A new horizon of DNP technology: application to in-vivo ¹³C magnetic resonance spectroscopy and imaging

Abstract

Dynamic nuclear polarization (DNP) is an emerging technique for increasing the sensitivity (>10,000-fold) of magnetic resonance spectroscopy and imaging (MRSI), in particularly for low-γ nuclei. DNP methodology is based on polarizing nuclear spins in an amorphous solid state at low temperature (ca. 1 K) through coupling of the nuclear spins with unpaired electron spins that are added to the sample via an organic free radical. In an amorphous solid state, the high electron spin polarization can be transferred to the nuclear spins by microwave irradiation. While this technique has been utilized in solid-state research for many years, it is only recently that dissolution methods and the required hardware have been developed to produce the high nuclear polarization provided by DNP to produce injectable hyperpolarized solutions suitable for in vivo studies. It has been applied to a number of ¹³C-labeled cell metabolites in biological systems and their real-time metabolic conversion has been imaged. This review focuses briefly on the DNP methodology and the significant molecules investigated to date in preclinical cancer models, in terms of their downstream metabolism in vivo or the biological processes that they can probe. In particular, conversion between hyperpolarized ¹³C-labeled pyruvate and lactate, catalyzed by lactate dehydrogenase, has been shown to have a number of potential applications such as diagnosis, staging tumor grade, and monitoring therapy response. Strategies for making this technique more viable to use in clinical settings have been discussed.

Introduction

A fundamental principle of magnetic resonance is the unequal population of spin eigenstates. This is typically accomplished by high magnetic fields (B₀), which increase the energy differences between the eigenstates, and inequality is through Boltzmann distributions. In hyperpolarized magnetic resonance, populations are driven further away from thermal Boltzmann equilibrium by transferring magnetization from another source. This can be achieved by three different methods (Golman et al. 2003): (1) dynamic nuclear polarization (DNP), (2) para-hydrogen induced polarization (PHIP), and (3) optical pumping methods. In this review, we highlight the advances in the dissolution DNP technique and its wide potential applications to in vitro and in vivo studies of cellular metabolism.

An essential requirement of a useful hyperpolarized agent is that it possesses a long spin–lattice relaxation time (T₁) because relaxation causes the polarization to decay back to thermal equilibrium. Recent advances of DNP technology have generated both new hyperpolarized agents as well as an automated dissolution process that prepares them for injection into living subjects. These have made feasible real-time metabolic imaging in biological systems, such as cancer.

¹H MRS has long been used in the clinic to detect metabolic changes that accompany treatment responses, but there is an inherent lack of sensitivity that limits the spatial and temporal resolution for spectroscopic imaging of naturally abundant nuclei, such as ¹H or ³¹P. Spectroscopy of nuclei, such as ¹³C, which is only 1.1 % naturally abundant, is extremely difficult, even with high isotopic enrichment, which can be quite expensive to achieve in an in vivo setting. This situation may change with dissolution DNP, which can increase sensitivity of solution ¹³C MRS by a theoretical limit of 50,000-fold (Ardenkjaer-Larsen et al. 2003). The enormous gain in sensitivity means that, following injection of a hyperpolarized ¹³C-labeled cell substrate, there is sufficient signal to image the molecule in vivo, and, more importantly, its metabolic conversion can be measured in real-time (Golman et al. 2006a, b). In addition, while in vivo ¹H MRS measures static steady-state levels of tissue metabolites, hyperpolarization of ¹³C enables dynamic imaging of cellular metabolism that provides a direct access to physio/pathological changes at cellular level in specific disease or normal tissue. A major limitation of this approach, however, is a requirement of a long T₁. For example, the T₁ of [1-¹³C] pyruvate is ∼30–40 s in vivo. The useful monitoring time is limited to 5X T₁, which includes dissolution, injection, and imaging. Hence, for pyruvate, there is a maximum of 3 min of useful measuring time. Hence, the hyperpolarized ¹³C-labeled substrate must be distributed throughout the system, be taken up rapidly by the tumor cells, and subsequently metabolized within this time frame (Albers et al. 2008; Brindle 2008; Chen et al. 2007; Gallagher et al. 2009a). Nevertheless, the technique has already shown promise for detecting treatment response in animals (Day et al. 2007; Kurhanewicz et al. 2011; Witney et al. 2009) and humans (Kurhanewicz et al. 2011).

