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British Journal of Radiology (2003) 76, S118-S127
© 2003 British Institute of Radiology
doi: 10.1259/bjr/26631666
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Silvanus Thompson Memorial Lecture

Molecular imaging using hyperpolarized 13C

K Golman, PhD 1 L E Olsson, PhD 1 O Axelsson, PhD 1 S Månsson, PhD 2 M Karlsson, PhD 1 and J S Petersson, PhD 1

1 Amersham Health R&D AB, Medeon, SE-205 12 Malmö and 2 Department of Experimental Research, Malmö University Hospital, SE-205 02 Malmö, Sweden


   Abstract
Top
 Abstract
Introduction
 Methods and materials
 Results
 Discussion
 Conclusion
 References
 
MRI provides unsurpassed soft tissue contrast, but the inherent low sensitivity of this modality has limited the clinical use to imaging of water protons. With hyperpolarization techniques, the signal from a given number of nuclear spins can be raised more than 100 000 times. The strong signal enhancement enables imaging of nuclei other than protons, e.g. 13C and 15N, and their molecular distribution in vivo can be visualized in a clinically relevant time window. This article reviews different hyperpolarization techniques and some of the many application areas. As an example, experiments are presented where hyperpolarized 13C nuclei have been injected into rabbits, followed by rapid 13C MRI with high spatial resolution (scan time <1 s and 1.0 mm in-plane resolution). The high degree of polarization thus enabled mapping of the molecular distribution within various organs, a few seconds after injection. The hyperpolarized 13C MRI technique allows a selective identification of the molecules that give rise to the MR signal, offering direct molecular imaging.


   Introduction
Top
 Abstract
 Introduction
Methods and materials
 Results
 Discussion
 Conclusion
 References
 
Thermal equilibrium polarization and hyperpolarization
The underlying principle of MRI is based on the interaction of atomic nuclei with an external magnetic field. A fundamental property of the atomic nucleus is the nuclear spin, described by the spin quantum number I. Many atomic nuclei have a non-zero spin quantum number and can be studied with nuclear magnetic resonance (NMR), e.g. 1H, 3He, 13C, 15N and 129Xe. However, the clinical use of MRI has to date been restricted to 1H, for reasons of sensitivity. Not only does 1H have a higher sensitivity than any other nucleus in endogenous substances; it is also abundant in a very high concentration (about 80 M) in biological tissues.

Nuclei with spin quantum number (such as 1H, 3He and 13C) can orient themselves in two possible directions: parallel ("spin up") or anti-parallel ("spin down") to the external field. The net magnetization per unit volume, and thus the available NMR signal, is proportional to the population difference between the two states. Denoting the number of spins in the "up" and "down" directions N+ and N–, respectively, the polarization P is by definition given as:Go


{76S118E001}

If the two populations are equal, their magnetic moments cancel, resulting in zero macroscopic magnetization, and thus no NMR signal. However, under thermal equilibrium conditions, slightly higher energy is associated with the "down" direction, and N– will thus be slightly smaller than N+ (see Figure 1Go). For a nucleus with spin quantum number , the polarization P is given by:Go


{76S118E002}

where tanh refers to the hyperbolic tangens, B0 is the magnetic field strength, {gamma} the gyromagnetic ratio for the nucleus, T the temperature, kB the Boltzmann constant and {hslash} the Planck constant. The thermal equilibrium polarization is very low: even at a magnetic field of 1.5 T it is only 5 x 10–6 for 1H, and 1 x 10–6 for 13C (at body temperature). In other words, only about one of a million nuclei contribute to the measured NMR signal in a standard clinical MRI scanner.



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  		Figure 1. Pictorial description of the orientation of the nuclei at thermal equilibrium and in the hyperpolarized state. In the figure, the magnetic field (B0) is directed vertically upwards.

