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CHAPTER ONE

INTRODUCTION

1.0 General introduction

A key goal in pharmaceutical development of dosage forms is a good

understanding of the in vitro and in vivo performance of the dosage forms. The efforts to

improve drug effectiveness have led to developments in drug delivery technology. The

problems associated with systemic drug delivery include uneven bio-distribution

throughout the body, a lack of drug targeting specificity, the necessity of a large dose to

achieve high local concentration and adverse effects due to such high doses. There is now

a growing realization that innovative delivery of drugs would not only increase safety and

efficacy levels, but also improve the overall performance of the drug [1]. The therapeutic

benefits of new systems include increased efficacy of the drug site-specific delivery,

decreased toxicity/side effects, increased convenience, shorter hospitalizations, viable

treatments for previously incurable diseases, potential for prophylactic applications and

lower healthcare costs-both short and long term and better patient compliance.

Targeted drug delivery implies selective and effective localization of

pharmacologically active ingredients at pre-selected targets in therapeutic concentration,

while restricting its access to non-target areas, thus maximizing the effectiveness of the

drug. The carrier is one of the most important entities required for successful

transportation of the drug [2]. Colloidal drug delivery system is a rapidly developing area

that has contributed significantly to the progress in the field of controlled and targeted

drug delivery. Solid lipid particles have been proposed as a colloidal drug carrier

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therapeutic system for different administration routes such as oral, topical, ophthalmic,

subcutaneous and intramuscular injection, and particularly for parenteral administration.

Constant drug delivery is not always pharmacologically effective. Nearly, all

functions of the body show significant daily pattern variations, needing medical

treatments to need to be coordinated with those biological patterns. If the right drug can

be delivered at the right time, medical crises and side effects can be minimized and

eventually costs are lowered and compliance is improved. Generally, conventional

medicines are uniformly distributed to the whole body with the drug level in the blood

following a zig-zag profile. The drug level increases and decreases after each

administration. In combating bone cancerous tumors with chemotherapy, healthy cells

along with mutated ones are being eliminated, leaving the patient vulnerable to

infections. In controlled drug delivery systems designed for long-term targeted

administration, the drug level in the blood remains constant to an optimum for an

extended period of time.

Delivering drugs at specific delivery rates to a targeted organ can be achieved by

drug delivery systems or colloidal carriers [3]. They take different configurations such as

nanospheres, nanocapsules, microparticles, liquid crystals, reverse micelles, self-

assembly, microemulsions, macromolecular complexes and ceramic nanoparticles,

among others.

1.1 Drug delivery systems

There is increasing need to develop suitable drug carrier systems in order to

control, localise and improve drug delivery. Many different drug carriers can be used

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depending on the route of administration, the chosen drug properties and the intended

release profile.

1.1.1 Particulate drug delivery systems

Particulate drug carriers include microparticulate, nanocarriers, lipid based

carriers and colloidal carriers [2, 4, 5]. These are some umbrella terms under which

recently, many drug carrier systems were developed including: niosomes, dendrimers,

lipoplexes, pharmacosomes, nanocrystals, nanosuspensions, and ethosomes [6].

In recent years, the interest in micron and sub-micron systems in pharmacy has

surged. This is in part due to the advantages these systems provide over existing systems.

Designing a drug delivery system is challenging in terms of targeting the drug to specific

sites. Certain therapeutic agents that show success in vitro fail to produce the same effect

in the human body because of the limitation to target the designated area, which may

result in high concentrations being given to patients leading to intense side effects.

Dosage forms which conform themselves as surfactant spherical vesicles are often known

as vesicular systems. Typically, a colloid is a dispersion with particle size intermediate

between molecular range and coarse range [7]. Colloidal carriers are small particles of

100-400 nm in diameter, suspended in an aqueous solution. Micro, nano, vesicular,

colloidal and other lipid based carriers have the advantage of easy administration and

efficacy due to their long residence time and better targeting [2, 4].

1.1.2 Need for particulate drug delivery systems

Development of drug carriers as stated above is a novel area of science that

provides, with a new hope, the tools and technology to work at atomic, molecular and

supramolecular levels, leading to creation of devices and delivery systems with

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fundamentally new properties and functions. These carriers offer a number of advantages,

making them ideal drug delivery vehicles including:

1. Better drug delivery to certain stubborn or impermeable sites of the body.

2. Owing to their small size, chemistry and distribution these carriers have better bridged

the gaps between the structure and function of biomolecules.

3. Reaching the micron or nano range with these particles enables them to be highly

potential carriers for many biological molecules like proteins, DNA, viruses and

xenobiotics.

4. Better targeting to body tissues and sites where action is required, elimination of side

effects and adverse effects.

5. Owing to their size, nature and chemistry, these systems give better drug permeability

in biological membranes thus aiding in solubilization of some practically insoluble drugs

and solving bioavailability problems of many drugs.

6. They involve an overlap of biotechnology, nanotechnology, and information

technology, which might result in many important applications in life sciences including

areas of gene therapy, drug delivery, imaging, biomarkers, biosensors and novel drug

discovery techniques [8, 9, 10, 11].

7. They also offer an attractive solution for transformation of biosystems, and provide a

broad platform in several areas of bioscience [2, 12].

8. The surface properties of carriers can be modified for targeted drug delivery [13, 14]

for example small molecules, proteins, peptides, and nucleic acids loaded nanoparticles

are not recognized by immune system and efficiently targeted to particular tissue types

[2, 15].

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9. Targeted drug carriers reduce drug toxicity and provide more efficient drug

distribution [8, 16].

10. Drug carriers hold promise to deliver biotech drugs over various anatomic extremities

of the body such as blood brain barrier, branching pathways of the pulmonary system,

and the tight epithelial junctions of the skin etc [2, 17].

11. Drug carriers better penetrate tumors due to their leaky constitution, containing pores

ranging from 100—1000 nm in diameter.

1.1.3 Limitation of vesicular, colloidal, micro and nanocarriers

Drug delivery systems of fine particulate nature exhibit obvious difficulties in

preparation and handling. Among these limitations are:

1. Drug carriers exhibit difficulty in handling, storage, and administration because of

their susceptibility to aggregation.

2. They are unsuitable for drugs with low potency.

3. The key area of concern is related to their small size, as nanocarriers can gain access to

unintended environments with harmful consequences, example, they can cross the

nuclear envelope of a cell and cause unintended genetic damage and mutations [18].

1.2 Various carrier based dosage forms

1) Nanoparticles- Nanoparticles are roughly defined as submicron-sized colloidal

systems (varying in size from 10 to 1000 nm), biodegradable or not. Nanospheres have a

matrix like-structure, where active compounds can be firmly adsorbed at their surface,

entrapped or dissolved in the matrix. Nanocapsules have a polymeric shell and an inner

core. In this case, the active substances are not only dissolved in the core, but may also be

adsorbed at their surface [4, 5].

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Nanocapsules are sub-microscopic colloidal carrier systems composed of an oily

or an aqueous core surrounded by a thin polymer membrane. Two technologies can be

used to develop such nanocapsules: the interfacial polymerization of a monomer or the

interfacial nano-deposition of a preformed polymer. Solid lipid nanoparticles were

developed at the beginning of the 1990‘s as an alternative carrier system to emulsions,

liposomes, and polymeric nanoparticles [2, 4, 5]. They are used in particular in cosmetic

and pharmaceutical formulations. A novel nanoparticle-based drug carrier for

photodynamic therapy has been developed. This carrier can provide stable aqueous

dispersion of hydrophobic photo-sensitizers; yet preserve the key step of photo

generation of singlet oxygen, necessary for photodynamic action. Nanoparticles have also

found applications as non-viral gene delivery systems.

2) Solid lipid nanoparticles (SLNs) - SLNs made of solid lipids are submicron colloidal

carriers (50-1000 nm). These consist of a solid hydrophobic core having a monolayer of

phospholipids coating. The solid core contains drug dissolved or dispersed in the solid

high - melting fat matrix. The hydrophobic chains of phospholipids are embedded in the

fat matrix. Depending on the type and concentration of the lipid, 0.5 to 5% emulsifier

(surfactant) is added for the physical stabilization of the system. . Factors such as velocity

of lipid crystallization, lipid hydrophilicity, and influence of self-emulsifying properties

of the lipid on the shape of the lipid crystals (and hence the surface area) were found to

affect the final size of the SLN dispersions [4, 19].

3) Polymeric Nanoparticles-Colloidal carriers based on biodegradable and

biocompatible polymeric systems have largely influenced the controlled and targeted

drug delivery concept. Nanoparticles are sub-nanosized colloidal structures composed of

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synthetic or semi-synthetic polymers that vary in size from 10—1000 nm. Biodegradable

polymeric nanoparticles, typically consisting of polylactic acid (PLA), polyglycolic acid

(PGA), polylactic- glycolic acid (PLGA), and polymethyl methacrylate (PMMA) are

being investigated for the delivery of proteins, genes and DNA. Polymeric nanoparticle

suspensions have been prepared from inert polymer resins (Eudragit RS100, and RL100)

and loaded with drugs [6].

4) Ceramic Nanoparticles -These are the nanoparticles made up of inorganic (ceramic)

compounds such as silica, titania and alumina. Ceramic nanoparticles exist in size less

than 50 nm, which helps in evading the reticuloendothelial system (RES) of the body.

These particles provide the complete protection to the entrapped molecules such as

proteins, enzymes and drugs against the denaturizing effects of external pH and

temperature as they involve no swelling and porosity changes with change in pH (20).

5) Hydrogel Nanoparticles- Hydrogel nanoparticles form another polymeric system

involving the self-assembly and self aggregation of natural polymer amphiphiles such as

hydrophobized polysaccharides like cholesteroyl pullulan, cholesteroyl dextran and

agarose where cholesterol groups provide cross linking points in a non-covalent manner.

Cross-linked hydrogel nanoparticles (PVP-NP) (35—50 nm in diameter) composed of

natural polymers offer targeting to intracellular sites and good acceptability because of

higher water content [21, 22].

6) Copolymerized Peptide Nanoparticles - Another modification of a polymer-based

system is copolymerized peptide nanoparticles. It is a novel approach utilized for delivery

of therapeutic peptides as drug–polymer conjugates in which the drug moiety is

covalently bound to the carrier instead of being physically entrapped [23].

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7) Nanocrystals and Nanosuspensions - Nanocrystals are aggregates of around

hundreds or thousands of molecules that combine in a crystalline form, composed of pure

drug with only a thin coating comprised of surfactant or combination of surfactants. The

production technique of nanocrystals is known as ‗nanonisation‘. To produce

nanosuspensions, the drug powder is dispersed in an aqueous surfactant solution by high

speed stirring [4, 24].

Inorganic crystals that interfer with biological systems have recently attracted

widespread interest in biology and medicine [4]. Semiconductor nanocrystals, also known

as quantum dots (QDs), have become an indispensable tool in biomedical research,

especially for multiplexed, quantitative and long-term fluorescence imaging and detection

[25-28]. The basic rationale for using QDs arises from their unique and fascinating

optical properties that are not generally available for individual molecules or bulk

semiconductor solids. In comparison with conventional organic dyes and fluorescent

proteins, QDs have distinctive characteristics such as size-tunable light emission,

improved signal brightness, resistance against photobleaching and simultaneous

excitation of multiple fluorescence colors. Recent advances in nanoparticle surface

chemistry have led to the development of polymer-encapsulated probes that are highly

fluorescent and stable under complex biological conditions [29-31]. This new generation

of water-soluble QDs solved the problems of quantum yield decrease, chemical

sensitivity and short shelf-life previously encountered by the ligand exchange based-QD

solublization method [32]. As a result, these particles, linked with bio-affinity molecules,

have raised new opportunities for ultrasensitive and multicolor imaging of molecular

targets in living cells and animal models [33-35]. The success of using QDs in biological

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imaging, sensing and detection has encouraged scientists to further develop this

technology for clinical and translational research. One of the most important emerging

applications of QDs appears to be traceable drug delivery, because it has the potential to

elucidate the pharmacokinetics and pharmacodynamics of drug candidates and to provide

the design principles for drug carrier engineering. Due to concerns about long-term in

vivo toxicity and degradation, QDs are currently limited to cell and small animal uses.

Nevertheless, traceable delivery of therapeutics in cells and animals still has a big impact

on life science research, such as drug discovery, validation and delivery.

8) Nanotubes and Nanowires- Nanotubes and nanowires are the self-assembling sheet

of atoms arranged in the form of tubes and thread-like structures of nanoscale range.

Nanostructures that have gained much attention are hollow, carbon-based cage like

structures—nanotubes and fullerenes. Fullerenes are spherical structures, also known as

bucky balls. Soluble derivatives of fullerenes such as C60—a soccer ball shaped

arrangement of 60 carbon atoms per molecule show promise as pharmaceutical agents

[36].

9) Functionalized Nanocarriers - The combination of functionalities of biomolecules

and non-biologically derived molecular species used for special functions such as

markers for research in cell, molecular biology, biosensing, bioimaging and marking of

immunogenic moieties to targeted drug delivery are known as functionalized

nanoparticles. Organically, functionalized nanoparticles of catalytic active metals offer a

high surface area and unique size dependent chemical behavior. One approach is the

bioconjugate quantum dots as fluorescent biological labels. Quantum dots are crystalline

clumps of several hundred atoms with an insulating outer shell of a different material.

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Quantum dots can be attached to the biologicals such as cells, proteins and nucleic acids

[37].

10) Nanospheres –Nanospheres are solid metrical structures with drug molecules within

the matrices and/or adsorbed on the surfaces of the colloidal carriers [38].

11) Nanocapsules-Nanocapsules are small capsules with a central core surrounded by a

polymeric shell, where drug molecules may be dissolved in an oily core or adsorbed to a

surface interface [39].

12) Liposomes - Liposomes are microscopic vesicles composed of one or more

concentric lipid bilayers, separated by water or aqueous buffer compartments with a

diameter ranging from 25 nm to 100 μm. According to their size, liposomes are known as

small unilamellar vesicles (SUV) (10-100 nm) or large unilamellar vesicles (LUV) (100-

3000 nm). If more than one bilayer is present, then they are referred to as multilamellar

vesicles (MUV). Liposomes are formed when thin lipid films or lipid cakes are hydrated

and stacks of liquid crystalline bilayers become fluid and swell. During agitation,

hydrated lipid sheets detach and self associate to form vesicles, which prevent interaction

of water with the hydrocarbon core of the bilayer at the edges [2].

Liposomes consist of an outer uni - or multilamellar membrane and an inner liquid

core [2]. In most cases, liposomes are formed with natural or synthetic phospholipids

similar to those in cellular plasma membrane. Because of this similarity, they are easily

utilized by cells. Liposomes can be loaded with pharmaceutical or other ingredients

through two principal ways: Lipophilic substances can be associated with liposomal

membrane, and hydrophilic substances can be dissolved in the inner liquid core of

liposomes. To decrease uptake by the cells of the reticuloendothelial system and/or

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enhance their uptake by the targeted cells, the membrane of liposomes can be modified

by polymeric chains and/or targeting moieties or antibodies specific to the targeted cells.

Because they are relatively easy to prepare, biodegradable, and non- toxic, liposomes

have found numerous applications as drug delivery systems [40].

13) Lipid Emulsions (LEs) - Lipid emulsions are heterogeneous dispersions of two

immiscible liquids (oil-in-water or water-in oil) and they are prone to various instability

processes like aggregation, flocculation, coalescence and hence eventual phase separation

according to the second law of thermodynamics. LEs may be in the form of oil-in-water

(o/w), water-in-oil (w/o), micron, submicron and double or multiple emulsions (o/w/o

and w/o/w). The o/w type LEs colloidal drug carriers have various therapeutic

applications [2, 41].

14) Lipid Microtubules/Microcylinders- Lipid microtubules are a self organizing

system in which surfactants crystallize into tightly packed bilayers that spontaneously

form cylinders of less than 1 μm in diameter during a controlled cooling process [42].

15) Lipid Microbubbles- Lipid microbubbles consist of gas filled microspheres

stabilized by phospholipids, polymer or proteins and used as contrast enhancers in

ultrasonic diagnostics due to the low density and high elasticity of these bubbles. They

have few micron size ranges [43].

16) Lipospheres- Lipospheres were first reported by Domb (1995) [44], as water

dispersible solid micro particles with a particle size between 0.2-100 μm in diameter,

composed of solid hydrophobic fat core stabilized by a monolayer of phospholipid

molecules embedded in a microparticle surface. Lipospheres can contain a biologically

active agent in the core, in the phospholipids, or a combination of the two [45, 46].

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17) Lipopolyplexes- These are assemblies, which form spontaneously between nucleic

acids and polycations or cationic liposomes, and are used in transfection protocols. The

shape, size distribution, and transfection capability of these complexes depend on their

composition and charge ratio of nucleic acid to that of cationic lipid/polymer [47].

18) Ethosomes - Ethosomes are non-invasive delivery carriers that enable drugs to reach

the deep skin layers and/or the systemic circulation. Ethosomes contain phospholipids,

alcohol (ethanol and isopropyl alcohol) in relatively high concentration and water. Unlike

classical liposomes, ethosomes were shown to permeate through the stratum corneum

barrier and were reported to possess significantly higher transdermal flux in comparison

to liposomes. The synergistic effects of combination of phospholipids and high

concentration of ethanol in vesicular formulations have been suggested to be responsible

for deeper distribution and penetration in the skin lipid bilayers [48].

19) Multicomposite ultrathin capsules - The most important discovery in the field of

supramolecular science is the development of ―self-assembling ultrathin multilayered

capsules‖. Multicomposite ultrathin capsules are molecular assemblies of tailored

architecture having layer-by-layer adsorption of oppositely charged macromolecules onto

colloidal particles. Self-assembling ultrathin multilayered capsules (biomimic capsules)

are multilayer films of organic compounds on solid surface and these have been studied

for more than 60 years because they allow fabrication of multicomposite molecular

assemblies on tailored architecture. However, both the Langmuir-Blodgelt technique and

chemiosorption from solution can be used only with certain classes of molecules. An

alternative approach for fabrication of multilayers by consecutive adsorption of

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polyanions and polycations is far more general and has been extended to other materials

such as proteins or colloids [49].

20) Aquasomes - These are spherical 60 – 300 nm sized particles used for drug and

antigen delivery. The particle core is composed of non-crystalline calcium phosphate or

ceramic diamond, and is covered by a polyhydroxyl oligomeric film. Aquasomes are

prepared by self-assembling of hydroxyapatite by co-precipitation method and thereafter

preliminarily coated with polyhydroxyl oligomers (cellobiose and trehalose) and

subsequently adsorbed with bovine serum albumin (BSA) as a model antigen. BSA-

immobilized aquasomes were around 200 nm in diameter and spherical in shape and had

approximately 20-30 % BSA-loading efficiency ([50].

21) Pharmacosomes - This is the term used for pure drug vesicles formed by

amphiphilic drugs. Any drug possessing a free carboxyl group or an active hydrogen

atom (–OH, NH2) can be esterified (with or without a spacer group) to the hydroxyl

group of a lipid molecule, thus generating an amphiphilic prodrug. The amphiphilic

prodrug is converted to pharmacosomes on dilution with water [42].

22) Dendrimers - Dendrimers are macromolecular compounds that consist of a series of

branches around an inner core whose size and shape can be altered as desired. These

represent a unique class of polymers that are fabricated from monomers using either

convergent or divergent step growth polymerization. Dendrimers are made from Abn

type monomers, each layer or generation of branching unit doubling or tripling (n-2, n-3)

the number of peripheral functional groups. Generally, during dendrimer formation,

molecules emanate from a core and like a tree, they ramify with each subsequent

branching unit referred to as generation. Drug molecules can be loaded either in the

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interior, or can be adsorbed or attached to the surface groups. Hydrophilic dendrimers are

suitable as coating agents for protection and delivery of drugs to specific sites, thus

minimizing drug toxicity. The unique properties of dendrimers, such as their high degree

of branching, multivalency, globular architecture and well-defined molecular weight,

make them promising new scaffolds for drug delivery [51].

23) Colloidosomes- Colloidosomes are solid microcapsules formed by the self-assembly

of colloidal particles at the interface of emulsion droplets. ―Colloidosomes,‖ are hollow,

elastic shells whose permeability and elasticity can be precisely controlled [52].

24) Niosomes-Niosomes are non-ionic surfactant vesicles and, as liposomes, are

bilayered structures. Niosomes present low production cost, greater stability, and

resultant ease of storage. Niosomes are chemically stable, can entrap both lipophilic and

hydrophilic drugs either in aqueous layer or in vesicular membrane and present low

toxicity because of their non-ionic nature. Other advantages include flexibility in their

structural constitution, improvement of drug availability and controlled delivery at a

particular site, and, at last, niosomes are biocompatible, biodegradable and non-

immunogenic. Niosomes are present in a size range of 10 to 1000 nm. The colloidal drug-

loaded particles consist of macromolecular materials in which drugs are dissolved,

entrapped, encapsulated, and/or to which the drugs are adsorbed or attached [53].

25) Discomes - These are defined as non-ionic surface active agent-based discoidal

vesicles. The discomes are relatively large in size, 12-60 microns [54].

26) Proniosomes - These are dry formulations of surfactant-coated carrier, which can be

measured out as needed and rehydrated by brief agitation in hot water. Proniosomes (and

proliposomes) are normally made by spraying surfactant in organic solvent onto sorbitol

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powder and then evaporating the solvent. Because the sorbitol carrier is soluble in the

organic solvent, it is necessary to repeat the process until the desired surfactant loading

has been achieved. The surfactant coating on the carrier is very thin and hydration of this

coating allows multilamellar vesicles to form as the carrier dissolves [55, 56].

27) Microspheres- Microspheres or protein protocells are small spherical units, or

spherical particles composed of various natural and synthetic materials with diameters in

the micrometer range ([17, 57].

28) Microemulsions - Microemulsions are also termed ―transparent emulsions,‖

―miceller emulsions,‖ or ―swollen micellar emulsions.‖ Microemulsion is defined as any

multicomponent fluid made of water (or a saline solution), a hydrophobic liquid (oil), and

one or several surfactants resulting in systems that are stable, isotropic, and transparent

with low viscosity. Micro emulsions are thermodynamically stable colloidal dispersions

of water and oil stabilized by a surfactant and, in many cases, also a cosurfactant. Micro

emulsions offer an interesting and potentially quite powerful alternative carrier system

for drug delivery because of their high solubilization capacity, transparency,

thermodynamic stability, ease of preparation, and high diffusion and absorption rates

when compared to solvent without the surfactant system [58-60].

Microemulsions are excellent candidates as potential drug delivery systems

because of their improved drug solubilization, long shelf life, and ease of preparation and

administration. Three distinct microemulsions- oil external, water external, and middle

phase- can be used for drug delivery, depending upon the type of the dug and the site of

action [61]. In contrast to microparticles, which demonstrate distinct differences between

the outer shell and core, microemulsions are usually formed with more or less

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homogeneous particles. Microemulsions are used for controlled release and targeted

delivery of different pharmaceutical agents. For instance, microemulsions were used to

deliver oligonucleotides (small fragments of DNA) specifically to ovarian cancer cells

[62].

29) Polymeric micelles – These systems include amphiphilic block copolymers such as

Pluronics (polyoxyethylene-polyoxypropylene block copolymers) that self-associate in

aqueous solution to form micelles. Polymeric micelles offer a number of advantages in

terms of thermodynamic stability in physiological solution leading to their slow

dissolution in vivo. Because of their core–shell structure, these serve as suitable carrier

for water insoluble drugs; such drugs partition in the hydrophobic core of micelles and

outer hydrophilic layer aids in dispersion in aqueous media making it an appropriate

candidate for intravenous administration. Nanometric size range helps micelles to evade

the RES, and aids passage through endothelial cells [5, 63, 64].

30) Solid lipid microparticles (SLMs) - SLMs were developed recently and have so far

been considered a promising drug carrier system, especially with a view to giving the

incorporated active substance a sustained-release profile. Compared with liquid lipid

formulations, such as fat nanoemulsions, drug mobility is indeed lower in solid lipids

than in liquid oils. SLMs are in the micrometer size range and are composed of a lipidic

matrix that is in the solid state at room temperature. They seem to provide an alternative

drug carrier system to liposomes and polymeric nanoparticles. SLMs combine several of

those carriers‘ advantages while avoiding some of their disadvantages. The lipids used

are similar to physiological lipids, so toxicity is reduced [2]. SLMs are physicochemically

stable and can be produced relatively easily on a large industrial scale. In addition, raw

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materials and production costs are relatively low [2]. Their most important limitation is

that the drugs that have to be incorporated into SLMs must be lipophilic enough so as to

ensure high entrapment efficiency (EE). So far, SLMs have been studied for parenteral

and oral administration, and topical application [4, 65-70].

31. Colloidal based drug delivery systems

Colloids are extensively used for modifying the properties of pharmaceutical

agents. The most common property that is affected is the solubility of a drug. However,

colloidal forms of many drugs exhibit substantially different properties when compared

with traditional forms of the dosage forms. Another important pharmaceutical application

of colloids is their use as drug delivery systems. The most often used colloid- type

delivery systems include hydrogels, microspheres, liposomes, micelles, nanoparticles,

and nanocrystals.

a. Hydrogels

Hydrogel is a colloidal gel in which water is the dispersion medium. Natural and

synthetic hydrogels are used for wound healing, as scaffolds in tissue engineering, and as

sustained- release delivery systems. When used as scaffolds for tissue engineering,

hydrogels may contain human cells to stimulate tissue repair and since they are loaded

with pharmaceutical ingredients, hydrogels provide a sustained drug release.