For therapy monitoring, the exchange of hyperpolarized ¹³C label between [1-¹³C]pyruvate and lactate, in the reaction catalyzed by lactate dehydrogenase (LDH), was shown to decrease in a drug-treated murine lymphoma in vivo (Day et al. 2007). In this case, decreased label flux was attributed to a number of factors, including: a loss of LDH activity, a reduction in tumor cellularity, and depletion of the NAD(H) coenzyme pool (Day et al. 2007; Witney et al. 2009). Another study in the same tumor model using [1,4-¹³C₂]fumarate, showed that the rate of the fumarase-catalyzed conversion of fumarate to malate was a measure of subsequent drug-induced cellular necrosis (Gallagher et al. 2009b). It has been shown in human breast adenocarcinoma that a combination of hyperpolarized [1-¹³C]pyruvate and [1,4-¹³C₂]fumarate can be used to detect response to doxorubicin treatment before there is any detectable change in tumor size (Witney et al. 2010).

Background to DNP (ENDOR)

Sensitivity is a critical issue in NMR spectroscopy and imaging, the factor that often limits the success of various applications. One approach to enhancing the sensitivity in NMR experiments is to couple the nuclear spins to a reservoir with much higher polarization, such as unpaired electrons. High-frequency electron paramagnetic resonance (EPR) and electron nuclear double resonance (ENDOR) have been attractive spectroscopic techniques (Abragam and Goldman 1978). Dynamic nuclear polarization (DNP) is based on the transfer of the large electron spin polarization to nuclear spins (γ_e/γ_p > 657; γ_e and γ_p are the gyromagnetic ratios for electron and proton, respectively). This concept, originally proposed by Overhauser in 1953 (Overhauser 1953), was first experimentally demonstrated in metals and subsequently in liquids. Thus, DNP is not a new area of scientific endeavor, but rather one undergoing a transition from low to high magnetic fields and frequencies. In particular, when the paramagnetic centers are localized, the so-called solid-state effect (Jeffries 1957), cross-effect (Hwang and Hill 1967), and thermal mixing (Goldman 1970) dominate the polarization transfer, and couple the nuclear spin to one, two, or more electron spins, respectively. In the DNP process, nuclear polarization arises upon saturation of the allowed EPR transitions by a cross-relaxation in which the electron spin and nuclear spins undergo flip-flop motions.

The state of a nuclear spin with quantum number I=1/2 (such as ¹H, and ¹³C) is a superposition of the two eigenstates of the Zeeman Hamiltonian: parallel (“spin up”) or anti-parallel (“spin down”) to the external field (Fig. 1). The magnetic moment that gives rise to the NMR signal is proportional to the population difference between the two eigenstates. Denoting the number of spins in the “up” and “down” directions, N⁺ and N⁻, respectively, the polarization P is defined by:

1

where tanh refers to the hyperbolic tangent, B₀ is the magnetic field strength, γ the gyromagnetic ratio for the nucleus, T the temperature, k_B the Boltzmann constant, ħ=h/2π, and h is the Planck constant. At the hyperpolarized state of the nuclei, the spin population difference (N⁺ − N⁻) is increased by several orders of magnitude compared with the thermal Boltzmann equilibrium set by the B₀ (Fig. 1). Equation 1 shows that a very high magnetic field and extremely low temperature can collaborate to increase polarization. In the latest generation of hyperpolarizer, the attainable magnetic field and temperature are 4.64 T and 1.15 K respectively (Johannesson et al. 2009), and this has been further developed to use in the clinic (SpinLab; GE Healthcare). The thermal equilibrium polarization is very low: even at a magnetic field of 1.5 T, it is only 5 × 10⁻⁶ for ¹H, and 1 × 10⁻⁶ for ¹³C (at body temperature).

Fig. 1.