 
The polarization, and thereby the strength of the NMR signal, increases proportionally with the magnetic field [1], which has been the motivation for developing higher field MRI systems. Accordingly, 3 T instruments have been introduced for clinical whole body imaging [2]. Higher fields are technically achievable, but practical problems such as costs, radiofrequency penetration depths, and tissue contrast, increase dramatically with increasing field.

A conceptually different method to increase the polarization is to create an artificial, non-equilibrium distribution of the nuclei: the "hyperpolarized" state, where the population difference N+–N– is increased by several orders of magnitudes compared with the thermal equilibrium (Figure 1Go). The hyperpolarized state can be created in vivo by means of dynamic nuclear polarization (DNP) techniques, such as the Overhauser effect [3], in combination with a suitable contrast agent [4]. Alternatively, the hyperpolarized state of an imaging agent can be created by an external device, followed by rapid administration of the agent to the subject to be imaged. Examples of the latter approach include hyperpolarization of the noble gases 129Xe [5] and 3He [6] using optical pumping, and hyperpolarization of a wide range of organic molecules containing 13C, by either parahydrogen-induced hyperpolarization [7] or DNP hyperpolarization [8].

Imaging of hyperpolarized agents
The concentration of a hyperpolarized imaging agent may be 0.5 M in the injection syringe and decrease to 1–20 mM in vivo, due to dilution in the vascular system. This is far below the typical 1H concentration of 80 M, but since the hyperpolarization can enhance the signal up to 106 times, MRI can be extended to nuclei other than 1H, thereby permitting the visualization of changes in the molecular structure in a reasonable time frame, e.g. caused by metabolic processes, which were previously inaccessible.

The ability of conventional 1H MRI to differentiate between various soft tissues and detect pathology is based mainly on the inherently different relaxation times (T1, T2 and T2*) of different tissues. Even so, the achievable dynamic range is below 10 [9]. With the administration of contrast agents containing paramagnetic atoms (e.g. Gd3+, Mn2+), the relaxation rates (1/T1, 1/T2) will increase proportionally with the concentration of the agent. Depending on the imaging sequence used, the reduced relaxation time can result in either an increased or a decreased signal where the agent accumulates, thereby increasing the image contrast [10]. The mechanism is fundamentally different for hyperpolarized agents: the hyperpolarized nuclei generate the signal themselves rather than moderating the signal from adjacent protons. Consequently, hyperpolarized MRI has the advantage of completely lacking background signal, either because the nuclei are not naturally present in the body (noble gases) or because the natural abundance signal is negligible (13C). In this respect, hyperpolarized MRI behaves similarly to the modalities positron emission tomography (PET) and single photon emission computed tomography (SPECT), where the scanner detects the radiation from an injected contrast agent containing gamma-emitting nuclei, and where the signal strength is directly proportional to the concentration.

Hyperpolarization techniques
For practical purposes, four different methods exist to create a hyperpolarized state:

The "brute force" approach
From Equation (2)Go, it follows that the thermal equilibrium polarization increases with increasing magnetic field strength and decreasing temperature. A straightforward, "brute force" approach to increase the polarization in a sample consists of subjecting it to a very strong magnetic field at a temperature close to absolute zero. The polarization, which is in the parts per million (ppm) range at 1.5 T and body temperature, can, for example, be increased by a factor of 1000 by cooling down the sample to liquid helium temperature (4 K) at a field strength of 20 T. If the sample is brought to 1.5 T and 310 K rapidly (i.e. without losses of polarization), it is thus hyperpolarized at body temperature. To obtain polarization levels where the hyperpolarized signal exceeds the 1H signal of conventional MRI, the "brute force" method would require impractically low temperatures (in the mK range). Large-scale production of hyperpolarized noble gases (3He and 129Xe) has been proposed using this approach [11], but due to the great technical challenges and costs associated with these extremely low temperatures, the method has not yet been used for in vivo applications.

