Polymers and Drug Delivery Systems

Polymers and Drug Delivery Systems 

Gemma Vilar a,b , Judit Tulla-Puche a,b and Fernando Albericio a,b,c,* 

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

a Institute for Research in Biomedicine, Barcelona Science Park, Baldiri Reixac 10, 08028-Barcelona, Spain; b CIBER- 

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

10, 08028-Barcelona, Spain; c Department 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 particular 

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 interaction 

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, several 

types of polymers have been tested as potential drug delivery systems, including nano- and micro-particles, dendrimers, 

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, neurodegenerative, 

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 depends 

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 achievement 

of a therapeutic response. 

In general, the nature of conventional therapeutics, especially 

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 Biomedicine, 

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 properties, 

such as the mechanism of action, pharmacological response, 

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. Excipients 

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 absorption, 

and therefore the pharmaceutical form of these substances 

affects drug bioavailability. 

In this regard, recent years have witnessed intense research 

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 finally, 

and possibly most importantly, these systems will facilitate 

more patient-friendly administration, thus resulting in 

patient increased compliance and satisfaction. 

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

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, micelles, 

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 carriers 

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 obtaining 

the desired release conditions. Therefore, the polymeric 

structure should be: i) biodegradable, because the chemical 

bonds that make up its chemical structure break; ii) disassemblable, 

because the various pieces forming the polymer 

disassemble but the chemical bonds do not break; and iii) 

undisassemblable, because the chemical bonds do not disassemble 

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 excreted 

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 structure 

of the molecule [1]. Molecular weight is especially relevant 

for substances that are not biodegradable, and macromolecules 

with a molecular weight lower than the glomerular 

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 modulated 

depending on the chemical structure of the polymer 

used. . 

If the polymer is not biodegradable, the drug can be covalently 

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 polymers. 

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 correct 

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 effectiveness 

of drug therapy [2]. These systems modify several parameters 

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 pharmacodynamics 

of the drug are modified by these formulations. 

Controlled release offers numerous advantages over conventional 

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 degradation. 

Many different kinds of drugs can benefit from distribution 

or time-controlled delivery, such as anti-inflammatory 

agents [3], antibiotics [4], chemotherapeutic drugs [5], immunosuppressants 

[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 release 

over time, this formulation allows the drug to reach a 

therapeutic concentration in tissues. 

The usual dose administration of a drug can achieve immediate 

(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 usually 

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 portion 

of the curve. Drug absorption ceases with time once the 

bioavailable dose has been absorbed. Therefore, the elimination 

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 therapeutic 

effect. In contrast, the maximum safe concentration is 

drug concentration in plasma above which toxic or side effects 

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 approach 

can lead to variations in API concentrations, reaching 

levels that are lower or higher than therapeutic concentrations, 

in other words, concentrations that are neither effective 

nor toxic, respectively. 

Fig. (1.2). Drug concentrations in plasma after delivery by conventional 

injection in a chronic disease. Representation of the therapeutic 

window between the minimal therapeutic concentration (therapeutic 

level, dotted line) and the maximum therapeutic concentration 

(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 polymeric 

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 particularly 

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 support 

wire whose function is to maintain arteries, blood vessels 

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 controls 

the release of the drug, thus preventing the arteries from 

reclosing. Nowadays, several DES platforms have been developed 

and evaluated for clinical use. The difference between 

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 drugrelease 

kinetics is obtained. San Juan et al. [14] functionalized 

a biocompatible polymer for stent coating for local arterial 

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. Consequently, 

the release of the drug can be modulated depending 

on the type of polymer used. 

There are some DES which have been approved for clinical 

use such as Cypher [15] and Taxus [16] containing sirolimus 

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 pharmaceutical 

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 concentration 

(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 immersion 

in the water, the osmotic system is hydrated and swollen, 

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 characteristics. 

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 represents 

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. 

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 release 

insulin in response to increased blood glucose levels 

have been designed. For the treatment of diabetes, a pulsatile-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 developed 

and studied by Sadaphal et al. [21]. This formulation is 

constituted by an impermeable anionic polymer called Eudragit 

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 intestine, 

thus preventing its delivery to the stomach. 

Fig. (1.7). Representation of the concentration of the active ingredient 

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 alternatives 

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 administration 

and thus have a short life, or those with a narrow therapeutic 

window. Some drugs are unstable in certain environments, 

such as in plasma or in acidic conditions. Therefore, 

in these conditions, it is of interest to use a polymer to protect 

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 formulation. 

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 overexpressed 

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 interact 

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 released 

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 certain 

substances, such as glucose, urea and morphine. However, 

these will be discussed further in the section on smart 

polymers (Table 1.1). 

Table 1.1. Some Examples of Smart Polymers 

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 environment 

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 contains 

proteolytic and hydrolytic enzymes. 

The introduction of macromolecules into cells by endocytosis 

[28] involves the invagination of the plasmatic membrane 

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 secondary 

lysosome. Therefore, in this environment within the 

lysosome, the linkers that are labile to acids or digestive enzymes 

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 lysosomal 

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 oligosaccharides. 

Linker libraries have been established by VectraMed, 

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 Specific 

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 

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 macromolecules 

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 polymers 

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, dendrimers, 

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). Nanoformulations 

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 covalently 

bound to the polymer backbone. In these types of 

drug-polymer conjugates, the goal is to improve the mechanism 

of cellular internalization and cell specificity to achieve 

optimal release of the drug at the proposed target. In contrast, 

the systems in which the drugs are embedded or encapsulated 

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 biodegradable 

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 controlled 

distribution). This type of release is used by nanoparticles, 

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 systems 

must have a certain particle size, surface charge and 

hydrophobicity in order to overcome physiological barriers. 

These barriers comprise biological structures or physiological 

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, prolonged 

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 accumulate 

passively in cancerous tissue as a result of certain 

differences in the biochemical and physiological characteristics 

of healthy and malignant tissue. In general, small molecules 

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 macromolecules 

to cross. Other features of malignant tissue include 

poor lymphatic drainage and increased vascularity 

(angiogenic process), which both lead to the retention of 

macromolecules. This passive accumulation of macromolecules 

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 phagocytosed 

(internalization without the mediation of receptor) 

by malignant tissues cells to form a vesicle called the endosome. 

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 secondary 

lysosome that contains hydrolytic and digestive enzymes 

that are active in acidic environments (lysosome pH: 

4). Some studies show overexpression of certain enzymes in


Fig. (1.10). Schematic representation of the anatomical and physiological characteristics of healthy and tumor tissue with respect to the vascular 

permeability and retention of small and large molecules (EPR effect). Adapted from reference 26. 

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 release 

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 clinical 

development stage [37]. 

Some of these compounds, such as HPMA [38] copolymer-doxorubicin 

galactosamine (PK2), HPMA copolymer 

platinate (AP 5346), carboxymethyldextran polyalcoholbound 

exatecan (DE-310), poly(L-glutamic acid)-paclitaxel 

(Xyotac), PEG-SN38 (EZN-2208), linear cyclodextrinbound 

camptothecin (IT-101) and poly(L-glutamic acid)camptothecin 

(CT-2106) have progressed further to the clinical 

phase. 

However, clinical studies for several other polymer conjugates, 

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 copolymerdoxorubicin 

(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 release 

through the action of lysozyme. Clinical studies 

showed that the toxicity of the conjugate was lower than 

doxorubicin alone. These studies also demonstrated prolonged 

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 targeting 

ligand, namely galactosamine, which specifically targets 

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 diaminocyclohexane 

(DACH)-platinum moiety coupled to a biocompatible 

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. Furthermore, 

DACH has activity against cisplatin-resistant cancer 

cells. Therefore, the DACH-platinum moiety was bound 

to HPMA via a pH-sensitive linker, and consequently released 

in the extracellular space of tumors and/or the intracellular 

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 passive 

process and then release the drug. Partial responses in 

antitumor activity in metastatic melanoma and ovarian cancer 

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 carrier. 

Used in cancer chemotherapy, exatecan is camptothecin 

analog, a cytotoxic quinoline alkaloid that inhibits the DNA 

enzyme topoisomerase I. Camptothecin binds to topoisomerase 

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 susceptible 

to hydrolysis and is inactive in its open form. In other 

words, the open form does not inhibit topoisomerase I. To


H 2 

C 

O NH 

HN 

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

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 several 

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 chemotherapy. 