Environmentally sensitive hydrogels have the ability to sense changes in the pH,

temperature, or the concentration of a specific metabolite and release their load as a result

of such a change; these hydrogels can be used as site specific controlled drug delivery

systems with mean particle diameter of 0.5-20 µm. Alginate, gelatin, chitosan, and other

polymeric hydrogels are some good examples. Light-sensitive, pressure-responsive, and

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electro-sensitive hydrogels also have the potential to be used in drug delivery. The most

important challenges that should be addressed in designing useful environmentally

sensitive hydrogels include slow response time, limited biocompatibility, and

biodegradability. However, if the achievements of the past can be extrapolated into the

future, it is likely that responsive hydrogels with a wide array of desirable properties will

be forthcoming [71].

b. Microparticles

Microparticles are small loaded microspheres of natural or synthetic polymers.

Microparticles were initially developed as carriers for vaccines and anti-cancer drugs.

More recently, novel properties of microparticles have been developed to increase the

efficiency of drug delivery and improve release profiles and drug targeting ([72]. Several

investigations have focused on the development of methods of reducing the uptake of the

microparticles by the cells of the reticuloendothelial system and enhance their uptake by

the targeted cells. The mean particle diameter has been shown to lie in the range of 0.2-5

µm, with polystyrene and polyactide microspheres as representative systems. Functional

surface coatings of non-biodegradable carboxylated polystyrene or biodegradable poly

(D, L- lactide-co-glycolide) microspheres with poly(L-lysine)-g-poly (ethylene glycol)

(PLL-g-PEG) were investigated in attempts to shield them from nonspecific phagocytosis

and to allow ligand- specific interactions via molecular recognition. It was found that

coatings of PLL-g-PEG-ligand conjugates provided for the specific targeting of

microspheres to human blood-derived macrophages and dendritic cells while reducing

non-specific phagocytosis. Microparticles can also be used to facilitate non-traditional

routes of drug administration. It has been found that microparticles can be used to

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improve immunization using the mucosal route of administration of therapeutics. It was

found in this study that mucosal route of administration of therapeutics can translocate to

tissues in the systemic compartment of the immune system and provoke immunological

reactions [73].

c. Nano-emulsions

In contrast to microemulsions, nanoemulsions consist of very fine oil-in-water

dispersions, having droplets diameter smaller than 100 nm [74]. Compared to

microemulsions, they are in a metastable state, and their structure depends on the history

of the system. Nanoemulsions are very fragile systems and can find applications in skin

care due to good sensorial properties (rapid penetration, merging textures) and their

biophysical properties (especially their hydrating power).

d. Micelles

Micelles are similar to liposomes but they do not have an inner liquid

compartment. Therefore they can be used as water-soluble biocompatible micro

containers for the delivery of poorly soluble hydrophobic pharmaceuticals [5]. Similar to

liposomes, their surface can be modified with antibodies (immunomicelles) or other

targeting moieties providing the ability of micelles to specifically interact with their

antigens. Pluronic block copolymers, a type of micelle are recognized as pharmaceutical

excipients listed in the U.S and British Pharmacopoeia [75, 76]. They have been

extensively used in a variety of pharmaceutical formulations including delivery of low

molecular mass drugs, polypeptides, and DNA. Furthermore, pluronic block copolymers

are versatile molecules that can be used as structural elements of polycation- based gene

delivery systems (polyplexes).

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1.3 Overview of solid lipid microparticles

The carriers that have been the most often studied in the controlled release of the

incorporated substances are:

• liposomes

• polymeric nano- and microparticles

• cyclodextrins

• solid lipid nanoparticles (SLNs).

Liposomes are spherical particles composed of one or more concentric

phospholipids bilayers alternating with aqueous partition. This kind of structure makes it

possible to incorporate lipophilic drugs into lipid bilayers as well as hydrophilic drugs

into the aqueous compartment. Drug release from liposomes, stability and

pharmacokinetic profiles depend on liposome composition, size and surface charge, and

drug solubility [2]. Liposome formulations of many different drugs show a significant

increase in therapeutic activity compared with non-liposomal formulations [65].

Liposomes are biocompatible and biodegradable, but also have some disadvantages

including low stability, low encapsulation efficiency, high cost and difficulties for scaling

up production [65, 77-80].

Polymeric nano- and microparticles are general terms that include nano- and

microspheres (consisting of a polymeric matrix) as well as nano- and microcapsules

(reservoir systems composed of a solid or liquid core which can contain either dispersed

or dissolved drugs and which is surrounded by a thin polymer layer). Hydrophilic and

lipophilic drugs can be incorporated or entrapped into polymeric nano- and

microparticles with relatively high efficiency [39, 78, 82, 83]. These kinds of drug carrier

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systems have proved to be more physicochemically stable than liposomes both in vivo

and during storage. Their main disadvantages are that preparation methods generally

require organic solvents and that large-scale production is rather difficult [83]. Moreover

it is crucial to choose suitable polymers that have proved to be nontoxic, biodegradable

and nonimmunogenic [77-83]. Synthetic polymer matrix materials have also been

suspected to lead to detrimental effects on peptides and proteins incorporated during the

manufacturing process [1].

Cyclodextrins are cyclic oligosaccharides composed of six (α-cyclodextrin), seven

(β-cyclodextrin), eight (γ–cyclodextrin) or more glucopyranose units. They are known for

being able to include apolar molecules inside their hydrophobic cavities and provide

these guest molecules with better stability, higher water solubility and increased

bioavailibility and/or decrease undesirable side effects [84]. However, so far, no study

has established the ability of cyclodextrins to induce a controlled release of the included

drug in vivo [77, 84-86]

SLNs were developed in the early 1990s and have since been considered to be

promising drug carrier systems, especially with a view to giving the incorporated active

substance a sustained-release profile. Compared with liquid lipid formulations, such as

fat nanoemulsions, drug mobility is lower in solid lipids than in liquid oils. SLNs are in

the submicron size range (50 – 1000 nm) and are composed of a lipidic matrix that is in

the solid state at room temperature. They provide an alternative drug carrier system to

liposomes and polymeric nanoparticles.

The composition and properties of SLMs are equivalent to SLNs, except for the

size ranges. Given the similar compositions of SLNs and SLMs, SLMs may also be

22

considered as physiologically compatible, physicochemically stable and allowing a large-

scale production. The difference in the size range between SLNs and SLMs means that

their application domains and administration routes can be different. Nevertheless, SLMs

as well as SLNs, in their respective application fields, can both be considered as

promising drug delivery systems. However, so far, SLMs have remained rather untapped.

1.3.1 Solid lipid microparticle preparation techniques

Commonly used materials for SLM preparation are:

• Lipids, including fatty alcohols, fatty acids, fatty acid esters of glycerol (mono-, di-

and/or triglycerides), waxes, cholesterol etc.

• Surfactants: Many different surfactants can be used, including:

Poloxamer 188 [16, 87, 88], Poloxamer 407 [89], Polysorbate 40 [65],

Polysorbate 80 [90], Sorbitan monopalmitate [65] Sodium dodecyl sulphate [65, 87, 88,

91], Polyvinyl alcohol [1, 87, 88], Soya lecithin [65, 92], Egg phosphatidyl choline [93]

and

• Water

1.3.1.1 Preparation techniques

Studies have shown that simply mixing the ingredients is not sufficient to ensure

controlled-release SLMs formulation [94, 95]. Drug release cannot be prolonged, based

on a solid matrix where drug and lipids are just physically mixed. The production

technique must allow the drug to dissolve or to disperse into lipids.

1.3.1.2 Solvent evaporation method

The classical solvent evaporation method regularly used is described in Figure 1.

Lipids are first dissolved in an organic solvent (most often chlorinated solvents) and are

23

Figure 1: Schematic representation of solid lipid microparticles production by: (1)

solvent evaporation method; (2) O/W melt dispersion technique; and (3) W/O/W

double emulsion technique.

Modified from CORTESI R, ESPOSITO E, LUCA G et al. 2002.

24

then emulsified in an aqueous phase containing an emulsifying agent. The resulting O/W

emulsion is finally stirred for several hours under ambient conditions in order to allow for

solvent evaporation [96]. A modified solvent evaporation method has also been described

[1, 97, 98]. In this technique, the lipids are also first dissolved in an organic solvent. By

mixing, the drug is then incorporated into the organic phase either as a solid (S/O/W)

which has been first ground in a mortar in the presence of liquid nitrogen, or dissolved in

an aqueous solution (W/O/W). The obtained preparation is then emulsified into an

aqueous surfactant solution. The emulsion is poured into an ice-cooled aqueous phase

and stirred. Obtained microparticles are filtered, rinsed with water and dried in a

desiccator.

1.3.1.3 O/W melt dispersion technique (for lipophilic drugs)

This is also called hot melt microencapsulation technique (which can be carried out by

normal or phase inversion technique). The drug is dissolved in the melted lipid (the

melting temperature depending on the lipid used). The hot mixture is emulsified into an

aqueous surfactant solution that is heated above the lipid melting point. The O/W

emulsion can then be poured into a larger volume of ice-cooled aqueous phase [1, 88,

97]. The emulsion, which is obtained by mixing with a high shear device (e.g., Ultra-

Turrax® [IKA], or Silverson mixer), is finally allowed to cool either at room temperature

or in an ice bath (Figure 1) [1, 16, 88, 93, 96, 98, 99].

Hardened microparticles are filtered, rinsed with water and dried in a vacuum

desiccator.

25

1.3.1.4 W/O melt dispersion technique (for hydrophilic drugs)

This method is a variant of the O/W melt dispersion technique, but it is used for

water-soluble drugs. This process does not use water in order to avoid excessive drug

solubility into the external aqueous phase and thereby low drug loading in microparticles.

First, the drug is dispersed into the melted lipid together with the surfactant. A hot non-

aqueous continuous phase (e.g. silicone oil) is poured into the molten lipid phase. The

obtained dispersion is then rapidly cooled through cold oil addition and immersion in an

ice bath. Solidified microparticles are separated from oil by centrifugation and are finally

washed and dried [92].

1.3.1.5 W/O/W multiple emulsion technique for water-soluble drugs

A heated aqueous drug solution is emulsified into the melted lipid. The obtained

primary W/O emulsion is put into an external aqueous phase and stirred so as to get a

W/O/W emulsion. The latter is then cooled either in an ice bath [88] or at room

temperature under stirring [96] (Figure 1). Hardened microparticles are filtered, rinsed

with water and finally dried in a vacuum desiccator.

1.3.1.6 High-pressure homogenisation

The homogeniser can reduce particle size to the micro- or even the nanometre

range of size depending on composition and process parameters.

1.3.1.7 Hot homogenisation

A pre-emulsion is obtained by mixing a hot aqueous surfactant solution with the

drug-loaded lipid melt, using a high shear device. The high-pressure homogeniser is

preheated at a temperature above the lipid melting point [87, 90, 100]. The preemulsion is

26

put through the homogeniser once or several times. Formulations are then allowed to cool

at room temperature.

1.3.1.8 Cold homogenisation

The drug is dissolved into the melted lipid. After solidification, the mixture is

milled in liquid nitrogen or dry ice with the help of a mortar mill. Milled particles are

then dispersed into an aqueous surfactant solution heated at 5 – 10 °C below the lipid

melting point [90, 100]. Particles can be disrupted by putting them through the

homogeniser once or several times.

1.3.1.9 Microchannel emulsification technique

This technique is considered to be a novel method used to prepare monodisperse

O/W and W/O emulsions without high mechanical stress and at lower energy input

compared with conventional emulsification processes.

A silicon microchannel (MC) plate, which is fabricated by micromachining

technology, is used, and droplets are produced by forcing the dispersed phase into the

continuous phase through the MCs [101, 102]. The droplet size is precisely regulated by

the structure of the MCs. This manufacturing technique yields monodispersed droplets. A

SLM suspension is obtained after cooling the emulsion at room temperature.

1.3.1.10 Cryogenic micronisation

Lipid matrices, obtained either by melt dispersion (the drug is added to the molten

lipid under magnetic stirring, the melting temperature depending on the lipid used) or

solvent stripping (the drug and lipid are dissolved into a solvent mixture under stirring,

e.g., benzyl alcohol/ethanol [103], are stored at - 80°C and then micronised in a

customised apparatus supplying liquid nitrogen during the process. Obtained powders are

27

finally sieved in an automatic sieving apparatus. This technique can be used for particles

of ± 5 – 5000 μm in diameter according to the chosen sieves.

1.3.1.11 Spray congealing (also called spray chilling)

Lipophilic material is heated to a temperature above its melting point. The drug is

then dissolved into the melt. The hot mixture is atomised with a pneumatic nozzle into a

vessel that is stored in a carbon dioxide ice bath. Obtained particles are finally vacuum-

dried at room temperature for several hours [89, 92, 95, 105-109].

In the first variant of this technique, the melted mixture is atomised by ultrasound

energy into small droplets that fall freely and solidify by cooling at room temperature

[92, 106, 108]. Another variant of the spray chilling method, using a rotating disc, has

also been described [105]. With this method the melted mixture is dropped onto a high-

speed rotating disc. The rotation causes the molten mixture to spread and spray from the

disc periphery onto a chilled surface from which microparticles are collected.

1.3.1.12 Spray drying

Lipids and the lipophilic drug are dissolved simultaneously into an organic

solvent. The mixture is then spray dried in order to get solid lipid particles [104, 109,

110].

1.4. Solid lipid microparticles characterisation

1.4.1 Determination of particle size distribution

1.4.1.1 Laser diffractometry

Laser diffractometry (LD) size analysis is based on the principle that particles of a

given size diffract light through a given angle, which increases with decreasing particle

size. Two different diffraction theories can be used (Mie and Fraunhofer) to determine

28

the size distribution from the light intensity reaching the detectors. However, it is

important to note that the LD technique does not measure particle size in the strict sense,

but rather calculates size from light scattering effects.

The laser diffraction technique has the advantage of covering a broad size range

(from the nanometre to the lower millimeter range [4, 111] while being usable with wet

as well as dry samples. This makes LD to be one of the most convenient techniques for

SLM size determination: submicronic particles as well as aggregates can be identified in

microparticles populations.

The results can be expressed in terms of standard percentiles D (v, 0.9), D (v,

0.5) (= mass median diameter) and D (v, 0.1), which correspond to size values below

which 90, 50 or 10% of sample particles lie. The span value is the measurement of size

distribution width and is calculated as follows [103, 112]:

–

……………………………. (1)

1.4.1.2 Electrical zone sensing method

The electrical zone sensing method, also called electroresistance particle counting

method (with ‗Coulter counter multisizer‘ or ‗Elzones‘ instruments), is based on the

principle that when a particle suspended in a conducting liquid gets through a small

orifice, on either side of which are electrodes, a change in electric resistance occurs. A

known suspension volume is actually drawn through an orifice, which is the only

conducting path between two electrodes. The resistance between those electrodes is

monitored. When a particle gets through the orifice, a pulse increase in resistance

appears. The increase in resistance is proportional to the particle volume. As a result, the

29

distribution of pulse magnitudes can be used as a measurement of particle volume

distribution [112-114].

1.4.1.3 Scanning electron microscopy and optical microscopy

Both techniques are used to determine particle size, particle shape and surface

characteristics simultaneously. The main disadvantage of such techniques is that they can

only examine a rather small number of particles. Indeed, the number of particles that need

to be counted (300 – 500) to obtain a good distribution estimate causes the method to be

slow and tedious. In addition, the diameter is obtained from only two particle dimensions

(i.e., length and breadth). No estimation of particle thickness is available [112, 114].

It is generally considered that optical microscopy makes it possible to measure

particles of 1 – 100 μm in size. Electron microscopy can measure particles of 0.01 – 1

μm. Optical microscopy seems to be sufficient to determine SLM size if distribution is

monodispersed.

However, SLM populations often contain some submicronic particles that can

only be detected by using scanning electron microscopy.

1.4.1.4 Sieving analysis

This method uses a series of standard sieves in a range of standard diameters. A

given powder mass is placed on the first sieve (with the broadest mesh) in a mechanical

shaker. The powder is shaken for a given period of time and the material that gets

through one sieve and is retained on the next, finer sieve, is collected and weighed [114,

115]. Sieving is a straightforward technique able to produce a separated size fraction for

possible further studies. This technique is unsuitable for particles < 40 μm, fragile

30

particles, irregular particles such as elongated needles, particles sticking to the sieves or

forming clumps, and particles electrostatically charged [112].

1.4.1.5 Image analysis system

The image analysis system is a new technology developed to determine and

analyse particle size (0.7 – 2000 μm) and shape. This technology can be seen as a kind of

automated microscope: combining the precision and sensitivity of an ordinary

microscope with the statistical significance of the number of analysed particles – this

being carried out either in real time [97] or within a few minutes [116]. Its ability to

analyse particle shape provides users with high-quality, helpful information to

characterise materials (emulsions, suspensions or powders) completely [116]. As a result,

the image analysis system can be used in order to better understand material behaviour

(e.g., powder flowability). Morphological parameters determined by the software include

sieving diameter [97], mean diameter, convexity, roundness and elongation, among

others.

This technology is bound to become increasingly popular, although the apparatus

still remains rather expensive [112].

1.4.1.6 Determination of aerodynamic size distribution

Aerodynamic size analysis only concerns the inhalation field. The aerodynamic

diameter of particles or droplets is actually the most important parameter influencing

aerosol deposition. This parameter is defined as the diameter of a unit-density sphere with

the same settling velocity, generally in air, as the particle. This includes particle shape,

density and physical size, all of which influence the particle aerodynamic diameter [117].

The determination of aerodynamic size distribution is useful to determine the respirable

31

fraction [115]. Such determination is generally carried out with a cascade impactor. The

principle on which these impactors operate is based on the erodynamic behaviour of

aerosol particles. They use the principle of inertial separation to size particulate samples

from the gas stream. The impactor usually has several stages for particle size

determination, each of which gives a cut off point based on the particle aerodynamic

diameter [112, 115, 117].

1.4.2 Determination of solid lipid microparticle morphology

The general morphology of SLMs is most often determined by microscopy

(scanning electron microscopy or optical microscopy, see Section 1.4.1.3), but can also

be studied by using new image analysis technology (see Section 1.4.1.5). The shape of

SLMs can be significantly different from a spherical shape.

The surface characteristics of SLMs (smooth or rough, regular or not) can be

visualised by microscopy. Their surface morphology varies depending on the excipients

used [100].

X-ray photoelectron spectroscopy (XPS), also known as electron spectroscopy for

chemical analysis (ESCA), is a high resolution technique for the elemental analysis of

solid materials surfaces. Consequently, XPS can determine the atomic composition of the

particles surface. XPS is based on the emission of electrons from materials in response to

photon irradiation, with sufficient energy to cause the core level electron ionisation.

These electrons are emitted at energies characterizing the atoms from which they are

emitted. In view of the fact that photons have a low penetration energy, only electrons

belonging to surface atoms or just underneath surface atoms (up to 100 Å) escape and are

counted.

32

This technology is used to gather information on drug distribution in

microparticles; in particular, to know if the drug is present on the surface of particles or

really entrapped within them. XPS is still rarely used in the microencapsulation field [1,

92, 118, 119]. This technique can be used when the compound to be localised contains

atoms that can emit electrons after photon irradiation and are not present in carrier

materials (e.g., Cl, N).

1.5 Solid-state analysis of solid lipid microparticles

This characterisation step is necessary in order to detect possible modifications in

the physicochemical properties of the drug incorporated into SLMs and of the lipophilic

excipients. It has been shown that although particles are produced from crystalline raw

materials, the presence of emulsifiers, the preparation method and the high-shear

dispersion may result in changes in the crystallinity of matrix constituents compared with

bulk materials. This may lead to liquid, amorphous or only partially crystallised

metastable systems [120, 121].

It has also been shown that with lipid drug delivery systems polymorphic

transformations may occur during dosage form preparation and subsequent storage.

During the melt solidification, triglycerides and fatty acids in particular can crystallise

into different polymorphic forms (i.e., the thermodynamically unstable α-form, the β′-

form, the stable β-form) depending on lipid composition and cooling rates. Polymorphic

transformations may cause changes in active and auxiliary substances solubilities and

melting points. In particular, the conversion of one polymorph into another may change

the physical properties of the substance [1, 77, 82, 107].

33

Because the degree of lipid crystallinity and the possible modifications in the

lipid‘s solid state are correlated with drug incorporation and release rates, and that the

drug‘s solid-state form (amorphous or crystalline) in solid dispersions influences

dissolution rates, it is important to pay special attention to these parameters [65, 95].

The solid states of bulk materials, as well as solid states obtained from solid

dosage forms (SLMs), are generally analysed by means of the following different

techniques:

1.5.1 Differential scanning calorimetry

Differential scanning calorimetry (DSC) is one of the most widely used

techniques to study solid state, and especially to determine compound purity, stability and

polymorphism. This technique relies on the principle that solid-state modifications are

characterised by different melting points and melting enthalpies [65]. DSC measures

transition temperatures (solidification and melting temperatures, glass transition

temperature, and thermal degradation temperature) as well as transition enthalpies [122].

1.5.2 X-ray diffraction

X-ray diffraction is based on the principle that X-rays are diffracted by crystals,

considering that their wavelengths have about the same magnitude as the distance

between crystal atoms or molecules. This technique makes it possible to investigate a

crystal structure [114], assess the compound‘s possible amorphisation, elucidate some

polymorphic transformations and study interactions between active substances and

microparticle excipients [108].

34

1.5.3 Hot stage microscopy

Hot stage microscopy (HSM) is an analytical technique that combines the best

properties of microscopy and thermal analysis in order to carry out characterisation of the

physical properties of the material as a function of temperature. Combined with high-

resolution cameras and image manipulation software, this technique is often used to

confirm the transitions observed with other techniques [123]. The solid states of bulk

drugs (lipophilic excipients and active substances) as well as the solid state of obtained

SLMs can be characterised by this technique. The main advantages of HSM are the

possibility to identify which particles (characterised by their shapes and sizes) are first

concerned by state transition, and the possibility of distinguishing between the excipient‘s

behaviour and that of the drug.

1.5.4 Fourier transform, Raman and infrared spectroscopy

Fourier transform Raman spectroscopy and infrared spectroscopy are useful tools

for investigating the structural properties of lipids [65]. These techniques have proved to

be highly sensitive to structural differences in a molecule‘s functional groups that can

take place during crystallisation or polymorphic transformations [124]. As a result, they

can be used in the field of SLMs to study the solid-states of bulk materials or solid

dosage forms, and in particular to detect interactions between active substances and

lipophilic excipients in molten samples [89, 95].

1.6 Drug loading determination

The determination of drug loading (or drug incorporation) is an important tool to

evaluate a potential drug carrier system. It is desirable to produce microparticles with

35

high drug content in order to decrease the amount of microparticles to be administered,

whatever the administration route.

Drug incorporation into microparticles can be expressed in terms of theoretical

drug loading, real drug loading or entrapment efficacy:

• Theoretical drug loading is expressed in a percentage related to the lipidic phase (lipidic

matrix + drug).

= ……………… (2)

• (Real) drug loading or drug content is expressed as a percentage related to the lipidic

phase (lipidic matrix + drug).

………… (3)

• Encapsulation efficiency (entrapment efficacy or loading efficiency) (EE) is calculated

as a percentage related to the total amount of drug initially used.

………………………………… (4)

The drug loading and EE can be influenced by a large number of factors. The

most often quoted parameters are the following:

• The drug solubility in melted lipids should be high enough to obtain a sufficient drug

loading [88] and thereby a relatively higher EE.

36

• The chemical and physical state of the solid lipid matrix and of the drug to be

incorporated have an influence on EE.

• The choice of the preparation method can also influence the EE of the drug into SLMs.

For example, the melt dispersion technique generally gives higher encapsulation

efficiency than does the solvent evaporation technique [1, 97], whereas the cold

homogenisation technique generally gives higher drug loading than hot homogenisation

[90].

• The way the drug is initially dispersed into the lipid at the initial stage of the preparation

(i.e., in the solid state or as a solution) can also be considered as a relevant factor

influencing EE [97].

• Increasing SLM‘s particle size generally leads to a higher drug loading. This parameter

has been studied by determining and comparing drug loadings of SLM in different size

fractions of the sample [106, 108]. It has also been noticed that some of the smallest

particles are formed by pure excipients only (empty spheres) [106].

• The theoretical initial drug loading influences encapsulation efficiency, which generally

decreases when the theoretical loading increases [1, 97, 98]. In this case, it is important to

use relatively high theoretical drug loading in order to get sufficient drug content, but the

theoretical drug loading must also be limited to avoid a decrease in encapsulation

efficiency and a resulting waste of drug. Studies have reached contrasting conclusions

which could be accounted for by poor water solubility of the drug, and therefore by a

smaller relative drug loss with increasing theoretical loading [88].

• In some special cases, the external aqueous phase pH can influence SLM‘s drug

loading; for example, if the drug is hydrophilic e.g., when the drug is a peptide, such as

37

insulin [1, 98], the external aqueous phase pH can be adjusted in order to decrease drug

aqueous solubility and thereby enhance drug-loading efficiency.

Because all these parameters can influence encapsulation efficiency, various

formulations and processes have to be studied in order to maximise drug content in

microparticles. Drug content optimisation can be achieved with the help of experimental

design methodology, which makes it possible to study several parameters simultaneously

on one or many chosen responses; for example, drug content [125, 126]. It is also

important to notice that drug loading might lead to some changes in SLM size

distribution [127].

Drug loading and encapsulation efficiency are generally determined as follows.

SLMs are first isolated from the aqueous phase. The aqueous SLM suspension is either

filtered or centrifuged, or even ultrafiltered (for the smallest microparticles), to separate

SLMs from the aqueous phase. Particles are then rinsed with water in order to eliminate

the drug crystals that are not incorporated in SLMs. Finally, obtained particles are dried.