Dynamic nuclear polarization (DNP)
As seen in the previous section, low temperature and high magnetic field increases the polarization. Under moderate conditions, e.g. 1 K and 3 T, the nuclear polarization is still insufficiently low for 13C MRI (polarization <0.1%), but electrons are highly polarized (>90%) due to the much larger gyromagnetic ratio of the electron (c.f. Equation (2)). Using the DNP technique, the high polarization of the electron spins can be transferred to coupled nuclear spins [12]. In the method described by Ardenkjaer-Larsen et al [8], the material containing the nuclei to be hyperpolarized is doped with a single-electron substance and exposed to a magnetic field of ~3 T and a temperature of ~1 K (Figure 2Go). Microwave irradiation near the electron resonance frequency transfers the polarization from the unpaired electrons to the 13C nuclei, whereby the nuclear polarization in the solid material can be increased to 20–40%. By rapid melting and dissolving, the solid can be transformed into an injectable liquid, with small to negligible polarization losses.



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  		Figure 2. During the dynamic nuclear polarization process, polarization is transferred from the electrons of the doping material to the 13C nuclei by means of microwave irradiation.

 
Parahydrogen-induced polarization
In the parahydrogen-induced polarization (PHIP) method, nuclear polarization is increased via a chemical reaction involving parahydrogen; a state where the hydrogen nuclei are oriented such that their magnetic moments cancel (Figure 3Go). Bowers and Weitekamp initially predicted and verified that the PHIP effect arises in molecules catalytically hydrogenated with parahydrogen [13, 14]. A prerequisite is that the hydrogenation mechanism operates by transfer of the hydrogen molecule as a unit onto the substrate (Figure 4aGo). By means of a diabatic field cycling scheme, the non-equilibrium spin order of the parahydrogen molecule is then converted to nuclear polarization of the 13C nucleus in the substrate (Figure 4bGo) [15].



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  		Figure 3. The four possible orientations of the nuclear spin in the hydrogen molecule.

 


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  		Figure 4. (a) A substrate molecule containing 13C is hydrogenated with parahydrogen. (b) The spin order of the parahydrogen molecule is converted to nuclear polarization of the 13C nucleus, via a diabatic field cycling scheme.

 
Optical pumping methods
In 1960, Bouchiat et al [16] showed that angular momentum could be transferred from the electron spins of optically pumped Rb atoms [17] to the nuclear spins of 3He by spin-exchange collisions. The method could be extended to efficiently polarize 129Xe as well [18]. In a magnetic field, the electronic transition S1/2–P1/2 of the Rb atoms can be driven by circularly polarized laser light (795 nm), to selectively pump the ground-state Rb electrons entirely to the (or ) state. The electronic polarization of the optically pumped Rb atoms is transferred to the nuclei of the noble gas atoms via formation of loosely bound van der Waals molecules or via binary collisions. The process creates a non-equilibrium polarization of the noble gas nuclei.

Hyperpolarization of 3He can also be achieved by the method of metastability exchange, first reported in 1958 [19]. In a low-pressure 3He gas (about 1–2 mbar), 3S1 metastable atoms are formed by a weak electrical discharge. By using circularly polarized light, transitions 3S1->3P0 (1083.0 nm) are induced. Due to strong hyperfine coupling, the nuclei of metastable atoms become polarized. When a polarized metastable atom collides with an unpolarized ground-state atom, a high probability for exchange of metastability exists: the metastable and the ground-state atom exchange their electron configurations, while the nuclear polarization remains unaffected. Thus, the collision yields a polarized ground-state atom and a non-polarized metastable atom. The latter can once more undergo the optical pumping process [20–22]. A classic review of optical pumping methods has been given by Happer [23].

Although the theory of hyperpolarizing noble gases by optical pumping was known in the early 1960s, large-scale production has only recently been possible, owing to the development of high-power lasers [24, 25]. An overview of the methods for hyperpolarization of noble gases was recently published by Goodson [26].