This drug stimulates the assembly of microtubules 

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 characteristics 

of an ideal polymeric drug carrier, such as biodegradability, 

high drug loading, stability in circulation, and 

C 

H 2 

O 

O 

O NH 

Ph 

HN 

O 

O 

C 

H 2 

NH O NH 

R 

OH 

HN 

O 

O 

O O OH 

O 

OH 

NH 

Ph 

HN 

O 

O 

O 

NH 

OH 

O OH 

OH 

HPMA 

Linker 

Doxorubicine 

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 tolerated 

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 responded 

to taxol therapy. Phase III trials are currently underway 

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 structure, 

containing a lactone ring. Thus, SN38 has to be administrated 

with a formulation, such as by conjugation to a PEG 

polymer that protects the lactone ring and prevents its hydrolysis. 

The results in animal experiments showed a higher


half-life of the drug-polymer conjugate than the free drug, a 

higher drug exposure to tumor cells, and an increased solubility. 

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 composed 

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 favorable 

results prompted the initiation of a phase II trial designed 

to determine the side effects of this compound’s profile 

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 derivative 

of paclitaxel. It was studied with the aim to improve 

drug solubility with the subsequent controlled release of paclitaxel. 

Toxicity associated with paclitaxel, such as neurotoxicity, 

was observed in phase I trials; however, this drug 

showed antitumor activity. 

PNU 166148 [54] is composed by N-(2hydroxypropyl)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 Encapsulated 

Chemists working in synthesis seek to develop polymeric 

structures with physical and chemical properties that allow 

the specific release of the therapeutic agent under predetermined 

conditions. 

O 

O 

O 

O 

O O 

O 

O 

O 

O 

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 controlled 

release. Interactions between the polymer chains, 

which can be non-covalent (electrostatic and hydrogen 

bonds) or have covalent bonds (cross-linked polymers), confer 

stability to these structures. The integrity and the behavior 

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 polymeric 

structure and thus the release of the drug. 

Drug release can also be achieved through polymer degradation. 

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 degradation 

rate is constant and the non-degraded material (material 

inside the external biodegradable scaffold) maintains 

its chemical integrity during the process. Thus, the therapeutic 

agent is released once the material surface has been degraded 

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 regulate 

the release of the therapeutic agent and thus the profile 

of the final formulation (delayed or immediate release, pulsatile, 

etc.). However, the release occurs through a combination 

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]. 

N H 

O 

O O 

Polyanhydrides > Polycarbonates > Polyesters > Polyurethanes > Polyorthoesters > Polyamides 

O 

O 

O 

N 

H


(release by swelling control, by pure diffusion and by polymer 

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, microand 

nano-capsules, capsosomes and micelles. In the first 

case, the therapeutic agent is normally absorbed into the particles 

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 selfassembly 

in water and the drug is encapsulated in the hydrophobic 

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 hydrogen 

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 covalent 

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 involve 

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 affinity 

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 catalyzes 

the hydrolysis of neurotoxins, thus leading to an increase 

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 emission 

spectrum. Therefore, when the biosensor is incubated in 

an aqueous medium containing a neurotoxin such as paraoxon, 

at first the polymer swells, thus allowing diffusion of 

the water and toxin inside the polymer. Once the organophosphorus 

hydrolase enters into contact with toxin, the latter 

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 concentration 

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 embedded 

(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 polymer 

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 polymers 

for this kind of pharmaceutical formulation because 

they are non-toxic and their three-dimensional structure allows 

controlled drug release. 

Since hydrogels were first introduced into the field of 

biomedicine, they have consistently shown excellent characteristics 

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. Therefore, 

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 swelling, 

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 copolymers 

of PVA (Poly(vinyl alcohol)) or PEG (polyethylene 

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 backbone. 

Given the capacity of hydrogels to absorb large 

amounts of water and thus their capacity to swell, many studies 

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 prepared 

by free radical copolymerization of the macromonomers 

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 photoinitiator, 

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 concentrated 

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 crosslinked 

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 glutaraldehyde, 

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 polyethylene 

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 internalization 

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 conditions 

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 characteristics 

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 polymer 

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 pharmaceutical 

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 degradation 

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 formulation; 

however, example of micro- and nano-capsule synthesis 

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, monosaccharides, 

joined together by glycosidic bonds. Polysaccharide 

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, Llactic-co-glycolic 

acid) PLGA are among the most commonly 

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 concentrated 

drug solution with the preformed polymer carrier. 

The optimum incorporation strategy for efficient entrapment 

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 matrix, 

which is in turn related to the composition and molecular 

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 dextran 

ion-exchange microparticles as carriers for drug delivery. 

These microparticles are hydrophilic polymers whose 

sulfopropyl group confers a net negative charge. They were 

loaded with the drug doxorubicin (Dox) to evaluate the anticancer 

activity of the drug released in vitro. Dox was incorporated 

into the microparticles by incubating these with an 

aqueous solution at room temperature. With the loaded microparticles, 

studies revealed that the drug release was dependent 

on the salt concentration of the medium. 

Furthermore, there are many methods for incorporating 

drugs during polymer production but the three most commonly 

used processes are phase separation, solvent evaporation 

and spray drying [64]. 

Solvent evaporation methods have been widely used to 

prepare polymers loaded with various drugs. Solvent evaporation 

methods can be classified into two types, single emulsion 

or double emulsion, depending on the number of emulsions 

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 (immiscible 

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 removed 

by solvent evaporation or extraction. The coating material 

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 solvent 

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 external 

aqueous phase and evaporated. 

Solvent evaporation is a method used to encapsulate hydrophobic 

(single emulsion) or hydrophilic (double emulsion) 

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 preparation 

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 encapsulated 

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 encapsulated 

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 distribution. 

They studied the drug release behavior of the microparticles 

synthesized with the new method and compared 

it with that of the microparticles produced with the traditional 

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 dispersed 

in a solution of the coating polymer, which is in the 

liquid phase. Then, a phase separation of the polymer coating, 

caused by a change in temperature or by the addition of 

salt or a non-compatible polymer with the first, creates coacervate 

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). 

Fig. (1.13). Method of coacervation. 

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 researchers 

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 hydrophobic 

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 microspheres, 

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 biodegradability 

(Fig. 1.15).


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 biomedical 

applications. These polyesters have both advantages 

and disadvantages when used for the encapsulation of therapeutic 

macromolecules. Degradation by non-enzymatic hydrolysis 

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 

O 

O 

Fig. (1.15). Polyesters. 

PGA PLA 

n 

O 

O 

O 

half-life of encapsulated molecules, thereby allowing a sustained 

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 therefore 

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 dividing 

rapidly. However, their toxic effect is not confined to 

malignant cells as they also exert their action on rapidly proliferating 

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 polymers 

is of great interest. 

Huo et al. [73] summarized the encapsulation of cisplatin, 

a potent anticancer agent, by using a biodegradable 

polymer such as PLGA. A solvent evaporation method 

(commented on earlier) was used to perform this encapsulation. 

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 

O 

O 

n 

O 

O 

x y 

PCL PHB 

O 

n 

O 

O 

PLGA 

n 

O 

O 

O 

O


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 addressing 

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 nanospheres 

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 associated 

with amphotericin B, it was encapsulated into biodegradable 

nanoparticles of PLGA [79] and PCL [80] for release 

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 tuberculosis 

and leprosy. In classical treatments of mycobacterial diseases, 

the drug reaches the macrophages infected with mycobacteria 

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 mycobacterial 

diseases using drug carriers. For example, Pandey 

et al. [82] encapsulated the antibiotic streptomycin in nanospheres 

of poly-lactide-co-glycolide (PLG) using the multiple 

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 microspheres 

were administered into the lungs of an animal model 

and resulted in a reduction of bacteria in the inoculated animals. 

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 bacteria. 

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 therapeutic 

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 compounds 

in the gut. Thus, almost all drugs of a peptide nature 

are given frequently via parenteral route because of susceptibility 

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 formulations 

to minimize the problems involved in administration. 

In recent years, considerable efforts have been made 

towards the encapsulation of proteins and peptides into polymeric 

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 micronized 

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 secretion 

and it is administered by subcutaneous injection. After 

administration, bioactive rhGH is released from the microspheres 

into the subcutaneous environment initially by diffusion 

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 concentration. 

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 encapsulated 

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 released 

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 example, 

an overview of the work on PLGA formulations highlights 

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 referred 

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 behavior, 

thereby affecting their degradation. The more hydrophobic 

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 modified. 

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 

O 

Fig. (1.16). Polyanhydrides. 

