**3. Advantages of nanoparticles for drug delivery**

Nanoparticles bring a new level of engineering and control to the field of medicine by being able to modify parameters such as solubility, diffusivity, half-life, toxicity, pharmacokinetics and biodistribution of drugs and diagnostic agents. The applications of nanoparticles are very diverse and are expected to increase with the advancement of technology. In recent years, numerous studies have demonstrated their ability to act as sensors [52], drug carriers [53–55], and diagnostic agents [1, 56]. Recent efforts have managed to integrate treatments and diagnoses in a single application, giving rise to the procedures known as "theranostic".

The justification for the use of nanoparticles as drug delivery systems lies in at least three mechanisms: (i) Enhanced Penetration and Retention (EPR) of nanoparticles in solid tumors; (ii) The possibility of transporting insoluble drugs in the blood through stable colloidal systems and (iii) the controlled release thereof. In this section, the possible advantages of each of these points will be developed.

## **3.1 Enhanced permeability and retention (EPR)**

The term EPR was coined by Matsumura and Maeda in 1986 [13]. In their work, the researchers observed that the anticancer protein neocarzinostatin, conjugated to a polymeric matrix, exhibited greater accumulation in tumor tissues than free neocarzinostatin. By applying labeled macromolecules to tumor-bearing mice, they observed that their concentration was up to 5 times higher in tumor areas than in blood over a period of 19 to 72 hours [13]. The authors affirm that the passive accumulation of these macromolecules in tumors is due to the abnormal physiology associated with tumor masses: fenestrated hypervascularization with increased permeability to macromolecules (or nanoparticles) and poor recovery through blood vessels or lymphatic vessels [57]. Subsequently, it was shown that other plasma proteins greater than 40 kDa are capable of passively and selectively accumulating in tumor areas [58]. The EPR effect can be demonstrated in mice with the intravenous injection of the Evans Blue marker, which binds to plasma albumin forming a complex that demonstrates differential accumulation in tumor areas [59], as shown in **Figure 1**.

#### **Figure 1.**

*Image of a metastatic lung cancer originating from 26 colon tumors implanted in the dorsal skin of a mouse. The mouse was sacrificed 3 months after implantation and 10 hours before sacrificing, a solution of Evans blue (5%) was injected intravenously to allow the EPR effect to become visible. Albumin-Evans blue complex (70 kDa) preferentially accumulated in metastatic tumor nodules, as in primary tumors. Arrows point to metastatic tumor nodules. From reference [60] with permission of Elsevier.*

#### *Nanoparticles as Drug Delivery Systems DOI: http://dx.doi.org/10.5772/intechopen.100253*

For passive accumulation through the EPR effect to be important, different requirements are needed. On the one hand, the nanoparticles must remain in circulation for a time greater than 6 hours [60, 61]. This can generally be achieved by functionalizing the nanoparticles with polyethylene glycol (PEG) [62]. On the other hand, the mechanism also depends on the particles being small enough to penetrate biological membranes but large enough to be retained. Yuan et al. [63] measured the microvascular permeability of several macromolecules in human colon adenocarcinoma LSI74T transplanted in mice with immunodeficiency and the results indicated that the cut-off size of the pore is around 400-600 nm, depending on physicochemical properties such as charge and hydrophobicity of the nanoparticles. Regarding the minimum size, Maeda et al. [58] estimated that the nanoparticle size must be greater than 40 kDa to show significant retention in the tumor area.

Vascular extravasation is also highly dependent on the morphology and the specific type of tumor. Scanning electron micrographs of normal vascular epithelium and two epithelia associated with different tumors are shown in **Figure 2**. As can be seen, tumor-associated epithelia have significant pores (fenestrations) and their size depends on the type of tumor. Smith et al. [64] studied the extravasation capacity of quantum-dots (20-25 nm) and single-walled carbon nanotubes (2-3 x 200 nm) in tumors implanted in mouse ears. The surface of both types of nanostructures was modified by PEG to avoid differences in charge or surface chemistry and that the results were only due to the morphology of the particles. The authors found that spherical quantum-dots are capable of extravasation of the endothelium of LS174T tumors, whereas cylindrical nanotubes are capable of extravasation in U87MG tumors. Surprisingly, the authors were not able to see the extravasation of the nanomaterials in normal endothelium. This suggests that the morphology of the nanoparticles may be a determining factor for penetrating certain tumors, while healthy endothelium could prevent nanoparticle transfer.