Medical applications of hyperpolarized nuclei: a brief overview
Lung imaging
With the advent of the hyperpolarized noble gases 3He and 129Xe, a natural tool was provided for imaging of the lung, which is a difficult area for 1H MRI because of the low density of protons in the lung parenchyma and the strong susceptibility gradients at the gas–tissue interface [27, 28]. In 1994, the first MR images using hyperpolarized gas were demonstrated, showing excised mouse lungs filled with 129Xe [5]. The first 3He images depicting the lungs of a dead guinea pig were presented in 1995 [6]. These initial works were followed by the first human images using 3He [29–31] and 129Xe [32, 33]. For a review of the historical background of hyperpolarized lung imaging see Albert and Balamore [9].

The diagnostic potential of hyperpolarized gas imaging was first demonstrated in studies revealing various ventilation defects in patients, where non-ventilated regions were depicted as signal voids [34]. Rapid dynamic imaging has also demonstrated the potential to detect abnormal breathing patterns caused by lung disease [35–37] and to quantify ventilation [38]. An example of a three-dimensional (3D) ventilation quantification is shown in Figure 5Go. Owing to the large diffusion coefficient of gases (especially 3He), the image intensity will decrease in regions with elevated mobility of the gas [39, 40]. By measuring the apparent diffusion coefficient (ADC), it is thus possible to gain information of pathological lung structure, e.g. in emphysematous lungs [41]. From measurements of the T1 relaxation time, it has further been possible to calculate the regional oxygen partial pressure pO2 in the lungs based on the depolarizing effect of O2 [42–44]. The ventilation–perfusion ratio () is highly relevant for the diagnosis of abnormal lung function [45]. By combining 3He ventilation imaging with 1H perfusion imaging, initial attempts have been made to assess the parameter in animals [46–48] and in humans [49, 50].



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  		Figure 5. (a) Three-dimensional 3He image of a rat lung (central slice and surface rendering). (b) The corresponding quantitative ventilation map calculated according to the method described by Deninger et al [38].

 
Multichannel receiver systems on modern MRI instruments allow detection of several nuclei simultaneously. Simultaneous acquisition of 3He and 1H images is therefore possible, and — even more fascinating — 13C images. By injecting a hyperpolarized 13C-labelled substance intravenously, we have obtained images of the pulmonary arterial circulation (Figure 6Go). Images of hyperpolarized 3He and 13C can be acquired in <1 s, and it may be possible to study regional ventilation and perfusion in "real-time" if a 13C substance is injected simultaneously with administration of 3He. The location and functional importance of pulmonary emboli may thus be quantified.



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  		Figure 6. A series of trueFISP 13C images, showing the lungs of a pig after injection of a hyperpolarized 13C imaging agent. The imaging sequence was repeated with 1 s intervals.

 
Organs other than the lungs
Owing to its solubility in blood and tissues, inhaled 129Xe is distributed throughout the body, and can thus be used for applications other than lung imaging [32]. The resonance frequency of 129Xe dissolved in liquids is ~200 ppm higher than in the gas phase [51], and varies over a ~20 ppm range in tissues in vivo [52, 53]. Because the resonance frequency of 129Xe is sensitive to its local environment, numerous experiments involving spectroscopy or chemical shift imaging (CSI) have been presented, e.g. CSI of the chest and the brain [54, 55], spectroscopy of tumours [56] and probing of the pO2 in blood [57]. Using the large frequency shift between the gas phase and the dissolved phase 129Xe, methods have been proposed to monitor the dynamics of 129Xe when transported from the alveoli to the pulmonary blood [58–60], from which information about the diffusing capacity of the lung can be obtained.

The spectral information obtained from CSI is a strength of MRI compared with other modalities, since it informs about the molecular structure and the environment of the molecule. The chemical shift of 129Xe is based only on the different environments, whereas for 13C-labelled substances, it is mainly the molecular structure that determines the chemical shift. CSI has been used to image the localization of metabolites within the brain of 13C labelled substances (e.g. glucose, alanine etc.) [61]. Without hyperpolarization, very long scan times have been needed to generate such images. To image the metabolic processes in a clinically relevant time frame (seconds), hyperpolarization is needed.