O 

CPP 

O O 

O 

O 

O 

n 

O 

authors used two types of radioactive labeling, one with 

[ 14 C]SA and unlabelled CPP, and the other co-polymer with 

[ 14 C]CPP and unlabelled SA implanted in the brain of rats. 

These studies showed the rate of degradation of the copolymer 

and the metabolic deposition of the degradation products. 

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 depending 

on the composition of the polymer or co-polymer. 

Many studies have addressed the influence of the physico-chemical 

characteristics of polyanhydrides on drug release. 

It has been demonstrated that aliphatic polyanhydrides 

are degraded within days or weeks while aromatic polyanhydrides 

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 differences 

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 release 

of drugs such as rhodamine B, p-nitroaniline and piroxicom, 

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 encapsulation 

of nucleophilic drugs. Polyanhydrides can react with 

drugs containing nucleophilic functional groups, such as free 

amino groups, especially when a high temperature is required 

for encapsulation, thus limiting the type of drug that 

can be successfully incorporated into a polyanhydride matrix. 

CHP 

O 

O 

O 

SA 

O 

n 

O 

n


A wide variety of proteins [98] have been incorporated 

into polyanhydride matrices to improve in vivo protein stability. 

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 fluorescein 

isothiocynate (BSA-FITC) [101] into polyanhydride 

microspheres. 

Furthermore, antigens have been encapsulated. Thus, 

Carrillo-Conde et al. [102] encapsulated Yersinia pestis antigen 

into the co-polymers CHP with SA and co-polymer 1,8bis 

(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 encapsulation 

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 characteristics 

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 degradation 

rate is too slow for their practical use as biodegradable 

polymers. Therefore, the chemical structure of some polyamides 

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 backbone 

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 metabolized 

to non-toxic products. 

These polymers have been used to transport drugs encapsulated 

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 physico-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. Furthermore, 

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 microparticles 

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-Laspartate) 

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 different 

ratios of the stereoisomers of L- and D-alanine. The microspheres 

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 developed 

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 reaction 

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 addition 

of diol or polyol to diketene acetal. Cross-linked structures 

(hydrogels) can be obtained when the polymer is synthesized 

from a polyol with diketene acetal (Fig. 1.18). 

Fig. (1.18). POE II. 

O 

O O R 

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 O O R 

O 

n 

n


dation of these polymers is very slow due to their hydrophobicity. 

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 controlled 

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 polymer. 

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 decrease 

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 hexanetriol, 

thus obtaining POE III polymers. However, the 

development of this family was abandoned due to the difficulty 

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 linkages 

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 autocatalyzed 

compounds that contain latent acid in the backbone, 

where erosion rates can be adjusted by varying the 

concentration of this acid. The synthesis of POE IV polymers 

is straightforward and reproducible and their structure 

Fig. (1.20). POE IV. 

O 

O O 

R' 

O 

O 

O 

O 

O 

R 

n 

CH 3 

O CH 

C 

O 

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 autocatalyzed 

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 activation 

of the immune system. Therefore, it can induce a humoral 

response (antibody-mediated) and also cellular responses 

(cytotoxic response mediated by Tc lymphocytes). 

Therefore, this vaccine works by inserting the DNA of bacteria 

or viruses into human or animal cells. In many cases, the 

DNA is encapsulated into polymers such as polymeric microspheres 

[113] or liposomes in order to protect it from degradation. 

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) microspheres 

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. 

O 

O R' O O 

m 

O 

O 

O R 

n


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 Oglycoside 

bonds. Therefore, there are polymers with the 

same monosaccharide that differ in their O-glycoside bond 

configuration, or isomers, while others differ in the hydroxyl, 

which is part of the O-glycoside bond. Thus distinct 

configurational and constitutional isomers are obtained respectively. 

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 biocompatibility 

make it suitable for use in biomedical and pharmaceutical 

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 carrier. 

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, being 

internalized via endocytosis. In the intracellular environment, 

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 coacervation 

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 spectrophotometrically, 

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 proteins. 

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 satisfactory 

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 system, 

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, hydrogen 

bonds, hydrophobic interactions and electrostatic interactions, 

determine the degree of hydrogel swelling at equilibrium. 

Researchers modify these structural properties to design 

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 structure. 

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 controlled 

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 preformed 

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 interactions 

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 conversion 

of glucose to gluconic acid. The formation of gluconic 

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 temperaturedependent 

hydrophilic hydrogels, has been studied. These 

phase transitions are manifested in significant changes in the 

volume of hydrogels. 

Temperature sensitive hydrogels are classified into negatively 

thermosensitive, positively thermosensitive and thermally 

reversible gels. Some polymers increase their hydrophilicity 

as the temperature increases, leading to an expansion 

in their structure (positively thermosensitive) while others 

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 temperature 

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 polymer 

network shrinks. These observations lead to the conclusion 

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 example 

of these hydrogels are copolymers of Nisopropylacrylamide 

(PNIAAm) whose value of LCST is 

around 32 °C. However, a very useful temperature for biomedical 

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(Nvinylcaprolactam) 

(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 thermoreversible. 

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 porosity 

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 morphology 

of the microparticles was examined by field emission 

scanning electron microscopy. Finally, in drug delivery 

studies, it was observed that when the temperature was increased 

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 studied 

for use in body-temperature control. These systems depend 

on the patient’s temperature when the polymer is introduced. 

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 temperature. 

Once the temperature decreases to normal values, the 

output of the drug from the polymer matrix decreases dramatically 

as a result of the shrinkage of the polymer structure. 

Recent work by Curcio et al. [124] described the synthesis 

of a thermosensitive polymer based on hydrolyzed gelatin. 

One of the interesting features of this polymer is its transition 

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 behavior 

of these microspheres. 

Another example of this application is found in a study 

by Yoon Ki Joung et al. [125], who encapsulated indomethacin 

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 temperaturesensitive 

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 biodegradable 

outer layer and contain liposomes, which compartmentalize 

the interior of the capsule polymer. 

These systems show permeability of the outer membrane, 

thereby allowing small molecules to cross by passive diffusion. 

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 synthesized. 

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 polymers. 

Finally, the structure around which the core was synthesized 

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 physiological 

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 respectively. 

Therefore, encapsulated enzymes are accessible to 

their substrates; however, they cannot leave the liposomes 

because they are high molecular weight molecules and cannot 

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 enzymes 

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 membrane 

of liposomes. This was the first step towards the possibility 

of incorporating transmembrane proteins that can act 

as specific transport proteins or as enzymes that must be attached 

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 segment, 

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 molecular 

weight drugs are incorporated inside them, thereby improving 

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 mixture 

in aqueous medium. The drug-loaded polymeric micelles 

are formed as the solvent is replaced by water. Another 

method is to introduce the drug into the polymer by a 

method of oil in water emulsion (o/w). First, an aqueous solution 

of copolymer micelles is prepared, and an o/w emulsion 

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 carriers. 

Among these is their capacity to encapsulate hydrophobic 

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 filtration 

cause of their large size. As a result of long-term circulation, 

the drug activity continues for a long period after a 

single injection. These polymeric structures are slowly excreted 

by the renal route as a result of their composition and 

structure. The micellar structure, which is formed by noncovalent 

intermolecular interactions between individual linear 

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 pharmacokinetics 

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 segments 

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 moieties 

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 believed 

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. Parameters 

such as the maximum tolerated dose, dose limiting toxicities, 

and the recommended dose for phase II were determined. 

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 incorporation 

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 esophageal 

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 clearance 

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 copolymer 

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 nonsmall-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 properties. 

Another important feature of the dendrimer is that it 

contains specific surface functional groups that allow attachment 

of hydrophilic drugs by covalent bonds, electrostatic 

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 (dendrimer 

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 tumor 

[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 division 

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 polyamidoamine 

dendrimer. The PAMAM dendrimer was conjugated 

with FA to obtain a specific focus, and with fluorescein 

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 Sphase 

of the cell cycle. Once the dendrimer had been synthesized 

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 2,3dihydroxypropyl 

or acetamide groups. Given the absence of 

repulsive forces from the charged amines, these new dendrimers 

showed a more relaxed structure than G5-FITC-FA. 

In contrast, G5-FITC-FA dendrimers bound poorly to KB


H 2N 

H 2N 

H 

N 

Fig. (1.21). PAMAM dendrimer. 

O 

H 

N 

cells, showing minimal uptake, thereby indicating the lack of 

specific interaction with the receptor. This finding is explained 

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 2,3dihydroxypropyl- 

G5-FITC-FA showed high interaction and 

internalization in these cells. Therefore, these results demonstrated 

that the modification of the surface is crucial for improving 

the interaction of the dendrimer with the target cell. 

