Invasomes are highly fluidic, ultra-deformable, elastic liposomal vesicles, which consist of phosphatidylcholine, ethanol, and terpenes (Dragicevic-Curic et al. 2008). Due to their fluidic membrane, liquid-state vesicles exhibit superior drug penetration enhancement ability compared to conventional, rigid gel-state liposomes (El Maghraby et al. 2001). Liposomes are of spherical shape, whereas invasomes are often of deformed spherical shape due to their highly fluidic and deformable membrane. Visualization of invasomes revealed that they appear unilamellar and bilamellar, which is independent of the loaded agent (Chen et al. 2011).

Moreover, it could be shown that the combination of ethanol and terpenes can enhance skin penetration of drugs, e.g., the highly lipophilic drug midazolam (Ota et al. 2003), as well as the water-soluble diclofenac sodium (Obata et al. 1990, 1991). Therefore, invasomes promote the penetration of drugs with unfavorable partition coefficient (logP value) and molecular weight (MW), e.g., the lipophilic drug temoporfin (MW = 681 g/mol, logP = 9.3) (Dragicevic-Curic et al. 2008, 2009; Chen et al. 2011) and the hydrophilic model agent carboxyfluorescein (MW = 378 g/mol, logP = −1.5) (Chen et al. 2011). Two mechanisms are described in the literature, which explain how the interaction of elastic vesicles with the skin appears to enhance penetration. The first mechanism proposes that vesicles or vesicle constituents disorganize and disrupt intercellular lipids, affecting the ultrastructure of the stratum corneum by the creation or modification of pathways for possible drug penetration, thus leading to increased skin permeability. However, treatment with rigid vesicles did not affect stratum corneum ultrastructure or permeability (van den Bergh et al. 1999). The second mechanism postulates that ultraflexible vesicles can penetrate the skin unfragmented on areas with low penetration resistance, e.g., between neighboring corneocyte clusters with no lateral overlapping of the corneocytes (Schätzlein and Cevc 1998; Cevc et al. 2002). Moreover, it was found that ultraflexible vesicles penetrate intact into deeper stratum corneum layers; however, only small amounts were found in the deepest stratum corneum layers. These findings could not be observed after the application of rigid gel-state vesicles (Honeywell-Nguyen et al. 2002).

Core multishell (CMS) nanotransporters are chemical chameleons, which, simular to liposomes, can encapsulate both hydrophilic and lipophilic guest molecules. Additionally, they can be dissolved in a variety of solvents, ranging from water to toluene. The unimolecular architecture of CMS nanotransporters contains a hydrophilic core, followed by a hydrophobic middle layer, and a hydrophilic outer shell. This architecture mimics the structure of liposomes, which are lipid vesicles that comprise of an aqueous inner phase, enclosed by a phospholipid membrane and surrounded by an exterior aqueous phase. As described above, the liposome architecture was found to have a penetration enhancing effect. It has been suggested that liposomes adhere to the skin surface and destabilize, fuse, or mix with the lipid matrix, resulting in a loosening of the lipid structure, which lowers the barrier strength of the skin and enhances skin penetration of loaded drugs or agents (Elsayed et al. 2007; Kirjavainen et al. 1996). Due to their liposome-like architecture, CMS nanotransporters might exhibit a similar penetration enhancing efficiency and additionally, as polymer-based nanotransporters, an improved stability.

The terminal shell provides a good solubility of the particle in water as well as in organic solvents and a high degree of biocompatibility. Furthermore, as well as the core, it serves as a matrix for encapsulation of hydrophilic agents. The nonpolar inner shell allows incorporation of hydrophobic guest molecules (Radowski et al. 2007).

CMS nanotransporters comprise of small unimers, 5 nm in size, but if the critical aggregation concentration of 0.1 g/l is reached, larger aggregates are formed with a diameter of approximately 30–50 nm. β-Carotene loaded CMS nanotransporters form aggregates up to 144 nm in size (Radowski et al. 2007).

CMS nanotransporters were shown to enhance skin penetration of the lipophilic dye Nile red (Küchler et al. 2009b) and the hydrophilic dye rhodamine B (Küchler et al. 2009a) and thus appear to be of outstanding versatility. Furthermore, CMS nanotransporters are used for tumor targeting (Quadir et al. 2008), cellular copper uptake by yeast cells (Treiber et al. 2009), or as stabilizers for catalytic platinum nanoparticles (Keilitz et al. 2010).

Nanostructured lipid carriers (NLCs) consist of a solid lipid matrix formed by, e.g., glycerol tristearate (Dynasan^® 118; Cognis, Monheim, Germany) and are loaded with liquid lipids of, e.g., caprylic/capric triglycerides (medium chain triglycerides, Miglyol^® 812, Caelo, Hilden, Germany). NLCs are dispersed in water and stabilized by an emulsifying agent. However, the structure of NLCs is described differently in the literature. One theory describes NLCs as spherical nanoparticles with droplets of liquid lipid within the solid lipid matrix (Müller et al. 2002), whereas another group describes them as lipid platelets with oil spots located on the surface of the solid lipid matrix (Jores et al. 2004). Preparation of NLCs is performed by high-pressure homogenization (HPH), applying two to three cycles at 500 bar. HPH can be performed at high temperatures, e.g., 80 °C but also cold HPH is known (Müller et al. 2000). HPH at room temperature is particularly interesting for the loading of temperature-sensitive agents to NLCs. The disadvantage of cold HPH is the increase in particle size. Mean particle size exhibits a wide range from 50 to 1000 nm. The lipid content of a formulation/nanodispersion can be from 5 to 40 % (Schäfer-Korting et al. 2007). NLCs were used for the controlled release of the highly lipophilic drug clotrimazole (Souto et al. 2004). Moreover, topical application of NLCs results in the formation of an occlusive film on the skin which improves skin hydration (Müller et al. 2002). Nevertheless, NLCs failed to enhance the uptake of Nile red (Lombardi Borgia et al. 2005) and cyproterone acetate (Stecova et al. 2007) when compared to the uptake from solid lipid nanoparticles.