Pictorial description of the orientation of the nuclei with spin 1/2 at thermal equilibrium and in the hyperpolarized state. The magnetic field (B₀) is directed vertically upwards. (Adapted from Golman et al. 2003)

Microwave-driven DNP experiments are evolving as a broadly applicable approach to enhancing signals in solid state, solution NMR and imaging (Ardenkjaer-Larsen et al. 2003). The lower nuclear spin polarization compared to the electron spin (i.e. the ratios of their gyromagnetic ratios) ensures the nuclear spin alignment during the process. DNP can increase the polarization of NMR nuclei by 10³–10⁴, which can hence reduce the acquisition time by 10⁶–10⁸, as S/N increases by the √2 for each acquisition. Thus, studies of reaction dynamics become possible. Figure 2 shows the differences in NMR signal intensity obtained in 1 s with hyperepolarized urea, and acquired in 65 h from the same urea sample with the thermal equilibrium polarization at 9.4 T magnet (Ardenkjaer-Larsen et al. 2003).

Fig. 2.

a ¹³C spectrum of urea (natural abundance ¹³C) hyperpolarized by the DNP-NMR method. The concentration of urea was 59.6 mM, and the polarization was 20 %. This spectrum is acquired in 1 s. b Thermal equilibrium spectrum of the same sample at 9.4 T and room temperature. The signal is averaged during 65 h. (Adopted from Ardenkjær-Larsen et al. 2003)

Current status

Magnetic resonance imaging of different endogenous and exogenous molecules in tissue and living organisms holds the promise of resolving metabolic pathways and diagnosing early disease states, thus broadly affecting pharmaceutical development and molecular medicine. Hyperpolarized ¹³C magnetic resonance spectroscopy and imaging (MRSI) using the dissolution DNP method provides a >10,000-fold signal enhancement for detecting ¹³C probes of endogenous, nontoxic, nonradioactive substances to monitor metabolic fluxes through multiple key biochemical pathways including glycolysis, citric acid cycle, and fatty acid synthesis. Table 1 presents a list of ¹³C labeled compounds that have been successfully polarized and imaged in vivo.

Table 1.

¹³C labeled molecules that have been polarized, their level of polarization and in vivo applications

Molecules	Level of polarization	Probing/activity	References
[1-13C]pyruvate	25 – 30 %	Glycolysis	Day et al. 2007; Chen et al. 2007; Albers et al. 2008; Ward et al. 2010
		LDH activity
		Therapy response
		PDH activity
[1-13C]lactate	6–7 %	LDH activity	Chen et al. 2008
13C bicarbonate	12–15 %	CA activity	Gallagher et al. 2008a
		pH
[5-13C]glutamine	5 %	Glutaminase activity	Gallagher et al. 2008b
[1,4-13C2]fumarate	15–20 %	Fumarase activity	Gallagher et al. 2009a; Witney et al. 2009
		Necrosis
[1-13C]urea	35–37 %	Perfusion	Golman et al. 2003; Morze et al. 2011
[1- 13C]DHA	5–6 %	Oxidation & reduction of Vitamin C	Gallagher et al. 2009a; Keshari et al. 2011
[1-13C]ethyl pyruvate	25–28 %	Brain metabolism	Hurd et. al. 2010
		LDH
[2-13C]fructose	12 %	Hexokinase	Keshari et al. 2009
		Glycolysis

The Warburg effect is a metabolic feature of cancers that causes them to preferentially metabolize pyruvate via glycolytic pathway to lactate (Gatenby and Gillies 2004). The utility of the powerful metabolic imaging technique for cancer imaging was shown first by Golman et al. in 2006 (Golman et al. 2006a, b) and since then it has been broadly applied in a number of preclinical studies for detecting cancer presence, progression, and response to therapy (Hu et al. 2010; Ward et al. 2010). The types of cancers include lymphoma (Day et al. 2007) prostate (Hu et al. 2010), breast (Gallagher et al. 2009b), liver (Hu et al. 2011), and brain (Park et al. 2010), which have shown promise for hyperpolarized metabolic imaging.