SLMs are then either dissolved into an appropriate solvent or heated with a suitable

aqueous solvent in which the drug is soluble and shaken in order to extract the drug in the

solvent. The drug assay is carried out on the obtained solution, generally by means of a

spectrophotometrical technique.

The preparation of SLMs by spray congealing or spray drying does not use water,

which makes it possible to avoid the separation step between SLMs and the aqueous

phase.

As described in Section 1.4.2 of this work, XPS analysis can give further

information about the encapsulation of drugs into SLMs. This technique is used to

38

localise the drug inside the particle, so as to know whether the drug is present on the

surface of the particles or really entrapped within them.

A few drugs or peptides with various lipophilicity degrees have been incorporated

into SLMs; for example, carbamazepine [108], theophylline [88, 105, 106], fenbufen

[106], hydrocortisone, indomethacin, ketoprofen and ibuprofen [88], pseudoephedrine

HCl [99], fluorouracil [99, 127], ftorafur [99], insulin [1, 98], thymocartine [1],

gonadotropin release hormone [103], DNA [128], piribedil [90, 124],

medroxyprogesteron acetate [107], estradiol 17-β cypionate [104], somatostatin [97],

verapamil HCl [92] and felodipine [89, 95].

1.7 In vitro drug release studies – (factors affecting in vitro drug release)

As described, SLMs are mainly used to ensure that the incorporated drug release

is controlled. Therefore, a drug release study has to be carried out on obtained SLMs.

Drug release profiles are determined by an in vitro dissolution test. This test is generally

carried out according to the Pharmacopeia (USP or European Pharmacopeia) guidelines;

for example, by using a basket or paddle stirring apparatus. The dissolution medium is

chosen depending on the intended administration route. The sample can be put either into

a cell with two chambers (one chamber contains the sample, the other chamber is the

acceptor compartment) separated by a stainless steel sieve plate (with pores of a chosen

diameter) [97, 8], or into a dialysis tubing device [93]. In order to improve the wettability

of microparticles, a surfactant is generally added to the dissolution medium [105]. Drug

release is finally assayed spectrophotometrically. In a special case of topical

administration, the drug release study may be done with the help of the Franz diffusion

cell technique [129].

39

It has also been considered that a 24-h time interval is considered sufficient to

study the sustained-release potential of drug carrier systems (i.e., SLMs) [103]. Release

profiles can be further studied by determination of two statistical moments: i) the in vitro

mean dissolution time (MDTin vitro); and ii) the variance in dissolution time (VRin

vitro); and an associated statistical parameter, the concentration–time profile relative

dispersion (RD). These parameters has described the method for calculating the statistical

moment approach which has the advantage of allowing the dissolution curve to be

separated into stages and, therefore, to check for modifications in the release mechanism

during the dissolution test [92].

Some drug release is affected by several parameters. It is important that the

dissolution medium [90, 95, 126] and the dissolution method [130] be correctly chosen in

order to obtain a correct prediction of the in vivo drug release from microparticles [109].

The nature (hydrophobicity) of the excipient is considered to be the most important

parameter influencing drug release with more hydrophobic materials expected to reduce

the drug release rate [95, 105, 106, 109]. The choice of matrix materials influences the

release process rate. Another way to change the matrix hydrophobicity is by adding a

hydrophobic or hydrophilic excipient [89, 95, 105].

The preparation method of the SLMs can affect the drug‘s release rate by

influencing the matrix wettability properties [109].

The particle size is also considered a relevant parameter influencing drug release.

Drug release from smaller particles is higher than from larger ones because of the larger

specific surface area of smaller microparticles [95, 106].

40

A faster release is obtained with higher drug and/or adjuvant content (e.g.,

lactose) in SLMs because matrix diffusion is easier due to an increase in the number of

pores created during the release process [88, 95 106, 127]. The drug release increases

when the medium agitation rate in the dissolution apparatus increases [127].

Storage can induce polymorphic changes in SLMs and thereby modify the drug

release rate [95]. Consequently, a suitable choice of SLM formulation (in terms of

excipient nature, drug nature and drug loading) can bring about the intended in vitro

release profiles e.g., sustained release [109], enhanced release [108].

If SLMs are not rinsed after separation from the aqueous phase, the dissolution

profile shows a rapid release from the external drug fraction towards the dissolution

medium, followed by a phase of decrease in the release rate [90, 97, 98].

At the end of the release study, some of the drug may remain enclosed in the

particles (98), in particular if the drug is adsorbed onto the lipid matrix material [97].

1.8 Administration routes, in vivo drug release and biocompatibility studies

Despite their high potential as promising drug carrier systems, SLMs have been

rather unexploited. So far, only a few complete studies on SLMs have been published.

Consequently, little data is currently available on SLM in vivo administration, drug

release and biocompatibility. The section below presents an overview of tested SLM

administration routes and corresponding in vivo drug release and biocompatibility studies

carried out so far.

1.8.1 Peroral administration

The peroral route is the most often cited SLM administration route in the

literature [89-92, 105, 106]. It includes aqueous SLM dispersion, SLM tablets, pellets or

41

capsules. However, data on in vivo drug release and biocompatibility studies are most

often missing. Demirel et al., (2001) have nevertheless perorally administered SLM

suspensions to rabbits [90]; such suspensions were composed of Compritol® 888 ATO

(Gattefosse) and Labrasol® (Gattefosse) as a lipidic matrix, Tween

® 80 (ICI America) as

a surfactant and piribedil as the active substance. The bioavailability of piribedil-SLMs

was found to be higher than with pure piribedil.

Considering that SLM lipidic matrices are composed of physiological lipids and

that most surfactants have already been used perorally, there is no doubt on the

biocompatibility of the SLMs after oral administration.

1.8.2 Parenteral administration

SLMs could also be parenterally administered aside from the intravenous route,

owing to particle micronic size (in contrast to SLNs, which are often used for the

intravenous administration). Some studies have been carried out on the in vivo drug

release and biocompatibility of SLMs. Reithmeier et al., (2001a; 2001b) have tested the

biocompatibility of SLMs composed of a glyceryl tripalmitate (Dynasan®

116, Hüls AG)

lipidic matrix and polyvinyl alcohol as a surfactant by implanting SLMs subcutaneously

in mice [1, 97]. Polymeric microparticles composed of poly (D, L-lactide-co-glycolide), a

well known approved polymer often used for parenteral applications, were also implanted

and used as a reference. The study showed only a slight inflammation reaction in the

implantation area, for both SLMs and polymeric microparticles. It has been concluded

that studied SLMs showed comparable biocompatibility to polymeric microparticles that

have been approved and used for parenteral administration.

42

Del Curto et al., (2003) have produced SLMs composed of glyceryl

monobehenate (Compritol® E ATO, Gattefossé) and containing gonadotropin release

hormone (Antide) by co-melting process [109]. After subcutaneous injection in rats, the

SLMs proved to give the incorporated active substance a sustained release profile.

Therefore, Antide-SLMs are potentially useful as a depot formulation when prolonged

action is required.

1.8.3 Topical administration

SLM topical applications have been seldom used. However, Yener et al., (2003)

have studied SLMs prepared with beeswax as matrix material, polysorbate 80 (Tween®

80) as a surfactant and containing a UV absorber (octyl methoxy cinnamate, OMC)

[129]. Obtained SLMs were put into topical vehicles (oleaginous cream, carbopol gel and

o/w emulsion). OMC release from the SLMs and the OMC penetration rate and amount

were tested through application on excised rat skin. The results were as those expected: a

decrease in OMC release rate and amount (and therefore sustained action compared with

free OMC action), and a decrease in the penetration rate and amount.

1.8.4 Pulmonary administration

SLMs can be considered a promising drug carrier system for pulmonary

administration even if they have been rather unexploited so far [125, 130]. However, a

preliminary in vivo tolerance study has been carried out with rats in SLMs composed of

glyceryl behenate (Compritol 888 ATO) as a lipidic matrix and poloxamer 188 (Lutrol®

F68, BASF) as a surfactant. SLM dispersions in phosphate buffer saline were

administered intratracheally. Bronchoalveolar lavages were performed on the

anaesthetised rats. Total and differential cell counts (i.e., inflammatory cells) were then

43

carried out with the collected bronchoalveolar liquids. Results did not show significant

differences between placebo groups and SLM-treated rats. It has been concluded that the

studied SLMs seem to be well tolerated by the lower airways, but tolerance must still be

assessed after repeated administrations [16].

1.9 In vivo fate

The in vivo fate of SLMs has not been studied thoroughly so far. However, in view of

their similar composition, SLMs are expected to behave in the same way as SLNs in vivo.

Consequently, the in vivo fate of SLMs should depend on administration routes and

especially on enzymatic processes. Because SLM lipidic matrices are composed of

physiological lipids, they are bound to undergo metabolisation in vivo. Lipases should

then be the most involved enzymes in the degradation of SLMs. This type of enzyme,

which is present in various organs and tissues (notably in the gastrointestinal tract, at the

subcutaneous or intramuscular injection sites), works by splitting the ester linkage and

thereby forming partial glycerides or glycerol and fatty acids. It has been shown that

SLNs composed of glyceryltrimyristate, glyceryltripalmitate, glyceryltristearate and

cetylpalmitate, are decomposed by enzymes such as lipases, and that such degradation is

influenced by several parameters (i.e., surfactant composition and storage time) [97, 131,

132]. These conclusions could reasonably be extrapolated to SLMs, although they would

need to be confirmed by experimentation.

SLMs present several advantages: a physiological composition and thereby a

supposed limited toxicity; a possibility of producing them on a fairly large industrial

scale; and the relative low cost of their raw materials and production processes. Examples

of drawbacks are the drug to be incorporated into SLMs must preferably be lipophilic

44

enough in order to achieve high entrapment efficiency. The latter is also affected by

several other parameters such as the preparation method, the chemical and physical state

of the drug and excipients, and the size of particles obtained. One of the main difficulties

in using SLMs is the optimization of formulation parameters (excipients and drug nature,

initial theoretical drug loading etc.) and production techniques in order to obtain SLMs

that have simultaneously high entrapment efficiency, high drug loading, the intended size

according to the desired administration route and presenting the desired drug release

profile.

The in vitro drug release studies tend to prove the ability of SLMs to provide a

controlled release of the incorporated substances. Nevertheless, it must be taken into

account that the dissolution medium and the dissolution method are both critical

parameters, which must be suitably chosen in order to get a good correlation between the

in vitro and in vivo drug release studies. The main difficulty in studying the rate of drug

release from a carrier lies in mimicking as close as possible the expected in vivo

conditions. Especially in the case of SLMs, the presence of enzymes such as lipases exert

an important influence on drug release, but this parameter is difficult to mimic in the in

vitro dissolution tests. Owing to the lipidic nature of SLMs, the drug release studies also

require the use of surfactant in the dissolution medium in order to improve the

microparticle‘s wettability although the eventual influence of the addition of a surfactant

on the in vitro drug release rates has notr been studied. This is why the promising drug

release results obtained by in vitro experimentation must be confirmed by in vivo studies.

In general, SLMs have numerous advantages and interesting in vitro drug release

results as a promising drug carrier system, which could be used by different

45

administration routes (i.e., peroral, parenteral, topical and pulmonary routes). However,

drug release results obtained by in vitro experimentation and suggesting ability for SLMs

to provide a controlled release to the incorporated substances must be confirmed by in

vivo studies. Although, the biocompatibility and in vivo fate of SLMs are disregarded in

most studies, they should be considered and assessed even if the physiologically used

materials tend to suggest that SLMs are biocompatible.

1.10 Biopharmaceutics Classification System of Drugs

The oral route of drug administration is the route of choice for formulators and

continues to dominate the area of drug delivery technologies. However, though popular,

this route is not free from limitations of absorption and bioavailability in the milieu of the

gastrointestinal tract. These limitations are even more prominent with the advent of

protein and peptide drugs and the compounds emerging as a result of combinatorial

chemistry and the technique of high throughput screening.

The Biopharmaceutics Classification System (BCS) is a drug development tool

that allows estimation of the contribution of three fundamental factors including

dissolution, solubility and intestinal permeability, which govern the rate and extent of

drug absorption from solid oral dosage forms [132]. Drug dissolution is the process by

which the drug is released, dissolved and becomes ready for absorption. Permeability

refers to the ability of the drug molecule to permeate through a membrane in to the

systemic circulation. The in vivo performance of orally administered drug depends upon

its solubility and tissue permeability characteristics. Based on these characteristics, drug

substances are divided into four classes and the classification system is called

Biopharmaceutical Classification System. BCS is also a fundamental guideline for

46

determining the conditions under which in vitro – in vivo correlations (IVIVCs) are

expected [133]. It is used as a tool for developing the in vitro dissolution specifications

[132, 134]. The classification deals with drug dissolution and absorption model, which

considers the key parameters controlling drug dissolution and absorption [135, 136]. The

biopharmaceutical classification system acts as a guiding tool for development of various

oral drug delivery technologies [133].

1.10.1 Determination of solubility

The solubility of a substance is the amount of substance that has passed into

solution when equilibrium is attained between the solution and excess, i.e. undissolved

substance, at a given temperature and pressure. The objective of BCS approach is to

determine the equilibrium solubility of a drug substance under physiological pH

conditions.

A drug substance is considered highly soluble when the highest dose strength is

soluble in 250 ml or less of aqueous medium over the pH range of 1-7.5 [137]. The

volume estimate of 250 ml is derived form the typical volume of water consumed during

the oral administration of a dosage Form. This is about the minimum fluid volume

anticipated in the stomach at the time of drug administration. The pH solubility profile of

the drug substance is determined at 37 ± 10 oC in aqueous medium with pH in the range

of 1-7.5. A sufficient number of pH conditions should be evaluated to accurately define

the pH-solubility profile. The number of pH conditions for a solubility determination

depends upon the ionization characteristics of the test drug substance. A minimum of

three replicate determinations of solubility in each pH condition should be carried out.

Standard buffer solutions described in pharmacopoeias are considered appropriate for use

47

in solubility studies. If these are not suitable for physical or chemical reasons, other

buffer solutions can also be used provided the pH of these solutions is verified. Methods

other than shake-flask method are also used with justification to support the ability of

such methods to predict equilibrium solubility of test drug substance as exemplified with

acid or base titration methods. The concentration of drug substance in selected buffers or

pH conditions should be determined using a validated solubility-indicating assay that can

distinguish the drug substance from its degradation products. If degradation of drug is

observed as a function of buffer composition and/or pH, it should be taken into

consideration.

1.10.2 Determination of permeability

The permeability class boundary is based directly on the extent of absorption (fraction of

dose absorbed) of a drug substance in humans. The recommended methods not involving

human subjects include in vivo or in situ intestinal perfusion in a suitable animal model

(e.g. rats), and/or in vitro permeability methods using excised intestinal tissues, or

monolayer of suitable epithelial cells. In many cases, a single method may be sufficient

but when not suitable to conclusively demonstrate a permeability classification, two

different methods may be used. Chemical structure and/or certain physicochemical

attributes of a drug substance (e.g. partition coefficient in suitable systems) can provide

useful information about its permeability characteristics.

Fundamental to understanding of the nature of gastrointestinal permeability

limitations are methods and techniques to both screen and grade these characteristics.

These methods range from simple oil/water (O/W) partition coefficient to absolute

48

bioavailability studies. The methods that are routinely used for determination of

permeability include:

a. Pharmacokinetic studies in humans

Mass balance studies

Absolute bioavailability studies

Intestinal perfusion methods

b. In vivo or in situ intestinal perfusion in a suitable animal model

c. In vitro permeability methods using excised intestinal tissues

d. Monolayers of suitable epithelial cells e.g. Caco-2 cells or TC-7 cells

In mass balance studies, unlabelled, stable isotopes or radiolabelled drug

substances are used to determine the extent of drug absorption. However this method

gives highly variable estimates and hence other methods are sought for.

In absolute bioavailability studies, oral bioavailability is determined and

compared against the intravenous bioavailability as reference.

Intestinal perfusion models and in vitro methods are recommended for passively

transported drugs. The observed low permeability of some drug substances in human

could be due to the efflux of drug by various membrane transporters like p-glycoprotein.

This leads to misinterpretation of the permeability of the drug substance.

An interesting alternative to intestinal tissue models is the use of well-established

in vitro systems based on the human adenocarcinoma cell line Caco-2. These cells serve

as a model of small intestinal tissue. The differentiated cells exhibit the microvilli typical

of the small intestinal mucosa and the integral membrane proteins of the brush-border

enzymes. In addition, they also form the fluid-filled domes typical of a permeable

49

epithelium. Studies of Caco-2 cell lines have indicated their ability to transport ions,

sugars and peptides [138]. The directed transport of bile acids and vitamin B12 across

Caco-2 cell lines has also been reported [139, 140]. These properties have established the

Caco-2 cell line as a reliable in vitro model of the small intestine.

1.10.3 Applications of BCS in oral drug delivery technology

Once the solubility and permeability characteristics of the drug are known, it

becomes an easy task for the research scientist to decide which drug delivery technology

to follow or develop.

The major challenge in development of drug delivery system for class I drugs is to

achieve a target release profile associated with a particular pharmcokinetic and/or

pharmacodynamic profile. Formulation approaches include both control of release rate

and certain physicochemical properties of drugs like pH-solubility profile.

The systems that are developed for class II drugs are based on micronisation,

lyophilization, addition of surfactants, formulation as emulsions and microemulsions

systems as well as use of complexing agents like cyclodextrins.

Class III drugs require the technologies that address the fundamental limitations

of absolute or regional permeability. Peptides and proteins constitute a major part of class

III drugs and the technologies for handling such materials are on the increase [135].

Class IV drugs present a major challenge for development of drug delivery system

and the route of choice for administering such drugs is parenteral with the formulation

containing solubility enhancers [135].

The biopharmaceutics classification system was developed primarily in the

context of immediate release (IR) solid oral dosage forms. It is the scientific framework

50

for classifying drug substances based on their aqueous solubility and intestinal

permeability [132]. The BCS takes into account three major factors: dissolution rate,

solubility and intestinal permeability, all of which govern the rate and extent of oral drug

absorption from immediate release solid oral dosage forms. The interest in this

classification system is largely because of its application in early drug development and

in the management of product change through its life cycle. It was first introduced into

regulatory decision-making process in the guidance document on Immediate Release

Solid Oral Dosage Forms: Scale Up and Post Approval Changes [137, 141, 142].

1.10.4 Classification

Combined with the dissolution, the BCS takes into account the three major factors

governing bioavailability viz. dissolution, solubility and permeability. The classification

deals with drug dissolution and absorption model, which considers the key parameters

controlling drug dissolution and absorption as a set of dimensionless numbers: the

absorption number (defined as the ratio of the mean residence time to mean absorption

time), the dissolution number (defined as the ratio of mean residence time to mean

dissolution time), and the dose number (defined as the mass divided by the product of

uptake volume (250 ml) and solubility of drug) [132, 134].

The extent of solubilization and particle aggregation in the small intestine is

unknown and therefore, the solubility, dose, and dissolution number of a drug in vivo are

difficult to estimate precisely [132]. As the drug dissolution and intestinal permeability

are the fundamental parameters governing rate and extent of drug absorption, drugs could

be categorized into high/low solubility and permeability classes.

According to BCS, drug substances are classified as:

51

Class I : High Solubility – High Permeability

Class II : Low Solubility – High Permeability

Class III: High Solubility – Low Permeability

Class IV: Low Solubility – Low Permeability

Class I drugs exhibit a high absorption and dissolution numbers.The rate limiting step

is drug dissolution and if dissolution is very rapid then gastric emptying rate becomes the

rate determining step [132]. This group of drugs is expected to be well absorbed unless

they are unstable, form insoluble complexes, are secreted directly from gut wall, or

undergo first pass metabolism [134]. For immediate release products that release their

content very rapidly, the absorption rate will be controlled by the gastric emptying rate

and no correlation of in vivo data with dissolution rate is expected [132]. Dissolution

therefore, needs only to verify that the drug indeed is rapidly released from the dosage

form under mild aqueous conditions [134]. A dissolution specification of 85 % of drug

contained in immediate release in 15 mins may insure bioequivalence [137, 143]. The

FIP considers a formulation as very fast releasing when at least 80 % of the drug

substance is dissolved in about 20-30 mins under reasonable and justified test conditions.

The aforementioned dissolution time limits are based on typical gastric emptying times

for water in the fasted state.

When a class I drug is formulated as an extended release product in which the release

profile controls the rate of absorption, and the solubility and permeability of the drug is

site independent, correlation can be expected [144-146].

Examples of drugs in this class include metoprolol, diltiazem, verapamil, and

propranolol.

52

Class II drugs have a high absorption number but a low dissolution number. In vivo

drug dissolution is then a rate limiting step for absorption except at a very high dose

number. The absorption for class II drugs is usually slower than class I and occurs over a

longer period of time [132]. The limitation can be equilibrium or kinetic in nature. In the

case of equilibrium problem, enough fluid is not available in the GI tract to dissolve the

dose. For instance, a dose of griseofulvin requires 33.3 litres of fluid to be dissolved

[147]. As the total volume of fluid entering the GI tract within 24 h period is only about

5-10 litres, insufficient fluid would be available at any given time to dissolve the entire

dose of griseofulvin [134]. Griseofulvin exhibits a high dosing number and a low

dissolution number such that bioavailability and the fraction of the dose absorbed can be

improved by either decreasing the dosing number by reducing the dose, by taking more

water with the administered dose or by increasing drug solubility. On the basis of

pharmacokinetic/pharmacodynamic considerations, the dose of a drug is determined and

cannot be altered. The volume of water initially taken with the dosage will be limited by

patient compliance and the anatomical and physiological capacity of the stomach.

Therefore, for griseofulvin, only enhancement of the drug solubility through appropriate

formulation approach (i.e. solid dispersion) can lead to considerably reduced dose

number and increased drug bioavailability [148].

In the case of kinetics, the entire dose of the drug dissolves too slowly. It is shown

that bioavailability of digoxin depends on the particle size. Digoxin exhibits dissolution

rate limited absorption at particle sizes of greater than 10 µ in diameter [134]. These

agree with reports indicating that digoxin, in micronized form, and griseofulvin in

ultramicronized form, were almost completely absorbed [147].

53

For class II drugs, therefore, a strong correlation between dissolution rate and the in

vivo performance can be established [134]. Therefore, it is essential that in vitro

dissolution tests reflect in vivo situations. Dissolution media and methods that reflect the

in vivo controlling process are particularly important in this case if good correlations are

to be obtained. The dissolution profile for class II drugs requires multiple sampling times

and the use of more than one dissolution medium. Addition of surfactant to simulate in

vivo environment might be required. When a class II drug is formulated as an extended

release product and the solubility and permeability of the drug are site independent, some

level of correlation is expected [146]. However, once the permeability is site dependent,

little or no correlation is expected.

BCS classification together with the numerous compendial and physiological media

available could be employed as a fundamental guidance for designing appropriate

biorelevant dissolution conditions leading to a more meaningful prediction of in vivo

performances. For class I drugs, simple and mild aqueous dissolution media such as SGF

without pepsin is suitable, while milk as dissolution medium might be appropriate for

specific food/formulation interaction [149]. For neutral class II drugs, the fluid simulating

conditions in the proximal intestine in the fasted state reflects the dissolution in the upper

GI tract under fasted state conditions [149]. If a class II drug is a weak base, SGF could

be used to assess the drug dissolution in the stomach under fasted state conditions [149].

Comparison of dissolution results obtained under fasted conditions to those of fasted state

intestinal condition could be a good indicator of whether the formulation should be

administered before or after meals [149]. In the case of class II weak acids dissolution

could be performed in fasted state intestinal condition. Milk with its composition of lipids

54

and proteins or fasted state intestinal conditions containing high bile salt/lecithin levels

can be employed to simulate the fed state conditions [149, 150].

Examples of class II drugs include: phenytoin, danazol, mefenamic acid,

ketoconazole, glibenclamide, nifedipine. NSAIDs generally belong to this class [151].

Class III drugs are rapidly dissolving but permeability is rate-controlling step for drug

absorption [152]. Rapid dissolution is particularly desirable in order too maximize the

contact time between the dissolved drug and absorption mucosa. These drugs exhibit a

high variation in the rate and extent of drug absorption. Since the dissolution is rapid,

such that 85 % of drug dissolves in 15 min, the variation could be attributable to

physiology and membrane permeability in terms of GI transit, luminal contents, and

membrane permeation rather than dosage form factors [132, 153]. As drug permeation is

rate controlling, limited or no in vitro-in vivo correlation is expected.

Examples of drugs in this class are cimetidine, acyclovir, neomycin B, captopril as

well as proteins and peptides [154].

Class IV drugs are low solubility and low permeability drugs. This class exhibit

significant problems for effective oral administration. Inappropriate formulation of class

IV drugs, as in the case of class II drugs, could have an additional negative influence on

both the rate and extent of drug absorption. However, the class IV drugs are rarely

developed and reach the market. Nevertheless, a number of class IV drugs do exist. e.g.

Taxol [155].

Thus for all categories, it is anticipated that well-designed dissolution tests can be a

key prognostic tool in the assessment of both the drugs potential for oral absorption and

of the bioequivalence of its formulation [133].

55

1.11 Solubilization of poorly soluble drugs

Therapeutic effectiveness of a drug depends upon the bioavailability and

ultimately upon the solubility of drug molecules. Solubility is one of the important

parameters to achieve desired concentration of drug in systemic circulation for

pharmacological response to be elicited. Currently, only 8 % of new drug candidates have

both high solubility and permeability [156].