Vascular imaging with hyperpolarized 13C
Hyperpolarized 13C has only recently been available at polarization levels sufficient for MRI [8, 15, 62, 63]. At present, a polarization of 10–30% can be obtained, and the long relaxation times (T1 and T2 up to ~60 s and 5 s, respectively) makes "real-time" vascular imaging with 13C molecules a new tool to examine pathological conditions. The feasibility of hyperpolarized 13C for MR angiography has recently been investigated [64].

Molecular imaging applications
With the possibility to polarize 13C-labelled molecules to >20%, MRI may emerge beyond anatomical (e.g. angiography) and functional (e.g. perfusion and diffusion) visualization. Since hyperpolarized 13C MRI directly informs about the molecules, to which the hyperpolarized atoms are attached, investigation of tissue and cell viability (direct molecular imaging) may be feasible.

In the following, we present a study in which the PHIP method was employed to polarize a 13C-labelled, water-soluble molecule to ~30%, followed by in vivo imaging of the distribution of the 13C molecule, after intravenous injection in rabbits.


   Methods and materials
Top
 Abstract
 Introduction
 Methods and materials
Results
 Discussion
 Conclusion
 References
 
The molecular imaging agent
The 13C-labelled molecule (2-hydroxyethylacrylate, molecular weight ~120 g mol–1) was polarized using an automated version of the parahydrogen process described elsewhere [15]. A polarization level of ~30% was achieved and a volume of 2.5 ml 0.5 M solution was produced for each injection. The polarization level had been determined from a series of calibration experiments on a 7 T spectrometer (Varian, Palo Alto, CA), by comparing the integral of a 13C spectrum acquired from a hyperpolarized sample, and a second spectrum acquired from the same sample after the polarization had decayed to thermal equilibrium.

Animals
Four rabbits (male, Swedish loop, 2.1–4.0 kg) were anaesthetized intramuscularly with butorphanol (Torbugesic®; Fort Dodge Animal Health, Fort Dodge, USA), xylazine (Rompun® vet.; Bayer AG, Leverkusen, Germany) and ketamine (Ketalar®; Pfizer Inc., New York, NY). The imaging agent was administrated through a venflon catheter placed in an ear vein. During the imaging procedure, anaesthesia was titrated as necessary. All animals were sacrificed after completion of imaging. The study was approved by the local ethics committee (Malmö/Lunds djurförsöksetiska nämnd; appl. no. M92–01).

MRI equipment
All MRI measurements were performed in a 1.5 T whole-body scanner (Magnetom Sonata; Siemens Medical Solutions, Erlangen, Germany). The scanner needed no extra modification, except for the coil for 13C imaging. A custom built, dual tuned 1H/13C transmit/receive coil (Rapid Biomedical GmbH., Würzburg, Germany) was used. The 13C part of the coil was built as a birdcage (Ø=170 mm, length=170 mm), and consisted of two elements in quadrature. Tuning and matching of the coil was performed individually for each rabbit.

MRI examinations
Three out of the four rabbits (in the following referred to as Rabbit 1–3) were used for 13C imaging to evaluate the hyperpolarized imaging agent. The fourth rabbit (referred to as Rabbit 4) was used to acquire reference proton images, and was injected with the conventional Gd-based contrast agent OMNISCANTM (Amersham Health, Oslo, Norway). All animals were positioned supine and proton localizer images were acquired in order to fit the heart/lung and the kidneys in the expected field of view (FOV). During the 13C imaging procedure (Rabbit 1–3), the scanner frequency was set to the 13C resonance using a measured relationship between the 1H resonance frequency and the 13C frequency of the imaging agent.