The hyperpolarization of [1-¹³C]pyruvate has demonstrated the ability to not only detect pyruvate uptake but also the in vivo enzymatic conversion to ¹³C-lactate, ¹³C-alanine, ¹³C-carbon dioxide, and ¹³C-bicarbonate (Fig. 3). It has also been applied in a first Phase I clinical trial to probe these metabolic pathways in patients with prostate cancer (Kurhanewicz et al. 2011). Other hyperpolarized substrates, including bicarbonate, fumarate, lactate, dehydroascorbate (DHA), ¹⁵N-choline, urea, glutamine, and fructose, have also been investigated in preclinical studies to probe cancer metabolism and physiology more widely (Table 1).

Fig. 3.

¹³C-pyruvate and its metabolic products such as ¹³C-lactate, ¹³C-alanine, and ¹³C-bicarbonate

The major limitation of DNP is that the polarization decays relatively rapidly. This is determined by the spin lattice relaxation time (T₁) of the nucleus, which for ¹³C in a carboxyl group is of the order of 30–40 s. For other carbons especially those are directly bonded to ¹H due to dipolar interaction, the relaxation time may be shorter. For that reason, ²H substitution is commonly used to increase the T₁ values for directly bonded carbons. In addition, chemical shift anisotropy (CSA) is field-dependent and starts to impact T₁ relaxation at higher fields, given its square field dependence (de Graaf 2007). Nonetheless, good candidate molecules must be transported from polarizer to scanner, injected, circulated, and metabolized within a few minutes to succeed in hyperpolarized MRSI. Furthermore, the chemical shift differences between the resonances from the ¹³C-labeled substrate and its metabolites must be sufficiently large that they can be readily differentiated in vivo. The dissolution process inevitably results in a substantial dilution of the polarized substrate, and therefore a further requirement is that the candidate molecule has high solubility to allow it to be polarized at high concentration. These limitations mean that relatively few metabolites fulfill all the criteria necessary to be useful for imaging in vivo following DNP.

Methodology

Clinical magnetic resonance images the ¹H signal because of its high concentration (>100 M), and localized magnetic resonance spectroscopy of even ¹H in metabolites is challenging and slow. Localized MRS of ¹³C requires expensive isotopic labeling and, even so, is limited to only the most abundant metabolites. This signal limitation has spurred the development of hyperpolarization methods. DNP methodology in particular is very versatile, and different molecules have been polarized. The majority of in vivo work has used (1-¹³C)pyruvate due to its favorable chemical characteristics as well as its centrality to intermediate metabolism. Following injection, the hyperpolarized molecule can undergo metabolic reactions before the NMR signal returns to thermal equilibrium and becomes undetectable. This lifetime permits some important metabolic processes to be studied, but it is vastly shorter than the lifetimes associated with other molecular imaging modalities (e.g., ¹⁸F positron emission tomography, which decays in ∼2 h) (Gallagher et al. 2011).

In dissolution DNP, the sample is polarized in the solid state at very low temperatures (typically 1.2–1.4 K) and magnetic fields of 3.35–5 T, in the presence of a stable free radical, which provides the electron spin polarization that is transferred in part to the nucler spin. Following a period of polarization (typically 1 h), the sample can be rapidly dissolved and injected into an animal or cells placed in an NMR spectrometer or scanner. Signal enhancements arising from the DNP for ¹³C can be retained during the dissolution and transfer process. Applications of this method range from MR imaging of metabolites in vivo to studies of chemical reaction mechanisms in vitro. Currently, hyperpolarized MRI is investigated intensely in a number of applications in molecular and metabolic imaging.

Selection and preparation of hyperpolarized agents

The choice of a hyperpolarized substrate is based both on metabolic and MR properties. A long T₁ is required to maintain the polarization until the time of in vivo imaging. Pyruvate has chosen as one of the first hyperpolarized substrates in animal studies and also first Phase I clinical trial, and it is been the most studied compound to date. The DNP process requires the ¹³C-labeled probe to be in an amorphous (glassy) solid state with the appropriate free radical concentration. Glassing agents serve to prevent the compound from forming a highly regular crystal lattice. Crystallization would prevent radicals from accessing most of the ¹³C, whereas an amorphous glass allows access. It is important to find a good glassing agent to achieve high level of polarization.