1.11.1 Solubility definitions and parts of solvent required for one part of solute

(B.P., 2001)

very soluble < 1;

freely soluble 1 – 10;

Soluble 10 – 30;

sparingly soluble 30 – 100;

slightly soluble 100 – 1000;

very slightly soluble 1000 - 10,000;

insoluble > 10,000.

1.11.1.2 Process of solubilisation

The process of solubilisation involves the breaking of inter-ionic or

intermolecular bonds in the solute, the separation of the molecules of the solvent to

provide space in the solvent for the solute, interaction between the solvent and the solute

molecules or ions.

1.11.2 Factors affecting solubility

Solubility depends on the physical form of the solid, the nature and composition of

solvent medium as well as temperature and pressure of system.

56

1.11.2.1 Particle size

The size of the solid particle influences the solubility because as a particle becomes

smaller, the surface area to volume ratio increases. The larger surface area allows a

greater interaction with the solvent. The effect of particle size on solubility can be

described by

…………………………………… (5)

Where, S0 is the solubility of infinitely large particles, S is the solubility of fine particles,

V is molar volume, γ or g is the surface tension of the solid, r is the radius of the fine

particle, R is the Gas constant, T is the temperature.

1.11.2.2 Temperature

Temperature will affect solubility. If the solution process absorbs energy, then, the

solubility will be increased as the temperature is increased. If the solution process

releases energy, then the solubility will decrease with increasing temperature. Generally,

an increase in the temperature of the solution increases the solubility of a solid solute. A

few solid solutes are less soluble in warm solutions. For all gases, solubility decreases as

the temperature of the solution increases.

1.11.2.3 Pressure

For gaseous solutes, an increase in pressure increases solubility and a decrease in

pressure decreases the solubility. For solids and liquid solutes, changes in pressure have

practically no effect on solubility.

57

1.11.2.4 Nature of the solute and solvent

While only 1 gram of lead (II) chloride can be dissolved in 100 grams of water at room

temperature, 200 grams of zinc chloride can be dissolved. The great difference in the

solubilities of these two substances is the result of differences in their natures.

1.11.2.5 Molecular size

Molecular size will affect the solubility. The larger the molecule or the higher its

molecular weight the less soluble the substance. Larger molecules are more difficult to

surround with solvent molecules in order to solvate the substance. In the case of organic

compounds, the amount of carbon branching will increase the solubility since more

branching will reduce the size (or volume) of the molecule and make it easier to solvate

the molecules with solvent [7].

1.11.2.6 Polarity

Polarity of the solute and solvent molecules will affect the solubility. Generally, non-

polar solute molecules will dissolve in non-polar solvents and polar solute molecules will

dissolve in polar solvents. The polar solute molecules have a positive and a negative end

to the molecule. If the solvent molecule is also polar, then positive ends of solvent

molecules will attract negative ends of solute molecules. This is a type of intermolecular

force known as dipole-dipole interaction. All molecules also have a type of

intermolecular force much weaker than the other forces called London Dispersion forces

where the positive nuclei of the atoms of the solute molecule will attract the negative

electrons of the atoms of a solvent molecule. This gives the non-polar solvent a chance to

solvate the solute molecules [7].

58

1.11.2.7 Polymorphs

A solid has a rigid form and a definite shape. The shape or habit of a crystal of a given

substance may vary but the angles between the faces are always constant. A crystal is

made up of atoms, ions, or molecules in a regular geometric arrangement or lattice

constantly repeated in three dimensions. This repeating pattern is known as the unit cell.

The capacity for a substance to crystallize in more than one crystalline form is

polymorphism. It is possible that all crystals can crystallize in different forms or

polymorphs. If the change from one polymorph to another is reversible, the process is

called enantiotropic. If the system is monotropic, there is a transition point above the

melting points of both polymorphs. The two polymorphs cannot be converted from one

another without undergoing a phase transition. Polymorphs can vary in melting point.

Since the melting point of the solid is related to solubility, so polymorphs will have

different solubilities. Generally the range of solubility differences between different

polymorphs is only 2-3 folds due to relatively small differences in free energy.

1.11.2.8 Rate of solution

The rate of solution is a measure of how fast substances dissolve in solvents.

1.11.2.8.1 Factors affecting rate of solution

1.11.2.8.2 Size of the particles

When the total surface area of the solute particles is increased, the solute dissolves more

rapidly because the action takes place only at the surface of each particle. Breaking a

solute into smaller pieces increases its surface area and hence its rate of solution.

59

1.11.2.8.3 Temperature

For liquids and solid solutes, increasing the temperature not only increases the amount of

solute that will dissolve but also increases the rate at which the solute will dissolve. For

the gases, the reverse is true.

1.11.2.8.4 Amount of solute already dissolved

When there is little solute already in solution, dissolution takes place relatively rapidly.

As the solution approaches the point where no solute can be dissolved, dissolution takes

place more slowly.

1.11.2.8.5 Stirring

With liquid and solid solutes, stirring brings fresh portions of the solvent in contact with

the solute, thereby increasing the rate of solution.

1.11.3 Techniques of solubility enhancement

Up to 40 % of lipophilic drug candidates fail to reach market although exhibiting

potential pharmacodynamic activities [156, 157]. Meanwhile, some lipophilic drugs in

the market have to be administered at high doses. As a result, various formulation

strategies have been investigated to improve the solubility and the rate of dissolution and

hence the oral bioavailability of lipophilic drugs. These strategies include solubilization

and use of surfactants, use of different polymorphic/amorphic drug forms, the reduction

of drug particle size, the complexation (e.g., cyclodextrins) and the formation of solid

drug solutions/dispersions [158, 159].

There are various techniques available to improve the solubility of poorly soluble

drugs generally. Some of the approaches to improve the solubility are:

60

I. Physical modifications

A. Particle size reduction: (i). Micronization; (ii). Nanosuspension

B. Modification of the crystal habit: (i). Polymorphs; (ii). Pseudopolymorphs

C. Drug dispersion in carriers: (i). Eutectic mixtures; (ii). Solid dispersions (iii) Solid

solutions

D. Complexation: Use of complexing agents

E. Solubilization by surfactants: (i). Microemulsions; (ii). Self microemulsifying drug

delivery systems.

II. Chemical modifications

A. pH adjustment B. Salt formation C. Cosolvency D. Hydrotrophy E. Solubilizing

agents

I. Physical Modifications

A. Particle size reduction

Micro-/nanonization is one of the most promising approaches to improve the

bioavailability of lipophilic drugs by an increase in surface area and saturation solubility

via reduction of the particle size to less than 1 μm [66]. Such size reduction cannot be

achieved by the conventional milling techniques. Patented engineering processes have

come up based on the principles of pearl milling (NanoCrystals®), high-pressure

homogenization (DissoCubes®), solution enhanced dispersion by supercritical fluids

(SEDS), rapid expansion from supercritical to aqueous solution (RESAS), spray freezing

into liquid (SFL) and evaporative precipitation into aqueous solution (EPAS) [160].

61

Pearl milling: NanoCrystals® involves filling an aqueous suspension of drug into a pearl

mill containing glass or zirconium oxide pearls as milling media. The drug microparticles

are ground to nanoparticles (< 400 nm) in between the moving milling pearls over a few

days. The milling efficiency is dependent on the properties of the drug, the medium and

the stabilizer. Rapamune®, an immune suppressant agent, is the 13 first FDA approved

nanoparticle drug using NanoCrystals® technology developed by Elan Drug Delivery.

Emend® is another product containing 80 or 125 mg aprepitant formulated by this

technique. The limitation of the pearl milling process is the introduction of contamination

to the product from the grinding material, batch-to-batch variations and the risk of

microbiological problems after milling in an aqueous environment for a few days.

High pressure homogenization: DissoCubes® manufacture involves dispersing a drug

powder in an aqueous surfactant solution and passing through a high pressure

homogenizer to obtain nanosuspensions. The cavitation force experienced is sufficient to

disintegrate drug from microparticles to nanoparticles. The particle size is dependent on

the hardness of the drug substance, the processing pressure and the number of cycles

applied. The possible interesting features of nanosuspensions are [4]:

• Increase in saturation solubility and dissolution rate of drug

• Increase in adhesive nature, thus resulting in enhanced bioavailability

• Increase in the amorphous fraction in the particles, leading to a potential change in the

crystalline structure and higher solubility

• Possibility of surface modification of nanosuspensions for site specific delivery

• Possibility of large-scale production, the prerequisite for the introduction of a delivery

system to the market.

62

However, only brittle drug candidates can be broken up into nanoparticles by this

technique. Even for these, the following would have to be considered, such as chemical

instability of fragile drugs under the harsh production conditions, Ostwald ripening in

long-term storage, toxicity of surfactants, redispersibility of the dried powder, batch-to-

batch variation in crystallinity level and finally the difficulty of quality control and the

stability of the partially amorphous nanosuspensions.

Solution enhanced dispersion by the supercritical fluids (SEDS): The SEDS process

was developed and patented by the University of Bradford. The use of a coaxial nozzle

provides a means whereby the drug in the organic solvent solution mixes with the

compressed fluid CO2 (antisolvent) in the mixing chamber of the nozzle prior to

dispersion, and flows into a particle-formation vessel via a restricted orifice. Such nozzle

achieves solution breakup through the impaction of the solution by a higher velocity

fluid. The high velocity fluid creates high frictional surface forces, causing the solution to

disintegrate into droplets. A wide range of materials have been prepared as carriers of

microparticles and nanoparticles using the SEDS process. A key step in the formation of

nanoparticles is to enhance the mass transfer rate between the droplets and the antisolvent

before the droplets coalesce to form bigger droplets. In another study, a significant

decrease in the particle size is achieved by using the ultrasonic nozzle-based supercritical

antisolvent process [161, 162].

Rapid expansion from supercritical to aqueous solution (RESAS): This process

induces rapid nucleation of the supercritical fluid dissolved drugs and surfactants

resulting in particle formation with a desirable size distribution in a very short time. The

surfactants in the supercritical fluid stabilize the newly formed small particles and

63

suppress any tendency of particle agglomeration or particle growth when spraying this

solution (drug + surfactant + CO2) into an aqueous solution containing a second surface

modifier [163]. The low solubility of poorly water soluble drugs and surfactants in

supercritical CO2 and the high pressure required for these processes restrict the utility of

this technology in pharmaceutical industry.

Spray freezing into liquid (SFL): The SFL technology was developed and patented by

the University of Texas at Austin in 2003 and commercialized by the Dow Chemical

Company. This technique involves atomizing an aqueous, organic, cosolvent solution,

aqueous-organic emulsion or suspension containing a drug and pharmaceutical excipients

directly into a compressed gas (i.e. CO2, helium, propane, ethane), or the cryogenic

liquids (i.e. nitrogen, argon, or hydrofluoroethers). The frozen particles are then

lyophilized to obtain dry and free-flowing micronized powders [164]. Use of acetonitrile

as the solvent increased the drug loading and decreased the drying time for lyophilization.

The dissolution rate was remarkably enhanced from the SFL powder containing

amorphous nanostructured aggregates with high surface area and excellent wettability

[66, 165, 166].

Evaporative precipitation into aqueous solution (EPAS): The EPAS process utilizes

rapid phase separation to nucleate and grow nanoparticles and microparticles of lipophilic

drugs. The drug is first dissolved in a low boiling point organic solvent. This solution is

pumped through a tube where it is heated under pressure to a temperature above the

solvent‘s boiling point and then sprayed through a fine atomizing nozzle into a heated

aqueous solution. Surfactants are added to the organic solution and the aqueous solution

to optimize particle formation and stabilization. In EPAS, the surfactant migrates to the

64

drug-water interface during particle formation, and the hydrophilic segment is oriented

towards the aqueous continuous phase [159]. The hydrophilic stabilizer on the surface

inhibits crystallization of the growing particles and therefore facilitates dissolution rates.

B. Modification of polymorphs

Polymorphism is the ability of an element or compound to crystallize in more

than one crystalline form. Different polymorphs of drugs are chemically identical, but

they exhibit different physicochemical properties including solubility, melting point,

density, texture, stability, vapour pressure, morphology, density and bioavailability [81,

167, 168]. Broadly, polymorphs can be classified as enantiotropes and monotropes based

on thermodynamic properties. In the case of an enantiotropic system, one polymorphic

form can change reversibly into another at a definite transition temperature below the

melting point, while no reversible transition is possible for monotropes. Once the drug

has been characterized under one of these categories, further study involves detection of

the metastable form of the crystal. Metastable forms are associated with higher energy

and thus higher solubility. Similarly, the amorphous form of drug is always more suited

than crystalline form due to higher energy associated and increases surface area [169].

Generally, the anhydrous form of a drug has greater solubility than the hydrates.

This is because the hydrates are already in interaction with water and therefore have less

energy for crystal break-up in comparison to the anhydrous (i.e. thermodynamically

higher energy state) for further interaction with water. On the other hand, the organic

(non-aqueous) solvates have greater solubility than the non-solvates. Some drugs can

exist in amorphous form (i.e. having no internal crystal structure). Such drugs represent

the highest energy state and can be considered as super cooled liquids. They have greater

65

aqueous solubility than the crystalline forms because they require less energy to transfer a

molecule into solvent. Thus, the order for dissolution of different solid forms of drug is

Amorphous >Metastable polymorph >Stable polymorph

Melting followed by a rapid cooling or recrystallization from different solvents

can produce metastable forms of a drug.

Metastable forms are associated with higher energy with increased surface area,

subsequently solubility, bioavailability and efficacy [168, 170]. With regard to

bioavailability, it is preferable to change a drug from crystal forms into metastable or

amorphous forms. However, the possibility of a conversion of the high energy amorphous

or metastable polymorph into a low energy crystal form having low solubility cannot be

ruled out during manufacture and storage. It is preferable to develop the most

thermodynamically stable polymorph of the drug to assure reproducible bioavailability of

the product over its shelf-life under a variety of real-world storage conditions. For

instance, ritonavir is the active ingredient in Norvir®, a protease inhibitor used to treat

HIV/AIDS. It was launched by Abbott Laboratories in 1996 as an amorphous semisolid

dispersion consisting of medium chain triglycerides, polyoxyl 35, castor oil, citric acid,

ethanol, polyglycolyzed glycerides, polysorbate 80, propylene glycol and 100 mg of

ritonavir. The dissolution and the oral bioavailability were decreased due to

crystallization of amorphous ritonavir into an insoluble crystal form during storage. This

polymorph (form II) was 50% less soluble than the original form in the market, and

caused the drug to fail its regulatory dissolution specifications. Finally, the drug was re-

launched with the form II polymorph in a soft gelatin formulation that required

refrigeration. Therefore, it is important to note that the selection of a polymorph of a drug

66

should balance between solubility and stability to maintain its potency over the shelf-life

period.

C. Drug dispersion in carriers

Solid solutions/dispersions

Solid dispersion was first introduced to overcome the low bioavailability of

lipophilic drugs by forming of eutectic mixtures of drugs with water-soluble carriers

[171]. It was defined as the dispersion of one or more active ingredients in an inert carrier

matrix in solid-state prepared by melting (fusion), solvent or melting-solvent method

[172]. More than 500 papers have been published on the subject and various materials are

employed as drug carriers (158). Despite active research interest, the number of marketed

products arising from this approach is disappointing mainly due to physical and chemical

instability and scale-up problems [173, 174]. Only two commercial products, a

griseofulvin in polyethylene glycol 8000 solid dispersion (Gris-PEG, Novartis) and a

nabilone in povidone solid dispersion (Cesamet, Lilly) were marketed during the last four

decades following the initial work of Sekiguchi and Obi (1961) [171].

Production methods

Solid solutions/dispersions are generally produced either by a solvent method, whereby

the drug and carrier are dissolved in a common solvent and then the solvent is evaporated

under vacuum (coevaporate), freeze-drying [175], spray-drying and spray–freezing into

liquid [66, 165]; or by a melting method, whereby drug-carrier mixtures are co-melted

and cooled. An important prerequisite to manufacture solid solutions/dispersions by the

hot melt method are the miscibility of the drugs and the carriers in the melt forms.

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Another limitation to the melt method is the thermo-instability of the drugs and

carriers. However, with the development of new techniques such as hot melt extrusion

[176] and hot spin melting [177], the second limitation associated with the melting

method was partially solved. For solvent-based methods, the ecological and subsequent

economic problems associated with the use of toxic organic solvents became more and

more problematic. Therefore, hot melt extrusion is the current method of choice for

preparation of solid dispersions. Briefly, the blend of drug and carrier is processed with a

twin-screw extruder of the same type used in the polymer industry. The blend is

simultaneously melted, homogenized, then extruded and shaped as tablets, granules,

pellets, sheets, sticks or powder. An important advantage of the hot melt extrusion

method is that the blend is only subjected to an elevated temperature for about 1 min,

which enables drugs or carriers that are thermolabile to be processed.

Carriers

Many water soluble excipients were employed as carriers of solid solutions/dispersions.

Among them, polyethylene glycols (PEG, Mw 1500-20000) were the most commonly

used due to their good solubility in water and in many organic solvents, low melting

points (under 65°C), ability to solubilize some compounds and improvement of

compound wettability. The marketed Gris-PEG is the solid dispersion of griseofulvin in

PEG 8000. The others carriers include polyvinyl pyrrolidone (PVP), polyvinylalcohol

(PVA), polyvinylpyrrolidone-polyvinylacetate copolymer (PVP-PVA), hydroxypropyl

methylcellulose (HPMC), hydroxypropyl cellulose (HPC), urea, Poloxamer 407, sugars,

emulsifiers (SDS, Tween 80) and organic acids (succinic acid and citric acid). Because of

the more rapid dissolution of the water-soluble carriers than the drugs, drug-rich layers

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are formed over the surfaces of dissolving plugs, which prevented further dissolution of

drug from solid dispersions. Therefore, surface-active or self-emulsifying agents

including bile salts, lecithin, lipid mixtures, Gelucire 44/14 [176] and Vitamin E TPGS

NF were used as additional additives, acting as dispersing or emulsifying carriers for the

liberated drug to prevent the formation of any water-insoluble surface layer. In addition,

the release behaviors of many drugs are also improved by using water-insoluble polymers

such as crospovidone and enteric polymers such as hydroxypropyl methylcellulose

phthalate (HPMCP), cellulose acetate phthalate (CAP), Eudragit® L100 and S100 and

Eudragit® E [178].

Challenges to solid dispersions

Although there has been a lot of interest in solid dispersion in the past four decades, the

commercial utilization is very limited. Problems of solid dispersion involve (i) method of

preparation, (ii) reproducibility of its physicochemical properties, (iii) formulation into

dosage forms, (iv) Scale-up of manufacturing processes and (v) Physical and chemical

stability of drugs and vehicles.

Method of preparation: High melting temperature may chemically decompose drugs

and carriers. No report addresses how much residual solvent is present in solid

dispersions when different solvents, carriers or drying techniques are used.

Reproducibility of physicochemical properties: Various investigators observe that

heating rate, maximum temperature used, holding time at a high temperature, cooling

method and rate, method of pulverization and particle size distribution may influence the

properties of solid dispersions prepared by the melting method. In addition, the nature of

69

solvent used, ratios of drug/solvent or carrier/solvent, solvent evaporation method and

rate may significantly affect the physicochemical properties of solid dispersions formed.

Dosage form development: Very few reports address the difficulty of pulverization and

sieving of the solid dispersion, which are usually soft and tacky with poor flow and

mixing properties. Thus, poor compressibility, drug-carrier incompatibility and poor

stability of the related dosage forms result.

Scale-up of manufacturing processes: Most solid dispersions reported in literatures are

prepared at the lab-scale. The scale-up of the preparation methods can be very

challenging. The physicochemical properties and stability of solid dispersions may be

affected by scale-up because heating and cooling rates of solid dispersion in large scale

differ from small-scale. It is also not practical and would be highly costly to evaporate

hundreds and even thousands of liters of organic solvents to prepare solid dispersion for

kilogram quantities of drugs. Removal of residual toxic organic solvent may be difficult

because the solid dispersions are usually amorphous and may exist in viscous and waxy

forms.

Stability: In a solid dispersion prepared by the melt method, a certain fraction of the drug

may remain molecularly dispersed depending on its solubility in the carrier. The excess

drug existing may greatly depend on the manufacture method. It may form a

supersaturated solution, separate out as an amorphous phase or crystallize out. The

supersaturated and amorphous forms may, in turn, crystallize out on aging. Certain

carriers may also exist in thermodynamically unstable states in solid dispersions and

undergo changes with time. As reported, polyvinyl pyrrolidone acts as stabilizer in the

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solid dispersion by retarding crystallization of drug at a low humidity. Hydrogen bonds

between the drug and PVP restrain drug crystallization [179].

D. Complexation

Complexation is the association between two or more molecules to form a non-

bonded entity with a well defined stoichiometry. Complexation relies on relatively weak

forces such as London forces, hydrogen bonding and hydrophobic interactions. There are

many types of complexing agents and types of complexes.

Inclusion complexes

Cyclodextrins and their derivatives have been employed as complexing agents to

increase water solubility, dissolution rate and bioavailability of lipophilic drugs for oral

or parenteral delivery [180, 181]. The solubility enhancement factors of pancratistatin,

hydrocortisone, and paclitaxel are 7.5, 72.7 and 99000 by forming complexes with

cyclodextrin derivatives [182]. The lower the aqueous solubility of the pure drug, the

greater the relative solubility enhancement obtained through cyclodextrin complexation.

Pharmaceutical applications of cyclodextins in drug solubilization and stabilization [182],

in vivo drug delivery, toxicological issues and safety evaluation [183] and mechanisms of

cyclodextrins modifying drug release from polymeric drug delivery systems have been

previously reviewed [184].

Cyclodextrins are a group of cyclic oligosaccharides obtained from enzymatic

degradation of starch. The three major cylcodextins α-, β-, and γ- (CD) are composed of

six, seven, and eight D-(+)-glucopyranose units. These agents have a torus structure with

primary and secondary hydroxyl groups orientated outwards. Consequently,

cyclodextrins have a hydrophilic exterior and a hydrophobic internal cavity. This cavity

71

enables cyclodextrins to complex ‗guest‘ drug molecules and hence alters the properties

of the drugs such as solubility, stability, bioavailability and toxicity profiles [182, 185].

The forces driving complexation by cyclodextrins were attributed to (i) the exclusion of

high energy water from the cavity, (ii) the release of ring strain particularly in the case of

α-CD, (iii) van der Waals interactions, and (iv) hydrogen and hydrophobic bindings

(186). β-CD, the most widely used native cyclodextrins, is limited in its pharmaceutical

application by its low aqueous solubility (1.85 g/100 ml, 25°C), toxicity profile and low

aqueous solubility of the formed complexes. Accordingly, derivatives such as

hydroxypropyl-β-CD (HP-β- CD; Enapsin®) and sulphobutylether-β-CD (SE-β-CD;

Captisol®) have been developed to produce more water-soluble and less toxic entities.

Staching complexation

Staching complexes are formed by the overlap of the planar regions of aromatic

molecules. Non-polar moieties tend to be squeezed out of water by the strong hydrogen

bonding interactions of water. This causes some molecules to minimize the contact with

water by aggregation of their hydrocarbon moieties. This aggregation is favored by large

planar non-polar regions in the molecule. Stached complexes can be homogeneous or

mixed. The former is known as self association and latter as complexation. Some

compounds that are known to form staching complexes include: nicotinamide [187],

anthracene, pyrene, methylene blue, benzoic acid, salicylic acid, ferulic acid, gentisic

acid, purine, theobromine, caffeine, and naphthalene.

Higuchi and Kristiansen (1970)

proposed a model according to which, the

compounds capable of undergoing stacking can be classified into two (classes A and B)

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based on their structure [188]. The compounds in class A have higher affinity for

compounds in class B than for those in class A and vice versa [185].

Factors affecting complexation

A number of factors affect complexation and include: steric effects, electronic

effects (effect of proximity of charge to CD cavity; effect of charge density; effect of

charge state of CD and drug), temperature, additives and cosolvent effects [189].

E. Solubilization by surfactants: Surfactants are molecules with distinct polar and

nonpolar regions. Most surfactants consist of a hydrocarbon segment connected to a polar

group. The polar group can be anionic, cationic, zwitterionic or non-ionic. When small

apolar molecules are added, they can accumulate in the hydrophobic core of the micelles.

This process of solubilization is very important in industrial and biological processes.

The presence of surfactants may lower the surface tension and increase the solubility of

the drug within an organic solvent [190].

F. Microemulsion

The term microemulsion was first used by Jack H. Shulman in 1959 [191]. A

microemulsion is a four-component system composed of external phase, internal phase,

surfactant and cosurfactant. In other words, microemulsion is a thermodynamically stable

isotropical dispersion composed of oil, a polar solvent, a surfactant and a cosurfactant.

The formation of microemulsions is spontaneous and does not involve the input of

external energy. One theory considers negative interfacial tension, while another

considers swollen micelles. The surfactant and the cosurfactant alternate each other

forming a mixed film at the interface contributing to the stability of the microemulsion.

Microemulsions are potential drug delivery systems for poorly water soluble drugs due to

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their ability to solubilize the drugs in the oil phase, thus increasing their dissolution rate

[192]. Even if the microemulsions are diluted after oral administration below the critical

micelles concentration (CMC), the resultant drug precipitates have a fine particle size

allowing enhanced absorption [193]. Non-ionic surfactants, such as Tweens

(polysorbates) and Labrafil (polyoxyethylated oleic glycerides), with high hyrophile-

lipophile balances are often used to ensure immediate formation of oil-in-water droplets

during production [190].

Advantages of microemulsion over coarse emulsion include its ease of

preparation due to spontaneous formation, thermodynamic stability, transparent and

elegant appearance, increased drug loading, enhanced penetration through the biological

membranes, increased bioavailability [193, 194], and less inter- and intra-individual

variability in drug pharmacokinetics [195].

