Current Drug Delivery, 2012, 9, 000-000

Current Drug Delivery, 2012, 9, 000-000 1

1567-2018/12 $58.00+.00 © 2012 Bentham Science Publishers

Polymers and Drug Delivery Systems

Gemma Vilara,b, Judit Tulla-Puchea,b and Fernando Albericioa,b,c,*

aInstitute for Research in Biomedicine, Barcelona Science Park, Baldiri Reixac 10, 08028-Barcelona, Spain;

bCIBER-

BBN, Networking Centre on Bioengineering, Biomaterials and Nanomedicine, Barcelona Science Park, Baldiri Reixac

10, 08028-Barcelona, Spain; cDepartment of Organic Chemistry, University of Barcelona, Martí i Franqués 1-11,

08028-Barcelona, Spain

Abstract: In the treatment of health related dysfunctions, it is desirable that the drug reaches its site of action at a particu-lar concentration and that this therapeutic dose range remains constant over a sufficiently long period of time to alter the process. However, the action of pharmaceutical agents is limited by various factors, including their degradation, their in-teraction with other cells, and their incapacity to penetrate tissues as a result of their chemical nature. For these reasons, new formulations are being studied to achieve a greater pharmacological response; among these, polymeric systems of drug carriers are of high interest. These systems are an appropriate tool for time- and distribution-controlled drug delivery. The mechanisms involved in controlled release require polymers with a variety of physicochemical properties. Thus, sev-eral types of polymers have been tested as potential drug delivery systems, including nano- and micro-particles, dendrim-ers, nano- and micro-spheres, capsosomes, and micelles. In all these systems, drugs can be encapsulated or conjugated in polymer matrices. These polymeric systems have been used for a range of treatments for antineoplastic activity, bacterial infections and inflammatory processes, in addition to vaccines.

Keywords: Dendrimers, drug delivery, micelles, nano-micro-particle, nano-micro-spheres, polymers.

1. INTRODUCTION

1.1. Conventional Pharmacotherapy

Human health is threatened by autoimmune, neurodegen-erative, metabolic and cancer diseases, just to mention a few, which are difficult to treat with systemically, delivered drugs. Conventional pharmacotherapy involves the use of drugs whose absorption and therefore bioavailability de-pends on many factors, such as solubility, pKa, molecular weight, number of bonds per hydrogen atom of the molecule, and chemical stability, all of which can hinder the achieve-ment of a therapeutic response.

In general, the nature of conventional therapeutics, espe-cially their low molecular weight, confers them the capacity to cross various body compartments and access numerous cell types and subcellular organelles. Thus these drugs are suitable for the treatment of diseases. However, this form of indiscriminate distribution leads to the occurrence of side effects and to the need for higher doses of the drug to elicit a satisfactory pharmacological response. Rapid renal clearance as a result of the low molecular weight of these compounds, among other factors such as protein binding, lipophilicity, ionizability etc., implies frequent administration and/or a high dose to achieve a therapeutic effect.

Research is being conducted into new formulations that ensure a greater pharmacological response, which in turn would lead to lower doses and therefore the minimization of side effects. Thus, it is necessary to improve the

*Address correspondence to this author at the Institute for Research in Bio-medicine, Barcelona Science Park, Baldiri Reixac 10, 08028-Barcelona, Spain; Tel:/Fax: ??????????; E-mail: ????????????????????????

bioavailability of drugs. Bioavailability is affected by several factors, including the physical and chemical characteristics of the drug, the dose and concentration, the frequency of dosing, and the administration route.

Therefore, research into drug delivery systems seeks to improve the pharmacological activity of drugs by enhancing pharmacokinetics (absorption, distribution, metabolism and excretion) and also by amending pharmacodynamic proper-ties, such as the mechanism of action, pharmacological re-sponse, and affinity to the site of action.

Active pharmaceutical ingredients (APIs) are almost never administered alone but in dosage forms that generally include other substances called excipients. The latter are added to formulations in order to improve the bioavailability and the acceptance of the drug on the part of patients. Ex-cipients come in numerous forms, such as emulsifiers, dyes, lubricants, diluents, supporters, and chemical stabilizers. These substances were initially considered inert because they do not exert therapeutic action per se nor do they modify the biological action of a drug. However, it is currently upheld that excipients influence the speed and extent of drug absorp-tion, and therefore the pharmaceutical form of these sub-stances affects drug bioavailability.

In this regard, recent years have witnessed intense re-search on the modification of drug release and absorption. The development of new drug delivery systems will offer additional advantages to those mentioned above and may facilitate the launch of poorly soluble drugs. They may also allow an extension of patent protection for an API. And fi-nally, and possibly most importantly, these systems will fa-cilitate more patient-friendly administration, thus resulting in patient increased compliance and satisfaction.

2 Current Drug Delivery, 2012, Vol. 9, No. 4 Vilar et al.

The integral components of drug delivery systems are usually high molecular weight carriers, such as nano- and micro-particles, nano- and micro-capsules, capsosomes, mi-celles, and dendrimers, in which the drug is embedded or covalently bound.

Hence, the main function of polymeric carriers is to transport drugs to the site of action. Drugs are protected from interacting with other molecules which could cause a change in the chemical structure of the active ingredient causing it to loose its pharmaceutical action. Moreover, polymeric carri-ers avoid the interaction of the drug with macromolecules such as proteins, which could sequester the active ingredient preventing its arrival at the action place.

If a polymeric carrier is to be used, the next step is to design a type of polymeric structure that will permit obtain-ing the desired release conditions. Therefore, the polymeric structure should be: i) biodegradable, because the chemical bonds that make up its chemical structure break; ii) disas-semblable, because the various pieces forming the polymer disassemble but the chemical bonds do not break; and iii) undisassemblable, because the chemical bonds do not disas-semble or break, that is, the polymer remains unchanged. .. In the first two cases, micro-sized polymeric carriers could be used. However, if the polymeric structure of the polymer is not biodegradable or disassembable, then nano-sized polymers must be used for renal elimination.

A crucial feature of these polymers is the mechanism by which they are removed from the body. They may be ex-creted directly via kidneys (renal clearance) or biodegraded (metabolic clearance) into smaller molecules, which are then excreted. Passage through the renal glomerular membrane is limited to substances with a molecular weight under 50 KDa, although this value varies depending on the chemical struc-ture of the molecule [1]. Molecular weight is especially rele-vant for substances that are not biodegradable, and macro-molecules with a molecular weight lower than the glomeru-lar limit can be safely removed from the body by preventing their accumulation and therefore their potential toxicity.

In the case of biodegradable polymers, another option to consider in their design is the chemical structure of the polymer (degree of hydrophobicity, covalent bonds between monomers, etc.), since the speed and degradation condition, and therefore, the rate and site of drug release, can be modu-lated depending on the chemical structure of the polymer used. .

If the polymer is not biodegradable, the drug can be co-valently attached to the polymeric structure by a linker which can be degraded under different conditions such as in an acidic medium or by different enzymes. On the other hand, targets can be bound covalently to the surface which will help the directionality of the vector to the site of action.

Another strategy to follow would be to use smart poly-mers. These are polymers that release drugs when they are induced by a stimulation which causes a change in their structure. Therefore, the drug is released at the appropriate time and place.

Therefore, these new formulations seek to improve the pharmaceutical profile and stability of a drug, ensure its cor-rect concentration, achieve maximum biocompatibility,

minimize side effects, stabilize the drug in vivo and in vitro, facilitate the accumulation of the drug at a specific site of action, and increase exposure time in the target cell.

1.2. Controlled Release Methods

Controlled release systems aim to improve the effective-ness of drug therapy [2]. These systems modify several pa-rameters of the drug: the release profile and capacity to cross biological carriers (depending on the size of the particle), biodistribution, clearance, and stability (metabolism), among others. In other words, the pharmacokinetics and the phar-macodynamics of the drug are modified by these formula-tions.

Controlled release offers numerous advantages over con-ventional dosage forms. This approach increases therapeutic activity and decreases side effects, thus reducing the number of drug dosages required during treatment.

Controlled release methods offer an appropriate tool for site-specific and time-controlled drug delivery.

There are two main situations in which the distribution and time-controlled delivery of a drug can be beneficial. The first is when the natural distribution of the drug causes major side effects due to its interaction with other tissues, while the second is when the natural distribution of the drug does not allow it to reach its molecular site of action due to degrada-tion.

Many different kinds of drugs can benefit from distribu-tion or time-controlled delivery, such as anti-inflammatory agents [3], antibiotics [4], chemotherapeutic drugs [5], im-munosuppressants [6], anesthetics [7] and vaccines [8].

1.2.1. Time-Controlled: Modified-Release Formulation

This formulation refers to the release of the drug through a system that protects and retains it. Through controlled re-lease over time, this formulation allows the drug to reach a therapeutic concentration in tissues.

The usual dose administration of a drug can achieve im-mediate (short-term) therapeutic concentrations of the active ingredient [9] (Fig. 1.1).

Fig. (1.1). Curve of drug concentration in plasma over time after a single dose [10]. The figure shows that drug absorption [11] does not usu-ally stop abruptly when it reaches the maximum concentration,

Polymers and Drug Delivery Systems Current Drug Delivery, 2012, Vol. 9, No. 4 3

but may continue for some time along the downstream por-tion of the curve. Drug absorption ceases with time once the bioavailable dose has been absorbed. Therefore, the elimina-tion phase of plasma drug concentration depends on the rate of removal by metabolism or excretion.

The effective (therapeutic) concentration is the minimal drug concentration required to achieve the desired therapeu-tic effect. In contrast, the maximum safe concentration is drug concentration in plasma above which toxic or side ef-fects occur. The interval between these two concentrations is the therapeutic window. The concentrations that yield the desired new therapeutic effect in the absence of toxic effects fall within this window.

In a chronic treatment, doses are administered at regular intervals (multiple doses, shown in Figs. (1.2)). This ap-proach can lead to variations in API concentrations, reaching levels that are lower or higher than therapeutic concentra-tions, in other words, concentrations that are neither effective nor toxic, respectively.

Fig. (1.2). Drug concentrations in plasma after delivery by conven-tional injection in a chronic disease. Representation of the therapeu-tic window between the minimal therapeutic concentration (thera-peutic level, dotted line) and the maximum therapeutic concentra-tion (toxic level, solid line). Adapted from reference 10. Chronic treatment calls for adherence to a pattern of drug administration because an incorrect number of dosages may exceed the limit of therapeutic concentration and thus result in a toxic response. In general, in chronic treatments drug doses are separated by two times the half-life of the drug in blood. The time required for a compound that has reached the systemic circulation of the body to reduce the maximum concentration by half is called the half-life.

In contrast, a modified-release formulation is designed to deliver the active ingredient at a predetermined speed or at a site other than that of administration. These formulations can be classified on the basis of the systems that regulate drug release.

1.2.1.1. Extended-Release Formulation

The release of the active ingredient in these formulations is initially produced in a sufficient amount to produce a therapeutic effect and then to continue with a release, which

is not necessarily constant but extends the therapeutic action. There are two types of extended-release formulations. One is a polymeric matrix that contains the active ingredient, the delivery of which is controlled by diffusion through the po-lymeric system network. These formulations are presented in the form of hydrogels and nano- and micro-particles. The other type of extended-release formulation comprises API capsules with polymer coatings, forming a reserve deposit. The permeability or the degradation of the polymer coating regulates the release of the drug.

There are some commercial products such as Plenidil Er or Kapanol. Plenidil or Kapanol tablets provide extended release of felopine or morphine surfate, respectively. Both tablets are administered orally and the drug is released by diffusion through the system [12].

There are other commercially available products in which the degradation of the polymer coating regulates the release of the drug. An example of this formulation is Sinemet CR. The newest is a polymeric-based drug delivery system that controls the release of carbidopa and levodopa by slowly eroding the polymer´s coating, making the system particu-larly suitable in the treatment of Parkinson’s disease and syndrome [13].

Finally, it is important to make a few comments about the Drug-Eluting Stent (DES), which consists of supports coated with drug-laden polymers. Generally, a stent is a metal sup-port wire whose function is to maintain arteries, blood ves-sels or other anatomical ducts open. One interesting example is the cardiovascular stent, which is a device that is inserted into the coronary arteries when they are blocked to keep the arteries open and normal blood flow is resumed. Sometimes, this stent can be coated with a drug or polymer which con-trols the release of the drug, thus preventing the arteries from reclosing. Nowadays, several DES platforms have been de-veloped and evaluated for clinical use. The difference be-tween them has to do with the type of stent, anti-proliferative drug and polymer used to control drug release.

Depending on the type of polymer used, different drug-release kinetics is obtained. San Juan et al. [14] functional-ized a biocompatible polymer for stent coating for local arte-rial therapy. Therefore, metal stents were coated with a cationized pullulan hydrogel which was loaded with small interfering RNA for gene silencing in vascular cells. It was observed that the release of siRNA from the polymer was modulated by the presence of the cationic groups. Conse-quently, the release of the drug can be modulated depending on the type of polymer used.

There are some DES which have been approved for clini-cal use such as Cypher [15] and Taxus [16] containing si-rolimus and paclitaxel, respectively (Figs. 1.3 and 1.4).

1.2.1.2. Sustained-Release Formulation

The active substance is steadily released with a kinetic profile of zero-order. The therapeutic effect is maintained over a long period of time Fig. (1.5). The closest pharmaceu-tical formulations to this type of system are osmotic pumps. These devices control the outflow of drug solutions through osmotic potential gradients across semi-permeable polymer barriers. The pressurized chambers contain the aqueous solution


Fig. (1.3). A). Nano- or micro-capsules whose degradation allows the release of the drug. B) Nano- or micro-particles of hydrogels whose swelling structure allows the diffusion of the drug through the pores of the structure. Adapted from reference 10a.