Majoros et al. [141] explained the synthesis, the characterization 

and in vitro and in vivo studies of PAMAM dendrimers 

conjugated with FA, FITC and methotrexate. A partial 

acetylation of the amino groups of the dendrimer surface 

was carried out to neutralize the structure and to prevent undesired 

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 compound 

containing hydrazide groups on the surface which 

were conjugated with folate residues to carry methotrexate to 

tumor cells. 

O 

O 

N 

O 

N 

NH 

NH 

HN 

HN 

NH 2 

N 

O 

O 

NH 2 

H 2N 

O 

O NH 

N 

NH 

NH 2 

HN 

HN 

N 

O 

O 

N 

H 

O 

N N H 

O 

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 diseases 

are mediated by polyvalent interaction of carbohydrate 

and protein components with human cells. 

The main advantage of dendrimers as drugs is the number 

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. Dendrimer 

biocides that contain quaternary ammonium salts as 

functional end groups are an example. Quaternary ammonium 

compounds [146] are antimicrobial agents that disrupt 

bacterial membranes. PAMAM dendrimers have demonstrated 

significant antimicrobial activity because of their 

higher density of functional groups, which make the dendrimer 

more reactive to antimicrobial conjugation. The incorporation 

of antibacterial agents, such as sulfamethoxazole 

[147], silver salts [148], nadifloxacin and prulifloxacin 

NH 2 

NH 2


[149], and niclosmide [150], has demonstrated significant 

antimicrobial activity. 

Moreover, lysine-based peptide dendrimers [151] have 

shown antimicrobial activity against Gram-positive (Staphylococcus 

aureus) and Gram-negative (Escherichia coli) bacteria 

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 dendrimer 

and/or the glycodendrimer compete with the cell surface 

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, inhibited 

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, conferring 

high selectivity to interact with the fusion protein of 

viruses. One example is the PAMAM dendrimer [153] conjugated 

with sialic acid. These sialic acids interact with glycoproteins 

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 response 

unless their molecular weight is increased by polymerization 

or they are coupled to a high molecular weight carrier. 

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 antigen 

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 antigen-presenting 

cells. Their results showed that MAP structures 

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 falciparum 

[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] synthesized 

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 positive 

charge and thereby support adherence of the complex to 

the negatively charged cell surface. Therefore, aminoterminated 

dendrimers carry DNA to the membrane and contribute 

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 consisting 

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 central 

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 dendrimers 

are taken up into the cell by non-specific endocytosis. 

Therefore, this new technology for gene transfer offers 

significant advantages, such as high gene transfer efficiencies 

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 concentration 

at the site of action, thus leading to undesirable toxicological 

and immunological processes. 

For these reasons, research efforts are devoted to designing 

new formulations, being of particular interest polymeric


systems of drug carriers, to achieve a higher desired pharmacological 

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 elements 

, such as antibodies and carbohydrates, or by linking 

the drug to the polymer through a specific linker that is degraded 

under certain conditions, such as in an acidic environment 

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 formulation, 

thus obtaining the following: an extended release formulation 

(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 ingredient 

is released at a time other than at the time of administration 

(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. Several 

types of polymers have been tested as potential drug 

delivery systems, including nano- and micro-particles, dendrimers, 

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 encapsulated 

inside the particle and it is released by the degradation 

of the polymer coating. Several polymers have been 

exploited for this kind of formulation, such as polyesters, 

polyanhydrides, polyamides, poly(orthoesters) and polysaccharides. 

Another formulation is capsosomes. These polymeric 

structures have a biodegradable outer layer and contain 

liposomes, which compartmentalize the interior of the capsule 

polymer. These systems show interesting features such 

as outer membrane permeability, which allows small molecules 

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 covalent 

or non-covalent interactions. 

Finally, there is another promising formulation for drug 

delivery, dendrimers. The chemical structure of the dendrimer 

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 hydrophilic 

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 malignant 

tissues (EPR effect) while they can actively accumulate 

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 infections. 

Most intracellular infections are difficult to eradicate 

because bacteria inside phagosomes are protected from antibiotics. 

Thus, treatments for intracellular bacterial infection 

are hindered because most antibiotics have poor intracellular 

diffusion and some have a reduced activity in the acidic conditions 

of lysosomes. Therefore, the treatment of the diseases 

caused by these bacteria call for drug carriers such as dendrimers, 

which are endocytosed by macrophages, thus allowing 

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 degradation 

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 polymeric 

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 carriers 

also protect their drug cargo from degradation and prevent 

their side effects. 

CONFLICT OF INTEREST 

None declared. 

ACKNOWLEDGEMENT 

None declared.


REFERENCES 

[1a] Seymour, L.; Miyamoto, Y.; Maeda, H.; Brerton, M.; Strohalm, J.; 

Ulbrich, K.; Duncan, R. Eur. J. Cancer., 1995, 31A, 766-770. (b) 

Seymour, L.W.; Duncan, R.; Strohalm, J.; Kopecek, J. J. Biomed. 

Mater. Res. 1987, 21, 1341-1358. (c) Gombotz, W.R.; Pettit, D.K. 

Bioconjug. Chem. 1995, 6, 332-351. 

[2a] Langer, R. Nature 1998, 392-395. (b) Brouwers, J. R. B. Pharm. 

World Sci. 1996, 18, 153-162. 

[3a] Wagenaar, B. W.; Müller, B. W. Biomaterials 1994, 15, 49-54. (b) 

Conforti, A.; Bertani, S.; Lussignoli, S.; Grigolini, L.; Terzi, M.; 

Lora, S.; Caliceti, P.; Marsilio, F.; Veronese, F. M. J. Pharm. 

Pharmacol. 1996, 48, 468-473. (c) Kalala, W.; Kinget, R.; Van den 

Mooter, G.; Samyn, C. Int. J. Pharm. 1996, 139, 187-195. 

[4] Schierholz, J. M.; Rump, A.; Pulverer, G. Drug Res. 1997, 47, 70-74. 

[5] Walter, K.A.; Tamargo, R.; Olivi, A.; Burger, P. C.; Brem, H. 

Neurosurgery. 1995, 37, 1129-1145. 

[6] Katayama, N.; Tanaka, R.; Ohno, Y.; Ueda, C.; Houjou, T.; Takada, 

K. Int. J. Pharm. 1995, 115, 87-93. 

[7] Maniar, M.; Domb, A.; Haffer, A.; Shah, J. J. Control. Release 

1994, 30, 233-239. 

[8] McGee, J.P.; Davis, S. S.; O’Hagan, D. T. J. Control. Release 

1994, 31, 55-60. 

[9] Wang, B., Siahaan, T. Soltero, R.A. In: Drug Delivery: principles 

and applications, Ed. Wiley, Inc., 2005. 

[10a] Sampath Kumar, K. P.; Bhowmik, D.; Chiranjib; Chandira, M.; 

Tripathi, K. K. J. Chem. Pharm. Res., 2010, 2, 349-360. b) Sáez, 

V.; Hernáez, E.; Angulo, L. S. Revista Iberoamericana de 

Polímeros, 2002, 3. c) Cho, X.; Wang, X.; Nie, S.; Chen, Z. G.; 

Shin, D. M. Clin. Cancer Res., 2008, 14, 1310-1316. 

[11] Michael E. A. In: Translation of Pharmaceutics: The Science of 

Dosage Form Design, 2e., 2002. 

[12a] http://pharmacycode.com/fda/Plendil_Er (Accessed November 

2011) (b) http://www.racgp.org.au/cmi/gwckapan.pdf (Accessed 

November 2011) 

[13] http://packageinserts.bms.com/pi/pi_sinemet_cr.pdf (Accessed 

November 2011) 

[14] San Juan, A.; Bala, M.; Hlawaty, H.; Portes, P.; Vranckx, R.; 

Feldman, L. J.; Letourneur, D. Biomacromolecules, 2009, 10, 

3074-80. 

[15] http://www.cordislabeling.com/pdf/5462872_4.pdf (Accesed December 

2011) http://www.cypherstent.com/Pages/index.aspx (Accesed 

December 2011) 

[16] http://www.ptca.org/articles/taxus_profileframe.html (Accesed 

December 2011) 

[17] Verma, R.K.; Mishra, B.; Garg, S. Drug Dev. Ind. Pharm. 2000, 

26, 695-708. 