G. Self-emulsification

In the absence of external phase (water), the mixture of oil, surfactant,

cosurfactant, one or more hydrophilic solvents and cosolvent forms a transparent

isotropic solution that is known as the self-emulsifying drug delivery system (SEDDS).

This forms fine O/W emulsions or microemulsions spontaneously upon dilution in the

aqueous phase and is used for improving lipophilic drug dissolution and absorption [192].

The self-emulsification process is specific to the nature of the oil/surfactant pair,

surfactant concentration, oil/surfactant ratio and temperature at which self-emulsification

occurs. The ease of emulsification could be associated with the ease of water penetrating

into the various liquid crystalline or gel phases formed on the surface of the droplet

74

One of the advantages of SEDDS in relation to scale-up and manufacture is that

they form spontaneously upon mixing their components under mild agitation and they are

thermodynamically stable. The drawbacks of this system include chemical instabilities of

drugs and high surfactant concentrations. The large quantity of surfactant in self-

emulsifying formulations (30-60%) irritates the GIT making safety a concern. Moreover,

volatile cosolvents in the conventional self-emulsifying formulations are known to

migrate into the shells of soft or hard gelatin capsules, resulting in the precipitation of the

lipophilic drugs. As an example of self-emulsification, Neoral® is composed of ethanol,

corn oil-mono-ditriglycerides, Cremophor RH 40 and propylene glycol. It exhibits less

variability and better drug uptake compared to Sandimmune®.

II. Chemical Modifications

A. pH adjustment

pH adjustment is the simplest and most commonly used method to increase water

solubility of ionizable compounds but discredits unionized compounds. The formed salts

may also convert to respective acid or base forms in gastrointestinal-tract (GIT).

For organic solutes that are ionizable, changing the pH of the system may be the

simplest and most effective means of increasing aqueous solubility. Under proper

conditions, the solubility of an ionizable drug can increase exponentially by adjusting the

pH of the solution. A drug that can be efficiently solubilized by pH control should be

either weak acid with a low pKa or a weak base with a high pKa. Similar to the lack of

effect of heat on the solubility of non-polar substances, there is little effect of pH on non-

ionizable substances. Non-ionizable, hydrophobic substances can have improved

solubility by changing the dielectric constant (a ratio of the capacitance of one material to

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a reference standard of the solvent by the use of co-solvents) rather than the pH of the

solvent.

The use of salt forms is a well known technique to enhance dissolution profiles

[196]. Salt formation is the most common and effective method of increasing solubility

and dissolution rates of acidic and basic drugs [197]. An alkaloid base is, generally,

slightly soluble in water, but if the pH of the medium is reduced by addition of acid, the

solubility of the base is increased as the pH continues to be reduced. The reason for this

increase in solubility is that the base is converted to a salt, which is relatively soluble in

water (e.g. Tribasic calcium phosphate). The solubility of slightly soluble acid increases

as the pH is increased by addition of alkali, the reason being that a salt is formed (e.g.

Aspirin, theophylline, barbiturates).

B. Other techniques:-

1. Co-crystallisation: The new approach available for the enhancement of drug solubility

is through the application of the co-crystals. This is also referred to as molecular

complexes. If the solvent is an integral part of the network structure and forms at least

two component crystals, then it may be termed as co-crystal. If the solvent does not

participate directly in the network itself, as in open framework structures, then it is

termed as clathrate (inclusion complex) [198]. A co-crystal may be defined as a

crystalline material that consists of two or more molecular (and electrically neutral)

species held together by non-covalent forces [180].

Co-crystals are more stable, particularly as the co-crystallizing agents are solids at

room temperature. Only three of the co-crystallizing agents are classified as generally

recognised as safe (GRAS) and include saccharin, nicotinamide and acetic acid limiting

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the pharmaceutical applications [198]. Co-crystallisation between two active

pharmaceutical ingredients has also been reported. This may require the use of

subtherapeutic amounts of drug substances such as aspirin or acetaminophen [181]. At

least twenty have been reported to date, including caffeine and glutaric acid polymorphic

co-crystals [182]. Co-crystals can be prepared by evaporation of a heteromeric solution or

by grinding the components together. Another technique for the preparation of co-crystals

includes sublimation, growth from the melt, and slurry preparation. The formation of

molecular complexes and co-crystals is becoming increasingly important as an alternative

to salt formation, particularly for neutral compounds or those having weakly ionizable

groups.

2. Cosolvency:

Cosolvents are the mixtures of miscible solvents often used to solubilize

lipophilic drugs. Currently, the water-soluble organic solvents include polyethylene

glycol 400 (PEG 400), ethanol, propylene glycol, and glycerin. For example, Procardia®

(nifedipine) developed by Pfizer contains glycerin, peppermint oil, PEG 400 and sodium

saccharin in soft gelatin capsules. The water-insoluble solvents include long-chain

triglycerides (i.e. peanut oil, corn oil, soybean oil, sesame oil, olive oil, peppermint oil,

hydrogenated vegetable oil and hydrogenated soybean oil), medium-chain triglycerides

(Miglyol 812), beeswax, d-α- tocopherol (vitamin E) and oleic acid. Progesterone is a

water-insoluble steroid and is solubilized in peanut oil (Prometrium®) [183].

Most cosolvents have hydrogen bond donor and/or acceptor groups as well as

small hydrocarbon regions. Their hydrophilic hydrogen bonding groups ensure water

miscibility, while their hydrophobic hydrocarbon regions interfere with water‘s hydrogen

77

bonding network, reducing the overall intermolecular attraction of water. By disrupting

waters self-association, cosolvents reduce water‘s ability to squeeze out non-polar,

hydrophobic compounds, thus increasing solubility.

A different perspective is that by simply making the polar water environment

more non-polar like the solute, cosolvents facilitate solubilization [183]. Solubility

enhancement as high as 500-fold was achieved using 20 % of 2-pyrrolidone to solubilize

nine poorly soluble compounds in aqueous solution in comparison to other solubilizers

like glycerin, propylene glycol, polyethylene glycol 400 or ethanol [184].

3. Hydrotrophy

Hydrotrophy designates the increase in solubility in water due to the presence of

large amount of additives. The mechanism by which it improves solubility is more

closely related to complexation involving a weak interaction between the hydrotrophic

agents (sodium benzoate, sodium acetate, sodium alginate, and urea) and the solute [185].

An example is the solubilisation of theophylline with sodium acetate and sodium

alginate.

In general, a drug administered in solution form is immediately available for

absorption and is more efficiently absorbed than the same amount of drug administered in

a tablet or capsule form. Solubility is thus the most important parameter for the oral

bioavailability of poorly soluble drugs. Drug dissolution is the rate determining step for

oral absorption of poorly water soluble drugs, which can subsequently affect the in vivo

absorption of drug. Because of the solubility problem of many drugs, their bioavailability

is affected and hence solubility enhancement becomes necessary. It is now possible to

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increase the solubility of poorly soluble drugs with the help of the various techniques

mentioned above.

1.12 Drugs used in the Study

In this study, the selection of the drug candidates was based on the BCS and

Classes II (Piroxicam and glibenclamide) and III (cimetidine) were represented.

a. Piroxicam

Piroxicam, a non-steroidal anti-inflammatory drug (NSAID), is used in the

treatment of dysmenorrhea, various acute and chronic musculoskeletal disorders like

rheumatoid arthritis, osteoarthritis etc., and also as a potent analgesics [186]. However,

the use of piroxicam has been associated with a number of gastrointestinal disorders

[187]. Enhanced bioavailability in a targeted delivery system based on improvement of

solubility is an alternative form, but requires a formulation which ensures total

solubilization of piroxicam in the host material. Solid lipid microparticle is such a system

that can enhance the performance of piroxicam in vivo in a self-emulsifying manner

thereby controlling the rate at which this drug is released in vivo and will therefore cause

less adverse effects normally associated with the drug in conventional dosage forms.

Several researchers have successfully delivered piroxicam via alternative forms like

organogel [188], buccal gel [189], mucoadhesive system [190], microspheres based drug

delivery [45, 191-193], iontophoresis [193], cyclodextrin based enhancement [194] and

gel based formulation which transdermally delivered piroxicam across the skin [195,

196]. Other studies show that dermal delivery of piroxicam had better stability in

proniosomal formulation as compared to niosomes

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b. Glibenclamide

The oral hypoglycaemic drug, glibenclamide, stimulates release of insulin from

pancreatic acinar cells, probably by blocking an ATP-sensitive K+ channel located in the

plasma membrane [197]. It is also one of the most potent inhibitors of the cromakalim-

activated K+ channel in smooth muscle [198-203].

c. Cimetidine

Cimetidine is one of several histamine H2-receptor antagonists widely used in

conditions where inhibition of gastric acid secretion may be beneficial, such as duodenal

and gastric ulcers [204]. It reduces pepsin output and competitively inhibits the action of

histamine at the histamine H2- receptors of the parietal cells [205]. Cimetidine has a wide

therapeutic index [204].

Cimetidine is slightly soluble in water. Its aqueous solubility is 11.4 mg/mL at 37

oC at a final pH of 9.3 [206, 207]. The minimum solubility determined in the pH range 1–

8 at 37 oC is 6 mg/mL [205]. The n-octanol/water partition coefficient (log P) of

cimetidine was reported as 2.5 at pH 9.2 [204, 205]. Cimetidine is weakly basic with the

pKa values reported as 6.808 and 6.93 [204, 205]. It is thus, present, at least partly, in the

ionized form in the upper gastrointestinal (GI) tract.

Cimetidine is rapidly, yet incompletely absorbed after oral administration. Its

bioavalability is between 56 – 68 % in healthy subjects and about 70 % in patients with

peptic ulcer, in whom a much greater variation in absorption was observed [204, 205]. In

the fed state, the absorption of cimetidine is slightly delayed but the extent absorbed is

not significantly different to that in the fasted state. A bioavailability study in a patient

with a massive bowel restriction demonstrated reduced absorption of cimetidine, which

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was attributed to rapid transit of the drug through the GI tract [208]. Both the absorption

and clearance of cimetidine are linear in the therapeutic dosing range (204). After oral

administration in the fasted state, cimetidine usually shows erratic double peak or

multiple peak phenomena in plasma drug concentration-time profile [204].

1.13 Problems to be addressed by this work

This piece of work aims to address certain problems perculiar to the lipid

matrices, drugs, and disease conditions to be treated by the delivery dosage form, SLM.

Lipid drug delivery systems have caught pace in the last decades employing lipids

of all origin – natural, semi-synthetic and synthetic. Trying to adapt this recent research

trend to our immediate environment and also in terms of bioremediation, survey of our

local abattoirs was done to find out some animal fats that were of relative abundance.

Goat fat, tallow fat and pig fat (lard) were found but as a matter of choice, goat and

tallow fats were chosen for this study alongside a commercial lipid, Softisan®

142 (a

coco-glyceride). Because these bulk lipids were crystalline (perfect crystal arrangement),

their crystal arrangements needed to be disordered to create spaces to improve their drug

holding potentials. This was done by adding a phospholipid (Phospholipon® 90G, P90G)

to these lipids, a process we coined as P90Gylation, synonymous with PEGylation. This

is because since these lipids have different fatty acids with varying lengths and degrees of

unsaturation, their P90Gylation would disorder the crystal arrangement/packing making

the matrices imperfect so as to be able to accommodate the studied drugs.

The basis for selecting the drugs was derived from the BCS which is a drug

development tool that considers the solubility and permeability of poorly-soluble drugs.

Because the BCS class II drugs have high permeability but low solubility, piroxcam and

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glibenclamide were selected to represent this class. It iss believed that sufficient

solubilization of these drugs in appropriate lipid matrices would increase their in vivo

performances. However, cimetidine was selected to represent the class III BCS drugs

which were characterized by high solubility and low permeability. It is believed that since

cinmetidine is slightly soluble in water that its permeability would be reduced due to the

fact that the biological membrane of the body has limited permeability to water soluble

drugs. As a matter a result, sufficient entrapement of cimetidine in an appropriate lipid

matrix would enhance its transport across the absorptive membranes of the body.

In terms of the disease states to be treated, this work was set to address some

common chronic conditions like inflammation (with piroxicam as an NSAID); diabetics

(glibenclamide as an antidiabetic) and ulcer (cimetidine as an antiulcer drug). It is

believed that these conditions are management diseases and so require more patient-

friendly dosage forms that are at least taken once daily unlike the conventional dosage

forms. It is believed that improved solubilization of these drugs in the lipid matrices and

subsequent formulation into solid lipid microparticles would retard their rate of release

because their small sizes make for long residence time in the GIT and would

conveniently comply to once-daily dosing with much reduction in the GIT disturbances

known for some of these drugs especially piroxicam.

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1.14 Objectives of present study

This study aims at:

Utilizing tallow fat, goat fat and Softisan® 142 in the development of SLM

suspension for poorly water-soluble drugs - piroxicam, glibenclamide and

cimetidine.

To induce various disease conditions – inflammation, diabetics and ulcer in

intact experimental animals and

To assess the SLM formulations for improved performance in terms of in vivo

release of the incorporated drugs thereby improving bioavailability.

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CHAPTER TWO

2.0 MATERIALS AND METHODS

2.1 Materials

Phospholipon® 90G (P90G) (Phospholipid GmbH Köln, Germany), is a purified,

deoiled and granulated soy lecithin with phosphatidylcholine content of at least 90 %.

Piroxicam and glibenclamide were kind gifts from Juhel Pharmaceuticals Nigeria Limited

(Enugu, Nigeria). Cimetidine CEMTAB® (Fidson Drugs, Nigeria) Softisan

® 142

(Pastillen, Germany), sorbic acid, sorbitol (BDH, England), and polysorbate 80 Tween®

80 (Uniqema, Belgium) were used as procured from their manufacturers without further

purification. Homolipids (tallow fat and goat fat) were from batches prepared in the

Pharmaceutics Laboratory of the University of Nigeria, Nsukka. Distilled water was

obtained from the University of Nigeria, Nsukka (Lion water).

2.2 Extraction and purification of homolipids

Goat fat was extracted from the adipose tissue of Capra hircus according to an earlier

method [192]. Briefly, the adipose tissue was collected from freshly slaughtered goat,

manually freed of extraneous materials, crushed and boiled in distilled water for 45 min,

filtered through a muslin cloth and allowed to solidify at room temperature. The solid fat

was manually removed and bleached/deodourized by passing it through a mixture of

activated charcoal and bentonite (2:1) at 100 ºC at a ratio of 10 g of the fat to 1 g of the

column material.

The above procedure was repeated using tallow fat from Bos indicus to obtain tallow

fat.

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2.3 Formulation of lipid matrices

The lipid matrices used in the formulations were a 4:1 mixture of goat fat and P90G;

tallow fat and P90G; and Softisan® 142 and P90G. They were prepared by fusion as

earlier described [2]. The lipids were weighed with an electronic balance (Mettler H8,

Switzerland), melted together at 60 oC on a thermo-regulated water bath shaker (Heto,

Denmark) and stirred until solidification.

2.4 Preparation of binary lipid matrices

Binary mixtures of goat fat and Softisan® 142 in the ratios of 1:1, 1:2 and 2:1

were prepared by fusion as described in section 2.3.

The above procedure was repeated for various combinations of tallow fat and

Softisan®

142 as well as goat fat and tallow fat combinations.

2.5 Incorporation of Phospholipon® 90G into the binary lipid matrices

The various lipid matrices of 1:1; 1:2, and 2:1 (section 2.4 above) were further

mixed with Phospholipon® 90G in a 4:1 ratio such that they separately contained 25 %

(w/w) of P90G in each of the 1:1; 1:2 and 2:1, binary mixtures of all three binary solid

lipid solutions. The lipids were prepared by fusion prior to microparticle preparation.

2.6 Characterization of the lipid matrices

2.6.1 Differential scanning calorimetry (DSC) of lipid matrices

Melting transitions and changes in heat capacity of the pure goat fat; tallow fat

and Softisan® 142 as bulk materials, and as physically structured lipid matrices were

determined by DSC (NETZSCH DSC 204 F1, Germany). Approximately, 3 – 5 mg of

each lipid matrix was weighed (Mettler M3 Microbalance, Switzerland) into an

aluminum pan, hermetically sealed, and the thermal behaviour determined in the range of

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35-190 oC under a 20 ml/min nitrogen flux at a heating rate of 10 ºC/min. The baselines

were determined using an empty pan and all the thermograms were baseline-corrected.

2.7 Formulation of unloaded solid lipid microparticles (SLMs)

SLMs were formulated to contain 5 % w/w of lipid matrix (4:1 mixture of goat fat

and P90G; tallow fat and P90G; and Softisan® 142 and P90G), graded concentrations of

polysorbate 80 (0.0, 0.75, 1.5, and 2 % w/w), 4 % w/w of sorbitol, 0.1 % w/w of sorbic

acid and enough distilled water to make 100 % w/w. The hot homogenization method

was adopted.

In each case, the lipid matrix was melted at 60 oC and the water containing

polysorbate 80, sorbitol and sorbic acid at the same temperature, was added to the molten

lipid matrix with gentle stirring on a magnetic stirring device (SR 1UM 52188, Remi

Equip., India). The mixture was further dispersed with a mixer (Silverson L4R, Adelphi

Manufac., England) at 6200 rpm for different emulsification times (2, 5, and 10 min) to

produce the hot primary emulsion, which was collected in hot containers and allowed to

recrystallize at room temperature.

2.8 Formulation of drug-loaded SLMs using single-structured lipid matrices

By adding piroxicam (graded concentrations of 250, 500, 750 and 1000 mg %) to

the lipidic phase (tallow fat structured with P90G); and glibenclamide (graded

concentrations of 100, 200, 300, 400 and 500 mg %) to the lipidic phase (Softisan® 142

structured with P90G) and cimetidine (graded concentrations of 50, 100 and 200 g %) to the

lipidic phase (goat fat structured with P90G) and following the previously described

procedure, piroxicam-, glibenclamde-, and cimetidine-loaded solid lipid microparticles

were obtained.

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2.9 Formulation of drug loaded-SLMs using the binary structured lipid matrices

By adding piroxicam (1.0 % w/w) to the binary structured lipidic matrices of

tallow fat-Softisan® 142 (1:1; 1:2, and 2:1) and following the above described procedure

(section 2.7), piroxicam-loaded solid lipid microparticles were obtained. This process

was repeated for glibenclamide (1.0 % w/w) in the structured lipidic matrices of

Softisan®

142-goat fat as well as cimetidine (10 % w/w) in the lipidic matrices of tallow

fat-goat fat. In each case, three determinations were undertaken for each ratio

combination of the matrices and mean values noted.

2.10 Evaluation of SLMs

2.10.1 Differential scanning calorimetry (DSC) of drugs and drug-loaded SLMs

Melting transitions and changes in heat capacity of the physically structured drug-

loaded lipid matrices were determined using a calorimeter (DCS 204F1) connected to a

disc station (NETZSCH, Germany) as previously described.

Subsequently, the thermal properties of the pure drug (piroxicam, glibenclamide

and cimetidine) were ascertained by DSC at different scan ranges of 35 – 250 oC for

piroxicam and 35 – 190 oC for glibenclamide and cimetidine. The thermal behaviour of

their SLM-containing formulations was also determined.

2.10.2 Particle size analysis and morphology of SLMs

Particle size analysis was carried out on the SLMs after production using a digital

light microscope (Leica Diestar, Germany) and images captured with Moticam 1000

camera (Magnification 65x). The morphology (shape and surface) of the particles was

also noted. The SLM were also subjected to time-resolved particle size analyses for 12

months at 6 month intervals to check the effect of storage on the particle size.

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2.10.3 Drug encapsulation efficiency

Approximately 6 ml of the piroxicam-loaded SLMs was added into a

microconcentrator (5, 000 MWCO Vivascience, Germany). This was centrifuged (TDL-4

B. Bran Scientific and Instru. Co., England) at 3,000 rpm for 120 min. The supernatants

was analyzed by UV/Vis Spectrophotometer (Unico 2102, England) at 332 nm. The

amount of drug encapsulated in the microparticles was calculated reference to a standard

Beer‘s plot to obtain the % encapsulation efficiency (EE) using the formula below:

EE (%) = Real drug loading X 100 …………………….….. (6)

theoretical drug loading

The above procedure was repeated for the glibenclamide-loaded SLMs and the

supernatants analyzed at 300 nm; for cimetidine-loaded SLMs, the wavelength was 254

nm.

2.10.4 In vitro diffusion studies

Franz diffusion cells with a receiver compartment volume of 20 mL and effective

diffusion area of 2.84 cm2 were used to evaluate drug delivery characteristics from the

selected compositions. A Millipore membrane filter (0.22 µm), (Millipore Corporation,

Billerica, MA) was used. The receptor phase (phosphate buffer solution, PBS, pH 7.4)

was continuously stirred and kept at a temperature of 37 ± 0.5 °C during the experiments.

A 1 ml volume of the drug-loaded SLM formulations was placed in the donor

compartment. At appropriate time, 1 ml of the sample was withdrawn from the receiver

compartment and the same amount of fresh solution was added to keep the volume

constant. Each experiment was run in three independent cells. The samples were analyzed

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spectrophotometrically at a wavelength of 332 nm and the concentration of piroxicam in

each sample was determined from a standard curve. Each data point represented the

average of three determinations. The release study was carried out for 24 h period. Sink

conditions were maintained throughout the experiment.

The above procedure was also repeated for the various batches of glibenclamide-

loaded SLMs and cimetidine-loaded SLMs with determinations at wavelengths, 300 nm

and 254 nm respectively.

2.10.5 Anti-inflammatory investigation

2.10.5.1 Preparation of experimental rats

Clinically normal male Sprague-Dawley albino rats weighing 200 ± 10 g and

normal male albino Wistar mice weighing 20 - 25 g were used for the experiment. The

animals were kept and maintained under laboratory conditions of temperature, humidity

and light; and allowed free access to food (standard pellet diet) and water ad libitum. All

the animals were fasted for 16 h, but still allowed free access to water, before

commencement of the experiments. The mice were used for the antinociceptive

evaluation of the piroxicam-loaded SLM; while the rats were used for the anti-

inflammatory investigation of the drug-loaded SLMs.

2.10.5.2 Evaluation of antinociceptive activity

The hot - plate (thermal) test method was used in this study. This method was

modified from those described elsewhere [209, 210]. A 600 ml glass beaker was placed

on a hot-plate with adjustable temperature (Heidolph® MR 2002). The temperature of the

hot-plate was then regulated to 45 ± 1 ºC. Each mouse was placed in the glass beaker (on

the hot-plate) in order to obtain the animal‘s response to electrical heat-induced

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nociceptive pain stimulus (licking of the forepaw and eventually jumping out of the glass

beaker). Jumping out of the beaker was taken as an indicator of the animal‘s response to

heat-induced nociceptive pain stimulus. The time taken for each mouse to jump out of the

beaker (i.e. reaction time) was noted and recorded in seconds. Each mouse served as its

own control. Thus, before treatment, its reaction time was determined thrice at 1 h

intervals. The mean of these determinations constituted the ‗initial reaction time‘ that is

reaction time before treatment of the mouse. The mean reaction time for all the mice were

pooled to get the final, ‗control‘ mean reaction time (Tb).

Each of the test mice was thereafter treated with either orally administered

distilled water (DW), piroxicam-loaded (SLM 1-4) or non-loaded SLMs (SLM-0),

commercial piroxicam sample (Feldene®

) or pure piroxicam sample dispersed in distilled

water (DW-P) (i.p). Twenty minutes after i.p. treatment with piroxicam, and oral

treatment with SLM formulations, commercial piroxicam sample and distilled water, the

reaction time was again evaluated. This value was pooled for the mice used in each

treatment group, and the final ‗test‘ mean reaction time value (Ta) for each treatment

group was calculated. This final ‗test‘ mean reaction time value represented ‗after

treatment reaction time‘ (Ta) for each group of treated mice. This ‗test‘ mean reaction

time value (Ta) was subsequently used to determine percentage thermal pain stimulus

(TPS) relief or protection, by applying the formula:

Protection against TPS (%) = test mean – control mean

Control

protection% = b

ba

T

TTX 100 …………..…………… (7)

The piroxicam-loaded SLM (obtained from the structured tallow fat) was tested at

concentrations of 2.5, 5.0, 7.5, and 10 mg/kg p.o. respectively. The commercial sample

90

was used at a dose of 10 mg; the pure piroxicam powder at a dose of 10 mg/ml i.p. only,

while another group received distilled water 3 ml/kg p.o. only.

The test was repeated with piroxicam-loaded SLMs formulated with structured

binary mixtures of tallow fat and Softisan®

142. Each of the test mice was treated with

either orally administered distilled water, drug - loaded SLMs (SLM-1a; SLM-2a; SLM-

3a) or non - loaded (SLM-1b; SLM-2b; SLM-3b), commercial piroxicam sample (S) and

pure piroxicam solution (DW-P) (i.p). Twenty minutes after i.p. treatment with

piroxicam, and oral treatment with SLM formulations, commercial sample and distilled

water, the reaction time was again evaluated, as described previously.

2.10.5.3 Evaluation of anti-inflammatory property of the SLMs

The rats used were divided into eight groups (DW, DW-P, S, SLM-0 and SLM 1-

4) of five rats each. The SLM-0 group served as the untreated control receiving only the

blank SLMs (without piroxicam). Each of the DW group received distilled water (3 ml/kg

p.o.) only, while the rats in the S-group received 10 mg/kg of a commercial sample of

piroxicam. The rats in the SLMs 1-4 received graded doses (2.5, 5.0, 7.5, and 10 mg/kg)

of piroxicam-loaded SLMs respectively. The group marked DW-P received pure

piroxicam powder in distilled water (10 mg/kg i.p.) each.