Fig. (1.4). Representation of the concentration of active ingredient versus time for a prolonged release formulation. The graph shows the therapeutic window between the minimal therapeutic concentration (therapeutic level, dotted line) and the maximum therapeutic concentra-tion (toxic level, solid line). Adapted from reference 10a.

Fig. (1.5). Osmotic pumps. Adapted from reference 10b.

of the drug and the polymeric osmotic system. Upon immer-sion in the water, the osmotic system is hydrated and swol-len, thus causing an increase in pressure, which is relieved by the flow of the solution out of the delivery device through an orifice in the upper part [17].

Drug delivery through osmotic systems is affected by a number of factors such as solubility, osmotic pressure, size of the delivery orifice and membrane type and characteris-tics. An example of this type of product is Adalat OROS. The OROS is a 24-hour controlled release oral drug delivery system in the form of a tablet, which acts as an osmotic pump releasing nifedipine [18]. The tablet has a rigid water-

permeable jacket with one or more laser drilled small holes. As the tablet passes through the body, the osmotic pressure of water entering the tablet pushes the active drug through the opening in the tablet. Once the active ingredient in the Oros tablet has been released, the remnant is excreted in the feces (Fig. 1.6)

1.2.1.3. Pulsatile-Release Formulation

Another form of time- controlled release is a responsive delivery system in which drugs are released in a pulsatile manner only when required by the body. Thus, the drug is released after a well-defined lag time. In general, these drug


delivery systems comprise two components, namely a sensor that detects the environmental parameter that stimulates drug release, and a delivery device that releases the drug.

Fig. (1.6). Representation of the concentration of active ingredient versus time for a sustained-release formulation. The graph repre-sents the therapeutic window between the minimal therapeutic con-centration (therapeutic level, dotted line) and the maximum thera-peutic concentration (toxic level, solid line). Adapted from refer-ence 10a. Pulsatile drug delivery systems release the drug at the right time and place and in accurate amounts, and they are convenient for patients suffering from chronic problems such as diabetes, arthritis, asthma, and hypertension [19].

The pulsatile-release formulation implies the release of the drug only when required by the body, thus preventing it from being continuously in the biophase.

Insulin formulations currently require repeated injections daily, and hence responsive drug delivery systems that re-lease insulin in response to increased blood glucose levels have been designed. For the treatment of diabetes, a pulsa-tile-release formulation that uses the enzyme glucose oxidase as a sensor has been proposed [20]. When blood sugar levels rise, glucose oxidase converts glucose to gluconic acid, thereby lowering the pH. This decrease in pH causes insulin release because the pH-sensitive polymers (smart polymers) either swell or degrade in acidic environments.

Another example of the above is the formulation devel-oped and studied by Sadaphal et al. [21]. This formulation is constituted by an impermeable anionic polymer called Eu-dragit S100 which encapsulates theophylline. This system attempts to mimic the circadian rhythm of the disease by releasing the drug at the appropriate times (Fig. 1.7).

1.2.1.4. Delayed-Release Formulation

In this case, the release of the active ingredient does not coincide with the time of administration and its therapeutic action is not extended.

Normally, these formulations are enteric coatings that protect the active ingredient against adverse physiological media or protect certain parts of the body against the hostile action of the active ingredient. Typically, these coatings are sensitive to pH and allow drug release into the small intes-tine, thus preventing its delivery to the stomach.

Fig. (1.7). Representation of the concentration of the active ingredi-ent versus time for a sustained-release formulation. The graph shows the therapeutic window between the minimal therapeutic concentration (therapeutic level, dotted line) and the maximum therapeutic concentration (toxic level, solid line). Adapted from reference 10a. Therefore, modified-release formulations provide alterna-tives to the routine administration of the drug for improving its management and optimizing its therapeutic action. Thus, controlled release is highly beneficial for drugs that show poor bioavailability, in other words, those that are rapidly metabolized and eliminated from the body after administra-tion and thus have a short life, or those with a narrow thera-peutic window. Some drugs are unstable in certain environ-ments, such as in plasma or in acidic conditions. Therefore, in these conditions, it is of interest to use a polymer to pro-tect the drug. These formulations are also used in pathologies in which the degree of compliance is low because of difficult administration or unpleasant taste of the medicine.

Given that modified-release formulations have a series of drawbacks over standard systems, such as higher production, cost, lack of reproducibility, unpredictable in vitro/in vivo correlation and problems associated with improper handling, their development for an active ingredient is not justified unless they show a clear advantage over the standard formu-lation.

1.2.2. Controlled Distribution

The treatment of most diseases requires the distribution of the drug at a specific site. It is therefore important to use drug carriers with a controlled distribution system. Several approaches are currently available to overcome the obstacles posed by non-specific drug delivery.

One method is to functionalize the surface of nano- and micro-particles with receptor recognition elements that are present in diseased tissue. Thanks to molecular biology, a comprehensive database of molecular targets has been built and it is now known that some specific receptors are overex-pressed in malignant tissue. Therefore, the identification of these overexpressed targets is used to functionalize polymers with tracking devices, such as antibodies [22], carbohydrates [23], or simply an electrically charged species, which inter-act with these targets [24]. Consequentlty, these tracking devices help the polymer reach the tissues and regions of the body where the drug is to be released [25] (Fig. 1.8).


Fig. (1.8). Polymer functionalized to interact with specific targets. Adapted from reference10c. Another example of a controlled distribution system is drug release by diffusion through a hydrogel. To this end, the drug is encapsulated or absorbed in the hydrogel and re-leased in response to an external stimulus.

The stimuli that can be used in various systems to trigger drug release have been reviewed by Qiu et al. and Gupta et

al. [26] One interesting model is called the pH-dependent system, which is characterized by ionizable groups in its structure. These groups are ionized depending on the pH, and therefore an attractive interaction may occur between them through hydrogen bonding, or a repulsive interaction may be induced because of ionic groups with the same charge. Thus, a change in the structure of these groups opens or closes the route through which the drug is released. Other systems are controlled by electrical signals. This stimulus can change features in the structure of the formulation, such as pore size and permeability. Another approach is the use of hydrogels, the behavior of which differs in response to temperature. Also of note are those systems that are triggered by external stimuli in the form of variations in the concentration of cer-tain substances, such as glucose, urea and morphine. How-ever, these will be discussed further in the section on smart polymers (Table 1.1).

The controlled release of an active ingredient can also be achieved by connecting the drug to the polymer through a specific linker [27] that can be degraded in an acidic envi-ronment or by a specific enzyme. Thus, the drug is released from the polymer when the entire system is exposed to either of these two conditions, such as in tumor tissues where the pH is lower than in healthy tissues, in an endosomal acidic environment, and in a lysosomal environment, which con-tains proteolytic and hydrolytic enzymes.

The introduction of macromolecules into cells by endo-cytosis [28] involves the invagination of the plasmatic mem-brane to form a small vesicle called the endosome. This early endosome matures and finally fuses with the lysosome, which contains hydrolytic and proteolytic enzymes in an acidic environment, thus forming what is known as a secon-dary lysosome. Therefore, in this environment within the lysosome, the linkers that are labile to acids or digestive en-zymes are cleaved, causing drug release [29].

It is important to highlight that these linkers are stable in plasma and are degraded only in tumor tissues or in lysoso-mal or endosomal [30] microenvironments. Tumor tissues or early endosomes have a slightly acidic pH, where labile acid linkers such as cis-aconityl or hydrazone can be degraded [31]. For example, the proteins carried by the polymer and conjugated in it with an acid-labile linker are released into the endosome because peptide bonds are degraded in the lysosome. Once the linker has been biodegraded, the drug is released from the endosome or lysosome by diffusion to the cytosol (Fig. 1.9).

The linkers most frequently used are peptides and oligo-saccharides. Linker libraries have been established by Vec-traMed, based on enzyme cleavage specificity of a selected disease. One example is the enzyme family of Cathepsin, which includes lysosomal endoproteases, Cathepsin B being among these [32], a molecule localized at the invasive edges of human tumors and therefore overexpressed in many types of cancer. Other examples of proteins include Prostate Spe-cific Antigen (PSA), which is expressed by prostate tumors, and Proline peptidases, which hydrolyze proline sites. These peptidases are overexpressed in pulmonary hypertension and during inflammation.

2. POLYMERS USED FOR CONTROLLED-DRUG

RELEASE

Polymers used for controlled drug release are often called therapeutic polymers. However, these systems do not have

Table 1.1. Some Examples of Smart Polymers

Stimulus Polymer Drug Released

pH Poly(methacrylic-g-ethylene glicol) (p(MMA-g-EG) Insulin

Electric field Poly(methacrylic acid) (PMA) Pilocarpine and raffinose

Glucose concentration Poly(methacrylic acid-co-butyl methacrylate) Insulin

Temperature Layer of Chitosan Pluronic on PLGA microparticles Indomethacin

Morphine concentration Methyl vinyl ether-Co-anhydride maleic copolymer Naltrexone

Urea concentration Methyl vinyl ehter-Co-anhydride maleic copolymer Hydrocortisone


specific therapeutic activity and are considered excipients in pharmaceutical formulations.

The main interest in polymers for drug delivery purposes is in controlled-release systems. In these applications, drugs are embedded or conjugated in polymer matrices.

Fig. (1.9). Endocytotic pathway for the cellular uptake of macro-molecules and nanocarriers for drug delivery. Adapted from ref. 10c. From a polymer chemistry perspective, it should be noted that distinct mechanisms of controlled release require poly-mers with a variety of physicochemical properties [33]. Therefore, formulation for drug delivery includes numerous architectural designs in terms of size, shape, and materials. Several types of polymers have been tested as potential drug delivery systems, including nano- and micro-particles, den-drimers, nano- and micro-spheres, capsosomes and micelles.

The difference between nano- and micro-formulation is the size. The size of the former confers this approach a greater capacity to reach the smallest capillary vessels and to penetrate tissues either through paracellular (transport of substances between cells of an epithelium) or transcellular pathways (substances travel through the cell). Nanoformula-tions can be administered through various routes, including oral, pulmonary, nasal, parenteral, and intraocular: however, some routes cannot be used with microformulations, such as the intravenous route, because of possible obstruction of veins.

2.1. Drug-Polymer Conjugates

As mentioned earlier, therapeutic agents can be cova-lently bound to the polymer backbone. In these types of drug-polymer conjugates, the goal is to improve the mecha-nism of cellular internalization and cell specificity to achieve optimal release of the drug at the proposed target. In con-trast, the systems in which the drugs are embedded or encap-sulated in the polymer seek to enhance the distribution and serum stability and to attain a decrease in drug immuno-

genicity through a controlled distribution at the correct site and over time.

Polysaccharides, for instance, contain hydroxyl groups that allow direct reaction with drugs with carboxylic acid functions, thereby producing ester linkages that are biode-gradable and thus facilitate the release of the drug in the body.

Drug polymer conjugates are of interest when these drugs are attached to the polymer with linkers that are labile to certain digestive enzymes or acidic conditions (see con-trolled distribution). This type of release is used by nanopar-ticles, which have high mobility in the smallest capillaries, allowing for efficient uptake and selective drug accumulation at the target sites [34]. This type of nanoparticle is common in cancer therapy.

Depending on the desired application, nanodelivery sys-tems must have a certain particle size, surface charge and hydrophobicity in order to overcome physiological barriers.

These barriers comprise biological structures or physio-logical mechanisms that prevent nanoparticles from reaching their targets, thus compromising therapeutic efficacy. These biological barriers can be overcome when the nanoparticles show efficient extravasation through the vasculature, pro-longed vascular circulation time, improved cellular uptake, and endosomal or lysosomal escape.

In recent years, polymeric carriers of antineoplastic drugs have been extensively studied [35]. These polymers can ac-cumulate passively in cancerous tissue as a result of certain differences in the biochemical and physiological characteris-tics of healthy and malignant tissue. In general, small mole-cules diffuse through the endothelial cell wall to healthy and malignant tissues. However, macromolecules can reach only the latter because the walls of the endothelial cells of these tissues are fenestrated, resulting in pores that allow macro-molecules to cross. Other features of malignant tissue in-clude poor lymphatic drainage and increased vascularity (angiogenic process), which both lead to the retention of macromolecules. This passive accumulation of macromole-cules in tumor tissue, known as the EPR effect (enhanced permeability and retention effect), was first reported by Maeda et al. [36] (Fig. 1.10).

Therefore, nanoparticle-mediated drug delivery can be either an active or a passive process. The latter refers to transport through leaky capillary fenestrations into tumor cells by passive diffusion. Thus, there is an accumulation of nanoparticles in these cells due to the EPR effect. In contrast, an active process involves the use of peripherally conjugated targeting moieties for enhanced delivery to a specific site.

Once macromolecules reach the malignant tissue, they are endocytosed (receptor-mediated endocytosis) or phago-cytosed (internalization without the mediation of receptor) by malignant tissues cells to form a vesicle called the en-dosome. Given that the pH in these vesicles is lower (6.5-5.0) than that of the extracellular matrix (7.2-7.4), acid-labile linkers can be hydrolyzed, thus releasing the drug. Finally, the endosome is absorbed by a lysosome, resulting in a sec-ondary lysosome that contains hydrolytic and digestive en-zymes that are active in acidic environments (lysosome pH: 4). Some studies show overexpression of certain enzymes in


malignant tissue; this overexpression could be used to obtain a controlled release when a linker labile to these enzymes is used. Furthermore, the pH of tumor tissue is slightly lower than that of healthy tissue and therefore controlled-drug re-lease can also be obtained when an acid-labile linker is used; however, in this case the drug is released at the extracellular level.