[18] M. Anschütz, M.; Wonnemann, M.; Schug, B.; Toal, C.; Donath, 

F.; Pontius, A.; Pauli, K.; Brendel, E.; Blume, H. International J. 

Clin. Pharmacol. Thera., 2010, 48, 158-170. 

[19] Ravikumar Reddy J.; Veera, M.; Mohamed T. S.; Madhu C. S. J. 

Pharm. Sci. & Res. 2009, 1, 109-115. 

[20] Kost, J.; Horbett, T.A.; Ratner, B.D.; Singh, M. J. Biomed. Mater. 

Res. 1985, 19, 1117-1133. 

[21] Sadaphal, K.P.; Thakare, V.M.; Gandhi, B.R.; Tekade, B.W. International 

Journal of Drug Delivery, 2011, 3, 348-356. 

[22a] Rihova, B.; Kopecek, J. J. Controlled Release, 1985, 2, 289-310. (b) 

Seymour, L. W., Flanagan, P. A.; Al-Shamkhani, A.; Subr. V.; Ulbrich, 

K.; Cassidy, J.; Duncan, R. Sel. Cancer Ther., 1991, 7, 59-73. 

[23] Wedge, S.R.; Duncan, R.; Kopeckova, P. Br. J. Cancer, 1991, 63, 

546-549 

[24] Clegg, J. A.; Hudecz, F.; Mezo, G.; Pimm, M. V.; Szerkerke, M.; 

Baldwin, R. W. Bioconjug. Chem., 1990, 1, 425-430. 

[25a] Allen, T.M. Nat. Rev. Cancer, 2002, 2, 750-763. (b) Brumlik, M.J.; 

Daniel, B.J.; Waehler, R.; Curiel, D.T.; Giles, F.J.; Curiel, T.J. Expert 

Opin. Drug Deliv., 2008, 5, 87-103. (c) Miller, K.; Erez, R.; 

Segal, E.; Shabat, D.; Satchi-Fainaro, R. Angew. Chem. Int. Ed., 

2009, 48, 2949-2954. 

[26a] Qiu, Y.; Park, K. Adv. Drug Deliv. Rev., 2001, 53, 321-339 

(b)Gupta, P.; Vermani, K.; Garg, S. Drug Discov. Today, 2002, 7, 

569-579. 

[27] Brocchini, S.; Duncan, R. In: Polymer-drug conjugates: drug release 

from pendent linkers. In: Encyclopedia of controlled release. 

Mathiowitz E (Ed.). Wiley, NY, USA, 1999, 786-816. 

[28] Nori, A.; Kopecek, J. Adv. Drug Deliv. Rev. 2005, 57, 609-636. 

[29] Duncan, R. Nat. Rev. Drug Discov., 2003, 2, 347-360. 

[30] Ferruti, P.; Marchisio, M.A.; Duncan, R. Macromol. Rapid Comm., 

2002, 23, 332-355. 

[31] Ulbrich, K.; Subr, V. Adv. Drug Deliv. Rev., 2004, 23, 1023-1050. 

[32] Duncan, R. In: Targeting and intracellular delivery of drugs. In: 

Encyclopedia of Molecular Cell Biology and Molecular Medicine, 

Meyers, R.A., Ed., Weinheim, Germany.Wiley-VCH Verlag, 

GmbH & Co. KgaA., 2005, 163-204. 

[33] Uhrich, K. E.; Cannizzarro, S.M.; Langer, R.S.; Shakesheff, K. M. 

Chem. Rev. 1999, 99, 3181-3198. 

[34] Panyam, J., Sahooo, S. K., Prabha, S., Bargar, T., Labhasetwar, V. 

Int. J. Pharm., 2003, 262, 1-11. 

[35] Tong, R.; Cheng, J. Polymer reviews, 2007, 47, 345-381. 

[36] Matsumura, Y.; Maeda, H. Cancer Res. 1986, 46, 6387-6392. 

[37a] Vicent, M. J.; Duncan, R. Trends Biotechnol. 2006, 24, 39-47. 

(b)Haag, R.; Kratz, F. Angew. Chem. Int. Ed. Engl., 2006, 45, 

1198-1215. (c) Duncan, R.; Vicent, M. J. Adv. Drug Deliv. Rev., 

2010, 62, 272-282. 

[38] Lammers, T. Adv. Drug Deliv. Rev. 2010, 62, 203-230. 

[ 39] Thomson, A. H.; Vasey, P. A.; Murray, L. S.; Cassidy, J., Fraier, 

D., Frigerio, E., Twelves, C. J. Br. J. Cancer. 1999, 81, 99-107. 

[40a] Seymour, L. W.; Ferry, D. R.; Kerr, D. J.; Rea, D.; Whitlock, M.; 

Poyner, R.; Boivin, C.; Hesslewood, S.; Twelves, C.; Blackie, R.; 

Schatzlein, A.; Jodrell, D.; Bisset, D.; Calvert, H.; Lind, M.; Robbins, 

A.; Burtles, S.; Duncan, R.; Cassidy, J. Int. J. Oncol., 2009, 

34, 1629-1636. (b) Vasey, P.A.; Kaye, S.B.; Morrison, R. Clin. 

Cancer Res. 1999, 5, 83-94. 

[41a] Hopewel, J. W.; Duncan, R.; Wilding, D.; Chakrabarti, K. Hum. 

Exp. Toxicol., 2001, 20, 461-470. (b) Julyan, P. J.; Seymour, L. W.; 

Ferry, D. R.; Daryani, S.; Boivin, C. M.; Doran, J.; David, M.; Anderson, 

D.; Christodoulou, C.; Young, A. M.; Hesslewood, S.; 

Kerr, D. J. J. Control. Release, 1999, 57, 281-290. 

[42a] Campone, M.; Rademaker-Lakhai J. M.; Bennouna, J.; Howell, S. 

B.; Nowotnik, D. P.; Beijnen, J. H.; Schellens, J, H. Cancer Chemother. 

Pharmacol., 2007, 60, 523-33. (b) Rice, J. R.; Howell, S. 

B. Drugs Fut, 2004, 29, 561-566. (c) Kelland, L. Expert Opin. Investig. 

Drugs, 2007, 16, 1009-1021. 

[43] Nowotnik, D. P.; Cvitkovic, E. Adv. Drug Deliv. Rev. 2009, 61, 

1214-1219. 

[44] Shiose, Y.; Ochi, Y.; Kuga, H.; Yamashita, F.; Hashida, M. Biol. 

Pharm. Bull. 2007, 30, 2365-2370. 

[45] Soepenberg, O.; de Jone, M. J.; Sparreboom, A.; de Bruin, P.; 

Eskens, F. A.; de Heus, G.; Wanders, J.; Cheverton, P.; Ducharme, 

M. P.; Verweij, J. Clin. Cancer Res. 2005, 11, 703-711. 

[46a] Chipman, S.D.; Oldham, F.B.; Pezzoni, G.;, Singer, J.W. Int. J. 

Nanomed., 2006, 1, 375-383. (b) Shaffer, S.A.; Baker Lee, C.; 

Kumar, A.; Singer, J.W. Eur. J. Cancer, 2002, 38, 428. 

[47] Singer; J. W.; Shaffer, S.; Baker, B.; Bernareggi, A.; Stromatt, S.; 

Nienstedt, D.; Besman, M. Anticancer Drugs 2005, 16, 243-254. 

[48] Mita, M.; Mita, A.; Sarantopoulos, J.; Takimoto, C. H.; Rowinsky, E. 

K.: Romero, O.; Angiuli, P.; Allievi, C.; Eisenfeld, A.; Verschraegen, 

C. F. Cancer Chemother. Pharmacol., 2009, 64, 287-295. 

[49a] Singer, J. W.; Baker, B.; De Vries, P.; Kumar, A.; Shaffer, S.; 

Vawter, E.; Bolton, M.; Garzone, P. Adv.Exp. Med. Biol. 2003, 

519, 81-99. (b) Sabattini, P.; Brown, J.; Aghajanian, C.; Hensley, 

M. L.; Pezulli, S.; O’Flaherty, C.; Lovegren, M.; Funt, S.; Warner, 

M.; Mitchell, P.; Botton, M. G.; Spriggs, D.; Duggan, B. Proc. Am. 

Soc. Clin. Oncol., 2002, 21, (abst. 871). (c) Li, C.; Price, J. E.; Milas, 

L.; Hunter, N. R.; Ke, S.; Yu, D. F.; Charnsangavej, C.; 

Wallace, S. Clin. Cancer Res. 1999, 5, 891-897. 

