations of higher fields, is discussed. 2. SI, SNR and CNR The question of optimum field strength has been a subject of intense controversy for over a decade. The interest in higher fields stems from the fact that SNRs increase with field strength (v), where SI and noise have different magnetic field-dependencies. SI8(number of spins) C29(voltage induced by each spin) (1) As shown in Eq. (1), theoretically, the signal intensity from a MR experiment is proportional to the square of the static magnetic field (v 2 ) since both ‘number of spins’ that can be observed and ‘voltage induced by each spin’ increase linearly as magnetic field (v) increases. Noise is proportional to the static magnetic field (v), when all noise comes from a sample, resulting in an SNR that is proportional to v in the case. On the other hand, noise is proportional to one-quarter of v (v 1/4 ) when all noise comes from the RF coil, resulting in an SNR that is proportional to v 7/4 . Therefore, SNR can be expected to increase more than 2.7 (C30/4/1.5) times at 4 than at 1.5 T. If this is true, since the SNR scales as the square root of the number of image averages, the time needed to obtain the same SNR is reduced by a factor of 8. To confirm this theory, we imaged the brain in a subject at both fields. To make our comparison between the magnetic fields as direct as possible, the same sets of experiments in the same subjects were conducted at both 4 and 1.5 T on the commercially supplied whole-body MR scanners (Signa TM , General Electric Systems, Mil- waukee, WI) with the equipped head coils. Fig. 1 shows the T1-weighted images (top) and T2-weighted images (bottom) obtained in the same level of the brain of the same subject. Each image was obtained with a conven- tional spin echo sequence with the same imaging parameters at 1.5 and 4 T, respectively. These images showed different tissue contrast between the magnetic fields even though the images were acquired with the same set of imaging parameters. In the quantitative measurements of SI, we found that 4 T increased the SI in both white and gray matter (Fig. 1). In addition, those enhancement ratios were also different between the imaging parameters (T1-WI and T2-WI). Thus, 4 T provides a different tissue contrast compared with 1.5 T using the same set of imaging parameters, which might be inconsistent with theoretical values. 3. Relaxation times As discussed above, SNR in biological tissue was found to be in approximate proportion to field strength. However, the practically achievable SNR gain may be somewhat less since the above theory assumes that all parameters except the magnetic field are consistent. One reason for the discrepancy is the increase in T1 relaxa- tion time with increasing field strength. SI is a function of relaxation time that is, in turn, magnetic field- dependent [3]. In theory, T1 value increases in a magnetic field-dependent manner in most biological tissues of which the correlation time (t c ) of tissue water is :/10 C288 s [13], whereas T2 value does not change (Fig. 2). Comparisons of relaxation times in humans have been published in the literature. Jezzard et al. and Duewell et al. presented a comparison of T1 and T2 relaxation times in human subjects between 1.5 and 4 T in the brain and several peripheral regions [9,10] (Table 1). In any tissue, T1 relaxation times are prolonged at a higher magnetic field, while T2 relaxation times are somewhat shortening. Those results are consistent with previous reports (Fig. 2). To confirm this phenomenon, we conducted the same set of phantom experiments at both 4 and 1.5 T on the same whole-body MR scanners with head coils [14]. Phantoms included different con- centrations of Gd-complex aqueous solution with each phantom representing tissue with a different T1 relaxa- M. Takahashi et al. / European Journal of Radiology 46 (2003) 45C1/5246 tion time. In this study, the trains of spin echo images with varied TRs or TEs were obtained with the same commercial clinical scanners with the head coils de- scribed above. The relaxation times (T1, T2) for all phantoms were determined at both 1.5 and 4 T from the fitting curves. The results in this confirmatory study demonstrated that any T1 relaxation times were pro- longed (1.10C1/1.47 times) at 4 T compared with those at 1.5 T, while T2 values were identical or slightly shortened (Table 2). Further, a standard contrast-enhanced MR angio- graphic sequence (3D spoiled gradient recalled acquisi- tion or SPGR) sequence with the same imaging parameters was utilized to confirm changes in SI. Peak SNRs at 4 T increased at least 2.21 times higher compared