Several drug-polymer conjugates have entered the clini-cal development stage [37].

Some of these compounds, such as HPMA [38] copoly-mer-doxorubicin galactosamine (PK2), HPMA copolymer platinate (AP 5346), carboxymethyldextran polyalcohol-bound exatecan (DE-310), poly(L-glutamic acid)-paclitaxel (Xyotac), PEG-SN38 (EZN-2208), linear cyclodextrin-bound camptothecin (IT-101) and poly(L-glutamic acid)-camptothecin (CT-2106) have progressed further to the clini-cal phase.

However, clinical studies for several other polymer con-jugates, such as HPMA copolymer camptothecin (PNU 166148) and HPMA copolymer paclitaxel (PNU 166945), have been discontinued or suspended [34-51].

Developed by Kopecek and Duncan, HPMA copolymer-doxorubicin (PK1) was the first cancer drug conjugated with a polymer to be tested clinically [39]. Its structure contains a peptide linker (-Gly-Phe-Leu-) that allows doxorubicin re-lease through the action of lysozyme. Clinical studies showed that the toxicity of the conjugate was lower than doxorubicin alone. These studies also demonstrated pro-longed circulation of the conjugate in plasma. Moreover, the results of phase II trials showed antitumor activity for the treatment of breast and non-small cell lung cancer but no response in patients with colorectal cancer [40].

Later, HPMA copolymer-doxorubicin galactosamine (PK2) was synthesized Fig. (1.11) [41]. This compound has a similar structure to PK1 but incorporates an additional tar-geting ligand, namely galactosamine, which specifically tar-gets liver tissues (controlled distribution). The drawback is

that the receptor of the galactosamine ligand is expressed on both malignant and healthy hepatocytes.

HPMA copolymer platinate (AP 5346) [42] is another drug-polymer conjugate that has progressed into clinical trials. This compound consists of a cytotoxic diaminocyclo-hexane (DACH)-platinum moiety coupled to a biocompati-ble hydroxypropylmethacrylamide copolymer (HPMA). Platinum drugs (carboplatin and cisplatin) are used to treat a wide range of cancers and they react in vivo, binding to and causing crosslinking of DNA, thus triggering apoptosis. Fur-thermore, DACH has activity against cisplatin-resistant can-cer cells. Therefore, the DACH-platinum moiety was bound to HPMA via a pH-sensitive linker, and consequently re-leased in the extracellular space of tumors and/or the intra-cellular lysosomal compartment. Another interesting feature of this drug-polymer conjugate is its size (approx. 25 kDa), which not only allows its renal clearance but also enables it to reach tumor cells through fenestrated endothelial cells. Therefore, AP 5346 has a longer half-life relative to small platinum drugs that enables it to reach tumor cells by a pas-sive process and then release the drug. Partial responses in antitumor activity in metastatic melanoma and ovarian can-cer to AP 5346 were observed in a Phase I study. Thus, this drug has been shown to have a higher therapeutic index in multiple preclinical tumor models and it is entering phase II clinical testing under the new name of Prolindac [43].

Another drug-polymer conjugate in clinical trials is DE-310. This conjugate is composed of exatecan (DX-8951) and carboxymethyldextran polyalcohol (CM-Dex-PA) as a car-rier. Used in cancer chemotherapy, exatecan is camptothecin analog, a cytotoxic quinoline alkaloid that inhibits the DNA enzyme topoisomerase I. Camptothecin binds to topoi-somerase I and DNA, resulting in a stable ternary complex, thus preventing DNA re-ligation. This drug causes DNA damage, thus causing apoptosis. The structures of exatecan and camptothecin hold a lactone ring that is highly suscepti-ble to hydrolysis and is inactive in its open form. In other words, the open form does not inhibit topoisomerase I. To

Fig. (1.10). Schematic representation of the anatomical and physiological characteristics of healthy and tumor tissue with respect to the vas-cular permeability and retention of small and large molecules (EPR effect). Adapted from reference 26.


protect the closed form (active form), this drug is conjugated to a polymer such as DE-310. In this polymer, exatecan is covalently linked via a peptidyl spacer (Gly-Gly-Phe-Gly). This spacer is labile to the activity of cathepsin and is thus used to release the drug inside the cell. This release was studied by Shiose et al. [44], who addressed the relationship between cathepsin activity and the release of DE-310 in sev-eral types of tumor. Moreover, clinical trials in patients with advanced solid tumors were studied by Soepenberg et al. [45]. Those clinical trials determined drug toxicity and pharmacokinetics and therefore, the maximum tolerated dose in phase II. Exatecan was observed to show a slow release from DE-310, thus achieving prolonged exposure to this topoisomerase I inhibitor.

Another interesting drug polymer conjugate that has been studied in numerous clinical trials, including phase III trials, is poly(L-glutamic acid)-paclitaxel, PG-TXL, also called XYOTAX (CT-2103) [46]. Paclitaxel is used in cancer che-motherapy. This drug stimulates the assembly of microtu-bules from tubulin dimers, thus preventing depolymerization. Thus, it inhibits the formation of mitotic cell division by blocking mitosis. Paclitaxel is used to treat ovarian, breast and lung cancer. Paclitaxel is conjugated via an ester linkage through its 2’hydroxyl group to the carboxylic acid of poly(L-glutamic acid). PG-paclitaxel is a highly promising anticancer drug polymer conjugate because it has the charac-teristics of an ideal polymeric drug carrier, such as biode-gradability, high drug loading, stability in circulation, and

apparent lack of immunogenicity. In vivo, the biodistribution [47], in other words, the area under the curve obtained in the graph on tissue tumor concentration versus time (AUC), is greater when the drug is administered as PG-TXL than when it is administered in other formulations of paclitaxel. This enhanced distribution is caused by prolonged tumor exposure and limited systemic exposure compared to the active drug. Consequently, PG-TXL shows less toxicity. This conjugate can be administered by short infusion into peripheral veins, and preclinical studies have demonstrated that it is well tol-erated in this mode. The maximum tolerable dose (MTD) and pharmacology of PG-paclitaxel administered weekly to patients with solid tumors were studied [48]. Results of phase II studies [49] showed that the MTD of PG-TXL was higher than that of other formulations of paclitaxel and that PG-paclitaxel has activity in patients who have not re-sponded to taxol therapy. Phase III trials are currently un-derway to test the efficacy of this compound in patients with advanced non-small-cell lung cancer.

PEG-SN38 (EZN-2208) [50] is currently entering phase II clinical testing. This compound is composed by SN38 linked by a covalent bound to poly(ethylene glycol). SN38 is a camptothecin analog and therefore has the same function as camptothecin (topoisomerase I inhibitor) and a similar struc-ture, containing a lactone ring. Thus, SN38 has to be admin-istrated with a formulation, such as by conjugation to a PEG polymer that protects the lactone ring and prevents its hy-drolysis. The results in animal experiments showed a higher

Fig. (1.11). Structure of PK2, which has a targeting ligand (R): galactosamine (targeting ligand).

CH2

O NH

HN

O

Ph

O NH

HN

O

RO

CH2

NH NH

HN

O

OH

O

O

Ph

NH

HN

O

O

NHO

OH

O

OH

O OH

O

O

O OH

OH

H2

C

HPMA

Linker

Doxorubicine


half-life of the drug-polymer conjugate than the free drug, a higher drug exposure to tumor cells, and an increased solu-bility. The dose-limiting toxicity and dosing schedules were determined in phase I clinical trials.

IT-101 [51] is a novel approach to cancer chemotherapy that is currently under study in human trials. IT-101 is com-posed by camptothecin covalently attached through a glycine link to a cyclodextrin-based polymer (CDP), which in turn consists of -cyclodextrin and poly(ethylene glycol). This compound is administered by intravenous injection and demonstrates high antitumor activity because of extended circulation times, which favors its access to tumor tissues. Results from a phase I clinical trial showed a decrease in the side effects of patients who had been administered this drug and thus an enhancement in their quality of life. These favor-able results prompted the initiation of a phase II trial de-signed to determine the side effects of this compound’s pro-file as maintenance therapy in ovarian cancer patients.

Another polymer conjugate is poly(L-glutamic acid)-camptothecin (CT-2106). This compound is composed by camptothecin covalently attached to poly(L-glutamic acid) through a glycine linkage. CT- 2106 has entered phase I/II trials and clinical studies have been conducted by Honsi et al. and Springett et. al. CT- 2106 was studied in patients with solid tumor malignancies [52].

In addition, as mentioned previously, clinical studies of other polymer conjugates, PNU 166148 and PNU 166945, have been discontinued or suspended.

PNU 166945 [53] is a HPMA polymer-conjugated de-rivative of paclitaxel. It was studied with the aim to improve drug solubility with the subsequent controlled release of pa-clitaxel. Toxicity associated with paclitaxel, such as neuro-toxicity, was observed in phase I trials; however, this drug showed antitumor activity.

PNU 166148 [54] is composed by N-(2-hydroxypropyl)methacrylamide (HPMA) conjugated with camptothecin. The latter is modified at the C20 position of hydroxyl group with glycine and was conjugated with Gly-C6-Gly. In phase I trials, PNU 166148 was administered to patients with solid tumors. The results showed toxicity and absence of tumor targeting. Thus, clinical development of this compound was abandoned.

2.2. Carriers in which the Drug is Embedded or Encap-

sulated

Chemists working in synthesis seek to develop polymeric structures with physical and chemical properties that allow the specific release of the therapeutic agent under predeter-mined conditions.

The polymers used in the development of drug delivery systems are designed with the capacity to form polymeric supramolecular structures (matrices and capsules), which are suitable for retaining a therapeutic agent and allow con-trolled release. Interactions between the polymer chains, which can be non-covalent (electrostatic and hydrogen bonds) or have covalent bonds (cross-linked polymers), con-fer stability to these structures. The integrity and the behav-ior of these structures in physiological conditions allow the controlled release of the therapeutic agent. The literature describes several mechanisms of drug release [55].

One such mechanism is controlled release by swelling. Hydration of the polymer causes an increase in the volume of the polymeric structure and the resulting increase in pore size allows diffusion of the aqueous medium within the po-lymeric structure and thus the release of the drug.

Drug release can also be achieved through polymer deg-radation. Gradual hydration of the polymeric structure may cause hydrolytic degradation with release of its contents. This degradation depends on the stability of the polymer chain linkages in body fluids Fig. (1.12).

The release of a therapeutic agent depends on the rate of polymer degradation. Heterogeneous degradation occurs when the polymer is not fully degraded. This system requires two types of polymers, one degradable and located on the surface of the material to allow it to be in contact with the physiological environment, and the other one non-degradable and located inside the external polymer. In this case, the deg-radation rate is constant and the non-degraded material (ma-terial inside the external biodegradable scaffold) maintains its chemical integrity during the process. Thus, the therapeu-tic agent is released once the material surface has been de-graded by diffusion through the non-degradable material. In contrast, homogeneous degradation is a random deterioration that occurs in the entire mass of the polymer matrix. While the molecular weight of the polymer decreases continuously, the material maintains its original shape and retains the drug. When the polymer mass is highly degraded and reaches a critical molecular weight, solubilization of the polymer and the drug begins.

Finally, another drug release mechanism is through pure diffusion. The drug slowly diffuses through the voids of the polymeric device. These polymers are fabricated as matrices in which the drug is uniformly distributed.

Modulation of the chemical architecture of polymers (functional groups, interactions between chains…) can regu-late the release of the therapeutic agent and thus the profile of the final formulation (delayed or immediate release, pulsa-tile, etc.). However, the release occurs through a combina-tion of the three basic mechanisms mentioned previously

Fig. (1.12). Linkages depicted from highest to lowest lability against hydrolysis in aqueous media, published by W. R. Gombotz, et al. [56].

O

O

O

O

O O

O

O

O

O NH

O

O O

O

O

O

NH

Polyanhydrides > Polycarbonates > Polyesters > Polyurethanes > Polyorthoesters > Polyamides


(release by swelling control, by pure diffusion and by poly-mer degradation).

There are several pharmaceutical formulations in which the drug is embedded or encapsulated in the polymer. Among these, we find the micro- and nano-particles, micro- and nano-capsules, capsosomes and micelles. In the first case, the therapeutic agent is normally absorbed into the par-ticles and the release is caused by diffusion through the pores of the polymer. In contrast, in the case of microcapsules, the drug is encapsulated within these structures and generally its release requires the degradation of the polymeric matrix. Other interesting carrier polymers are capsosomes, which consist of a drug carrier with an external core and some liposomes inside, thus allowing various compartments inside its structure. Finally, micelles are composed of amphipathic linear polymers that are spontaneously formed by self-assembly in water and the drug is encapsulated in the hydro-phobic core.

Furthermore, dendrimers are polymers with interesting properties as drug carriers because of their specific surface functional groups, which allow the attachment of hydrophilic drugs by covalent bonding, electrostatic interactions or hy-drogen bonding.

2.2.1. Gels and Hydrogels

Gels and hydrogels are hydrophilic polymers that vary in their structures. The former have a linear structure while the latter are polymeric networks with three-dimensional cova-lent crosslinking.

Therefore, linear polymer gels are formed by long-chain monomer units linked by covalent bonds such as amides, esters, orthoesters, and glycosidic bonds. Once the linear polymers are formed, other types of interactions (hydrogen bonds or Van der Waals interactions) contribute to achieve the three-dimensional structure. In contrast, hydrogels in-volve monomers of distinct chains that are covalently linked. Covalent bonds between chains affect polymer properties and thus make these polymers convenient for use as drug carriers in the form of micro- or nano-particles.

