Читать книгу Green Nanomaterials - Siddharth Patwardhan - Страница 26

Targeted drug delivery uses a nanoparticle (for this book, specifically an inorganic nanoparticle) as a vehicle to transport a drug to a specific disease cell sites. Targeted drug delivery is a game-changer for global pharmacology. Most cancer drugs’ activity depends on their toxicity to cancer cells, but the obvious issue is that these are invariably toxic to healthy cells too. This means that previous therapies had terrible side effects, as the body had to cope with being poisoned overall in order to poison the cancer tumour. Targeted drug delivery means that the drugs are delivered to just the cancer cells, and thus are not administered to the whole body. This has two advantages: (1) the side effects of the treatment are dramatically reduced as the effect on healthy cells is now negligible; (2) the quantity of drug required is vastly reduced, as it is only used on the much smaller affected area. Targeting can be achieved passively or actively.

Passive targeting takes advantage of the fact that tumours have more ‘leaky’ vasculature and faulty drainage, so have enhanced permeability and retention (EPR). This means that tumours are susceptible to nanoparticles of between 12–500 nm accumulating preferentially within tumours, without any other ‘active targeting’. However, accumulation quantities can be significantly increased by active targeting also. Active targeting can take the form of specific antibodies, as discussed above, which will target just the specific cancer cells. There are also other physiochemical properties that differentiate cancer cells from healthy cells which can be taken advantage of, such as pH. Cancer cells are more acidic, and as such a pH-responsive nanoparticle coating can be designed to release the drug only at the pH of the cancer cell, targeting the drug release.

Magnetic nanoparticles can be further exploited for targeting by utilising a magnetic field gradient to move the particles remotely in vivo. Magnetite or maghemite SPIONs have been most widely used due to their biocompatibility and easy availability for iron precursors. A permanent magnetic field near the tumour can enable the field gradient to concentrate and retain the drug carrying magnetic nanoparticles at the tumour site. However, depending where the tumour is, this will be more difficult when in deep tissue or internal organs. In these cases the field gradient in an MRI instrument can be used to concentrate the magnetic particles in all three dimensions at any location in the body [22].

The most studied SPION–drug conjugate is those of the intercalating chemotherapy drug doxorubicin (DOX). Doxorubicin is a well know and characterised potent chemotherapy agent, but it also has two further advantages for drug delivery design and research. The first is that it has two functional groups that can be modified to conjugate to the drug delivery vehicle, and the second is the conjugated aromatic nature of the drug makes it fluorescent, so it can be easily tracked and monitored. Researchers have designed a range of SPION–DOX nanomedicines ranging from the simple to the complex. For example, simply trapping SPIONs and unmodified DOX in polymer micelles/vesicles, or trapping unmodified DOX within the polymer coating on a SPION. Such a polymer coating can be specifically designed to change conformation or degrade to release the DOX on under the influence of a specific stimulus, or gradually over time for controlled slow release (figure 2.9(ci)) [17].

Examples of more complex systems include attachment of the DOX to the SPION through either iron coordination chemistry [23] or by conjugation to a responsive synthetic or biological polymer such as a tumour-targeting aptamer [24]. More complex systems include multiple nanoparticles such as SPION with quantum dots, gold nanoparticles and nanoporous silica for imaging and drug entrapment respectively, forming theragnostics, described in section 2.4.1.4. The details above have provided a brief sample of the extensive research being performed, to give an idea of the developments of drug delivery using nanoparticles. This field is developing all the time, and as such the information above will date rapidly. For more in-depth reviews [11, 25] of this subject, please see recent reviews and up-to-date literature.

Magnetic hyperthermia therapies are emerging non-invasive therapies, currently undergoing clinical trials, which work by causing magnetically-induced heating when magnetic nanoparticles are subjected to an alternating magnetic field (at frequencies in the range of 100–150 kHz). This heating acts to ablate and kill the diseased tissue at a higher localised temperature (above 45 °C) while lower temperatures between 40 °C–44 °C will sensitise the cell making them more susceptible to chemotherapy drugs. Lower temperatures still (39 °C–42 °C) could also be used to trigger a heat-activated drug or release a drug conjugated through a heat-activated linker [26, 27].

The heating is induced by one of two mechanisms. The alternating field can either switch the individual spins of larger single-domain MNPs, or physically flip the whole particle for smaller superparamagnetic MNPs. This switching/flipping induces heating as thermal energy which is released as the spins/particles attempt to resist the alternating field. This heat dissipates into the surrounding tissue, damaging or killing the diseased tissue or activating smart drug treatments. The first mechanism is called ferromagnetic hyperthermia, and describes the thermal energy generated by the switching dipole moments in the crystal lattice by Néel rotations. In effect the heating can be described by the area of the magnetic hysteresis loop (see equation (2.7)) so is increased with the increased magnetic saturation and coercivity of the MNPs. The second mechanism is called magnetic fluid hyperthermia, and is described by the Brownian rotation when the aligning of the magnetic moment causes the whole particle to rotate in superparamagnetic MNPs, causing shear stress as the particle flips resulting in delivery of thermal energy to the solution, which can be described by equation (2.8).