Rat hind paw oedema was used as a model of acute inflammation. The rat hind

paw oedema was induced by intra-plantar injection of fresh egg albumin (0.5 ml/kg), as a

cheap phlogistic agent [211, 212]. Acute inflammation of the hind paw was induced in

each of the rats by injecting 0.5 ml/kg of fresh egg albumin into the subplantar surface of

the right hind paw. Pedal inflammation (oedema) was evident within 5-8 min following

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fresh egg albumin injection. Two different test methods were used to assess the degree of

inflammation and protection.

The test was also repeated using piroxicam loaded into structured binary matrices

of tallow fat and Softisan® 142. The rats used were divided into nine groups (DW, DW-P,

S, and SLM 1a; SLM-2a; SLM-3a and SLM-1b; SLM-2b; SLM-3b) of five rats each. The

SLM-1b; SLM-2b and SLM-3b groups served as the untreated control receiving the blank

SLMs (without piroxicam), 3 ml/kg p.o, while each of the DW group received distilled

water (3 ml/kg p.o.) only. The rats in the S-group received 10 mg/kg of a commercial

sample of piroxicam (Feldene®

). Each test rat in the groups marked SLM-1a; SLM-2a

and SLM-3a received 10 mg /kg of piroxicam-loaded SLM. The group marked DW-P

received pure piroxicam powder in distilled water (10 mg/kg i.p.). The anti-inflammatory

test was then carried out as earlier discussed.

2.10.5.3.1 Linear diameter measurement

The linear diameter of the injected paw was measured for 3 h at 30 min intervals

after the administration of phlogistic agent. Increases in the linear diameter of the right

hind paws were taken as an indicator of paw oedema. Oedema was assessed in terms of

the difference in the ‗zero time‘ (Co) linear diameter at time t, (Ct – that is 30, 60, 90,

120, 150, and 180 min) following fresh egg albumin administration. The increase in the

right hind paw diameters induced by injections of fresh egg albumin was compared to

those of the contra-lateral, non – injected left hind paw diameters [211-213].

Graded doses of piroxicam-loaded SLMs were separately administered to each of

the rats in the test groups SLM 1-4 (i.e drug-loaded SLMs), 20 min before inducing

inflammation with the fresh egg albumin. Rats in the reference comparative ‗test‘ group

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marked DW-P received piroxicam (10 mg/kg i.p.) in distilled water; while the group

marked SLM-0 received blank (untreated) SLM; the rats in the group marked S received

10 mg/kg p.o. commercial brand of piroxicam and DW-rats received distilled water (3

ml/kg p.o.) only.

Percentage inflammation (oedema) was calculated from the formula

100% XC

Coedema

t

o ………………………………. (8)

while perentage inhibition of the oedema was calculated from the formula:

% inhibition = o

t

o

C

XC

C100

……………………………….. (9)

where Co is the average inflammation (hind paw oedema) of the control SLM-0 (blank

SLM) at any given time, and Ct is the average inflammation of the control DW (distilled

water) – SLM 1-4 (piroxicam-loaded SLMs) – or S (Commercial brand of piroxicam) –

or DW-P (pure piroxicam powder treated rats) at the same time.

The test was also repeated using piroxicam-loaded SLMs (SLM-1a; SLM-2a;

SLM-3a) prepared from structured-binary mixtures of tallow fat and Softisan® 142. The

formulations were separately administered (10 mg/kg p.o.) to each of the rats in the test

groups, 20 min before inducing inflammation with the fresh egg albumin. Rats in the

reference comparative ‗test‘ group marked DW-P received piroxicam (10 mg/kg i.p.) in

distilled water; while the untreated rats received SLM-1b; SLM-2b and SLM-3b; the rats

in the group marked S received 10 mg/kg p.o. commercial brand of piroxicam, Feldene®

and DW-rats received distilled water (3 ml/kg p.o.) only.

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2.10.5.3.2 Volume displacement method

Here, the volume of water displaced from 7.4 ml measuring cylinder was

measured immediately before the administration of the phlogistic agent and at 30 min

intervals for 3 h thereafter. For routine drug targeting, the increase in volume of water,

displaced 3 h after administration of the egg albumin was adopted as the parameter for

measuring inflammation.

Thus inflammation was assessed as the difference between zero time volume

displacement and displacement after 3 h following egg albumin administration. Exactly 1

h prior to the administration of the egg albumin, the SLMs 1-4 rat groups received 2.5,

5.0, 7.5, and 10.0 mg/kg p.o. respectively. The control groups marked SLM-0 and DW

respectively received blank SLMs and distilled water, 3 ml/kg p.o. The DW-P group

received pure piroxicam powder (10 mg/kg i.p.) dispersed in distilled water while the S

group received a commercial brand of piroxicam 10 mg/kg p.o.

The anti-inflammatory properties of the SLMs prepared from structured-binary lipid

matrices of tallow fat and Softisan® 142 was also assessed. The drug-treated rat groups

received SLM-1a; SLM-2a; and SLM-3a each as 10 mg/kg p.o respectively. The drug-

untreated groups received SLM-1b; SLM-2b and SLM-3b and distilled water respectively

3 ml/kg p.o. The DW-P group received pure piroxicam powder (10 mg/kg i.p.) in distilled

water, while the S group received a commercial brand of piroxicam 10 mg/kg p.o.

Percentage inflammation was calculated for each dose using the formula

% inflammation = Av. Inflammation time (t) X 100

Av. Inflammation of control at (t) ………………….……… (10)

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The initial volume of paw displacement was measured as Vi. At subsequent 30 min

intervals, the paw displacement was measured as Vf. The percentage oedema variation

was calculated using the expression below:

Oedema increase = 100XMV

MV

i

f ……….. …… .. …. (11)

% oedema inhibition = 100XDWcontrol

treatedDWcontrol

………….…….. (12)

where control (DW) = distilled water treated group.

1.10.5.3.3 Treatment of inflammation using piroxicam-loaded in structured binary

matrices

The above treatment protocol was repeated using the piroxicam loaded into the

various structured binary lipidic matrices of tallow fat-Softisan® 142 (1:1; 1:2, and 2:1).

2.10.6 Antidiabetic study

2.10.6.1 Preparation of experimental rats

Clinically normal male Wistar rats weighing 200 ± 10 g were prepared for the

experiment. Ab initio, the rats were supplied dry chick‘s mash finisher for adult rats

twice a day and given free access to tap water. They were acclimatized to the new

experimental environment for two weeks, housed separately in metabolic cages and

their body weights, consumption of food and water, urine volume and the levels of

serum glucose measured before the induction of diabetes. The rats were divided into

nine groups of five rats each.

2.10.6.2 Induction of diabetes mellitus

The rats were fasted for 24 h prior to the induction of diabetes mellitus. Blood

was collected for baseline glucose determination. The SLM formulations were

95

administered to the rats and the blood glucose checked at predetermined time intervals of

0, 1, 3, 7, 9, 12, and 24 h.

Fresh solution of alloxan monohydrate (Sigma, USA) was prepared just prior to

injection. Alloxan solution was made by dissolving alloxan in normal saline (0.9 % w/v

NaCl) as vehicle at a concentration of 100 mg/kg. This was given intra-peritoneally after

which the blood glucose levels were measured frequently for days using a glucometer

(ACCU-CHECK, Roche, USA). Food consumption was measured in (g), water (ml), and

urine volume (ml) on a daily basis. Diabetes was confirmed 3 days post-alloxan

administration.

2.10.6.3 Oral administration of glibenclamide-loaded SLMs

Nine treatment groups of five animals per group were assessed using glibenclamide-

loaded SLMs formulated using single-structured lipid matrices of Softisan® 142. The rat

group marked SLM-0 received blank SLM (i.e. without glibenclamide, 2 ml p.o). The

group marked DW received distilled water only (2 ml p.o), while that marked DW-G

received pure glibenclamide in distilled water (5 mg i.p.) and the commercial sample was

given to the last group. The other groups (SLM-1, SLM-2, SLM-3, SLM-4 and SLM-5)

received graded doses (1, 2, 3, 4, 5, mg/ml) of glibenclamide-loaded microparticles

respectively.

Subsequently, the treatment of diabetes using glibenclamide-loaded SLMs

formulated with binary-structured lipid matrices of Softisan® 142-goat fat was

investigated. Nine treatment groups of five animals per group were assessed. Three rat

groups were given zero-drug SLMs (i.e. blank SLM, 2 ml p.o) corresponding to SLM-4,

SLM-5 and SLM-6 containing SLMs from structured lipid matrices of goat fat and

96

Softisan® 142 in 1:1; 1:2 and 2:1 ratio combinations respectively while the group marked

DW received distilled water only (2 ml p.o); all as control. The group marked DW-G

received pure glibenclamide dispersed in distilled water (5 mg i.p.) while the commercial

sample (Daonil®) was given to the last group. The other groups received glibenclamide-

loaded microparticles as SLM-1, SLM-2, and SLM-3 corresponding to 5 mg/ml

respectively loaded into various matrices (1:1, 1:2 and 2:1) from where they were

prepared.

2.10.6.4 Pathological findings

After death or euthanasia, one rat in each group was selected for necropsy. Also,

one normal rat was sacrificed to compare the pancreatic islets of Langerhans. The

samples were fixed in 10 % formalin solution, stained with Hematoxylin & Eosin and

examined by microscopy (Leica Galen III, Leica Inc., USA).

2.10.7 In vivo investigation of ulcer

The aspirin model was employed to induce ulcer in the experimental rats. Wistar

male albino rats weighing 220-250 g obtained from the animal house of Department of

Pharmacology and Toxicology, University of Nigeria, Nigeria were used. The rats were

placed on standard feed and housing conditions and fasted over night before the

experiment. Thirty-six fasted rats were divided into 9 groups of 4 rats each.

The first three groups received (1 ml p.o.) cimetidine-loaded SLMs containing 5,

10 and 20 g %, their corresponding SLM- zero-drug concentrations (1 ml p.o.) were

given to the next three groups. The 7th

group received a commercial sample CEMTAB®

(1 ml p.o.); the 8th

group received pure cimetidine drug powder sample dispersed in

distilled water, while the last group received distilled water 1 ml p.o.

97

Subsequently, structured-binary lipid matrices loaded with cimetidine were

investigated for anti-ulcer activity in another set of ulcerated rats. Groups 1-6 were given

2 ml of the six different batches of the SLM preparations (i.e. three drug-loaded and three

zero-drug SLMs corresponding to the 1:1; 1:2 and 2:1 structured matrices of goat fat and

tallow fat) formulated with or without cimetidine (10 g %). Groups 7 and 8 received I ml

p.o. of 200 mg of cimetidine (CEMTAB®

) dispersed in distilled water and 2 ml of

distilled water per oral respectively, while Group 9 received 200 mg of pure cimetidine

powder in distilled water orally. One hour post administration, all rats were given 200

mg/kg of Aspirin p.o. and two hours later, they were sacrificed using ether, their

stomachs isolated and cut along the greater curvature. The stomachs were washed and

viewed with an x10 magnifying lens.

Ulcer scores were calculated as thus: ≤1mm = 1; >1 mm but ≤ 2mm = 2; > 2mm = 3

The scores were summed, divided by X10 magnification and averaged by number

of animals to get the mean ulcer indices from where the percentage ulcer inhibition (UI)

was calculated as:

= ………….. (13)

2.11 Stability studies of the formulations

The physical stability of the microparticles was evaluated for 12 months under

different temperature conditions. Some 6 ml volumes of each microparticle were stored

in closed glass bottles and placed at 4-6 oC; 25 ºC, and 40

oC away from direct light.

Aliquots were withdrawn every 6 months to determine particle size and morphology as

earlier described.

98

2.12 Determination of injectability

Injectability, defined as the smallest needle guage that a microparticulate sample

could pass through, was determined according to the method of Toongsuwan et al. (2004)

[214] but with little modification. This was carried out by pushing 4 ml of sample from a

5-ml plastic disposable syringe through hypodermic needles ranging from 18 to 27 within

20 sec. The formulation was first tested using the smallest needle (27 G). If the entire

content of the sample passed through a 27 G needle, its injectability was recorded as 27.

The study was repeated using 25 G needle, followed by the next smaller guage needle.

2.13 Statistical analysis

All experiments were performed in replicates for validity of statistical analysis.

Results were expressed as mean ± S.D. ANOVA and student‘s t-test was performed on

the data sets generated using Predictive Analytics SoftWare (PASW Statistics 18.0, 2009)

formerly called SPSS. Differences were considered significant for p-values < 0.05–

0.001.

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CHAPTER THREE

RESULTS AND DISCUSSION

3.1 Characterization of lipid matrices

The melting endotherm for tallow fat was 54.5 oC with an enthalpy of -5.067

mW/mg (Fig. 2). The DSC thermogram of goat fat showed an endothermic peak at 53.7

oC with an enthalpy of -6.42 mW/mg (Fig. 3). This melting point value was slightly

different from that in the literature and the possible reason may be a question of

sensitivity of the DSC machines used or because this preparation utilized distilled water

instead of bidistilled water [215]. The lower melting peak would belong to unstable

modification, while the higher peak belongs to stable modification. The DSC trace of

Softisan®

142 was 46.8 oC with an enthalpy of -7.962 mW/mg (Fig. 4). This melting

point value deviated from what is found on the product sheet or certificate of analysis

(42-44 oC) probably due to variation in sensitivity of the DSC machine.

The higher the enthalpy of the transitions, the more crystalline the matrix and

consequently, the more difficult it may be for any drug to be encapsulated [216]. This is

because highly crystalline matrices have perfect crystals without much space to entrap

any drug. Comparatively, it can be said that Softisan®

142 is the most crystalline (highest

enthalpy) of the three lipid matrices followed by goat fat and then tallow fat.

The structuring of these bulk crystalline matrices with P90G otherwise termed

P90Gylation, generally produced matrices with lower melting endotherms as well as

100

Fig. 2: DSC thermogram of tallow fat

Fig. 3: DSC thermogram of goat fat

101

Fig. 4: DSC thermogram of Softisan® 142

Fig. 5: DSC thermogram of P90G structured- tallow fat lipid matrix

102

Fig. 6: DSC thermogram of P90G structured- goat fat lipid matrix

Fig. 7: DSC thermogram of P90G structured- Softisan® 142 lipid matrix

103

enthalpies. For instance, when tallow fat was structured with P90G, the melting peak and

enthalpy changed from 54.5 oC and -5.067 mW/mg to 52.2

oC and -5.501 mW/mg (Fig.

5); goat fat changed from 53.7 oC and -6.42 mW/mg to 50.8

oC and -2.813 mW/mg (Fig.

6) while that of Softisan®

142 was from 46.8 oC and -7.962 mW/mg to 43.3

oC and -4.892

mW/mg (Fig 7). This is because P90G is a good surface modifier for solid lipid particles

[217, 218] with resultant improvement in targeting and pharmacokinetics [219, 220]. The

phospholipids bilayer structure formed around the lipid core may increase the drug

loading capacity, as biologically important molecules can be anchored on the colloidal

particle surface, and surface-modification also enables stabilization of colloidal particles

especially when generation of the microparticles is carried out in an aqueous medium

[221].

The thermotropic phase behaviour of a lipid matrix changes on encountering

guest molecules such that the thermodynamic variables of melting temperature and

changes in enthalpy depend on the nature of interaction between the constituents [216].

Since the degree of lipid crystallinity and the possible modifications in the lipid‘s solid

state are correlated with drug incorporation and release rates, and considering that the

drug‘s solid-state form (amorphous or crystalline) in solid dispersions influences

dissolution rates, it is important to pay special attention to these parameters [65, 95].

The determination of the thermal profile of the individual starting materials was

necessary in order to detect possible modifications in the physichochemical properties of

the drugs intended to be incorporated into SLMs and of the lipophilic excipients. It has

been shown that although particles were produced from crystalline materials, the

presence of emulsifiers, the preparation method and the high-shear dispersion may result

104

in changes in the crystallinity of matrix constituents compared with the bulk materials.

This may lead to liquid, amorphous or only partially crystallized metastable systems

[120, 121].

In addition, Phospholipon® 90G used in this study mainly contains linoleic, oleic,

stearic and palmitic acids, which are fatty acids of different chain lengths and degrees of

saturation [222]. The interaction of these fatty acids with the diverse fatty acids present in

goat fat, tallow fat and Softisan®

142 may have resulted in the partly amorphous nature of

the lipid matrix containing the phospholipids [223]. The fatty acid present in goat fat is

C16:0, C18:0, and C18:1, somewhat similar to that of theobroma oil and tallow fat alike

[215], while Softisan® 142 is a hydrogenated (saturated) coco-glyceride and which is

more homogeneous and melts sharply as against goat fat and tallow fat, which remain as

liquid crystals (solid/liquid) over a wide temperature range indicated by their broad

endotherms (Figs. 2 and 3) compared with that of Softisan® 142 (Fig. 4). Figs. 5-7 show

the observed structured modifications imparted by P90G on the bulk crystalline lipid

matrices of tallow fat, goat fat and Softisan® 142 respectively.

Figs. 8-10 respectively show the collective thermograms obtained when tallow fat

was combined with Softisan® 142 (Fig. 8), goat fat combined with Softisan

® 142 (Fig. 9)

and tallow fat combined with goat fat (Fig. 10) in different ratio combinations of 1:1; 1:2,

and 2:1.

105

Fig. 8: DSC thermograms of binary mixtures of tallow fat – Softisan® 142 matrices

Fig. 9: DSC thermograms of binary mixtures of goat fat – Softisan® 142 matrices

2:1

1:2

1:1

2:1

1:1 1:2

106

Fig. 10: DSC thermograms of binary mixtures of goat fat - tallow fat matrices

1:1

2:1

1:2

107

Considering the ratio combinations of tallow fat and Softisan® 142 as depicted in Fig. 8,

there were generally lower endothermic values as well as enthalpies for all ratio

combinations compared to the bulk crystalline tallow fat thermal properties. In other

words, Softisan® 142 modified the properties of tallow fat according to its proportion in

the mixture. Likewise, the tallow fat such that at 1:1 (i.e. 50:50) combination, the hybrid

matrix had high temperature of melting (51.9 oC) far above that of Softisan

® 142 (46.8

oC) but tending towards that of tallow (54.5

oC) but in terms of the enthalpy (-7.187

mW/mg), it tended more to the high crystallinity exhibited by Softisan® 142 (-7.962

mW/mg) rather than tallow fat (-5.067 mW/mg). At a combination of 1:2 (i.e. 25:75) of

tallow fat and Softisan®

142, although the melting endotherm (48.7 oC) tended more to

the Softisan® 142 side (46.8

oC) than the tallow fat (54.5

oC), the enthalpy (-7.901

mW/mg) remained crystalline. On analysis of the 2:1 (75:25) combination of tallow fat

and Softisan® 142, the endothermic peak (50.6

oC) of the hybrid matrix somewhat tended

towards tallow as was further confirmed by the enthalpy of -6.905 mW/mg signifying a

less crystalline matrix. The disorder in crystalline arrangement decreased in the order of

2:1>1:1>1:2 ratio combinations of tallow fat: Softisan® 142.

Mixtures of lipids have been shown to possess varied and mixed transition peaks

and have been suggested as alternatives to lipid modification by chemical techniques as

the latter often leads to products of decreased in vivo tolerability [224]. Fig. 9 shows the

thermograms of the binary mixtures of goat fat and Softisan® 142 matrices at different

ratio combinations of 1:1, 1:2 and 2:1. All the hybrid matrices showed modifications in

terms of reduction in crystallinity. For instance, the 50:50 mix of both lipids (1:1) gave a

matrix with peak melting endotherm of 49.9 oC compared to goat fat (53.7

oC) or

108

Softisan® 142 (46.8

oC) and an enthalpy of -4.578 mW/mg below those of goat fat (-6.42

mW/mg), and Softisan®

142 (-7.962 mW/mg). Interestingly, when the Softisan® 142

portion of the mix was higher (1:2) than that of the goat fat i.e. 25:75, the endothermic

temperature of the hybrid (47.5 oC) tended towards that of Softisan

® 142 (46.8

oC) but the

enthalpy (-5.572 mW/mg) remained less than that of both reactants. But the 75:25 mix of

goat fat and Softisan® 142 (i.e. 2:1) gave a hybrd matrix which melted at 50.0

oC with an

enthalpy of -4.115 mW/mg. This hybrid matrix, though having the highest melting

endotherm among the three matrices, was the least crystalline from the point of view of

the enthalpy.

Figure 10 shows the details of the DSC analysis of the binary mixtures of goat

and tallow fats in the ratios: 1:1, 1:2 and 2:1. A 1:1 (i.e. 50:50) mix of both lipids yielded

a hybrid matrix with melting endotherm of 54.0 oC and an enthalpy of -5.803 mW/mg.

This tended towards those of tallow fat (54.5 oC and -5.067 mW/mg) indicating a

somewhat amorphous system. The 25:75 mix of goat and tallow fats yielded a matrix

which melted at 54.1 oC but had an enthalpy of -8.298 mW/mg suggesting a high

crystalline matrix which can result in the expulsion of the entrapped drug on storage. The

75:25 mix of goat and tallow fats also gave a hybrid which melted at 53.6 oC with an

enthalpy of -7.811 mW/mg, suggesting a less crystalline system.

Crystallinity of lipid matrices affects the functional properties of the SLMs

derived from them. Lipid mixtures can result in increased or decreased crystallinity.

Directly, after preparation, lipids crystallize partially in higher energy modifications (α,

β') with more imperfections in the crystal lattice [156, 225]. If however, a polymorphic

transition to β modification takes place during storage, any incorporated drug could be

109

expelled from the lipid matrix and it can then neither be protected from degradation nor

released in a controlled manner. To overcome such phenomenon, use of mixtures of

lipids which do not form highly ordered crystalline arrangement is performed. Such lipid

matrix could be achieved by using solid lipid and liquid lipid [226] or solid lipid mixtures

of complex nature such as mono-, di- or triglycerides of different chain lengths [227].

Mixture of lipids also modifies the polymorphic properties of the individual lipids, and

has been shown to generate lipid matrices of low crystallinity [228].

The addition of P90G to these binary-lipid matrices resulted in further

modification of their properties. For instance, structuring 50:50 (1:1) combinations of

tallow fat and Softisan®

142 with P90G gave a matrix which melted at 51.0 oC with an

enthalpy of -4.981 mW/mg. This implies modification suggestive of deformation in the

lattice structure of the lipid constituents which perhaps may favour drug loading. The

P90G-structured 25:75 (i.e. 1:2) mix of tallow fat and Softisan® 142 yielded a matrix

which melted at 50.1 oC (a bit higher than that of the binary mixture of tallow and

Softisan® 142 i.e. non-structured) with an enthalpy of -8.526 mW/mg suggesting possible

modification with possible implications on incorporated-drug expulsion during storage.

But interestingly, when the 75:25 (2:1) counterpart mix of tallow fat and Softisan® 142

were structured with P90G, the resultant matrix melted at a temperature of 49.4 oC which

was the lowest melting temperature of all three structured binary matrices. Its

corresponding enthalpy was -2.391 mW/mg (Fig. 11), indicating a significant reduction

in crystallinity (half of that recorded for the 50:50 structured counterpart). In terms of

crystallinity, the employability of these structured matrices of tallow fat and Softisan®

142 in SLM productions is in the order: 2:1>1:1>1:2.

110

Fig. 11: DSC thermograms of P90G-structured tallow fat – Softisan® 142 matrices

1:1

2:1

1:2

111

Similarly, the result of the P90G-structuring of goat fat and Softisan® 142

combinations of 50:50 mix yielded a matrix which melted at 50.0 oC with an enthalpy of

-6.781 mW/mg, which was crystalline when compared to that of the binary mixture (-

4.578 mW/mg) before physical structuring. However, the corresponding 25:75 mix was

less crystalline with endothermic peak of 48.8 oC and an enthalpy of -5.168 mW/mg,

while that of 75:25 mix was the best and the least crystalline of all with peak melting

temperature of 50.3 oC and enthalpy of -2.511 mW/mg (Fig. 12).

The results of the physical structuring of goat fat and tallow fat combinations as

presented in Fig. 12 were also compared. The matrix obtained when 50:50 mix of goat fat

and tallow fat were structured with P90G melted at a peak temperature of 51.4 oC with an

enthalpy of -2.52 mW/mg, the structured matrix containing 25:75 mix of goat and tallow

fats had an endothermic peak of 51.7 oC and enthalpy of -2.766 mW/mg while the 75:25

structured mix melted at 52.0 oC with an enthalpy of -4.433 mW/mg. All three structured

matrices were less crystalline compared with the earlier values of their binary mixtures

prior to physical structuring with P90G with the 2:1 being the least crystalline of the three

(Fig. 13).

Table 1 shows a summary of the DSC measurements of the lipid matrices.