Although the EPR model has been tested in rodents with large induced tumor masses [59, 65], these models differ widely in morphology and physiology of possible human tumors and, for these reasons, there is still much controversy regarding it [55, 60, 66]. Firstly, tumors of up to 10% of body weight have been reported in mice. If we make an analogy with a 70 kg human, the tumor would be the size of a basketball [67], when they actually have a size between millimeters and centimeters at the time of diagnosis and treatment [68]. Such tumor masses filter out a significant proportion of the injected drug dose and act as a reservoir, enhancing efficacy while mitigating toxicity. In addition to this, the

#### **Figure 2.**

*Scanning emission microscopy (SEM) of tumors and normal blood vessels. SEM images show pores in the U87MG and LS174T tumor vasculature at the apparent border between endothelial cells. No pores are seen at the border of the vasculature of a tumor-free mouse ear. Scale bars: U87MG (500 nm), Normal (1 μm) and LS174T (1 μm). Reprinted with permission from reference [64]. Copyright 2012 American Chemical Society.*

tumor microenvironment in humans presents important physiological differences compared to murine tumors: (i) lack of fenestrations in the tumor endothelium for the entry of nanoparticles, (ii) heterogeneity of blood flow through tissues, which causes the regions to become acidic or hypoxic [69], (iii) lower pericyte coverage, (iv) heterogeneous basement membrane and (v) higher and heterogeneous density of the extracellular matrix. This leads to high interstitial pressure and therefore the main mechanism of matter transport is by diffusion and not by convective transport, which is more efficient [69, 70].

For these reasons, it is not possible to directly transfer the results obtained in rodents to humans, mainly because cell penetration depends on the nanoparticles go from the point of application to the tumor mass and be able to interact with cells to be internalized. Currently, different methods are being investigated to increase the EPR effect. For example, Fang et al. [71] developed agents which can selectively generate vasodilator molecules (carbon monoxide) in tumor areas, achieving an increase in the concentration of the nanocarrier between 2 and 3 times higher in these, while an increase in tissues healthy was not detected. Similar results have been achieved with nanocarriers that can release nitric oxide [72, 73]. The increase in blood pressure results in an increase in the osmotic pressure, which promotes the filtration of the particles towards the tumor areas so that when angiotensin II is coadministered with the nanocarriers, an increase in the transfer and accumulation in the tumor areas can be observed [74].

In contrast to the passive accumulation of drug nanocarriers in tumor areas by the EPR mechanism, active targeting is presented, which is based on the functionalization of the nanoparticle surface with recognition molecules such as antibodies [75, 76] or ligands [77, 78] which can specifically bind to molecules overexpressed at the target site [79]. In the active targeting strategy, two cellular targets can be distinguished: (i) targeting cancer cells, which present overexpression of molecules such as transferrin, folate, epidermal growth factor receptor or glycoprotein receptors, and (ii) targeting tumor endothelium, which have overexpression of vascular endothelial growth factors (VEGF), αvβ3 integrins, vascular cell adhesion molecule-1 (VCAM-1), or matrix metalloproteinases [66, 80]. In some cases, both receptors are overexpressed in cancer cells and endothelium and can be exploited simultaneously [80]. In addition, the design of nanocarriers as active targeting systems may involve the coupling of recognition molecules as surface receptors which are able to initiate endocytosis, and hence to increase cell internalization in contrast to simple accumulation [81]. Not only would this increase the antitumor efficacy of many drugs, but it could also be used for the delivery of genetic material [82].

#### **3.2 Insoluble drug transport**

Most orally administered drugs that are soluble in water and capable of penetrating biological membranes during the passage of the gastrointestinal tract will eventually become bioavailable in the body. In contrast, water-insoluble drugs will generally not be bioavailable after oral ingestion as they cannot dissolve and pass through the gastrointestinal barrier. Along the same lines, due to their low solubility, they cannot be administered intravenously and parenteral administration does not always increase bioavailability [83]. It is estimated that 90% of drugs in development are insoluble in water, while only 40% of drugs on the market share this characteristic [84]. These statistics could indicate that many drugs in development do not reach their administration to patients due to their low solubility in water. This not only means less capital invested in research and development but also lost treatment opportunities. The development of a drug in 2011 was estimated at between 92 million and 1.8 billion dollars [85], lasting for a period of between 11.4

#### *Nanoparticles as Drug Delivery Systems DOI: http://dx.doi.org/10.5772/intechopen.100253*

and 13.5 years on average [86]. Considering these, we can see that low water solubility represents a formidable challenge and opportunity for nanotechnology.

