by O’Regan are also overestimated, especially at the lower end of the range where the MRI contents already exceeded the MRS fat contents by 0.2–1.1%. Another recent modification is the sampling of the water and fat signals at in-phase and partially opposed-phase acquisition, meaning that the second image is acquired at a sampling angle not equal to 180° [74]. Similar to the quadrature sampling method [75–77] and direct phase encoding [78], water and fat are separated as well as identified using the lagging phase relationship [74]. An alternative MRI method that is possibly more precise (with only small systematic error), is the slice-selective gradient echo technique with spectral-spatial excitation [79, 80]. In fat-selective MRI the fat contents are typically calculated directly, i.e. after the appropriate corrections, from a water-saturated single slab fat image in which the signal intensities of liver tissue and subcutaneous fat are compared. In phantom measurements the fat contents were shown to be highly reproducible and well correlated with the results of optical image analysis [79]. On the other hand, because of the signal-to-noise ratio of the images, the detection threshold in vivo was assessed to be as high as 2% of volume [80]. All of the above MRI methods will work if T1 and T2* relaxation are corrected for, but will become less reproducible in the case of very short T2* due to, for example, iron overload. Overall, it appears that the Dixon-based MRI techniques for determining liver fat content currently are more reliable and robust than other methods, and thus recommended for routine clinical application. Metabolism (MR spectroscopy) In MRS nuclei other than hydrogen, such as phosphorus (31P) and carbon (13C), are hampered by sensitivity (31P, 13C) and natural abundance (13C) [81]. Hydrogen (1H, proton) has the highest sensitivity of all radiofrequency excitable nuclei and is therefore most suited to be studied in MRS, especially on low-field MRI systems. However, 1H MRS detection of tissue metabolites such as choline- containing compounds (Cho) and creatine/phosphocreatine (Cr) at concentrations near 1 mM requires effective suppression of the abundant water and fat signals. For that reason, phosphorus (31P) MRS detection of phosphate metabolites (phosphomonoesters, PME; inorganic phos- phate, P ; phosphodiesters, PDE; phosphocreatine, PCr; i and adenosine triphosphate, ATP, including the signals of its analogues UTP, GTP and CTP), has been performed in most human liver studies published to date. It should be noted here that because of their extremely small T2 relaxation times large molecules, such as those present in membranes, proteins, bone and DNA, do not contribute to clinical MR spectra. The acquisition and postprocessing methods, in particular how to quantify the peak areas of overlapping metabolite signals, are critical. This leads to 2601 discrepancies between studies, even in healthy controls [82–84]. For example, in clinical practice one has to compromise between the degree of T1 saturation as a result of large pulse angles and short repetition times (TR) on the one hand, and the signal-to-noise-ratio obtained on the other [85, 86]. Even when high-field MRI equipment and/ or advanced techniques such as nuclear Overhauser effect enhancement and proton decoupling are used [87, 88], resulting in improved signal-to-noise ratios and better spectral resolution between peaks, the spatial resolution between adjacent MRS voxels should be to the order of a centimetre to allow for the detection of low-concentration metabolites. Liver MRS studies focus on the status of the liver as a whole or on the characterisation of focal hepatic lesions. In the latter category of MRS studies one obviously wants to be able to measure MRS voxels smaller than the lesions of interest. Diffuse liver disease Phosphorus MRS Several large studies feature the use of 31P MRS for studying the metabolic alterations in chronic liver disease, a disease caused by alcohol abuse, primary biliary cirrhosis, virus infections, primary sclerosing cholangitis etc. (Fig. 6). Cirrhotic patients have increased levels of PME [89, 90] and decreased PDE [91]. Correlations of the PME/PDE ratio with the results of liver biopsy [92] and with the degree of fibrosis [93] indicate that 31P MRS can be used for staging and grading of chronic liver disease. Decreases in PME/PDE were observed in response to antiviral treatment for hepatitis C- related liver disease [94]. Recently, however, it was shown in a study of 80 cases of nonalcoholic fatty liver disease that the observed elevations of PME relative to healthy controls lost their significance when the results were corrected for body weight [95]. Furthermore, patients suffering from cachexia (body weight loss) as a result of cancer elsewhere in the body had increased liver PME levels and decreased ATP compared with weight-stable cancer patients and healthy controls [96, 97]. Thus, host body weight appears to be an underestimated factor with a significant impact on 31P MR spectra of the liver, leading to the requirement of body-weight-matched control groups. 31P magnetisation transfer quantification of the rate of ATP synthesis is developing into a new method of liver characterisation [98]. Another new 31P MRS application is the diagnosis of rejection after liver transplantation. It was recently shown that chronic rejec- tion (n = 18) can be differentiated from good graft function (n= 48) by the measurement of increased PME/ATP and PDE/ATP ratios [99]. 1H, 13C and 19F MRS Significant correlations between liver glutamine/glutamate and Cho signals and the stage of