112

Fig. 12: DSC thermograms of P90G-structured goat fat – Softisan®

142 matrices

Fig. 13: DSC thermograms of P90G-structured goat fat – tallow fatt matrices

1:1

1:2

2:1

1:2 1:1

2:1

113

Table 1: Melting point and enthalpy measurements of the lipid matrices by DSC

Lipid matrix (mg) Composition Melting

point (oC)

Enthalpy

(mW/mg) Goat

Fat

Tallow

Fat

Softisan®

142

P90G

G - - - single 53.7 -6.42

- T - - single 54.5 -5.067

- - S - single 46.8 -7.962

G - - P Structured 50.8 -2.813

- T - P Structured 52.2 -5.501

- - S P Structured 43.3 -4.892

G T - - 1:1 54.0 -5.803

G T - - 1:2 54.1 -8.298

G T - - 2:1 53.6 -7.811

- T S - 1:1 51.9 -7.187

- T S - 1:2 48.7 -7.901

- T S - 2:1 50.6 -6.905

G - S - 1:1 49.9 -4.578

G - S - 1:2 47.5 -5.572

G - S - 2:1 50.0 -4.115

G T - P 1:1 51.4 -2.52

G T - P 1:2 51.7 -2.766

G T - P 2:1 52.0 -4.433

- T S P 1:1 51.0 -4.981

- T S P 1:2 50.1 -8.526

- T S P 2:1 49.4 -2.391

G - S P 1:1 50.0 -6.781

G - S P 1:2 48.8 -5.168

G - S P 2:1 50.3 -2.511

G means goat fat; T means tallow fat; S means Softisan

® 142, and P means Phospholipin

® 90G

114

3.2 Particle size analysis and morphology of the SLMs

The SLMs formulated were well formed, smooth and non-porous. They were also

stable and did not show sedimentation even after centrifugation (3000 rpm for 90 min).

The effects of production conditions on the SLM characteristics were optimized in terms

of the lipid matrix, surfactant concentration and emulsification time.

The particle size as presented in Table 2 shows that increase in polysorbate 80

concentration reduced the particle size. At 2 % w/w of polysorbate 80, the particle size

was difficult to determine probably due to the fact that they were no longer within

micrometer range. Higher emulsification time of 10 min generally produced smaller

microparticles which gelled in most cases making them unsuitable for oral drug

administration (Fig. 14a). The formulations containing 0.75 % w/w of polysorbate 80

generally had some un-emulsified entities at 2 min emulsification times. This made it

impossible for this concentration to be selected for subsequent production. However, the

formulation obtained with 0.75 % w/w of polysorbate 80 when the emulsification times

were increased to 5 and 10 min exhibited a phase separation over time probably due to

the small amount of the surfactant which inefficiently lowered the interfacial tension. For

all the SLM formulations, significant differences in particle size were observed after 10

min emulsification.

115

Table 2: Effects of different Polysorbate concentrations and emulsification times on

SLM mean diameters

Formula

tions

Polysorbate

80

percentage

(w/w)

Emulsificat

ion time

(min)

Particle mean diameter (µm) ± S.D.

Goat fat and

P90G

Tallow fat and

P90G

Softisan® 142

and P90G

SLM 1a 0.75 2 13.4 ± 1.3 13.9 ± 2.1 13.4 ± 2.0

SLM 1b 0.75 5 12.9 ± 1.0 12.8 ± 1.2 12.8 ± 1.2

SLM 1c 0.75 10 10.1 ± 0.75 10.0 ± 1.0 10.5 ± 1.0

SLM 2a 1.5 2 8.6 ± 2.0 8.0 ± 2.4 8.9 ± 2.2

SLM 2b 1.5 5 5.3 ± 2.5 5.5 ± 2.5 5.0 ± 2.5

SLM 2c 1.5 10 3.5 ± 2.2 2.0 ± 1.2 2.1 ± 1.0

SLM 3a 2.0 2 0.1 ± 0.01 0.09 ± 0.01 ND

SLM 3b 2.0 5 ND ND ND

SLM 3c 2.0 10 ND ND ND

ND implies not determined. Results are the mean of 3 measures ± S.D.

116

Fig. 14a: Photo showing a free flowing sample (upper) and a gelled sample (lower)

of SLMs

Free flowing sample

Gelled sample

117

Prolonged emulsification time decreased the particle size as a result of particle

coalescence leading to the particles having high kinetic energy [229]. This is in line with

the DLVO theory of colloidal particles in suspension, results in the collapse of all

repulsive forces giving room for particle collision with sufficient energy thereby

increasing the attractive forces, necessary to pull the particles into contact such that they

adhere strongly and irreversibly together. The particle sizes of the SLMs prepared with 2

% w/w concentration of polysorbate 80 could not be determined at emulsification times

of 5 and 10 min. SLMs with these characteristics may perform better as drug delivery

systems for topical or transdermal applications where particle size and particle size

stability may be overlooked.

The particle size analysis of the SLMs by light microscopy showed that there was

only slight variation in the size of the microparticles according to their lipid carriers.

There was however some variations within each carrier depending on the polysorbate

concentration and emulsification time (Table 2). The SLM formulations were further

observed physically to ascertain which had the best properties in terms of uniformity of

dispersion and fluidity, in addition to a more uniform size range. The microparticles

formed with the 1.5 % w/w of the polysorbate 80 were more uniform in size than those of

the 2 % w/w surfactant concentration at the emulsification time of 5 min (Fig. 14a). As a

result of all the foregoing, the 1.5 % w/w concentration of polysorbate 80 was selected

for subsequent studies because the SLMs obtained were uniformly dispersed without any

un-emulsified entities and had uniformly sized particles.

The assessment of the morphology and shape of the SLMs revealed smooth spherical

surfaces that are non-porous with more or less a ring of surfactant coat on the inner core

118

Table 3: Optimized working formula for SLM production

Ingredients Concentrations (%)

• Tallow fat; Goat fat; Softisan® 142 4.0 %

• Phospholipon 90G 1.0 %

• Polysorbate 80 1.5 %

• Sorbitol 4.0 %

• Sorbic acid 0.1 %

• Distilled water to 100 %

119

of the particles. Results of the particle morphology after 12 h of preparation and after

storage for 6 months at 25 °C are shown in Fig. 14b. Table 3 shows the optimized

working formula for subsequent SLM productions.

3.3 Morphology and particle size analysis of SLMs containing different

concentrations of piroxicam

The result of the particle size analysis of the piroxicam-loaded SLMs is shown in

Table 4. It shows that the size of the microparticles increased with increase in drug-

loading which agrees with reports by other workers [2, 16]. The SLMs increased in size

within the first 6 months of storage after which they maintained a steady particle size.

The photomicrographs of these microparticles (Fig. 15a) show a set of spherical

and smooth non-porous particles with a thick surfactant ring shielding the inner lipid

core. The chalky-appearances however depict some degree of lipid crystallization. Yet

the core of the microparticles maintained the pale-yellow colour of piroxicam. This was

evident in the samples stored at 4 – 6 oC (Fig. 15b). It was observed that increases in drug

loadings or O/W ratios caused increases in the sizes of the prepared microparticles

resulting in higher particle sizes. Upon 1 week of preparation, all the SLMs had a

120

A

Goat fat/P90G SLM 12 h after preparation

B

Goat fat/P90G SLM after 6 months storage

C

Tallow fat/P90G SLM 12 h after

preparation

D

Tallow fat/P90G SLM after 6 months

storage

E

Softisan® 142/P90G SLM 12 h after

preparation

F

Softisan® 142/P90G SLM after 6 months of

storage

Fig. 14b: Photomicrographs of SLM 2b (X100) within 1 week of formulation and

after 6 months storage, Magnification 65x.

121

Table 4: Properties of the SLMs loaded with graded concentrations of piroxicam

TF means Tallow fat; P90G means Phospholipon® 90 G; SLM-0 means blank SLM formulation without

piroxicam; SLM 1-4 means drug-loaded SLMs containing graded concentrations of piroxicam – 0.25, 0.5,

0.75 and 1.0 g w/v.

Formu-

lations

(TF/P9

0G)

Drug

compo

-sition

(g)

Average particle size

(µm)

Drug

encapsulatio

n efficiency

(%)

Injectability (Gauge) at 25 oC

After

prepa-

ration

After 6

months

storage

1 week

of

prepara

-tion

After 6

months

After 12

months

SLM-0 0.00 10.2 ±0.4 7.4±4.3 - 27 25 25

SLM-1 0.25 22.95±0.8 25.70±5.2 28.57 ± 10.30 27 25 25

SLM-2 0.50 50.50±0.9 153.90±28.3 50.00 ± 20.30 27 18 18

SLM 3 0.75 90.5±1.2 273.30±10.1 53.30 ± 23.20 27 23 23

SLM-4 1.0 106.5±3.7 378.70±25.7 57.14 ± 20.50 27 23 23

122

A

A´

B B´

C

C´

D D´

E E´

Fig. 15a: Photomicrographs of the SLM dispersions; (A) Blank SLM; (B) 0.25 %

w/w, (C) 0.5 % w/w, (D) 0.75 %w/w, and (E) 1.0 % w/w piroxicam-loaded SLM

after one week of preparation and their corresponding photomicrographs after

storage for six months denotated as A´, B´, C´, D´, and E´ respectively. (Mag. 65x)

123

A

B

C D

E

Fig. 15b: Stability studies of the piroxicam-loaded singly-structured TF/P90G SLMs

after 6 months storage at 4 oC (Magnification 65x)

[A. blank; B. 0.25 g; C. 0.5 g; D.0.75 g and E. 1.0 g w/w of piroxicam]

124

syringeability of 27 G but varied upon storage at room temperature. The SLMs were best

stored at 4-6 oC because the samples stored at this temperature remained syringeable with

27 G (Table 4). Moreover, when the particle morphology was re-investigated after six

months, it was found that there were particle growth in that the size of the particles

increased especially with higher drug loadings (Fig. 15a and 15b). However, this particle

growth does not make the oral delivery of piroxicam using P90G-structured tallow fat

matrix unsuitable since there is no strict limit in particle size and particle size stability in

oral delivery systems [216].

3.4 Morphology and particle size analysis of SLMs containing different

concentrations of glibenclamide

The particle size analysis of the microparticulate dispersion by light microscopy

showed mean particle size of 5.5 - 173.9 µm (Table 5). The photomicrographs of the

SLMs after one week of formulation illustrate the spherical shape of the solid lipid

microparticles entrapping the glibenclamide (Fig. 16a) and after 6 months of storage (Fig.

16b). It shows the homogeneous monolayer coating of surfactant at the periphery of the

microparticles surrounding the lipid core. However, increase in the size of the SLMs did

not affect their shapes.

3.5 Morphology and particle size analysis of SLMs containing different

concentrations of cimetidine

A similar particle size observation was seen in the cimetidine-loaded SLMs (Table 6) and

its morphology in Fig. 16c within one week of preparation and after 6 months of storage

The increase in particle size with increasing drug loading has been observed by other

authors [230]. Increasing the O/W ratio leads to a decrease in particle size whereas

coalescence of droplets can be prevented by a large amount of aqueous phase

125

Table 5: Properties of the SLMs loaded with graded concentrations of glibenclamide

Formu-

Lations

(SFT/P9

0G)

Drug

compo

sition

(g)

Particle size (µm) Drug

encapsulatio

n efficiency

(%)

Injectability (Gauge) at

25 oC

After

preparation

After 6

months

storage

1

week

old

After

6

month

s

After 12

months

SLM-0 0.00 5.5 ± 1.6 95.4±14.2 - 27 18 18

SLM-1 0.1 8.95 ± 1.51 50.9±8.6 8.33 ± 2.60 27 25 25

SLM-2 0.2 15.50 ± 2.18 205.6±25.8 41.67 ± 15.20 27 27 25

SLM 3 0.3 90.60 ± 15.23 278.30±30.7 55.56 ± 20.70 27 27 25

SLM-4 0.4 145.7 ± 18.45 369.60±30.7 58.33 ± 23.80 27 27 25

SLM-5 0.5 173.9 ± 19.30 450.80±40.5 60.58 ± 25.00 27 27 25

STF means Softisan® 142; P90G means Phospholipon

® 90 G; SLM-0 means blank SLM formulation

without glibenclamide; SLM 1-4 means drug-loaded SLMs containing graded concentrations of

glibenclamide – 0.1, 0.2, 0.3, 0.4 and 0.5 g w/v.

126

A B C

D E F

Fig. 16a: Photomicrographs of the SLM dispersions; (A) Blank SLM; (B) 0.1 %

w/w, (C) 0.2 % w/w, (D) 0.3 %w/w, (E) 0.4 % w/w and (F) 0.5 % w/w

glibenclamide-loaded SLM after one week of preparation. (Mag. 65x)

127

A B

C D

Fig. 16b: Photomicrographs of the SLM dispersions; (A) Blank SLM; (B) 0.1 %

w/w, (C) 0.2 % w/w, and (D) 0.5 % w/w glibenclamide-loaded SLM after

six months of preparation. (Magnification 65x)

128

Table 6: Properties of the SLMs loaded with graded concentrations of cimetidine

NS = not syringeable; GF means goat fat; P90G means Phospholipon® 90 G; SLM-0 means blank

SLM formulation without cimetidine; SLM 1-4 means drug-loaded SLMs containing graded concentrations

of cimetidine – 0.05, 0.10 and 0.20 g w/v.

Formu-

lations

(GF/P9

0G)

Drug

composi

tion (g)

Average particle size (µm) Drug

encapsulati

on

efficiency

(%)

Injectability (Gauge) at 25 oC

After

preparatio

n

After 6

months

storage

1 week

of

prepara

tion

After 6

months

After

12

month

s

SLM-0 0.00 3.50±0.9 5.34±0.3 - 27 23 18

SLM-1 0.05 4.23±1.2 5.97±1.6 22.54 ± 3.40 27 23 18

SLM-2 0.10 10.71±1.5 56.68±10.8 25.00 ± 7.50 27 18 NS

SLM 3 0.20 21.36±2.0 110.76±35.9 17.21 ± 4.90 27 23 NS

129

A

A΄

B

B΄

C

C΄

D

D΄

Fig. 16c: Photomicrographs of the SLM dispersions; (A) Blank SLM; (B) 0.05 %,

(C) 0.10 % and (D) 0.2 % w/w cimetidine-loaded SLM within one

week of preparation and (A΄-D΄) after six months of preparation. (Mag. 65x)

130

available for diffusion in the O/W emulsion and hence smaller particles were produced

[231]. Upon storage, the microparticles grew in size in line with the findings of other

authors [4].

3.6 DSC analysis of the drugs and SLM formulations

The thermograms showed that piroxicam was the most crystalline of the three

drugs followed by glibenclamide and then cimetidine. Piroxicam had a sharp endothermic

peak at 203.1 oC with an enthalpy of -6.354 mW/mg (Fig. 17), glibenclamide melted at

175.3 oC with an enthalpy of -4.696 mW/mg (Fig. 18) wheres as that of cimetidine was

145.3 oC and an enthalpy of -2.759 mW/mg (Fig. 19).

Firstly, P90G-structured tallow fat matrix was loaded with piroxicam or not

containing piroxicam as was the case with the zero-piroxicam batch. The piroxicam-

loaded SLMs showed different endothermic peaks independent of drug loading. SLM-1,

SLM-2, SLM-3 and SLM-4 containing 0.25 g; 0.5 g; 0.75 g and 1.0 g w/w of piroxicam

respectively, showed endothermic peaks at 109.8 ºC; 95.0 ºC; 114.8 ºC and two

endothermic transitions which occurred at 78.8 ºC and 106.9 ºC for the 1.0 g w/w

piroxicam-loaded SLM-4 (Fig. 20). This implies that drug loadings resulted in a shift of

the melting endotherm towards the lower temperature ranges except for the 0.75 g

piroxicam-loaded SLM which significantly had a shift to higher temperature of melting.

131

Fig. 17: DSC thermogram of pure piroxicam

Fig. 18: DSC thermogram of pure glibenclamide

132

Fig. 19: DSC thermogram of pure cimetidine.

133

This suggests that the piroxicam-loadings in the SLMs formulated with P90G-structured

tallow fat matrices generally produced less-ordered crystals or amorphous state, such that

the melting of the substances required less energy than the perfect crystalline substance

(blank SLM), which needs to overcome lattice forces. However, the decrease in melting

point is associated with numerous lattice defects and the formation of amorphous regions

in which the drug is located. DSC thermogram of SLM dispersion without piroxicam

formulated with structured tallow fat matrix (52.2 oC with an enthalpy of -5.501 mW/mg)

alone showed two endothermic transitions with peak minima at 104.8oC and 108.8

oC.

Figure 20 shows the collective thermograms of the SLMs formulated with the P90G-

structured tallow fat containing graded concentrations of piroxicam.

Figure 21 shows the thermograms when P90G-structured Softisan® 142 matrix

was employed as delivery carrier for glibenclamide, ceteris paribus.

The different drug concentrations of glibenclamide generally had lower melting

endotherms as well as enthalpies as compared to the zero-drug counterpart. However, the

zero-drug SLM had a peak melting endotherm of 104.3 oC with an enthalpy of -16.58

mW/mg whereas the drug-loaded batches starting with the 0.1, 0.2, and 0.5 % w/w

respectively had 77.8 oC; 73.8

oC; 59.7

oC endothermic peaks and -15.07; -16.51; -12.52

mW/mg enthalpies. This implies that there was a decline generally to the lower

temperature side with the glibenclamide loadings on the SLMs formulated from

structured Softisan® 142 matrices suggesting less crystalline matrices with the

consequence of enhanced solubilization and entrapment of the glibenclamide in the core

of the microparticles. The long term benefit is suggestive of a prolonged release carrier

system with improved bioavailability performance.

134

Fig. 20: DSC thermograms of SLM formulations with P90G-structured tallow fat

matrices containing graded concentrations of piroxicam

Fig. 21: DSC thermograms of SLM formulations with P90G-structured Softisan®

142 matrices containing graded concentrations of glibenclamide

1

2

3

1

4

5

1 2

3

4 5

1

135

Table 7 shows the DSC results of the drug-loaded SLMs formulated from the singly-

structured lipid matrices.

When the 1:1 structured tallow fat- Softisan®

142 (50:50 mix) was employed in

formulation, the resultant SLMs had a melting peak of 105 oC and enthalpy of -13.87

mW/mg while on drug loading, it melted at a lower temperature of 98 oC but a higher

enthalpy of -15.4 mW/mg (Fig. 22). With the 25:75 mix, the resultant SLM without drug

melted at 119.5 oC with an enthalpy of -11.11 mW/mg, while its corresponding drug-

loaded counterpart melted at a lower temperature of 95.8 oC and had an enthalpy of -

13.57. For the structured 75:25 mix, the SLM resulting from the formulation had a

melting endothermic peak at 116.5 oC and enthalpy of -11.11 mW/mg, while its drug

loaded counterpart had a melting peak of 106.9 oC and enthalpy of -13.38 mW/mg. The

increase in enthalpy confirms higher amounts of crystals upon storage due to delayed

crystallization from fractions of a cooled amorphous melt. There was a general shift to

the lower temperatures in all the drug-loaded samples (Fig. 22).

When goat fat and Softisan® 142 structured admixtures were used to formulate

glibenclamide SLMs, the features of the DSC profile changed (Fig. 23). With the

structured non-drug – loaded 50:50 matrix, the resultant SLM melted at 114.3 oC with an

enthalpy of -10.91 mW/mg, while its glibenclamide-loaded counterpart recorded an

endothermic temperature of 104.7 oC and enthalpy of -13.16. The SLM resulting from the

structured 25:75 mix had peak endotherm at 111.9 oC and enthalpy of -18.72 mW/mg,

while its drug-loaded counterpart melted at 120.5 oC with an enthalpy of -11.61 mW/mg.

When the 75:25 mix was employed, the resultant SLM melted at 115.6 oC with an

136

Table 7: DSC properties of the SLMs produced using singly-structured SLMs.

Drugs Lipid matrix (mg) Composition

(g) w/w

Melting

point

(oC)

Enthalpy

(mW/mg) Goat

Fat

Tallow

Fat

Softisan®

142

P90G

Cimetidine G - - P Blank SLM 119.8 -18.32

,, G - - P 0.05 104.7 -16.70

,, G - - P 0.1 80.75 -9.56

,, G - - P 0.2 114.7 -13.11

Piroxicam - T - P Blank SLM 104.8

108.8

-13.31

-14.67

,, - T - P 0.25 109.8 -16.12

,, - T - P 0.50 95.0 -9.782

,, - T - P 0.75 114.8 -12.17

,, - T - P 1.00 78.8

106.9

-9.155

-6.717

Glibenclamide - - S P Blank SLM 104.3 -16.58

,, - - S P 0.1 77.8 -15.07

,, - - S P 0.2 73.8 -16.51

,, - - S P 0.5 59.7 -12.52

G means goat fat; T means tallow fat; S means Softisan® 142; P means P90G.

137

Fig. 22: DSC thermograms of SLM formulations with P90G-structured

tallow fat and Softisan® 142 matrices containig piroxicam

Fig. 23: DSC thermograms of SLM formulations with P90G-structured

goat fat and Softisan® 142 matrices to containing glibenclamide

1

2

3

4

5

1

2

3

4

1

5

1

6

1

138

enthalpy of -13.09 mW/mg, while its drug-containing part melted at 117 oC with an

enthalpy of -12.5 mW/mg. The lower values of enthalpy here are clear indication of low

crystallinity and improved drug holding capacity as a result of lattice structural

deformation thereby creating spaces for the accommodation of the incorporated drug.

Figure 24 shows the result of the DSC analysis of the SLMs (with or without

cimetidine) produced using the structured matrices of goat and tallow fats. A plain

structured 50:50 mix of this matrix yielded an SLM formulation which melted at 118.1

oC with an enthalpy of -12.1 mW/mg, while its drug-containing counterpart traced two

endothermic transitions which occurred at 120 oC with an enthalpy of -15.67 mW/mg for

the lower peak and at 124 oC with an enthalpy of -16.43 mW/mg for the higher peak.

This high temperature of melting of the SLMs was quite closer to the melting point of the

incorporated drug rather than the lipid matrices. However, when the 25:75 mix of the

same matrix was used, the resultant SLM melted at 119.8 oC with an enthalpy of -18.32

mW/mg, while its drug-containing counterpart melted at 120.82 oC and had an enthalpy

of -17.52 mW/mg. On further varying the composition to 75:25, the resultant SLM

recorded a melting peak of 113.2 oC and -9.336 mW/mg value of enthalpy, while its

drug-loaded counterpart melted at 107.5 oC with an enthalpy of -11.77 mW/mg. It

follows that with lipid drug delivery systems, polymorphic transformations may occur

during dosage form preparation and subsequent storage. During the melt solidification,

triglycerides and fatty acids in particular can crystallize into different polymorphic forms

(i.e., the thermodynamically instable α-form, the β′- form, the stable β-form) depending

on lipid composition and cooling rates.

139

Fig. 24: DSC thermograms of SLM formulations with P90G-structured

Goat fat and tallow fat matrices containing cimetidine

1

2

3

4

5

140

Polymorphic transformations may cause changes in active and auxiliary substances

solubilities and melting points. In particular, the conversion of one polymorph into

another may change the physical properties of the substance [1, 97, 104, 107]. Table 8

shows the details of the DSC measurements of the drug-loaded SLMs formulated from

the binary-structured lipid matrices.

3.6 Drug encapsulation efficiency

Table 4 (section 3.3) shows the results of the efficiency of loading of graded

concentrations of piroxicam on the optimized SLMs formulated with P90G-structured

tallow fat. Drug loading efficiency increased with increased drug concentration. In other

words, the percentage loading efficiency increased with increased drug loading. The

determination of drug loading (or drug incorporation) is an important tool to evaluate a

potential drug carrier system. It is obviously desirable to produce microparticles with

high drug content in order to decrease the amount of microparticles to be administered,

whatever the administration route. The prerequisite to obtain optimal loading capacity is

a sufficiently high solubility of the drug in the lipid melt. The solubilizers (active and

passive) in addition to the lipids used as matrices promoted drug solubilization. The

chemical nature of the lipid is also important because lipids which form highly crystalline

particles with perfect lattice lead to drug expulsion [120]. P90G structured-tallow fat

matrix contains fatty acids of different chain lengths and thus formed crystals with many

imperfections which may have offered spaces to accommodate the piroxicam. SLMs

formed without P90G were gelled and showed some physical instability. Alternatively,

intensive characterization of the physical state of the

141

Table 8: DSC properties of the SLMs produced using structured binary lipid

matrices

Drugs Lipid matrix (mg) Composition

(g) w/w

Melting

point

(oC)

Enthalpy

(mW/mg) Goat

Fat

Tallow

Fat

Softisan®

142

P90G

Cimetidine G T - P 1:1 120

124

-15.67

-16.43

Blank G T - P 1:1 118.1 -12.1

Cimetidine G T - P 1:2 120.82 -17.52

Blank G T - P 1:2 119.8 -18.32

cimetidine G T - P 2:1 107.5 -11.77

Blank G T - P 2:1 113.2 -9.366

Piroxicam

-

T

S

P

1:1

98.0

-15.4

Blank - T S P 1:1 105.0 -13.87

Piroxicam - T S P 1:2 95.8 -13.57

Blank - T S P 1:2 119.5 -11.11

Piroxicam - T S P 2:1 106.9 -13.38

Blank - T S P 2:1 116.5 -11.11

Glibenclamide

G

-

S

P

1:1

104.7

-13.16

Blank G - S P 1:1 114.3 -10.91

Glibenclamide G - S P 1:2 120.5 -11.61

Blank G - S P 1:2 111.9 -18.72

Glibenclamide G - S P 2:1 117.0 -12.5

Blank G - S P 2:1 115.6 -13.09

142

lipid particles by DSC, NMR, X-ray and other techniques is highly essential for a

controlled optimization of drug incorporation and drug loading. However, only DSC

method could be used in this work.

The same was true for SLMs based on the P90G-structured Softisan® 142 matrix

used for the delivery of glibenclamide, Table 5 (section 3.4). The drug loading efficiency

increased with increase in the concentration of the drugs such that the maximum

percentage drug loading was 60.58 ± 25.0 % whereas the minimum percentage drug

loading was 8.33 ± 2.60 %. The trend was the same for the cimetidine – loaded SLMs,

Table 6 (section 3.5). The maximum loading efficiency for cimetidine was only 25.00 ±

7.5 %.