Three factors govern the speed and degree of absorption of orally administered drugs: (i) dissolution rate, (ii) solubility and (iii) intestinal permeability, which are grouped according to the biopharmaceutical classification system (BCS, Biopharmaceutical Classification System) in the categories [87]:

Class I: High Solubility - High Permeability. Class II: Low Solubility - High Permeability. Class III: High Solubility - Low Permeability. Class IV: Low Solubility - Low Permeability.

The criterion established by the BCS classifies a drug as soluble when it is capable of dissolving an entire therapeutic dose in 250 mL of water, being this volume equivalent to the average amount of water found in the stomach [87].

As can be deduced, the possibilities of entering the market for a class I drug are substantially greater than that of the rest of the categories, however, a possible solution to these problems lies in the development of drug carriers which can transport them in a stable colloidal dispersion and with particles capable of crossing biological membranes [88]. As an example, Atovaquone (Wellvone®) is an antibiotic used for the treatment of *Pneumocystis carinii*, leishmaniasis and *P. falciparum* malaria, however, its low solubility limits its absorption. By formulating a dispersion of nanoparticles of this drug, it was possible to increase absorption from 15 to 40% with a 3-fold lower drug dose [89]. Xie et al. [90] prepared curcumin-loaded silk fibroin nanoparticles (SFN) to increase the dissolution rate of the drug and the mass of the drug in dispersion. Recent results from our research group revealed that SFN are an excellent vehicle for the transport of the natural drug naringenin, with anti-cancer properties [91], which has low solubility in water. The results indicated that this drug loaded in the nanoparticles is 1.7 times more effective in reducing the viability of HeLa cells than by itself. These results can be attributed to the low solubility and slow dissolution of free naringenin which, when loaded in the SFN, remains stable in dispersion, increasing its cellular penetration and improving the dissolution profile.

#### **3.3 Controlled release**

Nanoparticles can be used as drug reservoirs for their controlled release over time, which offers numerous advantages compared to conventional administration of multiple doses. Among them, it can be highlighted the improvement in efficacy and reduction of toxicity and patient cooperation [92]. The former can be considered as the increase in therapeutic activity compared to the intensity of the side effects, while the latter offers the advantage of reducing the number of applications required during treatment.

Controlled release is especially beneficial for those drugs whose half-life in the blood is relatively low due to a high rate of metabolism and elimination by the body. This effect can be observed in **Figure 3**, where the concentration of a drug in blood applied by a conventional method (red line) is represented against a controlled release system (blue line). As can be seen, the drug administered in a conventional manner is only a fraction of the time in the zone considered therapeutic, while fluctuating between subtherapeutic concentrations and above the maximum tolerable level. On the other hand, the controlled release system takes longer to reach the therapeutic concentration window but remains stable within it. The goal of the system is to match the rate of clearance to that of release in the

**Figure 3.**

*Diagram of the blood concentration of a drug after multiple administrations as a conventional injection (red line) and as a controlled release system (blue line).*

therapeutic concentration zone. In the clinic, this translates into numerous benefits, for example, in the case of administration of analgesics, the concentration could be prevented from falling to subtherapeutic levels and therefore the patient feeling pain. This is transferable to a large number of drugs including anti-inflammatories, antibiotics, anesthesia, hormones, chemotherapeutics, etc.

There are different mechanisms by which polymer nanoparticles can allow controlled drug release. On the one hand, the release can be delayed by using a water-soluble polymer as a matrix, whose dissolution rate is slow and consequently releases the drug at the rate of dissolution of the polymer. In the case of insoluble polymers, they can act as a diffusion barrier, slowing down the release of the drug from inside the nanoparticle to the medium. The release can also be controlled by an osmotic flow generated by a semipermeable membrane, which is itself the nanosystem, as is the case with liposomes. Finally, a delivery system that responds to internal or external stimuli could be achieved, which would be very useful, for example, in diabetic patients in which the nanosystem would release insulin on demand of the blood glucose concentration [93]. Volpatti et al. [94] have succeeded in synthesizing nanoparticles whose insulin release is sensitive to glucose levels by adding glucose oxidase and catalase to them. These researchers demonstrated that a single subcutaneous injection provides 16 h of glycemic control in diabetic mice. Cheng et al. [95] developed SFN capable of loading the antitumor drug paclitaxel (3%) and delivering it sustainably for 14 days.

