## 1 Introduction  
Ultrasound diagnostic equipment, which has widely been used in clinical diagnosis, assumes that sound velocity has a uniform distribution, since it is based on pulse-echo imaging. However, in biological soft tissues, the sound velocity ranges from 1400 m/s to 1600 m/s [1]. For example, the sound velocity in mammalian fat tissues ranges from 1400 m/s to 1490 m/s [1]. Therefore, since the sound velocity for ultrasound imaging (US imaging) is usually assumed to be 1540 m/s, the imaging of a patient with a thick layer of subcutaneous fat is degraded due to variations in the sound velocity. Significant image degradation presumably reduces spatial resolution and the signal-to-noise ratio (SNR) in US imaging. If the sound velocity distribution is known before imaging, image degradation could be reduced by compensating for improper beamformation. Various methods for the in vivo measurement of the sound velocity distribution have been proposed [2; 3; 4; 5]. For example, [PERSON] et al. [6] and [PERSON] et al. [7] proposed methods for the measurement of the sound velocity in cartilage using magnetic resonance imaging (MRI) and US imaging. In the two methods, sound velocity is estimated based on the length of cartilage, as measured with MRI, divided by the time-of-flight, as measured with the ultrasonic pulse-echo technique. Since the estimated sound velocity is obtained at different times and locations in US imaging and MRI, the measurements are not accurate enough to correct beamforming. Thus, the aim of the current study was to evaluate the feasibility of a method of measuring the sound velocity distribution in vivo using a prototype simultaneous US imaging and MRI system. This method corrected beamforming to improve image quality in terms of the spatial resolution and SNR. In this study, magnetic-resonance-visible (MR-visible) fiducial markers were attached to an ultrasound probe to display a cross section of US imaging in MRI coordinates. The ultrasound probe was developed for use in the coils of an MRI scanner. When measuring the sound velocity distribution with a simultaneous US imaging and MRI system, the following issues need to be resolved: (1) the development of an ultrasonic array probe for use in an MRI gantry; (2) the development of MR-visible fiducial markers and the estimation of the location of a cross section in MRI coordinates; and (3) the suppression of crosstalk (electrical noise) between the MRI scanner and ultrasound equipment. [PERSON] et al. [8] imaged a phantom kidney and a rabbit kidney by mechanically scanning a single transducer. [PERSON] et al. [9] used a commercially available probe with attached MR-visible fiducial markers to simultaneously image a phantom. In the current study, linear array transducers made of 1-3 piezocomposite materials [10; 11] were embedded in an ultrasound probe for use in an MRI gantry. The probe was made of nonmagnetic materials, and MR-visible fiducial markers were attached to it. The probe used in the gantry was connected to ultrasound equipment via a connector that passed through the walls shielding the MRI room and the control room. Crosstalk between the ultrasound equipment and the MRI scanner was reduced by grounding the connector. Simultaneous imaging of an abdominal phantom and a human neck was performed to evaluate the feasibility of the proposed method. This was accomplished with US imaging equipment with 128 transmission/reception channels and a 1.5 T MRI scanner. Imaging of the phantom indicated that the error rate for the accuracy of sound velocity measurement was less than 6%, while the spatial resolution was 0.43 (the ratio of the lateral resolution before and after compensation), and the SNR was 8 dB. Simultaneous imaging of a human neck was performed to evaluate the compensation for image degradation due to the layer of subcutaneous fat. An acoustic standoff pad with a sound velocity similar to that of the fat tissue was inserted between the probe and the surface of the skin on the neck.