One of the main characteristics of hydrophilic polymers is their capacity to absorb large amounts of water. The affin-ity for water is attributed to the presence of hydrophilic groups in their structure. This interesting property has led to the use of these polymers in several fields of biotechnology, such as biosensors and drug carriers.

As an example, Pishko et al. [57] developed a fluorescent biosensor consisting of seminapthofluorescein conjugated with the enzyme organophosphorus hydrolase (SNALF-1) which is encapsulated in a PEG hydrogel. This enzyme cata-lyzes the hydrolysis of neurotoxins, thus leading to an in-crease in pH. In addition, the fluorescent compound, seminapthofluorescein, which is conjugated to the enzyme, is labile to changes in pH, thereby causing a change in its emis-sion spectrum. Therefore, when the biosensor is incubated in an aqueous medium containing a neurotoxin such as par-aoxon, at first the polymer swells, thus allowing diffusion of the water and toxin inside the polymer. Once the organo-phosphorus hydrolase enters into contact with toxin, the lat-ter is hydrolyzed, resulting in a change in the pH of the me-

dium and thus a change in the emission spectrum of the seminapthofluorescein.

Another interesting example is the study carried out by Syu et al. [58] on the design and synthesis of biosensors for urea concentration measurement. They immobilized urease enzymes in a core of carbon paper on which a polypyrrole matrix was deposited forming an external layer. Urea con-centration measurement is based on ammonia concentration, which is the product obtained from the catalysis reaction of urease enzymes.

Gels and hydrogels are also used as drug carriers. These are classified as matrices or capsules, depending on the way in which the drug is found within the polymer, either embed-ded (matrix) or encapsulated (capsules).

2.2.2. Nano- and Micro-particles

Nano- and micro-particles are formulations in which the drug is trapped or embedded within the polymer. The poly-mer serves to transport and protect the drug and to allow its controlled release. The drug carrier consists of hydrophilic polymers, gels and hydrogels being among these.

Hydrogels are the most widely used hydrophilic poly-mers for this kind of pharmaceutical formulation because they are non-toxic and their three-dimensional structure al-lows controlled drug release.

Since hydrogels were first introduced into the field of biomedicine, they have consistently shown excellent charac-teristics of biocompatibility, which is attributed to their physical properties that make them similar to living tissue, especially regarding their high water content and also soft consistency and elasticity.

2.2.2.1. Synthetic Hydrogels

Hydrogels are obtained by simultaneous polymerization and crosslinking of several polyfunctional monomers. There-fore, a crosslinking agent is required to obtain the hydrogel structure. Different types of chemistry have been used to generate cross-linked polymers, among these the attack of the nucleophile to an electrophile, or radical chemistry.

The characteristics of the monomers used and the degree of crosslinking determine hydrogel properties, such as swel-ling, and, therefore, its applicability.

A number of synthetic hydrogels have also been studied for drug delivery purposes, such as PHEMA (2-hydroxyethyl methacrylate), PMA (poly((metha)acrylic acid)) and co-polymers of PVA (Poly(vinyl alcohol)) or PEG (polyethyl-ene glycol) with acrylamides.

2.2.2.2. Drug Loading into a Polymer

Therapeutic agents can be physically entrapped in the polymeric matrix or covalently bound to the polymer back-bone. Given the capacity of hydrogels to absorb large amounts of water and thus their capacity to swell, many stud-ies have been devoted to this type of polymer, in which the therapeutic agent is embedded in the gel and released when the gel structure swells. The rate of drug release depends on crosslinking density (highly cross-linked systems provide slow drug release because of their small mesh size) and drug solubility.


The drug can be loaded by in situ polymerization or can be embedded after the polymer has been synthesized.

In the first case, microparticles of PEG hydrogels were synthesized by Peppas et al. [59]. These hydrogels were pre-pared by free radical copolymerization of the macromono-mers poly (ethylene glycol) dimethacrylate (PEGDMA) and poly (ethylene glycol) monomethacrylate (PEGMA) in a range of proportions, thus obtaining several degrees of crosslinking. This radical reaction required the use of a pho-toinitiator, a UV lamp, and distilled water as a solvent of the reaction. The therapeutic agent, diltiazem, was loaded by adding it to the reaction mixture prior to exposure to the UV light source.

As mentioned earlier, a drug can also be absorbed into preformed polymers. This requires the incubation of a con-centrated drug solution with the preformed polymer carrier.

Varsohaz et al. [60] used theophylline as a model drug to be encapsulated in a polymer cross-linked PVA. In this case, the polymer was synthesized by the nucleophilic attack of the alcohol groups in the PVA structure and the aldehyde groups in the glutaraldehyde molecule, thus forming a cross-linked structure. Encapsulation of the drug was achieved by exposing the polymer previously synthesized with a drug solution in NaOH. Several tests of drug encapsulation in the polymer were carried out by varying the drug concentration and polymer composition (proportion of glutaraldehyde). The results showed that the greater the proportion of glu-taraldehyde, the greater the extent of drug loading.

An article published by Nakamura et al. [61] cites other interesting loaded preformed microparticles. In this case, Nakamura et al. addressed the behavior and kinetics of budenoside release from microparticles of P(MAA-g-EG). The latter is a copolymer of polymethacrylic acid and poly-ethylene glycol. They studied various methods of absorption because budesonide is soluble in ethanol but this solvent hinders the swelling of the resin and therefore the internali-zation of the drug within the polymer. Thus, the loading of the drug into the polymer was examined using a number of aqueous solutions with varying amounts of ethanol. It was observed that the smaller the amount of ethanol used in the solution, the greater the incorporation of the drug into the polymer.

We can see from previous articles that absorption condi-tions used differ because of the distinct physicochemical properties of each polymer and drug. Thus, a drug can be soluble or unsoluble in an aqueous medium. The polymer may show distinct degrees of swelling depending on the pH of the medium or depending on the degree of crosslinking in the structure of the polymer. Therefore, all these characteris-tics may cause a variation in the partition coefficient of the drug between the polymer and the solution. Therefore, in these cases, the ideal is to obtain optimum conditions of poly-mer swelling and an acceptable degree of drug solubility.

2.2.3. Nano- and Micro-Capsules

Nano- and micro-capsules are formed by an outer shell that encapsulates the active agent. This kind of pharmaceuti-cal formulation is commonly used to transport drugs, such as proteins, that are rapidly degraded in body fluids. Thus the

drug is encapsulated in the particle and is released by degra-dation of the polymer coating.

In general, micro- and nano-capsules are synthesized with hydrophilic polymers. Gels are the most widely used hydrophilic polymers for this kind of pharmaceutical formu-lation; however, example of micro- and nano-capsule syn-thesis with hydrogels can be found.

Gels are made either from natural or synthetic polymers. Among the former, are the most widely used. These are polymers composed of repeated monomer units, monosac-charides, joined together by glycosidic bonds. Polysaccha-ride chitosan has been extensively used in controlled drug delivery. In addition, poly(D, L-lactic acid) PLA, poly(glycolic acid) PGA and their copolymer poly(D, L-lactic-co-glycolic acid) PLGA are among the most com-monly used synthetic polymers.

Another interesting point concerns drug entrapment. Physical drug loading can be performed by incorporating the drug while producing the particles or by incubating a con-centrated drug solution with the preformed polymer carrier.

The optimum incorporation strategy for efficient entrap-ment should be selected on the basis of the physicochemical characteristics of the drug-carrier pair. Thus, entrapment efficiency depends on drug solubility in the polymeric ma-trix, which is in turn related to the composition and molecu-lar weight of the polymer, drug-polymer interactions, and the presence of functional groups [62].

The bibliography cites many examples in which the drug is physically incorporated after the micro- or nano-capsule has been synthesized. One instance is the study published by Z. Liu et al. [63] who addressed the use of sufopropyl dex-tran ion-exchange microparticles as carriers for drug deliv-ery. These microparticles are hydrophilic polymers whose sulfopropyl group confers a net negative charge. They were loaded with the drug doxorubicin (Dox) to evaluate the anti-cancer activity of the drug released in vitro. Dox was incor-porated into the microparticles by incubating these with an aqueous solution at room temperature. With the loaded mi-croparticles, studies revealed that the drug release was de-pendent on the salt concentration of the medium.

Furthermore, there are many methods for incorporating drugs during polymer production but the three most com-monly used processes are phase separation, solvent evapora-tion and spray drying [64].

Solvent evaporation methods have been widely used to prepare polymers loaded with various drugs. Solvent evapo-ration methods can be classified into two types, single emul-sion or double emulsion, depending on the number of emul-sions produced during the preparation of the drug-polymer.

The single emulsion method (o/o or o/w) involves mixing a solution of the polymer dissolved in organic solvent with the drug. This solution is added on a continuous phase (im-miscible with the first), which can be mineral oil (o/o) or an aqueous solution (o/w) and that contains an emulsifier whose function is to stabilize the emulsion produced. This emulsion is stirred or sonicated and the organic solvent is then re-moved by solvent evaporation or extraction. The coating mate-rial then shrinks around the core material, encapsulating it.


In double emulsion (w/o/w), an aqueous drug solution is first emulsified in a solution of the polymer in organic sol-vent and this emulsion (w/o) is added to an aqueous phase that contains an emulsifier, forming w/o/w emulsion. The organic solvent is then removed by extraction into an exter-nal aqueous phase and evaporated.

Solvent evaporation is a method used to encapsulate hy-drophobic (single emulsion) or hydrophilic (double emul-sion) drugs in hydrophilic polymers such as PLA and PLGA.

One example of the use of this method is the study by Jalón et al. [65] who synthesized PLGA microparticles loaded with rhodamine using solvent evaporation techniques for topical drug delivery. Specifically, these microparticles were prepared by the o/w solvent evaporation technique.

Wada et al. [66] developed a new approach for the prepa-ration of biodegradable lactic acid oligomer microspheres. These were obtained by the solvent evaporation method o/o (oil in oil). The solvent used for the dispersed phase solution was a mixture of acetonitrile and water, while the continuous phase medium was cottonseed oil. Dox and insulin were en-capsulated in microspheres with an efficiency of 80-90%.

Iwata et al. [67] synthesized PLA/PLGA microspheres containing water-soluble drugs, using a multiple emulsion solvent evaporation technique. Some drugs cannot be encap-sulated with traditional solvent evaporation methods because of their high water solubility. This new method leads to greater drug loading efficiency into the polymer and to the formation of microparticles with a heterogeneous drug dis-tribution. They studied the drug release behavior of the mi-croparticles synthesized with the new method and compared it with that of the microparticles produced with the tradi-tional method. A burst of drug release was detected with both methods, but the microparticles synthesized with this new method showed a larger amount of drug released than corresponding conventional microparticles.

Phase separation and coacervation is a widely used method for the formation of microcapsules. Basically, this method consists of three steps. In the first, the drug is dis-persed in a solution of the coating polymer, which is in the liquid phase. Then, a phase separation of the polymer coat-ing, caused by a change in temperature or by the addition of salt or a non-compatible polymer with the first, creates coac-ervate droplets of the polymer which are absorbed on the surface of the drug. Finally, in the last step, these droplets of polymer are solidified and stabilized, forming microcapsules with the drug inside them (Fig. 1.13).

This method was used by Nihant et al., [68] who studied the microencapsulation of a protein by coacervation of poly(lactide-co-glycolide).

As discussed above, this coating technique proceeds through three steps. The first consists of a phase separation of the coating polyesters (dissolved in CH2Cl2) induced by silicone oil, in the second step adsorption of the coacervate droplets around the protein phase occurs, and finally, the third step involves the hardening of microparticles. The re-searchers explained that the viscosity of the coacervate, in direct connection with the CH2Cl2 content, determines the size distribution, surface morphology and internal porosity of the final microspheres. Specifically, internal porosity occurs when the coacervate droplets harden so fast that solvent droplets are retained within the polymer matrix. Thus, these solvent droplets will create the pores.

The spray drying technique is used in the pharmaceutical industry to encapsulate active ingredients but it is also used to wrap food, oils, etc.

This technique initially involves dissolving the hydro-phobic or hydrolytic polymer in a solvent. Core particles are then dispersed in the polymer solution and the mixture is sprayed into a hot chamber through a nozzle of a spray dry. Consequently, the polymer solidifies onto the core particles when the solvent evaporates, resulting in microspheres that are precipitated into the bottom collector (Fig. 1.14).

This technique is widely used for the synthesis of micro-spheres, such as those that contain chitosan and gelatin as polymer coating and budesonide as the encapsulated drug. This research was performed by Naikwade et al., who mixed a polymeric phase with a drug solution that contained the drug dissolved in methanol and water. This mixture was spray-dried to achieve microspheres [69].

2.2.3.1. Polymers used in the Formulation of Nano- and

Micro-spheres

Several polymers have been exploited for the formulation of capsule carrier systems.

2.2.3.1.1. Polyesters

Polyesters based on PLA, PGA, their copolymers PLGA and poly ( -caprolactone) (PCL) [70] have been extensively employed because of their biocompatibility and biodegrad-ability (Fig. 1.15).

Fig. (1.13). Method of coacervation.