Data acquisition

With hyperpolarized MRSI, the use of traditional pulse-sequences is precluded by the fact that the polarization is not preserved by them. In order to obtain 3D metabolic images throughout the animal, echo planar spectroscopic imaging (EPSI) techniques have been developed to detect localized differences in ¹³C-pyruvate and its metabolic products in tumors (Larson et al. 2010). This double spin echo sequence uses low tip angle to preserve the polarization as long as possible and echo planar sampling in one spatial dimension to reduce the number of RF excitations required where T₂-weighted high-resolution ¹H MR images in axial, sagittal, and coronal views are acquired to correlate with the anatomy of the animal. The idea is to acquire spectroscopic images that can be anatomically registered with conventional FSE (fast spin echo) images obtained just prior to or after the hyperpolarized acquisition (Larson et al. 2010). Additionally, hyperpolarized MR acquisition protocols include methods such as: non-localized dynamic ¹³C spectroscopy, 2D and 3D ¹³C fast MRSI with Cartesian encoding, modified to apply compressed sensing (Lustig et al. 2007), and also with multiband spectral-spatial excitation (Larson et al. 2011), both of which serve to preserve polarization by minimizing the RF pulse requirements. Non-Cartesian trajectories have also been applied to achieve rapid spectroscopic images (Mayer et al. 2009). Notably, progress has been in the development of echo planar readouts coupled with clever application of spectral-spatial pulses.

Hyperpolarized [1-¹³C]pyruvate MRSI for molecular and metabolic imaging

Most of the studies of metabolic pathways by ¹³C hyperpolarized MRI have been performed using ¹³C pyruvate with dissolution DNP. Pyruvate was chosen because it intersects with three major metabolic pathways (Fig. 3): being converted to lactate, alanine, or bicarbonate depending on the type and status of the cell. It is well known that the tumor cells metabolism responds to anoxic conditions by transforming more pyruvate into lactate rather than into alanine. The fast in vivo enzymatic conversion allows detection of ¹³C polarization in the products of its metabolic transformation. This was evidenced by imaging pyruvate, lactate, alanine, and bicarbonate in rats and pigs following injection of hyperpolarized [1-¹³C]pyruvate (Golman et al. 2006a). A quite extensive number of papers have appeared in the literature since 2006 studying cancer metabolism using DNP hyperpolarized ¹³C pyruvate MRSI. The very first study using a rat model with a P22 tumor was done by Golman et al. at Malmø, Sweden (Golman et al. 2006b). This exciting experiment showed a strong lactate signal inside the tumor within 30 s of injecting hyperpolarized [1-¹³C]pyruvate. Since then, tremendous progress has been made in pre-clinical studies for different cancer models, including TRAMP (transgenic adenocarcinoma of mouse prostate) mice (Chen et al. 2007). Figure 4 shows a ¹³C MR spectrum from a tumor voxel (top panel) and the images of lactate and pyruvate distribution (bottom panel) from TRAMP mice (Hu et al. 2011). Furthermore, the tumor histological grades have been differentiated non-invasively on the basis of lactate level (Albers et al. 2008). The hyperpolarized [1-¹³C]pyruvate has also been used for detection of tumor response to the treatment both in vitro and in mice implanted with xenografted tumors, such as murine lymphoma. The lactate/pyruvate ratio in the tumor before and after the treatment has been measured. The reduction of this ratio has been observed within 24 h of chemotherapy in lymphoma EL-4 tumors (Day et al. 2007).

Fig. 4.

