ABSTRACT
Liposomes incorporating magnetic species have been investigated as tissue-specific drug-delivery vehicles for chemotherapy.11-15In drug-delivery applications, magnetic liposomes loaded with pharmaceutical compounds have been utilized either by placing the affected site within an external magnetic field16,17 or through implanting a magnet within a tumor in investigational studies.14Such liposomes localize at the tumor site directed by the magnetic field and have been successful in terms of both increasing drug efficacy and reducing systemic toxicity of chemotherapeutic agents over free drug or liposomes without encapsulated magnet-ite.15 Extensive reviews of the use of magnetic species, including liposomes, in the drug-delivery realm are available.6,18-21 Other treatment-related applications for magnetic liposomes have included the use for gene delivery22,23 and suggestion as an affinity matrix with bilayer-embedded HIV receptor proteins to reduce circulating HIV-infected cells.24However, the utility of magnetic liposomes is not limited solely to therapeutic purposes. Liposomes incorporating paramagnetic species have also been investigated as reagents for magnetic resonance imaging (MRI).25-28 At a very basic level, in MRI, a static magnetic field is applied, followed by a radiofrequency (RF) pulse.29-32 The static magnetic field orients protons in aqueous components toward the field, with a slightly larger proportion of the protons oriented in the parallel, rather than antiparallel direction. While the net magnetic effect of the bulk of the antiparallel and parallel protons cancels out, there exists a magnetization remaining from the unpaired protons. The RF pulse at a specific frequency causes low-energy parallel-oriented protons to flip to a high-energy antiparallel state and reduces the overall magnetization in the direction of the applied magnetic field. During this process, protons also rotate about their axes, or precess, in the X-Y direction, yielding magnetization 90° from the original field as this rotation occurs in phase. After the RF pulse, the protons return from their high-energy orientation to their low-energy orientation parallel to the magnetic field over what is known as the T1, or longitudinal, relaxation time. After the RF pulse, protons also gradually return to an out-of-phase state, which decreases magnetization in the X-Y direction, over what is known as the T2, or transverse, relaxation time. T1 and T2 relaxation times vary between tissues and are measured with
receiver coils. The T1 relaxation time in water is relatively slow as the fast-moving water molecules do not efficiently transfer their energy back to their surroundings, whereas in tissues, transfer is more efficient, hence T1 is shorter. Such differences allow measurements of contrast between tissues. To improve contrast, paramagnetic MRI imaging agents, such as gadolinium (Gd), are employed. These agents contain unpaired electrons, which are influenced by the magnetic field and interact with excited protons. An improvement in transfer of energy to the surroundings results, effectively shortening T1 relaxation times. The term relaxivity refers to the inverse of the relaxation time as a function of contrast agent concentration. The above synopsis provides only a primer on the technology for the purpose of discussion. For a more thorough coverage of MRI imaging, readers are directed to several excellent sources.29-32
Liposomes loaded with gadolinium species offer improved MRI contrast, especially in tissues such as the liver and spleen that do not readily accumulate free imaging agents.33,34 Hydrophilic gadolinium molecules, such as gadolinium diethylenetriamine penta-acetic acid (DTPA, commercially as Magnevist), have been encapsulated within liposomal aqueous cores while hydrophobic conjugates, such as Gd-DTPA-distearyl ester, have been incorporated into their lipid bilayers.10,34-36 The signal due to contrast agent encapsulation within liposomes is affected by water permeability across the lipid bilayer, which is dependent on factors such as liposome size, lipid composition, and temperature.37-39 Increased water exchange across the bilayer yields interactions of protons with the unpaired electrons of Gd3+ and greater relaxivity. Chelate-based complexation of Gd3+ reduces its otherwise high toxicity, yet its paramagnetism, which stems from its seven unpaired electrons, remains intact. Liposome-based formulations add the benefits of a reduction in nonspecific systemic toxicity, as well as provide an option for increasing the amount of Gd through high levels of encapsulation and thus a high local concentration at the imaging site.10,34While the use of liposomes as paramagnetic medical imaging agents is well established, efforts to develop liposomes capable of dual-purpose imaging are of increasing interest.40-44 For example, the paramagnetic properties of hydrophobic Gd in the form of 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid (DOTA)
conjugated to N,N-distearylamidomethylamine (DSA) (Gd-DOTADSA) for MRI response with a lipid-based fluorescent label (DOPE-rhodamine) have extended the utility of appropriately formulated liposomes to provide both MRI and fluorescence cellular imaging (Fig. 8.1). Here, embedding the bilayer with paramagnetic species yielded a higher relaxivity than encapsulation hence was more favorable for improved MRI contrast.40,43 Specificity toward specific targets can be afforded using bilayer-tagged recognition elements, such as antibodies or peptides. In another approach, PEGylated liposomes were prepared with hydrophobic 125I and Gd-DTPADPPE to provide trimodal liposomes capable of yielding luminescent, positron emission tomography (PET), and MRI signals (Fig. 8.2).45