with those at 1.5 T. Moreover, peak CNRs at 4 T increased at least 1.59 times higher compared with those at 1.5 T in the range of Gd concentrations expected during clinical use. In addition, those enhance- ments of SNR and CNR were a function of a flip angle that we used. Based on those results, using higher Fig. 1. T1- and T2-weighted images of a human subject obtained at 1.5 and 4 Tesla. Each image was acquired with the same set of imaging parameters (TR/TE is indicated in the parentheses), respectively. Note that different magnetic fields provided different image contrast. Fig. 3. Cross-sectional T1-weighted image of a fixed excised spinal cord of the larval sea lamprey. Image was obtained at 9.4 T experimental machine; resolution was 9C29/9 mm resolution. See Ref. [27]. Fig. 2. Magnetic field dependency in T1 and T2 relaxation times, modified from Ref. [13]. M. Takahashi et al. / European Journal of Radiology 46 (2003) 45C1/52 47 magnetic fields seems to be beneficial in CNRs as well as in SNRs even without optimization of imaging para- meters at each magnetic field. A relationship between the SI of a gradient echo sequence, the relaxation time and the optimal flip angle (a o : Ernst angle), can be expressed as follows: SIC30bC215 [1 C28 exp(C28TR=T1)] C215 exp(C28TE=T2C31) C215 sin a 1 C28 exp(C28TR=T1) C215 cos a (2) and cos a o C30C28exp(TR=T1) (3) where b is the scaling factor and a is the flip angle. SI is determined by its relaxation times (T1 and T2*) in individual tissue conditions in any imaging sequence. This implies that the same intensity will not be obtained with the same set of imaging parameters due to the alternation of relaxation times at different magnetic field. Since T1 values at higher magnetic field are longer than those at lower magnetic field, the TR, presumably as well as the flip angle, should be longer (smaller for flip angle) to optimize the SNR of the same sample at the higher field. Using longer TR, the advantage in SI at a higher field would be less in unit time. In other words, since the primary limitation imposed by long T1 relaxation time at higher magnetic field strength is reflected in the TR, the SNR per unit time is optimized with an Ernst angle pulse and the shortest achievable value of TR/T1. The necessity of optimization of imaging parameters was presented in a previous work. Keiper et al. [15] compared the usefulness in the diagnosis of white matter abnormalities in multiple sclerosis patients following the optimization of imaging parameters between 1.5 and 4 T. Their results demon- strated that MR imaging at 4 T (512C29/256 matrix) could depict smaller lesions that could not be detected at 1.5 T (256C29/192 matrix), implying that the higher resolution at 4 T provides higher accuracy of diagnosis in the same patients with almost identical total scan time. Although T2 values were substituted for T2* in the phantom study because T2 and T2* values should be theoretically identical in phantoms in each magnetic field [16], it is considered to be different from the conditions in some tissues where the T2* value is much shorter than the T2 value in some tissues. A magnitude of susceptibility (g) is proportional to the magnetic field as shown in the following equation [17]: gC30 C18 Dx 2 C19C18 B 0 RG z C19 (4) where Dx is the difference in magnetic susceptibility of adjoining substances, B 0 (C30/v) is the static magnetic field, R is the cross section radius and G z is the read-out gradient. However, this effect on T2* depends on T2 in tissue since 1/T2* is a function of T2 and T2? (R2*C30/ R2C27/R2?) [18]. The shorter T2 and T2* values at a higher magnetic field may cause a larger decrease in the SNR and CNR than would be expected in some tissue, such as the lung. Previously, we found that the CNR increased in the central arteries of the lung, but did not increase in the pulmonary peripheral arteries at 4 T as the dose of contrast agent increased, ranging from 0.05 to 0.2 mmol/kg body weight [19]. Therefore, the optimal imaging parameters for the clinical application should be carefully considered, particular when an undesirable T2* effect may be involved. 