Anatomical image slice of the animal and hyperpolarized ¹³C spectrum recorded from the indicated voxel on the tumor (top panel). Imaging the distribution of hyperpolarized pyruvate and lactate (bottom panel). (Adopted from Hu et al.2011)

Recently, the ¹³C MRS of hyperpolarized pyruvate has been employed to access the downstream tumor metabolites such as lactate and alanine and monitor the treatment response in human P493 lymphoma xenografts based on LDH-A inhibition (unpublished). These tumors were grown subcutaneously in SCID mice. The hyperpolarization was achieved by DNP (Hypersense; Oxford Instrument), i.e. irradiating a mixture of [1-¹³C]pyruvic acid, 15 mM trityl radicals (OX63) at 1.4 K and 3.35 T field with 94.082 GHz microwaves for an hour. Before the injection into the mouse via a jugular vein catheter, the polarized substrate was quickly dissolved in Tris/ETDA, NACl and NaOH at 37⁰C, yielding 80 mM pyruvate at natural pH. At the start of each dynamic scan, 350 μL of the hyperpolarized solution was injected over a period of 12–15 s. ¹³C-spectra were obtained using an Agilent 7 T imaging scanner utilizing a dual tuned ¹H – ¹³C volume coil. In vivo studies in tumor-bearing mice show peaks for pyruvate and its metabolic product lactate and alanine. Data are acquired right after the hyperpolarized pyruvate injection with a T_R of 1 s and flip angle 9°. Figure 5 represents the ¹³C spectrum obtained from a 7-mm-thick slice across the tumor (P493 human lymphoma xenograft) after 20 s of injecting hyperpolarized pyruvate (unpublished).

Fig. 5.

Hyperpolarized ¹³C MRS acquired after delivering (i.v.) pyruvate into a mouse. A 7-mm-thick slice was selected across the tumor (subcutaneous xenograft of human P493 lymphoma; unpublished)

In another study on pancreatic tumor xenografts (HS766T and SU8686), the ¹³C MRS of hyperpolarized pyruvate has been examined (unpublished). Figure 6a, b shows the in-vivo dynamic MR spectra acquired from a 7-mm-thick slice across the HS766T and SU8686 tumor xenografts respectively over a period of 100 s. The appearance of peaks indicate pyruvate and its metabolic conversion to lactate via LDH activity. The overall signal-to-noise ratio of the spectra and the pyruvate-to-lactate conversion profile suggests that the uptake of hyperpolarized pyruvate and conversion to lactate in a HS766T tumor is somewhat better compared to a SU8686 tumor, although all the experimental conditions were kept similar in both cases (unpublished). This observation exemplifies the hyperpolarized substrate uptake by tumors and its conversion also varies over tumor types and tumor microenvironment. These studies have demonstrated the feasibility of providing noninvasive biomarkers for profiling metabolic activity in different tumors using ¹³C hyperpolarized MRS.

Fig. 6.

Pyruvate and Lactate signals acquired after administration (i.v) of hyperpolarized ¹³C pyruvate in a HS766T and b SU8686 tumors. Pyruvate and Lactate peaks are at 171 and 183 ppm, respectively. Dynamic spectra are shown in every 2 s (unpublished)

Other ¹³C compounds

Beside pyruvate, other molecules have been investigated in probing the cancer metabolism and tumor physiology, discussed below (see Table 1).

Bicarbonate to CO₂

Hyperpolarized ¹³C bicarbonate has been used to probe the extracellular pH in tumors (Gallagher et al. 2008a). Tissue pH is an important parameter since many pathological states are associated with pH changes. For example, in tumors, extracellular pH is lower than in normal tissues and can therefore be correlated to diagnosis and therapy monitoring. Despite the importance of pH and its relationship to disease, there is currently no clinical tool available to image the spatial distribution of pH in humans. pH has been evaluated in vivo from hyperpolarized MRS measurements of the relative signal intensities of hyperpolarized bicarbonate and carbon dioxides, since the pH is determined from the Henderson–Hasselbalch equation: pH= pK_a + log([HCO₃ ^-/CO₂] (Fig. 7). Brindle et al. have shown that after injection of a solution of hyperpolarized ¹³C-bicarbonate, hyperpolarized carbon dioxide is rapidly formed by the action of carbonic anhydrase (CA) and successfully detected by MRS. In vivo measurements in murine lymphoma tumors have shown low intratumoral pH calculated from the ratio between HCO₃ ⁻ and CO₂ (Gallagher et al. 2008a). If this technique can be translated to the clinic, then it could be used as a generic marker of disease given the wide range of pathological states that are associated with an acidic extracellular environment. Bicarbonate is abundant in tissue (∼25 mM) and is already infused into patients at the concentrations that would be needed for a hyperpolarized ¹³C imaging measurement of tissue pH.