#### **4. Nanoparticle design**

From the point of application to the site of action, nanoparticles face a host of challenges. In the first place, they are diluted in approximately 5 L of blood that circulates at 5 L/min through the circulatory system about 106 km long, where the velocity in each blood vessel can be between 1.5-33 cm/s [96] hindering the interaction between nanoparticles and the target tissue. Interstitial fluids have a much lower speed, just a few μm/s, where interactions would be favored. However, reaching them means crossing biological barriers, which is not an easy task. Finally, to all of the above, it is added that when nanoparticles enter the body they are treated hostilely by the immune system. For these reasons, different design principles are applied to nanoparticles to try to get around different obstacles depending on their final application.

#### *Nanoparticles as Drug Delivery Systems DOI: http://dx.doi.org/10.5772/intechopen.100253*

As mentioned above, as soon as the nanoparticles enter the body, they are exposed to the mononuclear phagocyte system which consists of a system of phagocytic cells, predominantly macrophages resident in the spleen, lymph nodes and liver, which sequester the nanoparticles immediately after administration [97]. This process begins with the opsonization of the nanoparticles based on the adsorption of plasma proteins, including albumin, complementary system proteins, pattern recognition receptors and immunoglobulins. This process is relatively fast and can occur in a period as short as 30 seconds [98]. This "natural functionalization" is known as the formation of the protein crown and clearly can alter the function or fate of nanoparticles by disturbing different parameters such as size, charge and surface chemistry, as well as hydrophobicity. This protein crown can even mask the receptors or ligands attached to the nanoparticles [99].

Different design strategies have been developed to avoid opsonization and subsequent clearance by the immune system. This evasion of the immune system tries to increase the circulation time of the nanoparticles in the body and, consequently, the chances that they find the target tissue while they circulate through the bloodstream. One of the easiest and most direct strategies is PEGylation, based on the functionalization with polyethylene glycol (PEG) molecules on the surface of the nanoparticles where the polymer units form very strong associations with the water molecules, generating a hydration layer and a steric barrier to opsonization [100, 101]. An alternative strategy may be to functionalize the nanoparticles with endogenous signals normally present in healthy cells. Rodríguez et al. [102] functionalized viral particles with the CD47 membrane protein, which acts as a "nonphagocytizing" signal [103], thus prolonging the circulation time. Another similar strategy is to cover the particles with biomimetic molecules such as cell membranes, to hide the particles from the immune system [104, 105]. Another way to increase circulation time is the one proposed by Nikitin et al. [106], which is based on a slight and transient suppression of the mononuclear phagocyte system through the administration of anti-erythrocyte antibodies. They were able to increase the circulating half-life of different nanosystems up to 32 times through the suppression of ca. 5% of hematocrits.

Silk fibroin exhibits unique low immune response properties, allowing it to evade the immune system. This can be exemplified by the study by Catto et al. [107], who implanted tubular matrices based on silk fibroin in mice, detecting few macrophages labeled with anti-ED1 antibodies, which was indicative of a low inflammatory response. The absence of T lymphocytes (anti-CD4 antibodies) demonstrated that there was no cell-mediated immune response. Recently, under a state-of-the-art design, Tan and colleagues [108] have designed a doxorubicin delivery nanosystem using silk fibroin as a Trojan horse. The researchers synthesized drug-loaded amorphous calcium carbonate nanoparticles and coated them with silk fibroin. It prevents the premature release of doxorubicin and helps evade the immune system. Thanks to the EPR mechanism, nanoparticles are accumulated in cancerous tissues and, finally, internalized by lysosomes. The acidic pH of the latter promotes the generation of CO2 from calcium carbonate, resulting in the bursting of the lysosome due to the expansion of the gas and the release of doxorubicin inside the target cell. Results in mice revealed that silk fibroincoated nanoparticles are more effective in reducing tumor mass and preventing side effects in mice compared to free doxorubicin or uncoated calcium carbonate nanoparticles. In addition, the immunotoxicity tests indicated that the nanoparticles did not initiate an immune response by not increasing the amount of T cells (CD4<sup>+</sup> and CD8<sup>+</sup> ) or IgM, IgG and IgA compared to the control group. More information on intracellular drug release can be found in the review by Fenghua et al. [109].