EPR spectroscopy allows distinguishing between environments differing in polarity (Kempe et al. 2010). Smirnov et al. (1995) could show TEMPO partitioning between the aqueous and lipid domains of liposomes. Previously, Haag et al. (2011a) studied ultraflexible liposomes (invasomes) and nanostructured lipid carriers, which had also been prepared with TEMPO. Partitioning of TEMPO between lipid and aqueous phases could be observed and quantified. Changes in partitioning after application to the skin resulted in a sustained release of TEMPO to the skin. In Fig. 12.5, EPR spectra of TEMPO-loaded invasomes (solid, black line) recorded at different microwave frequencies are shown. Sensitivity is directly related to the square of operating frequency, which means the higher the frequency, the higher the measuring sensitivity of the spectrometer. Due to the low magnetic field strength at L-band frequency (1.3 GHz), spectra from phases originating from different polar environments appear as a single three-line spectrum, because the spectra from the different environments overlap. By increasing the frequency to X- and Q-band, spectra from the different phases are separated more efficiently. Measurements at W-band frequency (94 GHz) can completely separate the spectra from environments differing in polarity due to the high magnetic field strength, which results in two spectra. In this case one from the phosphatidylcholine/lysophosphatidylcholine phase and one from the aqueous phase of the invasomes. Since we can separate the spectra at W-band, the magnetic parameters can be fully derived from the spectra in each phase (i.e., g-value indicates the position of the spectra within the magnetic field), ¹⁴N hyperfine coupling (a_iso, distance between low field and central line, high values indicate a hydrophilic environment and vice versa), and the rotational correlation time (the time it takes for a spin probe to turn once around its own axis). This is a prerequisite for spectral simulation of spectra recorded at lower frequencies. By entering these parameters into computer software, e.g., EasySpin the spectra can be simulated at each frequency as shown in Fig. 12.5 (gray, dashed line).

The L-band low frequency measurement shows a broadened low and high-field line. The spectrum at X-band frequency shows a partitioned high-field line. When measuring at Q-band frequency, two lines of the lipophilic TEMPO EPR spectrum already appear and are shifted down-field due to higher g-values. The measurement at W-band clearly resolves three lines from each of the two phases. As demonstrated in Fig. 12.5, the resolution increases with increasing frequency. At W-band frequency the distribution between phases of different polarity can be determined, and most importantly quantified. This is especially important for the development of lipid-based carrier systems for a specific purpose, e.g., in the surfactant shell of the nanoparticle for burst release or in the core of a liquid lipid emulsion for sustained release. At W-band frequency, the distribution can be easily determined and quantified..

The use of different spectrometers operating at diverse microwave frequencies allows comprehensive studies of pharmaceutical formulations. At high frequency W-band (94 GHz) the localization of a spin probe or a spin-labeled drug within a carrier matrix becomes feasible allowing the observation of drug–carrier interactions. This can be the localization in a certain polar environment in a two-phase system like invasomes, or the localization of a spin probe in the core of a lipid nanoparticle or within the surfactant shell. Especially interesting for the investigation of topical dermatics is the low frequency L-band (1.3 GHz) EPR spectroscopy, because measurements of formulations containing spin probes applied to the skin ex vivo and in vivo are feasible. This allows the determination of penetration efficiency after skin application of spin probe loaded carriers. Moreover, it can be determined whether a carrier system releases its payload slowly in a depot manner.

Measurements to investigate the stability of delivery systems were further performed by Yucel et al. (2012), who investigated the distribution of the hydrophobic nitroxide 4-phenyl-2,2,5,5-tetramethyl-3-imidazoline-1-oxyl (PTMIO). PTMIO was homogenized with eicosane and a sodium caseinate solution. After cooling, solid lipid nanoparticles were formed, because eicosane is crystalline at 21.5 °C. Moreover, a further emulsion-based delivery system with tetradecane as the lipid phase was produced the same way. The only difference is that tetradecane forms liquid droplets at 21.5 °C. For both delivery systems, two different sizes were produced, fine droplets with a diameter of 0.2 μm and coarse droplets with 1.3 μm in diameter.. The distribution of the probe can be obtained from analysis of the spectra, and the chemical stability by measurements of reduction kinetics (e.g., vitamin C, which turns PTMIO EPR silent) as demonstrated in Fig. 12.9 where the degradation of TEMPO by the antioxidant system of the skin is shown. Analysis of the resulting EPR spectra revealed populations of the spin probe in two discrete environments (i.e., aqueous and lipid). PTMIO is largely hydrophobic with 77 and 70 % present in the coarse and the fine liquid lipid droplets of tetradecane, respectively. In the solid droplets (i.e., eicosane droplets) the probe was completely excluded from the droplets into the aqueous environment. Yucel et al. (2012) presumed that during the crystallization process the probe moves out of the solid lipid to the aqueous environment where it is more affected by the presence of the aqueous portions of the interfacial protein (sodium caseinate solution). Braem et al. (2007) showed that the EPR active cholestane is only located in the surfactant shell of solid lipid nanoparticles but not in the solid lipid core.