All imaging was performed using a trueFISP [65] sequence with a high flip angle ({alpha}=160°), which has been found superior for MRI of hyperpolarized 13C substances [64]. The trueFISP sequence was preceded by a {alpha}/2 preparation pulse, and a centric encoding scheme was used for the phase encoding gradient. The sequence is shown schematically in Figure 7Go. The imaging volume consisted of coronal slices with a thickness larger than the animal (projection imaging). The in-plane spatial resolution was increased in Rabbit 3 (1.0 x 1.0 mm2 pixel size, compared with 2.5 x 2.5 mm2 in Rabbits 1–2). The echo time (TE) and repetition time (TR) were the shortest possible for the selected in-plane spatial resolution (TR/TE=4.9/2.5 ms for Rabbits 1–2, TR/TE=8.9/4.5 ms for Rabbit 3). Due to the low gyromagnetic ratio of the 13C nucleus, the gradient performance was the limiting factor for achieving short TR. In all 13C experiments, the volume and concentration of the injected molecular imaging agent was 2.5 ml and 0.5 M, respectively. The scanning was started 2 s, 4 s and 6 s after the end of the injection.



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  		Figure 7. Schematic drawing of the trueFISP sequence.

 
The Gd contrast agent injected in Rabbit 4 was diluted 1:5 to obtain an injection volume equal to the 13C injections in Rabbits 1–3. A dynamic 3D 1H study with a temporal resolution of 0.86 s was performed. The spatial resolution was 1.5 x 1.5 mm2. The image information in the partitions covering the animal was added to a single slab per time point, resulting in one projection image for each time point.

To quantitatively evaluate the images, a region of interest (ROI) was located over specific regions or organs. Signal-to-noise ratios (SNRs) were calculated from the mean signal in a ROI divided by the mean noise value in a ROI outside the animal and free from imaging artefacts.


   Results
Top
 Abstract
 Introduction
 Methods and materials
 Results
Discussion
 Conclusion
 References
 
During the experiments, a new syringe with the imaging agent could be produced with 2 min intervals. With the polarization level of ~30%, it was possible to perform 13C imaging of live rabbits with a spatial resolution of 1 x 1 mm2. The signal strength from the molecular imaging agent was sufficient to trace the distribution of the agent into several organs.

The distribution of the 13C imaging agent at different time points is depicted in Figure 8Go. Two seconds after the injection, the imaging agent is mainly located in the heart and the lungs, but there is also signal from the kidneys (Figure 8aGo). There is a considerable amount of image artefacts, especially ringing artefacts emanating from the aorta and the heart. 4 s after the injection, the signal in the lungs and the aorta has decreased (Figure 8bGo). The signal from the kidneys has increased, and structures within the kidney parenchyma are distinguishable. A definition of the stomach wall is now clearly visible, contrary to the Gd-enhanced 1H image, which did not show this anatomic structure. 6 s after the injection, the intestines are visible, in addition to the stomach walls (Figure 8cGo). This again differs from what could be seen with the Gd-based contrast agent.



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  		Figure 8. Images depicting the distribution of the injected hyperpolarized 13C imaging agent at different times after the injection. The delay between injection and imaging is indicated at the top of each image.

 

   Discussion
Top
 Abstract
 Introduction
 Methods and materials
 Results
 Discussion
Conclusion
 References
 
In this study, the first MR images of live animals injected with a 13C substance, hyperpolarized with the PHIP method, are presented. With the present polarization levels (~30%), in vivo images with SNR between 20 and 60 in various organs were achieved after injection of 2.5 ml 0.5 M molecular imaging agent in rabbits. Images could be acquired with a spatial resolution of 1.0 x 1.0 mm2. Initially, the distribution of the 13C substance was similar to that of the conventional Gd-based contrast agent, a finding supported by a quantitative ROI evaluation of the SNR in the heart, lung, kidney and aorta, but already 4 s after the injection other structures, such as the stomach wall, was seen in the 13C images. This could not be observed in the corresponding Gd-enhanced 1H images.

During the first few seconds after the injection, most of the 13C imaging agent is still present in the blood and the highest vascularized organs, i.e. the heart, the lungs and the kidneys. The 13C imaging agent is a small molecule (molecular weight ~120 g mol–1) and leaks out rapidly into the extravascular space; already after a few seconds, there is a significant uptake in soft tissues.