4. Relaxivities of Gd-complex The R1 relaxivity of MR contrast agent is dependent upon various parameters, such as the type of contrast agent [20], temperature and tissue environment as well as magnetic field strength [11,12]. R1 relaxivity of a paramagnetic contrast agent is higher at lower field strength [11]. R2 and R2* values should be theoretically identical in phantoms in each magnetic field [16]. In the phantom study described above, the authors attempted to compare the effects of contrast agent. For an accurate determination of the efficacy of Gd-complex (R1, R2 and R2*), only some of the relaxation times Table 1 Comparison of T1 and T2 relaxation times in human subject [9,10] Tissue T1 (s) T2 (ms) 1.5 T 4 T 1.5 T 4 T Brain a Gray matter 0.9C1/1.3* 1.72 77C1/90 63 White matter 0.7C1/1.1* 1.04 62 C1/80 50 Muscle b 0.98 1.83 31 26 Fat b 0.31 0.39 47 38 Bone marrow b 0.29 0.42 47 42 a Lezzard et al. [9]. b Duewell et al. [10]. * From previous literature. Table 2 Comparison of T1 and T2 relaxation time in gadolinium doped water solution at room temperature, modified Ref. [14] Gd concentration (mmol/l) T1 (ms) T2 (ms) 1.5 T 4 T 1.5 T 4 T 0 2556 3636 1643 1504 0.125 1067 1566 911 862 0.5 419 562 348 351 1.25 191 253 160 160 2.5 123 142 84 83 5 67814342 At room temperature. M. Takahashi et al. / European Journal of Radiology 46 (2003) 45C1/5248 (T1, T2) that could be excellently fitted to the curve(rC21/ 0.995) were reciprocally plotted against the concentra- tions of Gd at both 4 and 1.5 T. As a result, R1 and R2 relaxivity values were determined to be 2.95 and 4.82 (lC215/ s C281 C215/mmol C281 ) at 4 T and 3.89 and 4.67 (lC215/s C281 C215/mmol C281 ) at 1.5 T, respectively. R1 at 4 T was lower (:/25%) than R1 at 1.5 T, while the R2 at 4 T was almost that at 1.5 T (Table 3). Hence, we found that R1 relaxivity decreases as the magnetic field strength increases, while R2 relaxivity does not change as much, which is consistent with previous reports [16]. Unlike Gd-complex, R2 and R2* might be consider- ably changed depending upon the type of contrast agent (e.g. super paramagnetic iron oxide: SPIO), application root and/or tissues. This suggested that we should also consider the use of the MR contrast agent, though it is not clear whether this change is substantially effectivein current clinical usage at higher magnetic field. 5. RF coil The application of higher magnetic field strengths to MR imaging (particular in whole body imaging) is more demanding because of the difficulty in building RF coils since the penetration of radio frequency into the tissue becomes harder [3,21]. It is necessary to understand the relationship between SNR and RF coil, since an incomplete RF coil may sacrifice the advantage in SNR at increased magnetic field strength. RF coil characteristics, especially a receive coil, significantly impact SNR. SNR increases with decreasing coil diameter. Thus, the coil sensitivity of the head coil is :/3-fold higher than that of the body coil. The surface coil with smaller diameter gains more sensitivity, whereas the SNR drops off very rapidly with increasing depth from the surface. To cover these difficulties, an array of surface coils must be developed. Reported by Wright et al. [22], another idea to increase coil sensitivity and further improve SNR is to reduce coil temperature, thus lowering its resistance and thermal noise voltages and increasing its Q, while keeping the sample at room temperature. The cryogenic SNR gain would be greatest for coil and sample configurations having Q L /Q U close to 1. 6. Safety consideration Theoretical calculations of the interaction of high magnetic fields with human subjects havebeenreviewed. To date, no hazardous physical or physiological phe- nomena have been shown. The mechanism considered included orientation of macromolecules and mem- branes, effects on nerve conduction, electrocardiograms and electroencephalograms, and blood flow. The most current clinical MR imagers at lower magnetic field (5/1.5 T) equip up to 25 mT/m. If higher magnetic fields are to be used to archive higher spatial resolution, the gradient strength must increase. In the combination of higher statistic magnetic field and gradients, strength may be an issue in some applications due to limitations in the current FDA guidelines for specific absorption rate (SAR). SAR is defined as follow: SARC30 sjEj 2 2r C18 t TR C19 N P N S (5) where s is conductivity, E is the electric field, r is tissue density, t is pulse duration and N P and N S are number of pulses and image slices, respectively. Since E is proportional to static magnetic field, SAR greatly increases at higher magnetic field, which may limit the application in number of slices, selection of flip angle, etc. Additionally, RF energy is absorbed more effec- tively at higher frequencies; RF absorption, as expressed by SAR, must be carefully monitored. This could be a major concern in any application at high field strength as Bottomley et al. previously suggested [21]. 7. MR microscopy In using a higher magnetic field, the investigators expect images with higher spatial r