Fig. (1.14). Spray drying technique Polyesters are a large group of polymers characterized by the presence of esters bonds in the main chain. Their interest as biomaterials resides in the fact that the ester groups are hydrolytically degradable. Among the most commonly used are PLA and PLGA, which have been approved by the FDA for the development of drug delivery systems and other bio-medical applications. These polyesters have both advantages and disadvantages when used for the encapsulation of thera-peutic macromolecules. Degradation by non-enzymatic hy-drolysis of PLA and PLGA may lead to an accumulation of acidic monomers, causing a decrease in local pH. Thus, when the encapsulated therapeutic agent is a protein it may denaturalize [71]. However, these two polyesters are the most widely used because of their non-toxic degradation products and adjustable degradation rate.

In the case of PCL polymers, their biodegradation is slower than that of PLGA. This slowness leads to an increased

half-life of encapsulated molecules, thereby allowing a sus-tained drug release over prolonged periods.

Many studies have used combinations of polymers [72] to modify surface properties and to change the speed of polymer degradation and thus the drug release rate.

Considerable attention has been given to micro- and nano-spheres as drug carriers. The synthesis, characterization and in vitro and in vivo behavior of these carriers, and there-fore the profile of the drug release, have been extensively studied.

Polyesters have been used for the encapsulation of many types of therapeutic agents.

• Cancer

Cytostatic drugs inhibit the disordered growth of cells, disrupting cell division and destroying the cells that are di-viding rapidly. However, their toxic effect is not confined to malignant cells as they also exert their action on rapidly pro-liferating tissues such as skin, mucous membranes, bone narrow and so on. Therefore, to reduce the side effects of these drugs, their encapsulation in controlled release poly-mers is of great interest.

Huo et al. [73] summarized the encapsulation of cis-platin, a potent anticancer agent, by using a biodegradable polymer such as PLGA. A solvent evaporation method (commented on earlier) was used to perform this encapsula-tion. Recently, another interesting work was published by Li et al. [74]. They proposed to reduce the toxicity of cisplatin by loading it onto Fe5 (Fe3+ doped hydroxyapatite at atomic ratio of Feadded/Caadded=5%) and/or encapsulating the latter or cysplatin alone into PLGA microspheres using oil in water single emulsion. Once the nanoparticles were obtained, studies about drug release profile and degradation behaviors were performed and the results demonstrated that Fe5/PLGA composite has a favorable drug release profile with a short initial burst release and a long release time.

Another relevant intercalating agent is doxorubicin, which blocks the synthesis and transcription of DNA and

Fig. (1.15). Polyesters.

O

O

O

O

O

O

O

O

O

O

O

O

O

O

O

O

nn x y

n

n

PGA PLAPLGA

PCL PHB


also inhibits the enzyme topoisomerase II. This therapeutic agent has been encapsulated in PLA microspheres in order to obtain lower drug release and sustained action. Experiments of these formulations achieved a significant reduction of side effects in mice [75]. Ike et al. [76] conducted a study ad-dressing the treatment of malignant pleural effusion with this type of formulation and observed that local administration of microspheres produced an low systemic drug concentration, therefore improving patient quality of life.

5-Fluorouracil is a potent anti-metabolite that is used in some cancer therapies. This drug acts on DNA synthesis and causes cell death. Nagarwal et al. [77] synthesized nano-spheres of PLA polymer encapsulating 5-flourouracil. These carriers were synthesized to topical ocular applications.

• Bacterial and Parasitic infections

Nanotechnology offers drug delivery systems that are easily endocytosed by phagocytic cells because of the nature of the polymer [78]. This idea is relevant when the target of the transported drug is the pathogen inside the infected phagocytic cell. Polymeric nanoparticles are efficient carriers and after their in vivo administration they are endocytosed by phagocytic cells, thereby releasing their drug load.

Therefore, this feature has been used in various studies such as those performed by Italia et al. and Espuelas et al. on the in vitro anti-leishmanial activity of biodegradable nanoparticles. Leishmaniasis is a disease caused by obligate intracellular parasites. To minimize the adverse effects asso-ciated with amphotericin B, it was encapsulated into biode-gradable nanoparticles of PLGA [79] and PCL [80] for re-lease at the intracellular level in macrophages. Another study on the treatment of the same disease performed by Nahar et al. [81] involved encapsulating amphotericin into PLGA particles with anchored mannose, which is a macrophage target.

Nanoparticles have been examined for the treatment of other intracellular bacterial infections, such as those caused by mycobacteria. These types of bacteria include pathogens that cause serious diseases in mammals, including tuberculo-sis and leprosy. In classical treatments of mycobacterial dis-eases, the drug reaches the macrophages infected with my-cobacteria in low amounts and sometimes it does not persist long enough to develop the desired antimycobacterial effect. Therefore, the treatment involves prolonged large systemic doses, which are associated with unwanted side effects. As a result, some studies have addressed the treatment of myco-bacterial diseases using drug carriers. For example, Pandey et al. [82] encapsulated the antibiotic streptomycin in nano-spheres of poly-lactide-co-glycolide (PLG) using the multi-ple emulsion technique. This formulation was administered orally to mice and an increase in the bioavailability of the encapsulated antibiotic was observed. Moreover, Suarez et al. [83] synthesized PLGA microspheres with rifampicin, encapsulated via the spray drying technique. These micro-spheres were administered into the lungs of an animal model and resulted in a reduction of bacteria in the inoculated ani-mals.

Other antibiotics, such as cefazolin [84] and nafcillin [85], have been encapsulated into biodegradable nano- and micro-spheres. The former was encapsulated into PLGA

microspheres and the latter into PLGA nanospheres. Their activity was evaluated with pathogenic staphylococcus bac-teria. Other drugs, such as ciprofloxacin [86] and rifampicin [87], have also been encapsulated into this kind of PLGA system.

• Peptide Hormones

Numerous studies have focused on polymer carriers that carry peptide hormones.

The administration of peptides and proteins for therapeu-tic purposes is a challenge because of the enzymatic barriers, drastic chemical conditions in parts of the digestive system (such as the stomach), and low absorption of these com-pounds in the gut. Thus, almost all drugs of a peptide nature are given frequently via parenteral route because of suscepti-bility to degradation by enzymes. Therefore, as a result of the short in vivo half-lives of proteins, multiple injections are required to achieve a desirable therapeutic effect.

The low bioavailability of peptide compounds has prompted the pharmaceutical industry to introduce new for-mulations to minimize the problems involved in administra-tion. In recent years, considerable efforts have been made towards the encapsulation of proteins and peptides into po-lymeric particles such as PLGA and PLG because these polymers maintain the integrity and activity of the peptide or protein.

Nutropin Depot® [88] is a polymeric formulation that is available in the market. This formulation consists of mi-cronized particles of rhGH (growth hormone) embedded in biocompatible and biodegradable PLG microspheres. This system is indicated for the long term treatment of growth failure caused by the lack of adequate endogenous GH secre-tion and it is administered by subcutaneous injection. After administration, bioactive rhGH is released from the micro-spheres into the subcutaneous environment initially by diffu-sion and then by both polymer degradation and diffusion. Therefore, the in vivo profile is characterized by an initial rapid release followed by a slow decline in GH concentra-tion.

Recently, Rafi et al. [89] encapsulated rhGH into PLGA microparticles. The researchers studied the release behavior of the therapeutic agent and the results showed that the in vitro release of the hormone in porous particles is quicker than in non-porous particles; on the other hand, the hormone maintained a high average life in the body, as shown by the presence of bioactive rhGH in the serum of animals 30 days after one single injection.

Nafarelin, another therapeutic agent of interest, was en-capsulated into polymeric microspheres. Nafarelin is a potent hormone responsible for releasing gonadotropin. The study was conducted by Sanders et al. [90] who encapsulated the therapeutic agent into PLGA microspheres. Once these were administered as an injectable suspension, the drug was re-leased by a diffusional mechanism, which resulted in a rapid pharmacological response followed by an abrupt termination.

Another objective is to improve the bioavailability of peptides and proteins that are administered orally. For exam-ple, an overview of the work on PLGA formulations high-lights some studies devoted to enhancing the bioavailability


of hormones such as insulin [91]. However, most studies on this hormone are carried out with smart polymers, which will be discussed in the next section.

Other polymers, such as polyanhydrides, polyamides, poly(ortho esters) and natural or modified polysaccharides, are also used as drug carriers.

2.2.3.1.2. Polyanhydrides

Polyanhydrides are bioabsorbable [92] materials that are useful in drug delivery because they are biocompatible [93] and degraded in vivo into their non-toxic diacid counterparts, which are easily removed from the body.

The structure of polymers is characterized by repeated units linked by anhydride groups. These repeated units, called monomers, are diacids, which can have the same or a different chemical structure; in the latter case, they are re-ferred to as copolymers.

Most studies on polyanhydrides are based on sebacic acid (SA), p- (carboxyphenoxy)propane (CPP), p-(carboxyphenoxy)hexane (CHP) and their copolymers. Kumar et al. [94] reviewed the development, synthetic methods, structures and characterization of polyanhydrides (Fig. 1.16).

Depending on their chemical structure, the monomers confer the polymer differing degrees of hydrophilic behav-ior, thereby affecting their degradation. The more hydropho-bic the monomer, the more stable the anhydride bond is to hydrolysis. Also, by changing the ratio of monomers in the copolymers, their physical properties can be altered, thus polymer degradation and drug release behavior can be modi-fied.

In general, polyanhydrides are hydrolytically unstable polymers. The degradation of microspheres is pH-sensitive, being enhanced at high pH and becoming more stable in acidic conditions. Studies of polyanhydride degradation were carried out by Domb et al. [95] who used the co-polymer poly[1,3- bis(p-carboxyphenoxypropane):sebacic acid] in a ratio of 20:80 [P(CPP:SA)] 20:80] for this purpose. These

authors used two types of radioactive labeling, one with [14C]SA and unlabelled CPP, and the other co-polymer with [14C]CPP and unlabelled SA implanted in the brain of rats. These studies showed the rate of degradation of the copoly-mer and the metabolic deposition of the degradation prod-ucts. Therefore, the copolymer is biodegradable and can be used for drug delivery, resulting in a sustained release as a result of polymer degradation over a period of time depend-ing on the composition of the polymer or co-polymer.

Many studies have addressed the influence of the phys-ico-chemical characteristics of polyanhydrides on drug re-lease. It has been demonstrated that aliphatic polyanhydrides are degraded within days or weeks while aromatic polyanhy-drides take months or even years. Kipper et al. [96] used microspheres of three bioerodible polyanhydrides, namely poly[1,6-bis (p-carboxyphenoxy) hexane] (poly (CHP)), poly(sebacic anhydride) (poly(SA)) and the copolymer poly(CHP-co- SA), to study the release of p-nitroaniline. The release profiles shown by the polymers differed. These dif-ferences are attributed to polymer erosion rates and to the different distribution of the drug in the polymer.

On the other hand, the release profile also depends on the polarity of the drug, and therefore, on its different solubility in aqueous solution. Berkland et al. [97] examined the re-lease of drugs such as rhodamine B, p-nitroaniline and pi-roxicom, each of these with different water solubility and encapsulated them into the surface of erodible polymer poly (sebacic anhydride) PSA. Researchers observed that each drug exhibited different release mechanisms and therefore distinct release profiles.

It is important to indicate that this kind of polymer shows high reactivity, thus limiting its application for the encapsu-lation of nucleophilic drugs. Polyanhydrides can react with drugs containing nucleophilic functional groups, such as free amino groups, especially when a high temperature is re-quired for encapsulation, thus limiting the type of drug that can be successfully incorporated into a polyanhydride ma-trix.

Fig. (1.16). Polyanhydrides.

O

O

O O

O

CPP

O

O

O

n

n

O

O

O

O

O

n

SA

CHP


A wide variety of proteins [98] have been incorporated into polyanhydride matrices to improve in vivo protein sta-bility.

Recent studies have addressed the encapsulation of a range of proteins, such as insulin [99], human serum albumin (HSA) [100], and bovine serum albumin labeled with fluo-rescein isothiocynate (BSA-FITC) [101] into polyanhydride microspheres.

Furthermore, antigens have been encapsulated. Thus, Carrillo-Conde et al. [102] encapsulated Yersinia pestis anti-gen into the co-polymers CHP with SA and co-polymer 1,8-bis (p-carboxyphenoxy) -3,6-dioxooctane (CPTEG) with CHP . They studied the effect of the anhydride monomers, which are the degradation products, on the protein structures (antigen). The results showed that the environment provided by the CPTEG monomer was crucial for the preservation of the structure and antigenicity of the protein. Tetanus Toxoid (TT) [103] has also been used as model antigen for encapsu-lation into biodegradable polyanhydride microspheres to modulate the immune response mechanism. The copolymer CHP and SA were used for this purpose. In vivo studies in mice demonstrated that microspheres provided a prolonged exposure to antigen for sufficient time to induce primary and secondary immune response.

2.2.3.1.3. Polyamides

Polyamides are polymers whose repeated units contain amide groups. Polyamides have some very attractive charac-teristics with respect to their use as biodegradable materials. First, these molecules have excellent mechanical properties afforded by the presence of amide groups and hydrogen bonds, and consequently display a highly polar behavior. Second, polyamides are metabolized into non-toxic products. However, the amide bond is hydrolyzed much more slowly than the ester linkages, so in general, the polyamide degrada-tion rate is too slow for their practical use as biodegradable polymers. Therefore, the chemical structure of some polyam-ides is modified to accelerate degradation. Acceleration can be achieved through the copolymerization of monomers with other monomers to introduce a hydrolysable bond in the main chain. Thus, the amide bonds are combined in different proportions with other groups that are more susceptible to degradation.

Polymers that contain ester and amide bonds in the back-bone chain are the most interesting class of polyamides for controlled release. Generally, polyamides have been used to deliver low molecular weight drugs. These polymers are generally hydrophilic and their degradation rates depend on the hydrophilicity of the amino acids [104]. They are me-tabolized to non-toxic products.