It can be observed that the 13C images have lower signal than expected in the aorta and other large vessels. The low signal in these vessels may be explained by flow losses, since the echo times were rather long, despite the shortest possible settings being used. Because the gyromagnetic ratio for 13C is four times lower than for 1H, longer echo times are compelled since the gradient strength is finite (40 mT m–1). This explanation is further supported by the fact that large vessels were clearly depicted when a hyperpolarized 13C substance was imaged with echo times <2 ms on a different type of MR scanner with 200 mT m–1 gradients [64].

Since the natural abundance of 13C is far below the detection limit of a MRI scanner, the background signal in 13C MRI is reduced to the noise produced by the imaged object and the detection system. This suggests that the contrast-to-noise ratio will be high and, at least initially, equals the SNR. Accordingly, thick imaging slices, or even projection techniques, may be used to visualize the distribution of the molecular imaging agent.

Hyperpolarized MRI differs from conventional MRI, in the sense that any magnetization used up by the imaging process cannot be recovered. The obvious drawback is that the imaging protocol gets restricted [66, 67], and the time window for imaging is limited. Typically, several images with high temporal resolution are needed to study the distribution of a contrast agent. In this study, only one image was acquired for each injection. A technique to preserve the magnetization between subsequent images has been presented [64], but was not available on the scanner used in this study.

The introduction of an injectable, hyperpolarized 13C substance opens a new field of MRI. With the coil and the receiver system of the scanner tuned to the resonance frequency of the hyperpolarized nucleus, only signals from the injected substance will be detected. The signal strength is a linear function of the concentration and the polarization level of the nucleus in question. This is not the case for a conventional contrast agent (e.g. Gd-chelates and other paramagnetic molecules/particles) that operates by altering the relaxation times of water protons in surrounding tissues [68–70]. When irradiated with a radiofrequency wave, the injected 13C nuclei emits a radiofrequency signal, contrary to the tracer substance used in PET or SPECT imaging, which emits {gamma}-rays. More information can be obtained from an NMR-active nucleus, since its resonance frequency is a function of its chemical and physiological environment (e.g. chemical shielding, viscosity and mobility). Thereby, it is possible to separate the signals from 13C nuclei within different molecules. This feature is exploited in the field of analytical NMR spectroscopy and in vivo MR spectroscopy. While PET and SPECT are only capable of mapping the distribution of the nuclei, regardless if they are still contained within the injected molecules or not, NMR is capable of distinguishing signals from the tracer nuclei (e.g. 13C) present in different molecules. This specificity on a molecular level is the basis for the clinical use of MR spectroscopy [71–73]. Due to SNR limitations, this has been restricted to protons, 19F and 31P, and to the use of large image voxels (~1 cm3). The hyperpolarization procedure used in the present work overcomes these restrictions, but still retains the specificity of the NMR technique, and thus makes it possible to perform direct molecular imaging. Consequently, distribution patterns may be mapped by injection and imaging of several hyperpolarized 13C molecules simultaneously, delivering valuable information about membrane structure and permeability.


   Conclusion
Top
 Abstract
 Introduction
 Methods and materials
 Results
 Discussion
 Conclusion
References
 
We have demonstrated that it is feasible to image hyperpolarized 13C-labelled molecules with MRI. The technique is in its infancy, and improvements with respect to polarization and imaging techniques seem possible. MRI as an image modality offers, in connection with a 13C-labelled image agent the possibility of obtaining information about molecular behaviour in vivo. The choice of substance to be hyperpolarized will influence the medical information obtained.


   Acknowledgments
 
The authors wish to express their gratitude to Ms B-M Lilja and Mr C Johansson for their excellent technical assistance.


   References
Top
 Abstract
 Introduction
 Methods and materials
 Results
 Discussion
 Conclusion
 References
 

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