[50a] Enzon Pharmaceuticals, Inc. www.enzon.com. (Accessed May 

2011) (b) Pastorino, F.; Loi, M.; Sapra, P.; Becherini, P.;Cilli, M.; 

Emionite, L.; Ribatti, D.; Greenberger, L. M.; Horak, I. D.; Ponzoni, 

M. Clin. Cancer Res. 2010, 16, 4809-4821. 

[51a] Davis, M. E. Adv. Drug Deliv. Rev. 2009, 61, 1189-92. (b) Schluep, 

T.; Hwang, J.; Cheng, J.J. Clin. Cancer Res. 2006, 12, 1606-1614 

[52] Homsi, J.; Simon, G. R.; Garrett, C. R.; Springett, G.; De Conti, R.; 

Chiappori, A. A.; Munster, P. N.; Burton, M. K.; Stromatt, S.; Allievi, 

C.; Angiuli, P.; Eisenfeld, A.; Sullivan, D. M.; Daud, A.I. 

Clin. Cancer Res., 2007, 13, 5855-5861. (b) Springett, G.M.; 

Takimoto, C.; McNamara, M. J. Clin. Oncol. 2004, 22, 3127. 

[53] Meerum Terwogt, J. M.; ten Bokkel Huinink, W. W.; Schellens, J. 

H.; Schot, M.; Mandjes, I. A.; Zurlo, M. G.; Rochetti, M.; Rosing, 

H.; Koopman, F. J.; Beijnen, J. H. Anticancer Drugs, 2001, 12, 

315-323. 

[54a] Bisset, D.; Cassidy, J.; de Bono, J. S.; Muirhead, F.; Main, M.; 

Robson, L.; Fraier, D.; Magnè, M. L.; Pellizzoni, C.; Porro, M. G.;


Spinelli, R.; Speed, W.; Twelves, C. Br. J. Cancer, 2004, 91,50-55. 

(b) Schoemaker, N. E.; Van Kesteren, C.; Rosing, H.; Jansen, S.; 

Swart, M.; Lieverst, J.; Fraier, D.; Breda, M.; Pellizzoni, C.; 

Spinelli, R.; Grazia Porro, M.; Beijnen, J. H.; Schellens, J. H.; ten 

Bokkel Huinink, W. W. Br. J. Cancer, 2002, 87, 608-14. (c) 

Duncan, R. Adv. Drug Deliv. Rev. 2009, 61, 1131-1148. 

[55] Uhrich, K. E.; Cannizzaro, S. M.; Langer, R. S.; Shakesheff, K. M. 

Chem. Rev., 1999, 99, 3181-3198. 

[56] Gombotz, W. R.; Pettit, D. K. Bioconjug. Chem., 1995, 6, 332-351. 

[57] Russell, J. R.; Pishko, M. V.; Simonian, A. L.; Wild, J. R. Anal. 

Chem. 1999, 71, 4909-4912. 

[58] Syu, M.J,; Chang, Y.S. Biosens Bioelectron., 2009, 24, 2671-2677 

[59] Peppas, N. A.; Keys, K. B.; Torres-Lugo, M.; Lowman, A. M. J. 

Control. Release., 1999, 62, 81-87. 

[60] Varshosaz, J.; Koopaie, N. Iranian Polymer Journal, 2002, 11, 

123-131. 

[61] Nakamura, K.; Maitani, Y.; Lowman, A. M.; Takayama, K.; Peppas, 

N. A.; Nagai, T. J. Control. Release 1999, 61, 329-335. 

[62a] Govender, T.; Riley, T.; Ehtezazi, T.; Garnett, M. C.; Stolnik, S.; 

Illum, L.; Davis, S. S. Int. J. Pharm. 2000, 199, 95-110. (b) 

Panyam, J.; Williams, D.; Dash, A.; Leslie-Pelecky, D.; Labthasetwar, 

V., J. Pharm. Sci. 2004, 93, 1804-1814. 

[63] Liu, Z.; Cheung, R.; Wu, X. Y.; Ballinger, J. R.; Bendayan, R.; 

Rauth, A. M. J. Control. Release, 2001, 77, 213-224. 

[64a] Sáez, V.; Hernáez, E.; Angulo, L. S.; Katime, I. Revista 

Iberoamericana de Polímeros, 2004, 5, 87-101. (b) Yeo, Y,; Baek,; 

Park, K. Biotechnol. Bioprocess Eng., 2001, 6, 213-230. 

[65] de Jalón, E. G.; Blanco-Príeto, M. J.; Ygartua, P.; Santoyo, S. Int. 

J. Pharm. , 2001, 226, 181-184. 

[66] Wada, R.; Hyron S.H.; Ikada, Y. J. Pharm. Sci. 1990, 79, 919-924. 

[67] Iwata, M.; McGinity, J. W. Pharm. Res., 1993, 10, 1219-1227. 

[68] Nihant, N.; Stassen, S.; Grandfils, C.; Jérome, R.; Teyssié, P.; 

Goffinet, G. Polymer International, 1994, 34, 289-299. 

[69] Naikwade, S.; Bajaj, A. Sci. Pharm., 2009, 77, 419-441. 

[70] Vaida, C.; Mela, P.; Kunna, K.; Sternberg, K.; Keul, H.; Möller, M. 

Macromol. Biosci., 2010, 10, 925-933. 

[71a] Bittner, B.; Ronneberger, B.; Zange, R.; Volland, C.; Anderson, 

J.M.; Kissel, T. J. Microencapsul., 1998, 15, 495-514. (b) Uhrich, 

K.E.; Cannizzaro, S.M.; Langer, R.S.; Shakesheff, K.M. Chem. 

Rev., 1999, 99, 3181-3198. 

[72a] Xiao, R. Z. ; Zeng, Z. W.; Zhou, G. L. ; Wang, J. J. ; Li, F. Z. ; 

Wang, A. M. Int. J. Nanomedicine, 2010, 5, 1057-1065. (b) Parajó, 

Y.; d’Angelo, I,; Horváth, A.; Vantus, T.; György, K.; Welle, A.; 

Garcia-Fuentes, M.; Alonso, M. J. Eur. J. Pharm. Sci., 2010, 41, 

644-649. 

[73] Huo, D.; Deng, S.; Li, L.; Ji, J. Int. J. Pharm., 2005, 289,63-67. 

[74] Li, Y.; Lim, S.; Ooi, C. P. Pharm. Res. 2011 (DOI 10.1007/s11095- 

011-0600-9). 

[75] Ike, O.; Shimizu, Y.; Ikada, Y.; Watanabe, S.; Natsume, T.; Wada, 

R.; Hyon, S-H.; Hitomi, S. Biomaterials 1991, 12, 757-762. 

[76a] Ike, O.; Shimizu, Y.; Hitomi, S.; Wada, R.; Ikada, Y. Chest, 1991, 

99, 911-915. (b) Ike, O.; Hitomi, S.; Wada, R.; Watanabe, S. Drug 

Delivery Systems, 1990, 5, 23-27. 

[77] Nagarwal, R. C.; Singh, P. N.; Kant, S.; Maiti, P.; Pandit, J. K. J. 

Biomed. Nanotechnol., 2010, 6, 648-657. 

[78] Sharma, G.; Valenta, D. T.; Altman, Y.; Harvey, S.; Xie, H.; Mitragotri, 

S.; Smith, J. W. J. Control Release, 2010, 147, 408-412. 

[79] Italia, J. L.; Yahya, M. M.; Singh, D.; Ravi Kumar, M. N . Pharm. 

Res., 2009, 26, 1324-1331. 

[80] Espuelas, M.S.; Legrand, P.; Loiseau, P.M.; Bories, C.; Barratt, G.; 

Irache, J.M. J. Drug Target., 2002, 10, 593-599. 

[81] Nahar, M.; Jain, N. K. Pharm. Res., 2009, 26, 2588-2598. 

[82] Pandey, R.; Khuller, G.K. Chemotherapy., 2007, 53, 437-441. 

[83] Suarez, S.; O’Hara, P.; Kazantseva, C.; Newcomer, E.; Hopfer, R.; 

Mc Murray, D. N.; Hickey, A. J. J. Antimicrob. Chemother., 2001, 

48, 431-434. 

[84] Chen, G.; He, C. X.; Xu, D. H.; Liu, L.; Zhang, F. J.; Xu, J. Y.; 

Zheng, C. H.; Gao, J. Q.; Yan, M. Pharmazie, 2009, 64, 284-286. 