These polymers have been used to transport drugs encap-sulated into a micro- or nano-sphere. Moreover, because their main structure is formed by amino acids, these can have a side chain that can provide sites for the attachment of drugs or pendant groups and can thus be used to modify the phys-ico-chemical properties of the polymer or to transport the conjugated drug. For example, the amino acid lysine of the copolymer poly(lactic acid-co-lysine) (PLAL) provides an amino group that allows modifications of the PLAL systems. Capponetti et al. [105] synthesized and characterized biode-

gradable microparticles with a PLAL backbone. Further-more, in order to modify the physico-chemical properties of the polymer, pendant groups such as poly (L-lysine), poly (D, L, alanine) or poly (L-aspartic acid) were attached to the lysine residues on the backbone. Researchers studied the encapsulation and release of rhodamine B, a low molecular weight drug model, to demonstrate the capacity of micropar-ticles to serve as carriers in controlled drug delivery devices.

Hrkach et al. synthesized other types of copolymers such as poly(L-lactic acid-co-L-lysine), poly (L-lactic acid -co-L- aspartate) and poly(L-lactic acid-co-D, L -alanine) [106].

Furthermore, Vera et al. synthesized some biodegradable poly(ester amide) composed of SA, dodecanodiol and differ-ent ratios of the stereoisomers of L- and D-alanine. The mi-crospheres were loaded with dicofenac sodium salt, triclosan and clofazimine to study the release rate in vitro [107].

2.2.3.1.4. Poly (Ortho Esters)

Over the last 40 years, poly(ortho esters) have been de-veloped for drug delivery purposes. These esters comprise four families: POE I, POE II, POE III and POE IV [108]. These differ in the type of monomers used and the synthesis of the polymer, which affect the physico-chemical features of the polymer and thus drug release.

POE I polymers were prepared by transesterification re-action of diols and diethoxy-tetrahydrofuran. The drawback with this family is that degradation products are acids and these catalyze polymer degradation (Fig. 1.17).

Fig. (1.17). POE I. In contrast, POE II polymers were synthesized by addi-tion of diol or polyol to diketene acetal. Cross-linked struc-tures (hydrogels) can be obtained when the polymer is syn-thesized from a polyol with diketene acetal (Fig. 1.18).

Fig. (1.18). POE II. POE II polymers were developed in order to prevent autocatalytic degradation because the degradation products are not acidic and therefore do not react with the ortho ester groups catalyzing their decomposition. These polymers were achieved by changing the initial acids to neutral products (diketene acetals). The problem in this case is that the degra-

O

O O R

n

O

O O

O

O O R

n


dation of these polymers is very slow due to their hydropho-bicity. Therefore, acid additives were added to the polymer matrix to accelerate degradation. Heller published a paper on the synthesis and use of various excipients to obtain con-trolled drug release of therapeutic agents that are physically dispersed in polymers. Heller [109] developed a polymeric system which is stable at physiological pH (7.4), but unstable at lower pH as a result of the excipients carried in the poly-mer. Therefore, polymer degradation can be controlled by these excipients and will take place only when the polymer and excipients are in contact with water, which causes a de-crease in pH. Then, if the polymer is highly hydrophobic, only excipients that are on the surface layer will be exposed to water and thus polymer degradation will occur only on the surface layer.

However, this approach proved only partially successful and POE II polymers were not developed for commercial use.

Later on, the transesterificacion reaction was used again, but in this case, using trimethylorthoacetate and 1,2,6 hex-anetriol, thus obtaining POE III polymers. However, the development of this family was abandoned due to the diffi-culty in obtaining polymers with a specific molecular weight (Fig. 1.19).

Fig. (1.19). POE III. Finally, the last family studied was POE IV [110], which consists of polymers synthesized from polyols and diketene acetals. However, in this case, latent acid monomers, such as fast hydrolysable glycolic or lactic acid, are introduced into the polymer backbone to obtain ester linkages. These link-ages are rapidly hydrolyzed and their acid reaction products can catalyze autocatalytic degradation (Fig. 1.20).

The amount of latent acid is small in order to maintain the hydrophobic nature of the polymer and it allows control over erosion rates. Therefore, POE IV polymers are auto-catalyzed compounds that contain latent acid in the back-bone, where erosion rates can be adjusted by varying the concentration of this acid. The synthesis of POE IV poly-mers is straightforward and reproducible and their structure

gives good mechanical and thermal properties, and allows control over the rate of degradation and therefore a regulated drug release rate. As a result of these features, this family of polymers has been commercialized.

A number of studies have addressed the encapsulation of various drugs in poly (ortho esters) and their release kinetics. Bai et al. [111] encapsulated bovine serum albumin into a range of self-catalyzed poly (ortho esters). The experimental results suggested that drug release rates can vary depending on the composition of the polymer. Also, the researchers reported that the more hydrophilic the polymer, the faster the drug release. This is because the release occurs through two mechanisms: surface erosion and diffusion of the therapeutic agent. On the other hand, Chia et al. [112] proposed that the degradation of the polymer and therefore the release of the drug, which is encapsulated into microspheres of auto-catalyzed poly (ortho esters), depends on the composition of the polymer (its hydrophobicity), its molecular weight and its morphology.

Microspheres of poly (ortho esters) have been designed for enhancing the in vivo efficacy of newly developed DNA vaccine. This DNA vaccine is a plasmid that contains the DNA fragment of interest and an expression vector. This DNA encodes a protein antigen of interest that induces acti-vation of the immune system. Therefore, it can induce a hu-moral response (antibody-mediated) and also cellular re-sponses (cytotoxic response mediated by Tc lymphocytes). Therefore, this vaccine works by inserting the DNA of bacte-ria or viruses into human or animal cells. In many cases, the DNA is encapsulated into polymers such as polymeric mi-crospheres [113] or liposomes in order to protect it from deg-radation. The composite polymer-DNA is phagocytosed by the cells and DNA is released into the cell as a result of the lower pH of the phagosome. Therefore, cells then begin to express proteins of foreign DNA and the immune system recognizes the protein and initiates a response. Wang et al. [114] described genetic vaccination, which consists of the encapsulation of a plasmid DNA into poly(ortho esters) mi-crospheres in order to protect DNA from degradation. The results of the tests performed in mice indicated an activation of adaptive immunity with humoral and cytotoxic responses, leading to the suppression of the growth of tumor cells, which had phagocytosed the DNA-polymer system and therefore expressed protein antigen.

Other therapeutic agents, such as 5-fluorouracil [115] (antineoplastic agent) and dehydroepiandrosterone [116] (steroid hormone), have been encapsulated into microspheres of poly(ortho esters). Their encapsulation and release has been studied.

Fig. (1.20). POE IV.

O

OR'

OR

n

O O

O

O

O

O CH

CH3

C

O

O R' O O

O

O

O

O R

mn


2.2.3.1.5. Natural or Modified Polysaccharides

Natural and modified polysaccharides are biodegradable and biocompatible polymers. Polysaccharide polymers differ in their structure and conformation. The polymers can be formed by several monomers or by various types of O-glycoside bonds. Therefore, there are polymers with the same monosaccharide that differ in their O-glycoside bond configuration, or isomers, while others differ in the hy-droxyl, which is part of the O-glycoside bond. Thus distinct configurational and constitutional isomers are obtained re-spectively.

These polymers also differ in the length of the chain and the number of branches in it.

Chitosan is an interesting polymeric carrier. Properties such as biodegradability, low toxicity and good biocompati-bility make it suitable for use in biomedical and pharmaceu-tical formulations. Its relevance is due to its cationic nature, which favors the encapsulation of compounds with opposite charge, therefore making it a good candidate as a gene car-rier. Cationic polymers condense DNA into a nanosized polymer by a self-assembly process, which consists of the electrostatic interaction of the positively charged polymer with the negatively charged DNA. These drug carriers may interact with the negatively charged cellular membrane, be-ing internalized via endocytosis. In the intracellular envi-ronment, the carriers are located in the endosomes. These gene carriers have to leave the endosome before it fuses with lysosomes and before exogenous DNA is degraded by lysosomal enzymes. Once out of the endosome, the gene carrier will be translocated to the nucleus if the free DNA is located near this structure.

Masotti et al. [117] studied the encapsulation of DNA into chitosan, obtaining nanoparticles of a definite size and shape. The release of DNA from nanoparticles showed an initial burst release within 3 hours of incubation and then, DNA was released constantly up to 72 hours. On the other hand, Özbas-Turan et al. [118] discussed the possibility of encapsulating two plasmids, pMK3 and pPGL2, containing

-galactosidase and luciferase as reporter genes, in the same microsphere structure as chitosan. Double plasmid-loaded chitosan microspheres were prepared by complex coacerva-tion to obtain high encapsulation efficiency. They conducted studies on the in vitro and in vivo release of the plasmid. The in vitro release of the plasmid was studied spectrophotomet-rically, observing continuous drug release. In contrast, in

vivo release was determined by the expression of the -galactosidase and luciferase proteins. After transfections, these microspheres showed a high production of these pro-teins.

Another key property of the polysaccharide chitosan is its mucoadhesivity; for this reason, it has been studied as a drug carrier in the pulmonary route [119].

2.2.4. Smart Polymers

Recent years have witnessed increasing attention to smart polymers as drug carriers because their properties can be tailored to fit the requirements required for drug delivery.

Hydrogels are a good choice for their use in controlled drug release because they are biocompatible and display sat-isfactory swelling properties in aqueous media.

The main interest in studying these materials is to control the swelling by specific stimuli that will allow the controlled release of bioactive molecules. Stimuli such as pH, electrical signals, and temperature can change the features of the sys-tem, such as pore size and permeability. Therefore, once this stimulus causes polymer swelling, the drug is released through a diffusion mechanism.

The volume of the hydrogel or the degree of swelling depends on the balance between the attractive and repulsive interactions present in the network. Thus, the combination of molecular interactions such as van der Waals forces, hydro-gen bonds, hydrophobic interactions and electrostatic inter-actions, determine the degree of hydrogel swelling at equilib-rium. Researchers modify these structural properties to de-sign these intelligent polymers.

The smart polymers that received most attention are those that respond to stimuli of temperature and pH because these are inherent variables of physiological systems.

2.2.4.1. pH-Sensitive Hydrogels

These polymers are characterized by the presence of acid groups, such as carboxylic acids or sulfonic acids, or also basic groups, such as amines, in their structures. These groups can accept or donate protons depending on the pH of the medium, leading to ionization variations in their struc-ture.

In the case of polymers [120] that contain acidic groups, when the pH of the medium rises, the degree of polymer ionization increases. In contrast, in the case of polymers that have amino groups in their structure, when the pH of the medium rises, the degree of polymer ionization decreases. This change in the degree of ionization causes an alteration in the swelling as a result of electrostatic repulsion forces between the charges present. Therefore, a way to decrease the repulsion between charges is to ensure polymer swelling, thereby maintaining the charges as far apart as possible.

pH-sensitive hydrogels have been used to develop con-trolled release formulations for oral administration. Lowman et al. [121] studied the use of a hydrogel that responds to a change in pH for the transport of orally administered insulin. This drug is a peptide labile to proteolytic degradation in the acidic stomach. Therefore, the researchers examined a drug carrier that would protect the peptide in this organ and allow its release in the intestine where the pH is slightly alkaline. This drug carrier was poly(methacrylic-g-ethylene glycol) (P(MMA-g-EG)) and the peptide was absorbed in the pre-formed polymer. Thus this pH-responsive carrier protected insulin in the acidic environment of the stomach as a result of the intermolecular interaction that prevented the hydrogel from swelling. But once the microparticles reached alkaline and neutral environments, namely the intestine, the interac-tions that occurred previously were lost and the pore size of the hydrogel increased, thus allowing insulin release.

pH-sensitive hydrogels have also been used to produce biosensors. Generally, hydrogels are loaded with enzymes that change the local pH of the microenvironment inside the


gels. For example, there is an enzyme that catalyzes the con-version of glucose to gluconic acid. The formation of glu-conic acid leads to a decrease in local pH, thereby resulting in a change in the swelling of the hydrogel. This mechanism is widely used for the release of insulin [122].

2.2.4.2. Temperature-Sensitive Hydrogels

In the last decade, phase transition in temperature-dependent hydrophilic hydrogels, has been studied. These phase transitions are manifested in significant changes in the volume of hydrogels.

Temperature sensitive hydrogels are classified into nega-tively thermosensitive, positively thermosensitive and ther-mally reversible gels. Some polymers increase their hydro-philicity as the temperature increases, leading to an expan-sion in their structure (positively thermosensitive) while oth-ers decrease their hydrophilicity when there is a temperature increase, leading to a collapse of the polymer and polymer shrinkage (negatively thermosensitive).

To obtain thermosensitivity properties, the polymer must have a correct balance of hydrophilic and hydrophobic chains in its structure. For instance, a polymer at room tem-perature in water forms hydrogen bonds; therefore, the polymer is swollen because its structure is hydrated. When the temperature increases, hydrogen bonds are weakened, while hydrophobic interactions are intensified and the poly-mer network shrinks. These observations lead to the conclu-sion that if the proportion of hydrophobic groups on the structure of the polymer rises, the temperature of transition between the shrunken to the swollen state increases.