PFM=μ0f∮HdM(2.7)

PSPM=μ0πfχ″H2(2.8)

where P is the power of the heat generated, μ₀ is the permeability of free space = 2.566 3706 × 10⁻⁶ m kg s⁻² A⁻², f is the frequency and H is the magnitude of the alternative magnetic field. The integral of H dM is the area of the hysteresis curve, while χ″ is the out-of-phase component of the magnetic susceptibility of the SPM MNP (effectively how much M lags behind H at that frequency). Within each regime the heating is largely dependent on the magnetic saturation/susceptibility and field frequency. While there is coercivity dependence in single domain ferromagnetic hyperthermia, the effect of the magnitude of the field is squared in magnetic fluid hyperthermia, so can lead to much greater heating. Furthermore, it should be noted that both mechanisms can operate in the same sample where the particles are small enough to rotate and have some coercivity, giving them both Néel and Brownian relaxation [27].

The heating power of a particle is determined by calculation of its specific absorption rate (equation (1.1)), which calculates the heating power (P) of the particles based on the mass of the MNPs in the sample (m_mnp). Each SARs value is for a specific field and frequency, so the intrinsic loss parameter (ILP) is the term normalised for field and frequency, both of which are shown in equations (2.9) and (2.10) respectively:

SAR=ΔTΔtcMFe(2.9)

ILP=SARH2f.(2.10)

As with MRI and MPI the potential for MNPs in nanomedicine is far reaching, but particle size along with field and frequency must be considered, as all these factors will affect the heating power. For example, at the same frequency of 100 kHz, 8 nm iron-oxide MNPs gave the best increase in temperature of 9.3 °C in lower fields (9.6 kA m⁻¹), whereas at higher fields (23.9 kA m⁻¹) the 8 nm particles showed an improved increase in temperature of 25 °C, but MNPs of 24 nm in diameter proved more effective with increases in temperature of 55 °C. This is probably due to the combined effect of both heating mechanisms for these larger particles [28].

There are numerous examples in the literature of MNPs used to treat cancer cells from in vitro tissue culture to mouse model studies, both as independent therapeutics and in combination with drugs that are improved when cancer cells are heat sensitized [29]. SPM nanoparticles within polymersomes have been extensively probed for this purpose and ILP in the range of values between 1.6 and 5 nH m² kg⁻¹ have been extensively shown to be functional for hyperthermic therapies [30]. Similarly, bacterial magnetosomes have shown excellent promise for magnetic hyperthermia with the highest recorded ILP of 23.4 nH m² kg⁻¹ [31, 32], and with vastly increased SARs values compared to comparatively sized synthetic magnetite nanoparticles at lower fields [33]. It is not clear why these biomineralised MNP have such enhanced properties, but is has been speculated that it could be due to the monodispersity or the slightly reduced nature of the magnetite core. Magnetosomes have thus been investigated in mouse models in vivo and it has been found that magnetic hyperthermia using chains of magnetosomes has a much greater impact for tumour shrinkage than SPION control particles (figure 2.8(Dii)) [18].

Again, the field moves fast, so recent reviews [29] and research papers in the literature are the best sources of information.

Photothermia is another type of localised hyperthermia that takes advantage of the energy a metallic nanoparticle receives from an electromagnetic wave (see section 2.3.1), where this absorbed photon energy is transferred to phonons in the lattice and dissipated as heat. A relatively new phenomenon, the potential of photothermal therapies has only been realised over the last 15 years. As such it has not yet been subjected to the same level of theoretical treatment as magnet hyperthermia. The electromagnetic wave is provided by a near-infra-red laser (typically an 808 nm laser powered at 0.3–5 W cm⁻²), which crucially is more transparent to body tissue and can penetrate deeper into the tissue than optical light. While the science in this area is still at the research stage, several inorganic nanomaterials are proving to be ideal materials for photothermal therapies. The front runner is gold nanoparticles (see recent reviews [8, 34]), as well as carbon nanotubes (CNTs) and MNPs such as SPION and larger magnetite MNPs. Once again, bacterial magnetosomes have proved to be excellent candidates with increases of temperature of up to 50 °C (under 1 W cm⁻² laser power) [31]. Meanwhile, synthetically produced magnetite MNPs coated in macrophage biological coating have proved to be excellent photothermal agents. The macrophage coating renders them invisible to the immune system, offering them longer circulation time and also targeting them to the cancer cells. Furthermore, the magnetic properties of the particles mean they can be magnetically targeted to the tumour in vivo, with photothermal therapy resulting in tumours five times smaller by weight compared to treatment with non-macrophage coated MNPs [10].