[85] Pillai, R.R.; Somayaji, S.N.; Rabinovich, M.; Hudson, M.C.; Gonsalves, 

K.E. Biomed. Mater., 2008, 3, 034114. 

[86] Jeong, Y.I.; Na, H.S.; Seo, D.H.; Kim, D.G.; Lee, H.C.; Jang, 

M.K.; Na, S.K.; Roh, S.H.; Kim, S.I.; Nah, J.W. Int. J. Pharm., 

2008, 352, 317-323. 

[87] Esmaeili, F.; Hosseini-Nasr, M.; Rad-Malekshahi, M.; Samadi, N.; 

Atyabi, F.; Dinarvand, R. Nanomedicine., 2007, 3, 161-167. 

[88] Drugs information on line, Drugs.com. Web page: 

http://www.drugs.com/pro/nutropin-depot.html . (Accessed May 

2011) 

[89] Rafi, M.; Singh, S. M.; Kanchan, V.; Anish, C. K.; Panda, A. K. J. 

Microencapsul., 2010, 27, 552-560. 

[90] Sanders, L. M.; Mc Rae, G. I.; Vitale, K. M.; Kell, B. A. J. Control. 

Release, 1985, 2, 187-195. 

[91] Cui, F.D.; Tao, A. J.; Cun, D. M.; Zhang, L. Q.; Shi, K. J. Pharm. 

Sci., 2007, 96, 421-427. 

[92] Katti, D. S.; Lakshmi, S.; Langer, R.; Laurencin, C. T. Adv. Drug 

Deliv. Rev. 2002, 54, 933-961. 

[93] Laurencin, C.; Domb, A.; Morris, C.; Brown, V.; Chasin, M.; 

McConnell, R.; Lange N.; Langer, R. J. Biomed. Mater. Res., 1990, 

24, 1463-1481. 

[94] Kumar N.; Langer, S. R.; Domb, A. J. Adv. Drug Deliv. Rev., 2002, 

54, 889-910. 

[95] Domb, A. J.; Rock, M.; Schwartz, J.; Perkin, C.; Yipchuck, G.; 

Broxup, B.; Villemure, J. G. Biomaterials 1994, 15, 681-688. 

[96] Kipper, M. J.; Shen, E.; Determan, A.; Narasimhan, B. Biomaterials 

2002, 23, 4405-4412. 

[97] Berckland, C.; Kipper, M. J.; Narasimhan, B.; Kim, K. K.; Pack, D. 

W. J. Control. Release. 2004, 94, 129-41. 

[98] Petersen, L. K.; Sackett, C.K.;Narasimhan, B. J. Comb. Chem., 

2010, 12, 51-56. 

[99] Manoharan, C.; Singh, J. J. Pharm. Sci., 2009, 98(11), 4237-4250. 

[100] Sun, L.; Zhou, S.; Wang, W.; Su, Q.; Li, X.; Weng, J. J. Mater Sci: 

Mater Med. 2009, 10, 2035-2042. 

[101] Determan, A. S.; Trewyn, B. G.; Lin, V. S.; Nilsen-Hamilton, M. 

N.; Narasimhan, B. J. Control. Release, 2004, 100, 97-109. 

[102] Carrillo-Conde, B.; Schiltz, E.; Yu, J.; Chris Minion, F.; Phillips, 

G. J.; Wannemuehler, M. J.; Narasimhan, B. Acta Biomater., 2010, 

6, 3110-3119. 

[103] Kipper, M. J.; Wilson, J. H.; Wannemuehler, M. J.; Narasimhan, B. 

J. Biomed. Mater. Res. A. 2006, 76, 798-810 

[104a] Marck, K. W.; Wildevuur, C. H.; Sederel, W. L.; Bantjes, A.; Feijen, 

J. J. Biomed. Mater. Res. 1997, 11, 405-22. (b) Martin E. C.; May, P. 

D.; McMahon, W. A. J. Biomed. Mater. Res. 1971, 5, 53-62. 

[105] Caponetti, G.; Hrkach, J. S.; Kriwet, B.; Poh, M.; Lotan, N.; Colombo, 

P.; Langer, R. J. Pharm. Sci.. 1999, 88, 136-141. 

[106] Hrkach, J. S.; Ou, J.; Lotan, N.; Langer, R. Hydrogels and Biodegradable 

Polymers for Bioapplications. 1996, 8, 93-102. 

[107] Vera, M.; Puiggalí, J.; Coudane, J. J. Microencapsul., 2006, 23, 

686-697. 

[108a] Heller, J.; Barr, J.; Ng, S. Y.; Abdellavoi, K. S.; Gurny, R. Adv. 

Drug Deliv. Rev., 2002, 54(7), 1015-1039. (b) Heller, J.; Barr, J. 

Biomacromolecules, 2004, 5, 1625-1632. 

[109] Heller, J. Annals of the New York Academy of Sciences. Macromolecules 

as Drug and as Carriers for Biologically Active Materials. 

1985, 446, 51-66. 

[110] Heller, J.; Barr, J. Polymeric Drug Delivery II, 2006, 3, 29-43. 

[111a] Bai, X. L.; Yang Y. Y.; Chung, T. S.; Ng, S.; Heller, J. J. Applied 

Polymer Science. 2001, 80, 1630-42. (b) Bai, X. L.; Yang Y. Y.; 

Chung, T. S.; Ng, S.; Heller, J. Processing and fabrication of advanced 

materials, VIII. 2001, 177-184. 

[112] Chia, H. H.; Yang, Y. Y.; Chung, T. S.; Ng, S.; Heller, J. J. Control. 

Release 2001, 75, 11-25. 

[113] Nguyen, D. N.; Raghavan, S. S.; Tashima, L. M.; Lin, E. C.; Fredette, 

S. J.; Langer, R. S.; Wang, C. Biomaterials 2008, 29, 2783- 

2793. 

[114] Wang, C.; Ge, Q.; Ting, D.; Nguyen, D.; Shen, H. R.; Chen, J.; 

Eisen, H. N.; Heller, J.; Langer, R.; Putman, D. Nat. Mater. 2004, 

3, 190-196. 

[115] Lin, Y. H.; Vasavada, R. C. J. Microencapsul. 2000, 17, 1-11. 

[116] Bouchemal, K.; Briançon, S.; Chaumonz, B.; Fessi, H.; Zydowicz, 

N. J. Microencapsulation 2003, 20, 637-651. 

[117a] Masotti, A.; Bordi, F.; Ortaggi, G.;Marino, F.; Palocci, C. In: Modern 

Research and Educational Topics in Microscopy; Mendez-Vilas A., 

Díaz J., Eds.; 2007. (b) Masotti, A.; Bordi, F.; Ortaggi, G.; Marino, 

F.; Palocci, C. Nanotechnology 2008, 19, 055302/1-055302/6. 

[118] Özbas-Turan, S.; Aral, C.; Kabasakal, L.; Keyer-Uysal, M. Akbuga, 

J. J. Pharm. Pharm. Sci. 2003, 6, 27-32. 

[119] Grenha, A.; Al-Quadi, S.; Seijo, B.; Remunan-Lopez, C. J. Drug 

Deliv. Sci. Tech., 2010, 20, 33-43. 

[120] Lee, E.; Kim, K.; Choi, M.; Lee, Y.; Park, J. W.; Kim, B. Drug 

Deliv., 2010, 17, 573-580.


[121] Lowman, A. M.; Morishita, M.; Kajita, M.; Nagai, T.; Peppas, N. 

A. J. Pharm. Sci., 1999, 88, 933-937. 

[122a] Ito, Y.; Casolaro, M.; Kono, K.; Yukio, I. J. Control. Release, 

1989, 10, 195-203. (b) Ishihara, K.; Kobayashi, M.; Shinohara, I. 

Polymer J., 1984, 16, 625-631. 

[123] Na, K.; Jung, J.; Lee, J.; Hyun, J. Langmuir 2010, 26, 11165- 

11169. 

[124] Curcio, M.; Puoci, F.; Spizzirri, U. G.; Iemma, F.; Cirillo, G.; Parisi, 

O. I.; Picci, N. AAPS. PharmSciTech., 2010, 11, 652-662. 

[125] Joung, Y. K.; Choi, J. H.; Park, K. M.; Park, K. D. Biomed. Mater., 

2007, 2, 269-273 . 

[126a] Srivastava, R.; Brown, J. Q.; Zhu, H.; McShane, M. J. Macromol. 