Negative temperature-sensitive hydrogels have a lower critical solution temperature (LCST) and they show an on/off drug release with on at a low temperature and off at a high temperature, thus allowing pulsatile drug release. One exam-ple of these hydrogels are copolymers of N-isopropylacrylamide (PNIAAm) whose value of LCST is around 32 °C. However, a very useful temperature for bio-medical applications is close to that of body temperature, and therefore an adjustment of the LCST of PNIPAAm can be achieved by copolymerizing with other monomers which will change the degree of hydrophobicity and consequently the LCST value of the resulting polymer. Other polymers which can be used by delivery of therapeutic molecules are poly(N, N-dietylacrylamide) (PDEAAm) and poly(N-vinylcaprolactam) (PVCL).

In contrast, positive temperature-sensitive hydrogels have an upper critical solution temperature (UCST) and swell as a result of increased temperature. Some examples of these are poly(acrylic acid) (PAA), polyacrylamide (PAAm) and poly(acrylamode-co-butyl methacrylate).

Other temperature sensitive hydrogels are thermoreversi-ble. Some of these are Pluronics, Tetronics and poloxamer).

This type of polymer lends itself to many applications, one of the most important one being as drug carriers. Na et al. [123] conducted a study on the changes in the structure of a hydrogel versus temperature and how this change affects drug release. Specifically, they studied the changes in poros-ity on the surface of elastin polypeptide (ELP) microspheres. The transition temperature (Tt) of the polymer was deter-

mined by differential scanning calorimetry (DSC) and this coincided with a change in volume of the polymer. The mor-phology of the microparticles was examined by field emis-sion scanning electron microscopy. Finally, in drug delivery studies, it was observed that when the temperature was in-creased above the transition temperature, drug release was much faster. Therefore, when the temperature is higher than the transition temperature, the polymer structure, and thus the pores, change, resulting in drug release.

Temperature-sensitive hydrogel systems are widely stud-ied for use in body-temperature control. These systems de-pend on the patient’s temperature when the polymer is intro-duced. When the body temperature increases and exceeds a certain value, the hydrogel swells, leading to the release of the drug, whose function is to reduce the patient’s tempera-ture. Once the temperature decreases to normal values, the output of the drug from the polymer matrix decreases dra-matically as a result of the shrinkage of the polymer struc-ture.

Recent work by Curcio et al. [124] described the synthe-sis of a thermosensitive polymer based on hydrolyzed gela-tin. One of the interesting features of this polymer is its tran-sition temperature, which is slightly above body temperature. Diclofenac sodium salt, a non-steroidal anti-inflammatory, was encapsulated and used as a model drug to study the be-havior of these microspheres.

Another example of this application is found in a study by Yoon Ki Joung et al. [125], who encapsulated indometha-cin into temperature-sensitive polymers. This drug reduces fever, pain and swelling. The researchers encapsulated the therapeutic agent into PLGA microparticles, which were prepared by the oil-in water (O/W) emulsion method. Once the microparticles were synthesized, they were embedded into a chitosan-pluronic hydrogel matrix, a temperature-sensitive hydrogel. One objective of introducing a layer of chitosan-pluronic hydrogel into PLGA microparticles was to reduce the drug-release rate of microparticles synthesized exclusively with PLGA, while the other was to control this release by the stimulus of temperature.

2.2.5. Capsosomes

Capsosomes are polymeric structures that have a biode-gradable outer layer and contain liposomes, which compart-mentalize the interior of the capsule polymer.

These systems show permeability of the outer membrane, thereby allowing small molecules to cross by passive diffu-sion. Moreover, the liposomes in these systems divide the interior of the capsule into compartments, which can contain the same molecule or different ones. At the same time, these encapsulated molecules can be hydrophobic, and therefore will bind to the liposome membrane, or hydrophilic, which will be located in the aqueous core of liposomes.

Capsosome synthesis is complex. First, a structure called the core is required, on which the capsosome will be synthe-sized. A layer of a polymer called the precursor is added onto this core and the first layer of liposomes is immobilized on it. After adding this first layer of liposomes, a second polymeric layer called the separation layer is added, and onto this, a second layer of liposomes is added again. After add-


ing the last layer of liposomes, the outermost layer of the capsosomes is built, which consists of biodegradable poly-mers. Finally, the structure around which the core was syn-thesized is removed by degradation to obtain the capsosome structure.

One of the most studied subjects in capsosomes is the enzyme catalysis reaction. For this purpose, the enzymes were encapsulated into the different compartments (liposomes). The semi-permeable nature of the outermost layer of the capsosomes allows small substrates and products to diffuse inside these capsules. Furthermore, the lipidic membrane of the liposomes has a transition temperature phase that is normally moderately higher than the physio-logical temperature, so the structure of the lipidic membrane changes slightly at this temperature and facilitates the entry and exit of compounds into and from the liposome respec-tively. Therefore, encapsulated enzymes are accessible to their substrates; however, they cannot leave the liposomes because they are high molecular weight molecules and can-not diffuse through membranes.

Enzymes such as glucose oxidase [126], peroxidase, chymotrypsin [127], catalase [128] and urease [129] have been encapsulated into capsosomes and these systems have been reported to show catalytic activity. Kreft and colleagues [130] published a paper about the synthesis of a capsosome with two coupled enzymatic reactions, in which the two en-zymes were found in separate compartments. The catalyst system was composed of glucose oxidase and horseradish peroxidase.

Finally, in a study by Hosta [131] and colleagues, a model peptide (thiocoraline) was introduced into the mem-brane of liposomes. This was the first step towards the pos-sibility of incorporating transmembrane proteins that can act as specific transport proteins or as enzymes that must be at-tached to the membrane to retain their function.

Therefore, the development of systems with compartments and the success in the encapsulation of various enzymes or proteins with diverse functions illustrate the potential of these systems as drug carriers and also as microreactors.

2.2.6. Micelles

These copolymers are composed of individual linear polymers that contain a hydrophobic and hydrophilic seg-ment, which confers them the capacity to form micellar structures in aqueous media. These amphipathic polymers are arranged so that the hydrophobic part is located inside the structure, while the hydrophilic part is located outside, in contact with the aqueous medium. The block-copolymer micelles form spontaneously by self-assembly in water when the concentration of the amphiphilic block copolymer is above the critical micellar concentration (CMC) [132].

Micelles are used as drug carriers, in which low molecu-lar weight drugs are incorporated inside them, thereby im-proving therapeutic efficacy. Therefore, the hydrophobic core of polymeric micelles facilitates the incorporation of hydrophobic drugs either through covalent or non-covalent interactions.

Drug loading into micelles can be achieved several ways [133]. One approach is to add the drug and the copolymer in

a solvent miscible with water, and then to dialyze the mix-ture in aqueous medium. The drug-loaded polymeric mi-celles are formed as the solvent is replaced by water. An-other method is to introduce the drug into the polymer by a method of oil in water emulsion (o/w). First, an aqueous so-lution of copolymer micelles is prepared, and an o/w emul-sion is formed by adding the drug solution in chloroform. The copolymer emulsion stabilized o/w is stirred until the chloroform is evaporated and the micelles are loaded with the drug.

Micellar systems have several advantages as drug carri-ers. Among these is their capacity to encapsulate hydropho-bic drugs, thereby allowing circulation through the blood as a result of the high solubility and stability of the complex in physiological medium. Moreover, the drug is protected from metabolization by enzymes or other bioactive species of the physiological medium. Finally, the release rate is controlled by the stability of the micelles, the hydrophobic micellar core, and the links or interactions used to attach the drug to the polymer backbone.

Another advantage of micelles is their long half-life in the body because they cannot be removed via kidney filtra-tion cause of their large size. As a result of long-term circu-lation, the drug activity continues for a long period after a single injection. These polymeric structures are slowly ex-creted by the renal route as a result of their composition and structure. The micellar structure, which is formed by non-covalent intermolecular interactions between individual lin-ear polymers, is in equilibrium with free-form linear ones. Thus, the polymer can be removed as an individual linear polymer from the micelle structure by the kidney because the polymer chain molecular weight is lower than the critical value for kidney filtration.

The outer layer plays a crucial role in polymeric micelles. It is responsible for interactions with biostructures, such as proteins and cells, and this interaction determines the phar-macokinetics and biodistribution of the drug. The outer layer often consists of a polar poly(ethylene oxide) (PEO) that forms the shell of the carriers and protects its core.

Therefore, in vivo drug release is controlled by the seg-ments of the outer layer, which can improve the interaction with cells, together with the polarity of the micellar core, which controls the drug release rate, depending on the degree of hydrophobicity polymer core.

Polymer micelles have been studied for the delivery of poorly soluble and toxic chemotherapeutic agents to treat cancer. These micelles can reach tumors by a passive mechanism (EPR effect) or they can be have targeting moie-ties conjugated to reach tumors by an active mechanism for enhanced delivery to a specific site [134].

Some polymer micelle formulations of anticancer drugs have been reported to reach clinical trials.

Kataoka’s group developed a micelle carrier (NK 911) [135] for anticancer therapy. This is the first candidate of an antitumor drug on polymer micelles to have progressed to phase I clinical trials in Japan. NK911 is based on poly(ethylene oxide)-b-poly(aspartic acid) (PEO-b-PAsp) block copolymers conjugated with doxorubicin. PEO is be-lieved to form the outer shell of the micelles and the doxoru-


bicin conjugated poly(aspartic acid) chain is hydrophobic and forms the hydrophobic core of the micelles in aqueous media. The hydrophobic core enables NK911 to entrap a sufficient amount of doxorubicin. NK911 expresses higher in vivo antitumor activity than free doxorubicin because of the enhanced permeability and retention (EPR) effect. Parame-ters such as the maximum tolerated dose, dose limiting tox-icities, and the recommended dose for phase II were deter-mined.

Another micelle formulation in clinical trials is SP1049C [136]. This is an anticancer agent that contains doxorubicin and two nonionic pluronic block copolymers. Doxorubicin is mixed with these systems, resulting in a spontaneous incor-poration of the drug into micelles. In preclinical studies, SP1049C demonstrated increased efficacy compared to free doxorubicin. The toxicity profile, dose limiting toxicity, maximum tolerated dose and pharmacokinetic profile were determined in phase I.

SP1049C was evaluated in phase II studies in patients with metastatic adenocarcinoma of the esophagus and eso-phageal junction. This agent showed a better response than when single drugs such as cisplatin, paclitaxel and docetaxel were used. In 2008, Supratek Pharma obtained FDA clear-ance for an investigative application for SP1049C for the treatment of metastatic adenocarcinoma. This formulation has been assigned to a pivotal phase III study protocol under the Special Protocol Assessment process.

Finally, another biodegradable micelle, Genexol-PM, is in clinical trials [137]. Genexol is a micellar formulation characterized by a poly(ethylene glycol)-poly(D,L-lactide) copolymer that contains paclitaxel in the core. The co-polymer residue increases the water-solubility of paclitaxel and allows the delivery of higher doses than free paclitaxel. The efficacy and safety of Genexol-PM and cisplatin were evaluated in phase II for the treatment of advanced non-small-cell lung cancer [138].

2.3. Drug Carriers in which the Drug can be Embedded or Linked Covalently: Dendrimers

The chemical structure of a dendrimer consists of a core or central structure that contains several branches, which in turn branch off, leading to a three-dimensional globular structure of concentric layers. Therefore, the dendrimeric structure is characterized by layers between each focal point (or cascade point) called generations, that is, a generation 4 dendrimer is a system that has 4 focal points between the core and the surface. The nucleus is called generation zero because it does not have a focal point.

The dendrimeric structure contains two key features that make them of interest as drug carriers. One is its interior hollows, which have precise chemical characteristics and allow the encapsulation of drugs with hydrophobic proper-ties. Another important feature of the dendrimer is that it contains specific surface functional groups that allow at-tachment of hydrophilic drugs by covalent bonds, electro-static interactions or hydrogen bonds.

The literature cites a variety of polymer configurations (lipid, polysaccharide, peptide, etc.). These dendrimers can have distinct functional groups, such as polyamines

(dendrimer PPI), a mixture of amines and polyamides (den-drimer PAMAM) Fig. (1.21), and carboxylic acids. They may even hold more hydrophobic subunits such as poly(aryl ethers).

Dendritic structures can be synthesized by two strategies, divergent or convergent synthesis. The former consists of synthesizing the dendrimer from the nucleus, the starting point on which the first layer will be joined. This results in a first-generation polymer that will be joined by the next layer and so the dendrimer will grow successively generation to generation. In the convergent synthesis, the segments of the dendrimer are synthesized separately, and once obtained, they are coupled to the core. The convergent synthesis starts building the dendrimer from the surface and ends up in the nucleus. Finally, a structure is obtained which is formed by a nucleus surrounded by a protective shell, since it forms a multivalent surface with a high number of reactive sites. Therefore, as mentioned above, the interior of a dendrimer is an ideal place to encapsulate guest molecules; alternatively, they can be conjugated or interacted with the multivalent surface of the dendrimer, which contains a high number of functional groups.

Dendrimers are used for many treatments, including tu-mor [139], bacterial infection, antiviral, vaccine and gene therapy.

Synthetic nanoscale dendrimers are multifunctional molecules capable of interacting specifically with tumor cells and they cause apoptosis. These controlled release drugs allow the administration of lower concentration doses and a therefore achieve a decrease in toxicity to healthy cells.

One way of carrying dendrimers into malignant cells is through their conjugation with folic acid (FA) because the FA receptor is over-expressed in tumor cells. This is because FA is an essential ingredient for the replication of DNA and therefore, is required in large amounts for the rapid cell divi-sion feature of cancer. The conjugation of dendrimers with FA also facilitates their internalization via receptor-mediated endocytosis.