Fig. 7.

Imaging of tumor ph in-vivo (Adopted from Gallaghar et al. 2008a)

Fumarate to malate

Recently [1,4-¹³C₂]fumarate has been hyperpolarized and its metabolic conversion to [1,4-¹³C₂] malate has been detected in vivo. Drug-treated tumors have demonstrated that the malate signal correlates with cellular necrosis (Gallagher et al. 2009b). Monitoring the fumarte-to-malate conversion and their quantification have shown potential to detect treatment response in different tumor types (Gallagher et al. 2009b)

Urea

This was one of the first molecules to be polarized and a high level of polarization was achieved (Ardenkjaer-Larsen et al. 2003). Although urea will not be metabolized, recent study shows it is useful to assess perfusion which may be a biomarker for perfusion deficits and therapy response (von Morze et al. 2011).

Co-polarized substrates

The method of simultaneous polarization of multiple ¹³C-labeled compounds has been successfully developed. For instance, the co-polarization of ¹³C sodium bicarbonate and [1-¹³C] pyruvate was achieved in the same sample (Fig. 8; unpublished) and their T₁s and solution-state polarizations retain values similar to those recorded for the individual compound. For in vivo use, the co-polarized ¹³C sodium bicarbonate and [1-¹³C] pyruvate was injected into mice (TRAMP Model) and measured the tumor pH and lactate/pyruvate ratio from the single MR scan (Wilson et al. 2010). The technique was extended to polarize four ¹³C-labeled substrates namely [1-¹³C] pyruvic acid, ¹³C sodium bicarbonate, [1,4-¹³C₂] fumaric acid, and ¹³C urea to providing information on pH, metabolism, necrosis, and perfusion, respectively. These studies demonstrated the feasibility of simultaneously measuring in vivo pH and tumor metabolism using nontoxic, endogenous species, and the potential to extend the multi-polarization approach to include up to four hyperpolarized probes providing multiple metabolic and physiologic measures in a single MR acquisition (Wilson et al. 2010).

Fig. 8.

Hyperpolarized MRS of co-polarized ¹³C bicarbonate and ¹³C pyruvate simultaneously. Spectra are shown in every 4 s (unpublished)

Glutamine to glutamate

[5-¹³C] glutamine has been hyperpolarized using DNP, and polarization up to 5 % has been achieved. The real-time metabolism of hyperpolarized [5-¹³C] glutamine to [5-¹³C] glutamate by intra-mitochondrial glutaminase has been demonstrated in cultured human hepatoma cells (Gallagher et al. 2008b). Low signal intensity and considerably shorter T₁ of glutamine limit its extensive in vivo use.

Dehydroascorbate (DHA) to vitamin C(AA)

The hyperpolarization of [1-¹³C]ascorbic acid (AA) and [1-¹³C]dehydroascorbic acid (DHA), the reduced and oxidized forms of vitamin C, respectively, has been reported recently (Bohndiek et al. 2011). Redox status in murine lymphoma tumor has been evaluated from the conversion of ¹³C dehydroascorbate (DHA) to ¹³C-labeled vitamin C. Hyperpolarized [1-¹³C] DHA was found to have a T₁ of 57 s at a clinically relevant magnetic field strength (3 T), and facile chemical reduction to [1-¹³C] Vitamin C by NaBH₃CN with a 3.8 ppm downfield chemical shift. In vivo, its rapid conversion to [1-¹³C] Vitamin C in kidneys, liver, and tumor in a transgenic adenocarcinoma of the mouse prostate (TRAMP) model, as well as in a normal rat brain, has been observed (Keshari et al. 2011). These results demonstrated the utility of hyperpolarized [1-¹³C] DHA as a probe for redox chemistry in living biologic systems.