Biosci., 2005, 5, 717-27.(b) Z.-l. Zhi, Z. I.; Haynie, D. T. Chem. 

Commun., 2006, 147-149. 

[127a] Petrov, A. I.; Volodkin, D. V.; Sukhorukov, G. B. Biotechnol. 

Prog. 2005, 21, 918-925. (b) Tiourina, O. P.; Antipov, A. A.; Sukhorukov, 

G. B.; Larionova, N. I.; Lvov, Y.; Mohwald, H. Macromol. 

Biosci., 2001, 1, 209-214. 

[128a] Caruso, F.; Trau, D.; Mohwald, H.; Renneberg, R. Langmuir, 2000, 

16, 1485-1488. (b)Wang, Y.; Caruso, F. Chem. Commun.(Camb), 

2004, 7, 1528-1529. 

[129a] Antipov, A.; Shchukin, D.; Fedutik, Y.; Zanaveskina, I.; 

Klechkovskaya, V.; Sukhoroukov, G.; Mohwald, H. Macromol. 

Rapid Commun., 2003, 24, 274-277. (b) Yu, A.; Gentle, I.; Lu, G.; 

Caruso, F. Chem. Commun. (Camb), 2006, 20, 2150-2152. 

[130a] Kreft, O.; Prevot, M.; Möhwald, H.; Sukhorukov, G. B. Angew. 

Chem. Int. Ed. Engl., 2007, 46, 5605-5608. (b) Mak, W. C.; Bai, J.; 

Chang, X. Y.; Trau, D. Langmuir, 2009, 25, 769-775. 

[131] Hosta, L.; Städler, B.; Yan, Y.; Nice, E. C.; Heath, J. K.; Albericio, 

F.; Caruso, F. Adv. Funct. Mater. 2010, 20, 59-66. 

[132] Kwon, G. S.; Naito, M.; Kataoka, K.; Yokoyama, M.; Sajurai, Y.; 

Okano, T. Colloids Surf. B., 1994, 2, 429-434. 

[133] Kwon, G.; Naito, M.; Yokohama, M.; Okano, T.; Sakurai, Y.; 

Kataoka, K. Pharm. Res., 1995, 12, 192-195. 

[134] Boddu, S. H. S.; Jwala, J.; Chowdhury, M. R.; Mitra, A. K. J. Ocul. 

Pharmacol. Ther., 2010, 26, 459-468. 

[135] Matsumura, Y.; Hamaguchi, T.; Ura, T.; Muro, K.; Yamada, Y.; 

Shimada, Y.; Shirao, K.; Okusaka, T.; Ueno, H.; Ikeda, M.; Watanabe, 

N. Br. J. Cancer 2004, 91, 1775-1781. 

[136] Danson, S.; Ferry,D.; Alakhov, V.; Margison, J.; Kerr, D.; Jowle, 

D.; Brampton, M.; Halbert, G.; Ranson, M. Br. J. Cancer 2004, 90, 

2085-2091. 

[137] Kim, T-Y.; Kim, D-W.; Chung, J-Y.; Shin, S. G.; Kim, S-C.; Heo, 

D. S.; Kim, N. K.; Bang, Y-J. Clin. Cancer Res. 2004, 10, 3708- 

3716. 

[138] Kim D-W.; Kim, S-Y.; Kim, H-K.; Kim, S-W.; Shin, S. W.; Kim, 

J. S.; Park, K.; Lee, M. Y.; Heo, D. S. Annals of Oncology., 2007, 

18, 2009-2014. 

[139] Baker, J. R. Jr. Hematololy Am Soc Hematol Educ Program., 2009, 

1, 708-719. 

[140] Quintana, A.; Raczka, E.; Piehler, L.; Lee, I.; Myc, A.; Mahoros, I.; 

Patri, A. K.; Thomas, T.; Mulé, J.; Baker, J. R. Pharm. Res., 2002, 

19, 1310-1316. 

Received: ??????????? Revised: ??????????? Accepted: ??????????? 

[141] Majoros, I. J.; Williams, C. R.; Becker, A.; Baker, J. R. Jr. Wiley 

Interdiscip Rev. Nanomed. Nanobiotechnol., 2009, 1, 502-510. 

[142] Kono, K.; Liu, M.; Fréchet, J. M. Bioconjug. Chem., 1999, 10, 

1115-1121. 

[143] Malik, N.; Evagorou, E. G.; Duncan, R. Anticancer Drugs, 1999, 

10, 767-776. 

[144] Myc, A.; Patri, A. K.; Baker, J. R. Jr. Biomacromolecules, 2007, 8, 

2986-2989. 

[145] Zhuo, R. X.; Du, B.; Lu, Z. R. J. Control. Release, 1999, 57, 249- 

257. 

[146] Chen C. Z.; Cooper, S. L. Biomaterials, 2002, 23, 3359-3368. 

[147] Abeylath, S. C.; Turos, E.; Dickey, S.; Lim, D. V. Bioorg. Med. 

Chem., 2008, 16, 2412- 2418. 

[148] Balogh, L.; Swanson, D. R.; Tomalia, D. A.; Hagnauer, G. L.; 

McManus, A. T. Nano Lett., 2001, 1, 18-21. 

[149] Cheng, Y.; Qu, H.; Ma, M.; Xu, Z.; Xu, P.; Fang, Y.; Xu, T. Eur. J. 

Med. Chem., 2007, 42, 1032-1038. 

[150] Devarakonda, B.; Hill, R. A.; Liebenberg, W.; Brits, M; de Villiers, 

M. M. Int. J. Pharm., 2005, 304, 193-209. 

[151a] Klajnert, B.; Janiszewska, J.; Urbanczyk-Lipkowska, Z.; 

Bryszewska, M.; Shcharbin, D.; Labieniec, M. Int. J. Pharm., 2006, 

309, 208-217. (b) Klajnert, B.; Janiszewska, J.; Urbanczyk- 

Lipkowska, Z.; Bryszewska, M.; Epand, R. M. Int. J. Pharm., 

2006, 327, 145-152. 

[152a] Witvrouw, M.; Fikkert, V.; Pluymers, W.; Matthews, B.; Mardel, 

K.; Schols, D.; Raff, J.; Debyser, Z.; De Clercq, E.; Holan, G.; 

Pannecouque, C. Mol. Pharmacol., 2000, 58, 1100-1108. (b) 

Witvrouw, M.; Weigold, H.; Pannecouque, C.; Schols, D.; De 

Clercq, E.; Holan, G. J. Med. Chem., 2000, 43, 778-783. 

[153] Landers, J. J.; Cao, Z.; Lee, I.; Piehler, L. T.;Myc P. P.; Myc, A.; 

Hamouda, T.; Galecki, A. T.; Baker, J. R. Jr. J. Infect. Dis., 2002, 

186, 1222-1230. 

[154a] Lo-Man, R.; Bay, S.; Vichier-Guerre, S.; Dériaud, E.; Cantacuzène, 

D.; Leclerc, C. Cancer Res., 1999, 59, 1520-1524. (b) Bay, S.; Lo- 

Man, R.; Osinaga, E.; Nakada, H.; Leclerc, C.; Cantacuzène, D. J. 

Pept. Res., 1997, 49, 620-625. 

[155] Ota, S.; Ono, T.; Morita, A.; Uenaka, A.; Harada, M.; Nakayama, 

E. Cancer Res., 2002, 62, 1471-1476. 

[156a] Moreno, C. A.; Rodriguez, R.; Oliveira, G. A.; Ferreira, V.; 

Nussenzweig, R. S.; Moya Castro, Z. R.; Calvo-Calle, J. M.; 

Nardin, E. Vaccine, 1999, 18, 89-99. (b) Joshi, M. B.; Gam, A. A.; 

Boykins, R. A.; Kumar, S.; Sacci, J.; Hoffman, S. L.; Nakhasi, H. 

L.; Kenney, R. T. Infect Immun., 2001, 61, 4884-4890. 

[157] Kukowska-Latallo, J. F.; Bielinska, A. U.; Johnson, J.; Spindler, 

R.; Tomalia, D. A.; Baker, J. R. Jr. Proc. Natl. Acad. Sci.U.S.A., 

1996, 93, 4897-4902. 

[158] Navarro, D. G.; Tros de Ilarduya, A. M. An. R. Acad. Farm. 2008, 

74, 229-256. 

[159] Tang, M.X.; Redemann, C.T.;Szoka, F.C. Jr. Bioconjug. Chem. 

1996, 7, 703-714.

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