Quintana et al. [140] synthesized a generation 5 poly-amidoamine dendrimer. The PAMAM dendrimer was conju-gated with FA to obtain a specific focus, and with fluo-rescein isothiocynate (FITC), which is used as an imaging agent to track the localization of the dendrimer. And finally, it was also conjugated to methotrexate, a drug that inhibits FA metabolism. This drug specifically acts during DNA and RNA synthesis; its cytotoxic activity occurs during the S-phase of the cell cycle. Once the dendrimer had been synthe-sized and conjugated, in vitro studies were performed with the KB cell line, which corresponds to human epidermoid carcinoma, a cell line that over-expresses folate receptors. These tests were performed with various dendrimer models. Amino groups that are on the surface of a structure of the generation 5 PAMAM, namely fluorescein and folic acid (G5-FITC-FA), and had not been conjugated by FA or FITC were capped with chemical groups such as carboxy or 2,3-dihydroxypropyl or acetamide groups. Given the absence of repulsive forces from the charged amines, these new den-drimers showed a more relaxed structure than G5-FITC-FA. In contrast, G5-FITC-FA dendrimers bound poorly to KB


cells, showing minimal uptake, thereby indicating the lack of specific interaction with the receptor. This finding is ex-plained by the fact that some FA moieties are hidden inside the dendrimer, resulting in a reduced interaction with the receptor. In contrast, acetamide- G5-FITC-FA and 2,3-dihydroxypropyl- G5-FITC-FA showed high interaction and internalization in these cells. Therefore, these results demon-strated that the modification of the surface is crucial for im-proving the interaction of the dendrimer with the target cell.

Majoros et al. [141] explained the synthesis, the charac-terization and in vitro and in vivo studies of PAMAM den-drimers conjugated with FA, FITC and methotrexate. A par-tial acetylation of the amino groups of the dendrimer surface was carried out to neutralize the structure and to prevent un-desired reactions with molecules present in the body. The results of this study showed that dendrimers conjugated with FA inhibited the growth of KB cells carried out and showed a decrease in toxicity.

Kono et al. [142] synthesized a polyether dendritic com-pound containing hydrazide groups on the surface which were conjugated with folate residues to carry methotrexate to tumor cells.

Furthermore, dendrimers have been used as carriers for other anticancer drugs, such as cisplatin [143], BH3 peptides [144] (which induce apoptosis), and 5-fluoro-uracil [145], the latter conjugated by an acid-labile linker to the surface of the dendrimer.

A wide range of bacteria and yeast that cause human dis-eases are mediated by polyvalent interaction of carbohydrate and protein components with human cells.

The main advantage of dendrimers as drugs is the num-ber of potential sites for the attachment of functional surface groups that can grow exponentially generation to generation. The multiple surface groups on a single entity (multivalency or polyvalency) allow their association with pathogens. Den-drimer biocides that contain quaternary ammonium salts as functional end groups are an example. Quaternary ammo-nium compounds [146] are antimicrobial agents that disrupt bacterial membranes. PAMAM dendrimers have demon-strated significant antimicrobial activity because of their higher density of functional groups, which make the den-drimer more reactive to antimicrobial conjugation. The in-corporation of antibacterial agents, such as sulfamethoxazole [147], silver salts [148], nadifloxacin and prulifloxacin

Fig. (1.21). PAMAM dendrimer.

N

N

O NH

N

O

NH

NH2

O

NH

H2N

O

HN

N NH

NH2

HN

O

O

NH2

HN

N

O

HN

NH2

HN

H2N

OO

NH

N

O

NH

NH2

HN

H2N

O

O


[149], and niclosmide [150], has demonstrated significant antimicrobial activity.

Moreover, lysine-based peptide dendrimers [151] have shown antimicrobial activity against Gram-positive (Staphy-lococcus aureus) and Gram-negative (Escherichia coli) bac-teria as well as against fungal pathogens (Candida albicans).

Antiviral dendrimers attempt to mimic the anionic cell surface, so they are designed with surface anionic groups such as sulfonate residues or sialic acid, acidic carbohydrates that are present on the cell surface. Thus, the anionic den-drimer and/or the glycodendrimer compete with the cell sur-face by binding to the virus, leading to less likelihood of infection by the virus to the cell.

PAMAM dendrimers [152] with a surface that had been modified covalently with naphthyl-sulfonate residues, inhib-ited HIV attachment and fusion by binding on the surface of the virus, thus preventing its binding to healthy cells. They also suppressed the early stages of viral replication.

Furthermore, glycodendrimers, which are dendrimers that present multiple copies of carbohydrates on their surface, are widely used to address carbohydrate-protein interaction, con-ferring high selectivity to interact with the fusion protein of viruses. One example is the PAMAM dendrimer [153] con-jugated with sialic acid. These sialic acids interact with gly-coproteins localized on the surface of the influenza A virus, thus preventing viral adhesion to the sialic acid healthy cell surfaces. However, many viruses, such as influenza A virus and HIV, present a high degree of mutation, which leads to major changes in envelope protein glycosylation patterns when comparing different strains. This feature explains why it is so difficult to develop a vaccine against these viruses.

Dendrimers are also used in vaccines. Haptens are low molecular weight molecules that induce a weak immune re-sponse unless their molecular weight is increased by polym-erization or they are coupled to a high molecular weight car-rier. Dendrimers are an excellent tool for vaccines because of their high capacity to combine a large number of haptens on the cell surface, thus inducing an immune response.

MAP is the most widely used dendrimer for this purpose. This molecule has a multi-branched lysine core onto which haptens can be conjugated. MAPs have been developed as vaccines against cancer. Carbohydrate antigens, which are exposed on the surface of tumor cells, are potent targets for anticancer vaccines. Therefore, a multiple antigenic-linked dendrimer (MAG) that was synthesized carrying the Tn anti-gen linking to the carbohydrate tumor target, was developed. Hence, the dendrimer had the CD4+ T cell epitope to activate immune responses (lymphocytes T and B) specific for some types of carcinomas [154].

In other experiments, Ota et al. [155] showed that MAP structures are internalized by dendritic cells, which are anti-gen-presenting cells. Their results showed that MAP struc-tures are processed in the same way as an intracellular pathogen (cross presentation), leading to a peptide presented by a major histocompatibility complex (MHC) class I.

MAP structures have also been used to transport protein antigens against microorganisms such as Plasmodium falci-parum [156], which causes malaria in humans. To this end,

epitopes of the major surface protein in this parasite were conjugated to the dendrimer. The results of experiments in mice showed an immune response.

Finally, another interesting application of dendrimers is in gene therapy. Amino-terminated PAMAM or PPI are the dendrimers most commonly used as non-viral transfection agents of DNA to the nucleus. Kukowska-Latallo [157] syn-thesized several types of PAMAM dendrimers that differed in their core and number of generations. These were studied to improve transport and DNA transfection in a variety of mammalian cell lines. Various reporter genes were used to test the efficiency of transfection. In general, the results showed a highly efficient transfection in several cell lines with minimal cytotoxicity in eukaryotic cells.

The capacity of dendrimers to transfect cells depends on their size, shape and number of primary amine groups on the polymer surface. Transfection is higher in dendrimers with an excess of primary amines on their surface than with DNA phosphate groups, thus giving the complex an overall posi-tive charge and thereby support adherence of the complex to the negatively charged cell surface. Therefore, amino-terminated dendrimers carry DNA to the membrane and con-tribute to the transfection process by inducing the disruption of the membrane because, in theory, naked DNA, which is negatively charged, is repelled by the cell surface, which is also negatively charged.

Navarro et al. [158] have studied PAMAM dendrimers as carriers for cancer gene therapy. PAMAM is used to protect DNA from degradation by nucleases and to deliver genetic material to tumor cells.

Superfect [159] is a commercial transfection reagent con-sisting of activated dendrimer molecules that can carry larger amounts of genetic material than modified viruses for gene therapy. The structure of the dendrimer is formed by a cen-tral core and charged amino groups on the surface. These features allow the interaction of these charged groups with the negatively charged phosphate groups of nucleic acids and enhance their adherence to the cell surface. Thus the den-drimers are taken up into the cell by non-specific endocyto-sis.

Therefore, this new technology for gene transfer offers significant advantages, such as high gene transfer efficien-cies and minimal toxicity, and it can be used with a broad range of cell types.

3. CONCLUSION

In the treatment of a pathophysiological process, it is desirable that the drug reaches its site of action at a particular concentration. This therapeutic dose range must be constant over a sufficiently long period of time to alter the process.

The action of the drugs is limited by their degradation, interaction with other cells, and inability to penetrate tissues due to their chemical nature. Therefore, administration of relatively high doses is required to obtain the desired concen-tration at the site of action, thus leading to undesirable toxi-cological and immunological processes.

For these reasons, research efforts are devoted to design-ing new formulations, being of particular interest polymeric


systems of drug carriers, to achieve a higher desired pharma-cological response. These polymeric systems are suitable for site-specific and time-controlled delivery of drugs.

The treatment of most diseases requires drug distribution at a specific site. This can be achieved by functionalizing the surface of polymeric particles with receptor recognition ele-ments , such as antibodies and carbohydrates, or by linking the drug to the polymer through a specific linker that is de-graded under certain conditions, such as in an acidic environ-ment or by a specific enzyme. Other controlled-distribution systems use polymers, whose properties, such as swelling, change in response to an external stimulus such as pH.

Time-controlled drug delivery can be achieved depending on the features of the polymer. Modulation of the chemical architecture of the polymer can regulate the release of the therapeutic agent and thus the profile of the final formula-tion, thus obtaining the following: an extended release for-mulation (swelling of polymeric matrix or degradation of polymeric capsules); a sustained-release formulation, in which the active substance is released with a kinetic profile of order 0 (osmotic pumps); a pulsatile-release formulation, in which drugs are released only when required by the body; and a delayed-release formulation, in which the active ingre-dient is released at a time other than at the time of admini-stration (polymers sensitive to stimuli).

From a polymer chemistry perspective, it is important to appreciate that the mechanisms of controlled-release require polymers with a variety of physico-chemical properties. Sev-eral types of polymers have been tested as potential drug delivery systems, including nano- and micro-particles, den-drimers, nano- and micro-spheres, capsosomes and micelles. In these systems, drugs can be encapsulated or conjugated into polymer matrices.

Nano- and micro-particles are formulations in which the drug can be conjugated, trapped, or embedded within the polymer. Hydrogels such as PHEMA, and copolymers of PVA or PEG with acrylamides, are the most widely used. The three-dimensional structure of these hydrogels allows a controlled drug release.

Nano- and micro-capsules are comprised of an outer shell that encapsulates the active agent. Usually, this kind of pharmaceutical formulation is used to transport drugs that are rapidly degraded in body fluids. Thus the drug is encap-sulated inside the particle and it is released by the degrada-tion of the polymer coating. Several polymers have been exploited for this kind of formulation, such as polyesters, polyanhydrides, polyamides, poly(orthoesters) and polysac-charides.

Another formulation is capsosomes. These polymeric structures have a biodegradable outer layer and contain liposomes, which compartmentalize the interior of the cap-sule polymer. These systems show interesting features such as outer membrane permeability, which allows small mole-cules to cross the membrane by passive diffusion. Moroever, the division of the capsule into compartments by liposomes allows them to hold the same or distinct molecules.

Micelles are composed of individual linear polymers that contain a hydrophobic and a hydrophilic segment, which have the capacity to form micellar structures in aqueous me-

dia. The hydrophobic core of polymeric micelles facilitates the incorporation of hydrophobic drugs either through cova-lent or non-covalent interactions.

Finally, there is another promising formulation for drug delivery, dendrimers. The chemical structure of the den-drimer consists of a core or central structure that contains several branches, which in turn branch off, leading to a three-dimensional globular structure with concentric layers. The dendrimeric structure allows the encapsulation of drugs with hydrophobic properties in its interior cavities while functional groups found on their surface can attach hydro-philic drugs by covalent bonds, electrostatic interactions and hydrogen bonds.

These polymeric systems have been used for various purposes including treatments for antineoplastic activity, bacterial infection and inflammatory process as well as for vaccines.

Polymeric carriers of antineoplastic drugs can passively accumulate in cancerous tissues because of differences in the biochemical and physiological features of healthy and ma-lignant tissues (EPR effect) while they can actively accumu-late in the same tissues because they have been conjugated to targeting moieties. Furthermore, antineoplastic drugs are highly toxic and these delivery systems can prevent their degradation and interaction with other tissues.

These formulations are also used against bacterial infec-tions. Most intracellular infections are difficult to eradicate because bacteria inside phagosomes are protected from anti-biotics. Thus, treatments for intracellular bacterial infection are hindered because most antibiotics have poor intracellular diffusion and some have a reduced activity in the acidic con-ditions of lysosomes. Therefore, the treatment of the diseases caused by these bacteria call for drug carriers such as den-drimers, which are endocytosed by macrophages, thus allow-ing drug release within them.

Finally, another relevant application of drug delivery systems is vaccines. Vaccines are classified into two groups, namely proteins and nucleic acids. In both cases they require polymeric carriers because they are susceptible to degrada-tion by peptidases or nucleases. The DNA vaccine is a newly developed system. The DNA in this vaccine system encodes a protein antigen of interest that induces activation of the immune system. This DNA can be encapsulated into polym-eric carriers, thereby protecting it from degradation. It is then released into the phagosome, thus allowing it to reach the cell nucleus and express the foreign protein.

In summary, the chemical nature of the many types of polymeric drug carriers available facilitates the distribution and interaction of drugs with their target tissue. These carri-ers also protect their drug cargo from degradation and pre-vent their side effects.

CONFLICT OF INTEREST

None declared.

ACKNOWLEDGEMENT

None declared.


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