Secondary polarization through chemical reaction

Not all molecules can be polarized to a high degree or at high concentrations, but an interesting approach to overcome this problem is to transfer polarization from a molecule that polarizes to a high degree to another that polarizes less well but may be of greater biological interest. This transfer, which we term ‘secondary polarization’, can be mediated enzymatically or non-enzymatically. For example, a recent proof-of-concept paper has shown that [1,1-¹³C₂]acetic anhydride can be polarized and that the hyperpolarized ¹³C can be chemically transferred to other molecules because of the preferential reaction of acetic anhydride with amine nucleophiles; this was demonstrated with a number of hyperpolarized [1-¹³C]N-acetylated amino acids (Wilson et al. 2009). The new methods for creating secondary polarization may extend the possibility to polarize more molecules with DNP.

Other nuclei for polarization

Other nuclei that have been polarized using DNP include ¹⁵N, ⁶Li, ⁸⁹Y, and ²⁹Si. ¹⁵N-labeled choline has been polarized and it has a very long T₁ (∼4 min). This makes it very promising for imaging (Gabellieri et al. 2008). However, the small chemical shift difference between the ¹⁵N resonances of choline and phosphocholine (∼0.2 ppm) would make them difficult to differentiate in vivo at clinical magnetic field strengths. ⁶Li also has a long T₁ (∼120 s) and ⁶LiCl has been polarized using DNP (van Heeswijk et al. 2009). ⁸⁹Y is a spin half nucleus that is difficult to detect at thermal polarization levels because of its small magnetic moment. ⁸⁹Y-labeled complexes have been polarized using DNP and its pH dependent T₁ (>60 s) has been measured (Merritt et al.2007). ²⁹Si nanoparticles have been hyperpolarized using DNP and measured T₁ ≈ 200 s (Aptekar et al. 2009). None of these nuclei have so far been used significantly in in vivo applications.

Concluding remarks and future directions

There are significant advancements taking place in developing DNP methodology and hyperpolarized MRI for translating it into a clinical tool for evaluation of cancer and to monitor the therapy response. More hyperpolarized agents for clinical uses have to be evaluated. Also, different biophysical and/or biochemical pathways and associated biomolecules with practical T₁ values for real-time imaging have to be determined. The methods being developed are expected to have general applicability to a wide variety of cancers including prostate, breast, brain, and liver cancer models. Further work is needed to develop and standardize data analysis methods, design high-resolution fast imaging sequences, and to find methods that could extend the relatively short polarization lifetime (Warren et al. 2009).

It is widely expected that this new metabolic imaging technique will have a major clinical impact where it offers a new functional imaging approach to detect early tumor responses to treatment without radiation exposure and may become a complimentary clinical technique to PET imaging (Gallagher et al. 2011). Translation of DNP-MRSI into a routine clinical imaging tool may help to identify the patient groups that are responding to a certain therapy at the beginning.

The innovative hyperpolarization methodologies (dissolution DNP), aimed at overcoming the low sensitivity of the NMR/MRI signal are opening new horizons to diagnostic imaging through the real-time visualization of metabolic processes that occur at cellular level. Much interest is being currently focused on the search for more ¹³C- and ¹⁵N-labeled hyperpolarized agents in which the heteronucleus to be observed is characterized by a reasonably long T₁ and also on the advancement of hyperpolarization techniques for routinely clinical use. The high sensitivity of hyperpolarized MRS and its ability to monitor pyruvate and its metabolic product (lactate and alanine) has the potential to improve the characterization of different type of cancers and to evaluate the different therapeutic outcomes.

All these approaches are potentially applicable to a wide range of important NMR experiments in chemistry, physics, biology, and medicine, and consequently their successful development will have an enormous impact on the field. Thus, collaborative efforts among researchers from chemistry, physics, biology, and medical disciplines will be required to optimize DNP for applications in high-field NMR and MRI. Although the method used to generate the hyperpolarized molecules and their successful delivery to the subjects is challenging, the potential of this technique is so high that it deserves all the efforts that are currently devoted to make it a viable methodology in clinical settings.