A number of factors affect the loading efficiency of drug in the lipid. Among

them are solubility of the drug in the melted lipid; miscibility of drug melt and lipid melt;

chemical and physical structure of solid lipid matrix as well as polymorphic state of lipid

material (4). Low loading efficiency may result from crystallization of the matrix which

differs from crystallization of the SLM. Lipid microparticles recrystallize at least partially

in the α-form, whereas bulk lipids tend to recrystallize preferentially in the β΄-

modification and transforming rapidly into the β-form (121). With increasing formation

of the more stable modifications the lattice gets more perfect and the number of

imperfections decreases, implying that formation of β΄/ βi – modifications promotes drug

expulsion. In general, the transformation is slower for long-chain than short-chain

triglycerides [232].

The concentration of the drugs yielding the highest loading efficiency was

employed further when the lipids were mixed amongst themselves, further templated or

143

P90Gylated and used to deliver the drugs. Thus, piroxicam was loaded at a concentration

of 1.0 g w/w; glibenclamide at 0.5 g w/w and cimetidine was 10 g w/w into the respective

structured ratio combinations of each binary lipid admixture.

Tables 9-11 show the result of the above studies and all the SLMs generally

recorded higher percentage drug loadings. This is probably due to the fact that mixtures

of these lipids were so imperfect that cavities existed much that allowed for more drug

particles to be entrapped. Table 9 shows the result of the properties of the piroxicam-

loaded SLMs formulated with P90G-structured tallow fat – Softisan® 142 matrices. The

result also corresponds to the earlier findings from the solid state characterization using

DSC. The 2:1 matrix was less crystalline and thus allowing the drug to be highly

entrapped up to 68.50 ± 10.30 % followed by the 1:1 matrix which had an entrapment

efficiency of 58.47 ± 11.50 %. The 1:2 matrix was the most crystalline and so had the

least drug incorporation of 50.30 ± 13.20 %.

The smaller the unit dose of a drug the higher its encapsulation efficiency

provided it is lipophilic enough. This was the case for glibenclamide (Table 10) with a

dose of 5 mg which at 1.0 g w/w, had loading efficiency of 70.35 ± 7.73 % w/w in the

structured 2:1 goat fat-Softisan® 142 structured matrix. This was the highest

encapsulation efficiency amongst all the formulations. The corresponding 1:2 and 1:1

structured matrices had 64.89 ± 10.21 % and 58.23 ± 5.64 % respectively. This agrees

with the result of the DSC analysis which showed the 1:1 matrix as the most crystalline in

terms of the enthalpy.

144

Table 9: Properties of the piroxicam-loaded SLMs formulated with P90G-structured

tallow fat and Softisan®

142

*Tallow fat: Softisan® 142

refers to the lipid matrices containing 25 % of P90G.

Formula

tions

*Tallo

w fat:

Softis

an®

142

Drug

composi

tion

(g) w/w

Average particle sizes

at different storage

temperatures (µm)

Drug

encapsulatio

n efficiency

(%)

Injectability (Gauge)

at 25 oC

1

wee

k

After

6

month

s

After

12

month

s

4 oC 25

oC 40

oC

SLM-1a 1:1 1.0 9.30 17.80 26.50

58.47 ± 11.50 27 25 25

SLM-2a 1:2 1.0 10.50 11.70 20.50

50.30 ± 13.20 27 25 25

SLM-3a 2:1 1.0 8.50 18.10 27.60

68.50 ± 10.30 27 25 25

SLM-1b 1:1 0.0 - 1.70 3.30

- 27 25 25

SLM-2b 1:2 0.0 - - 4.7

- 27 25 25

SLM-3b 2:1 0.0 - - 5.5

- 27 25 25

145

Table 10: Properties of the glibenclamide-loaded SLMs formulated with P90G-

structured goat fat and Softisan® 142

*Goat fat: Softisan® 142

refers to the lipid matrices containing 25 % of P90G.

Formul

ations

*Goat

fat:

Softisa

n® 142

Drug

compo

sition

(g)

w/w

Average particle

sizes at different

storage

temperatures (µm)

Drug

encapsulation

efficiency

(%)

Injectability (Gauge) at

25 oC

1

week

After

6

month

s

After 12

months

4 oC

25 oC

40 oC

SLM-1 1:1 1.0 7.0 10.7 13.0 58.23 ± 5.64 27 25 25

SLM-2 1:2 1.0 7.8 7.5 9.1 64.89 ± 10.21 27 25 25

SLM-3 2:1 1.0 9.0 11.7 12.3 70.35 ± 7.73 27 25 25

SLM-4 1:1 0.0 1.0 0.7 2.3 - 27 25 25

SLM-5 1:2 0.0 1.0 1.7 3.4 - 27 25 25

SLM-6 2:1 0.0 - 2.3 6.4 - 27 25 25

146

Table 11: Properties of the cimetidine-loaded SLMs formulated with P90G-

structured goat fat and tallow fat

*Goat fat: tallow fat refers to the lipid matrices containing 25 % of P90G

Formula

tions

*Go

at

fat:

Tall

ow

fat

Drug

com

posit

ion

(g)

w/w

Average particle

sizes at different

storage

temperatures

(µm)

Drug

encapsulatio

n efficiency

(%)

Injectability (Gauge) at

25 oC

1

week

After

6

month

s

After 12

months

4 oC 25

oC

40 oC

SLM-1 1:1 10.0 5.2 5.9 15.7 45.37 ± 9.26 27 25 18

SLM-2 1:2 10.0 1.20 42.9 14.0 40.30 ± 10.20 27 25 18

SLM-3 2:1 10.0 1.0 90.5 32.2 38.87 ± 3.90 27 25 18

SLM-4 1:1 0.0 1.0 1.8 2.6 - 27 25 18

SLM-5 1:2 0.0 0.7 4.8 0.9 - 27 25 18

SLM-6 2:1 0.0 - 5.5 1.7 - 27 25 18

147

Table 11 shows the result of the properties of the cimetidine-loaded SLMs

formulated with P90G-structured goat fat – tallow fat matrices. Although the

thermograms of the structured matrices (Table 1) suggested low crystallinity

(amorphosity) which means that the matrices had a lot of spaces for entrapment, the drug

loading efficiency was the least of all the drugs used in this study. This is probably

because cimetidine is a high dose drug with usual dosage of 200 or 400 mg. It would

have been difficult to entrap the 10 g w/w of this heavy drug in the 5 g w/w of the lipid

matrix. The 1:1 matrix gave the highest loading efficiency of 45.37 ± 9.26 % whereas

the 1:2 gave 40.30 ± 10.20 % while the 2:1 gave 38.87 ± 3.90 %.

A few drugs or peptides with various degrees of lipophilicity have been

incorporated into SLMs; for example carbamazepine [108] theophylline [105], non-

steroidal anti-inflammatory drugs, NSAIDs (ibuprofen, ketoprofen) [88], gonadotropin

releasing hormone [103], DNA [128], steroids (estradiol, medroxyprogesterone acetate)

[104], insulin [97, 98], vasoconstrictors [91], and antitumor agents [97, 99].

3.8 In Vitro drug release studies

The in vitro dissolution rate of pure drug samples (piroxicam, glibenclamide and

cimetidine) was compared to their release rates from SLMs containing equivalent

concentrations of the drugs along side their representative commercial samples

(Feldene®, Daonil

® and Cemtab

®).

Fig. 25 shows the release of piroxicam from the structured matrix containing

tallow fat and Softisan®

142, the pure drug dispersion in distilled water and a

conventional capsule dosage form (Feldene®). The result shows that the dissolution of

pure piroxicam was complete within 3 h and that the release rates from the SLMs were

148

generally high. The drug release was highest (87.53 ± 7.83 %) at 7 h in the 2:1 structured

lipid matrices of tallow fat-Softisan® 142, followed by the 1:1 batch with maximum drug

release of 80.14 ± 3.6 % at 4 h and finally the 1:2 batch, which showed a maximum drug

release of 66.61 ± 2.7 % at 6 h. It was also observed that the piroxicam release rates from

the commercial Feldene® sample was higher than the release of the corresponding

piroxicam-loaded SLMs even though the release of piroxicam from the SLMs was

sustained.

A similar observation was seen in the release profile of glibenclamide from

structured matrices of goat fat and Softisan® 142 as shown in Fig. 26. The outstanding

impression was that of burst release encountered with the structured lipid matrix

corresponding to 1:2 combinations of goat fat and softisan® which initially released 22.66

± 1.7 % of the glibenclamide within the first 30 min. It might be that during particle

production by the hot homogenization technique, the drug partitioned between the liquid

oil phase and the aqueous phase due to its slight aqueous solubility. During the cooling of

the produced O/W microemulsion, the solubility of the drug in the water phase decreases

continuously with decreasing temperature of the water phase, which means a re-

partitioning of the drug into the lipid phase occurs. When reaching the recrystallization

temperature of the lipid, a solid core starts forming including the drug which is present at

this temperature in the lipid phase. Reducing the temperature of the dispersion further,

reduced drug solubility in water and results in further re-partition into the lipid phase.

Since it would no longer be possible for the drug to dissolve in the crystallized core it

concentrates in the still liquid

149

0

10

20

30

40

50

60

70

80

90

100

0 2 4 6 8 10 12 14

% d

rug r

elea

se

Time (h)

Fig. 25: In vitro piroxicam release from SLMs formulated from

admixtures of structured tallow fat-Softisan 142

SLM 1:2 SLM 2:1 SLM 1:1 Pure drug Feldene

SLM-2a

SLM-1a

SLM-3a

150

0

20

40

60

80

100

120

0 5 10 15 20 25 30

% d

rug

rel

ease

Time (h)

Fig. 26: In vitro release studies of glibenclamide from SLM

formulated with P90G-structured lipid matrices containing goat fat

and softisan 142.

SLM 2:1 SLM 1:1 SLM 1:2 Daonil Pure sample

SLM-1a

SLM-2a

SLM-3a

151

outer shell of the SLMs and/or on the surface of the particles. The amount of drug in the

outer shell and on the particle surface is released in the form of a burst.

A prolonged drug release was also obtained for the structured lipid matrix

corresponding to 1:1 and 2:1 combinations of goat fat and Softisan® 142 (Fig. 26). This

demonstrated the suitability of the SLM system for prolonged drug release. The profiles

showed prolonged release without any burst. In the initial 0.5 h, the drug release was less

than 10 % probably because of slow diffusion of drug from the lipid core. Afterwards, the

drug release rate increased with time until 10 h, followed by a steady-state release. The

prolonged drug release could be attributed to embedment of drug in the solid lipid matrix.

The structured matrix of 2:1 goat fat and Softisan® 142 showed maximum glibenclamide

release of 56.99 ± 3.2 % which was sustained till 16 h, while the corresponding 1:1 lipid

composition attained maximum drug release of 20. 82 % at 12 h. Comparing the drug

release from the SLMs, the release of glibenclamide was slower from the 1:1 structured

matrix, 21 ± 1.2 % at the end of 24 h compared with 56.98 % from the 2:1 structured

matrices of goat fat and Softisan® 142.

Correlating the result of the in vitro drug release with that of DSC, it could be

seen that the structured 2:1 goat fat and Softisan® 142 matrices had better thermal

properties. Generally, they had the highest melting peak at each level of assessment (as

binary mixtures and/or ternary mixtures with P90G alone or in formulation) than their

counterpart matrices but in spite of this high temperature of melting, they had the least

enthalpy value than any other matrix. The lower melting enthalpy suggests less-ordered

lattice arrangement or amorphous state, which is associated with numerous lattice defects

and the formation of amorphous regions in which the drug was located [233].

152

The release profile could be affected by particle size. The dominant factors that

may affect the shape of the profiles are the production parameters (surfactant

concentration, temperature) and also the nature of the lipid matrix. With the dissolution

medium (PBS, pH 7.4) rightly chosen to reflect possible in vivo environment, correct

prediction could be easily drawn [90, 95, 126, 130]. More hydrophobic materials are

expected to reduce the drug release rate [95, 103, 105, 106]. The choice of matrix

materials influences the release process rate. Another way to change the matrix

hydrophobicity is by adding a hydrophobic or hydrophilic excipient [89, 95, 105].

The drug‘s physicochemical characteristics (its water solubility) also play a part

[106]. The release rate and the amount of drug released from SLMs increase with drug

hydrophilicity. This explains why cimetidine-loaded SLMs had poor performance due to

the drug‘s slight aqueous solubility. The preparation method of the SLMs could also

affect the drug‘s release rate by influencing the matrix wettability properties [103]. The

particle size is also considered as a relevant parameter influencing drug release.

Generally, the small size of the particles especially in the glibenclamide- and piroxicam-

loaded SLMs could be responsible for the high release since drug release from smaller

particles is higher than release from larger ones due to larger specific surface area of

smaller microparticles [95, 105].

In addition, lipid mixtures may alter the crystal arrangement of the individual

lipids after melting and solidification, which may increase their drug holding capacity, as

it is known that highly ordered crystalline lipid matrices lead to drug expulsion upon

crystallization of the previously molten matrices [215]. On another note, these lipids have

different fatty acid locations in their triglycerides and may have crystallized into lose

153

packing after melting, producing highly unordered lipid structures or disordered

imperfect lipid matrix structures offering spaces for drug molecules and amorphous

clusters of drugs [215].

The in vitro drug release profile of the SLMs containing cimetidine in P90G-

structured goat fat and tallow fat matrices was different from those of the other drugs

(Fig. 27). The poor release observed may be due to the fact that cimetidine is slightly

soluble in water. The highest release profile of 15.72 mg % was seen at 9 h in the batch

corresponding to 1:2 structured combinations of goat fat and tallow fat. The release

profile traced an erratic nature.

The batch corresponding to the 2:1 matrix had the highest release of 13.60 mg %.

The 1:1 batch had the least release, achieving a maximum of 7.07 mg % after 6 h. The

amount of drug released was generally lower than the theoretical drug loading.

154

0

2

4

6

8

10

12

14

16

18

0 2 4 6 8 10 12

% d

rug r

elea

se

Time (h)

Fig. 27: In vitrorelease profiles of cimetidine from SLMs formulated

with P90G-structured lipid matrices containing goat and tallow fats

SLM 1:1 SLM 1:2 SLM 2:1

SLM-1a

SLM-2a

SLM-3a

155

The release was generally erratic with plateau and troughs until it continuously decreased

in quantity of drug released. Yet the quantities released were very small. This may

probably be due to the fact that cimetidine is a high-dose drug (200 mg or 400 mg) and

encapsulating up to 10 % w/w of it in 5 % w/w of the P90G-structured goat-tallow fat

matrices might have led to supersaturation coupled with the fact that the drug is only

slightly water soluble making controlled delivery of the drug difficult. It has earlier been

reported that the amount of drug partitioning to the water phase will increase with the

solubility of the drug in the water and will defeat the meaning of controlled release [4]. It

can therefore be said that SLMs are better drug delivery system for low dose poorly-

water soluble drugs.

The release of piroxicam and glibenclamide from the SLMs was further

analysed using Fickian diffusion model to determine the mechanism of release. To

understand the mechanism of release, the release rate was described with the following

equation:

M

Mt

n

Kt … … … … …………………… (14)

Log M

Mt = Log K n log t ….. ……………….. (15)

where M

Mt is the fraction of released drug at time t, K is a characteristic constant that

incorporates the structural and geometric characteristics of the mechanim of release, n.

As the K value becomes higher, the drugs are released faster. The n value of 1

corresponds to zero-order release kinetics; 0.5 < n < 1 means a non-Fickian (anomalous)

release model and n=0.5 indicates Fickian diffusion [219]).

156

Table 12 shows the kinetic parameters n and K derived from the plots of Log

M

Mt Vs log t (Figs. 28 and 29). The n values of all piroxicam-loaded SLMs in the

structured tallow fat-Softisan® 142 and glibenclamide loaded in the 2:1 structured goat fat

– Softisan®

142 matrix were between 0.5 – 1. This indicated that their release followed a

non-Fickian diffusion model (anomalous behaviour). However, values for piroxicam

release approached unity and could be said to have exhibited almost zero-order kinetics.

The k values for all unformulated drug samples and the commercial representatives

(Feldene®and Daonil

®) were relatively high indicating fast release. This is also true for

piroxicam and explains why glibenclamide could be better delivered as SLMs than

piroxicam.

157

Table 12: The release kinetic parameters

Formulations Glibenclamide Formulations Piroxicam

N K N K

2:1 TF/SFT 0.6621 0.4355 2:1 GF/SFT 0.9409 1.1429

1:2 TF/SFT 0.1962 0.4584 1:2 GF/SFT 0.9869 1.0732

Daonil®

-13.142 1.2211 Feldene®

0.6855 1.0671

Glibenclamide powder -0.3667 1.3498 Piroxicam powder 0.6752 1.0704

The negative signs in the glibenclamide drug samples imply no sustained release in comparison to the

equipotent concentrations encapsulated in the SLMs.

158

y = 1.142x

R² = 0.940

y = 1.073x

R² = 0.986

y = 1.070x

R² = 0.675

y = 1.067x

R² = 0.685

0

0.5

1

1.5

2

2.5

0 0.5 1 1.5 2

Log M

t/M

α

Log t

Fig. 28: Log-log plot of the amount of piroxicam released from

structued tallow fat - Softisan 142

2:1 matrix 1:2 matrix Pure piroxicam

Feldene Linear (2:1 matrix) Linear (1:2 matrix)

Linear (Pure piroxicam) Linear (Feldene)

159

y = 0.435x

R² = 0.662

y = 0.458x

R² = 0.196

y = 1.221x

R² = -13.1

y = 1.349x

R² = -0.36

-1

-0.5

0

0.5

1

1.5

2

0 0.2 0.4 0.6 0.8 1 1.2 1.4

Lot

Mt/

Mα

Log t

Fig. 29: Log-log plot of the amount of glibenclamide released from the

structured goat fat - Softisan 142 matrices

2:1 matrix 1:2 matrix Daonil

Drug powder Linear (2:1 matrix) Linear (1:2 matrix)

Linear (Daonil) Linear (Drug powder)

160

3.9 In vivo release studies of piroxicam-loaded SLMs

3.9.1 Antinociceptive property

The piroxicam-loaded SLM formulations produced from the structured tallow fat

showed a dose-related significant (p<0.05 – 0.001) nociception in mice (Table 13). The

piroxicam-loaded SLM in a dose-dependent manner delayed the reaction times of the

mice to electrical heat-induced pain.

3.9.2 Evaluation of anti-inflammatory properties of SLMs

Subplantar injection of fresh egg albumin (0.5 ml/kg) provoked marked time-

related increases in the hind paw diameters of the rat control group that received blank

SLM (SLM-0). Although pedal inflammation (oedema) was always evident within 5-8

min following fresh egg albumin injection, maximal swelling and/or oedema occurred

approximately 90 min following fresh egg albumin administration.

The piroxicam-loaded SLM (SLM 1-4) produced significant reductions (p<0.05 –

0.001) in the fresh egg albumin-induced acute inflammation of the rat hind paw (Table

14). The blank SLM (2 ml/kg p.o.) neither modified responses to nociceptive stimuli in

mice, nor the rat hind paw oedema induced by fresh egg albumin administration.

The findings of this experiment indicate that the piroxicam-loaded SLM

formulations (SLM 1-4) possess antinociceptive and anti-inflammatory properties in the

mammalian laboratory animal models used (Tables 14 and 15). These findings are in

agreement with an earlier work [234].

161

Table 13: Effect of piroxicam-loaded SLMs on electrical heat-induced pain

(nociception)

Formulations Dose Mean reaction time (s) % protection

SLM-0 2 ml/kg 10.57 ± 1.32 0.75 NS

DW 3.0 ml/kg 10.65 ± 1.40 0.00

SLM-1 2.5 mg/kg 15.2 ± 1.42 42.72

SLM-2 5.0 mg/kg 16.82 ± 1.65b 57.93

b

SLM-3 7.5 mg/kg 18.64 ±1.70b 75.02

b

SLM-4 10 mg/kg 20.50 ± 2.30a

92.49a

DW-P 10 mg/kg 21.25 ± 2.0a 99.53

a

Sample 10 mg/kg 21.23 ± 2.0a 99.34

a

SLM-0 means blank formulation; DW means distilled water; SLM 1-4 means piroxicam-

loaded SLMs containing 0.25, 0.5, 0.75 and 1.0 g w/v respectively; DW-P means pure

drug dispersed in distilled water; sample means Feldene®.

Each value represents the mean (± SEM) of five observations. NS= p<0.005; ap<0.001 Vs

control; bp<0.01.

162

Table 14: Linear diameter measurements from SLM-treated oedematous rats

Treat

Ment

Dose Time (min) and paw diameter (mm) %

inhibit

ion

30 60 90 120 150 180

SLM-0 2 ml/kg 10.36±0.2 12.50±0.4 15.27±0.5 13.50±0.4 12.40±0.4 11.45±0.4 0.25

DW 3 ml/kg 10.36±0.4 12.40±0.3 15.42±0.5 13.6±0.40 12.35±0.4 11.42±0.4 -

SLM-1 2.5 mg/kg 10.±0.39 11.0±0.35 12.67±0.5 10.76±0.3 9.3±0.30 8.24±0.37 28.25b

SLM-2 5.0 mg/kg 9.3±0.33 9.58±0.4 11.35±0.3 9.47±0.26 8.30±0.28 7.1±0.25 38.03b

SLM-3 7.5 mg/kg 8.15±0.25 6.8±0.25 5.2±0.34 3.39±0.31 2.30±0.21 1.85±0.23 84.30a

SLM-4 10 mg/kg 7.3±0.35 4.10±0.30 3.53±0.35 2.10±0.30 0.9±0.06 0.4±0.04 96.33a

DW-P 10 mg/kg 5.14±0.28 2.5±0.51 0.42±0.01 - - - 100.00a

Sample 10 mg/kg 5.10±2.0 3.0±0.45 0.35±0.25 0.35±0.25 - - 96.84a

Each value represents the mean (±SEM). bp<0.05;

ap<0.01 Vs control

163

Table 15: Volume displacement measurements from SLM-treated eodematous rats

Treat

Ment

Dose Time (min) and paw diameter (mm) %

inhibition 30 60 90 120 150 180

SLM-0 2 ml/kg 2.0 ± 0.5 2.3 ±0.6 2.2 ± 0.5 2.2 ± 0.4 2.1 ± 0.7 2.1 ± 0.7 -5

DW 3 ml/kg 2.0 ± 0.4 2.2 ± 0.3 2.2 ± 0.4 2.1 ± 0.3 2.2 ± 0.4 2.0 ± 0.4 -

SLM-1 2.5 mg/kg 1.9 ± 0.5 1.9 ± 0.3 1.7 ± 0.6 1.6 ± 0.5 1.5 ± 0.5 1.45 ±0.6 27.5 ± 0.31

SLM-2 5.0 mg/kg 1.9 ± 0.7 1.9 ± 0.6 1.6 ± 0.4 1.5 ± 0.5 1.4 ± 0.3 1.25 ± 0.2 37.5 ± 0.26

SLM-3 7.5 mg/kg 2.0 ± 0.6 2.3 ± 0.5 1.7 ± 0.5 1.5 ± 0.4 1.0 ± 0.5 0.3 ± 0.2 85.0 ± 0.23

SLM-4 10 mg/kg 2.2 ± 0.5 2.0 ± 0.4 1.9 ± 0.5 1.0 ±0.6 0.1 ± 0.5 0.08 ± 0.2 96. 0 ± 0.03

DW-P 10 mg/kg 1.6 ± 0.6 1.0 ± 0.5 0.01 ± 0.4 0.01 ± 0.3 0.01 ± 0.2 - 99.5 ± 0.01

Sample 10 mg/kg 2.1 ± 0.8 1.0 ± 0.3 0.08 ± 0.5 0.08 ± 0.5 - - 96.0 ± 0.02

Each value represents the mean (±SEM).

164

3.9.3 In Vivo piroxicam release studies from SLMs formulated with structured

tallow fat-Softisan® 142

Using percentage inflammation as criterion for comparison, it could be seen that

the percentage inflammation for all the test and control groups decreased with time after

reaching maximum one or two hours after induction of inflammation (Fig. 30). The zero-

drug loaded formulations (SLM-1b; 2b and 3b) had higher inflammation than the distilled

water (negative control). The drug-loaded SLMs were generally comparable to the

positive controls (pure piroxicam powder and Feldene®

) (Fig. 30). For the 1st three hours,

the drug-loaded formulations showed lower percentage inflammation than the pure

piroxicam powder and the commercial sample (Feldene®

). However, by the 4th

hour,

Feldene®

exhibited lower percentage inflammation than two of the drug-loaded

formulations (SLMs-1a and 2a) but still surpassed SLM-3a. The percentage inflammation

trend was also determined by the oedema rate measuremaent (Figs. 31-33).

When the percentage inhibition of inflammation was used as basis for evaluation,

the trend was same as above. However, using the results for the 3rd

hour as the basis of

comparison, [211, 212], it was observed that SLM-3a had the best performance at 3 h,

surpassing Feldene®, pure piroxicam powder, and SLMs- 1a and 2a. This is in

consonance with the results of the encapsulation efficiency, DSC and in vitro release

studies.

The percentage inhibition of inflammation results show that for the 1st three

hours, the bioavailability from the drug-loaded SLMs was superior to that of the pure

165

Fig. 30: Plot of average right hind-paw volume against time

SlM-2a (drug) SLM-2b (zero drug) SLM-3a (drug)

SLM-3b (zero drug) SLM-1a (drug) SLM-1b (zero drug)

(Pure drug) Feldene® Distilled water

166

Fig. 31: Plot of percentage inflammation against time

SlM-2a (drug) SLM-2b (zero drug) SLM-3a (drug)

SLM-3b (zero drug) SLM-1a (drug) SLM-1b (zero drug)

(Pure drug) Feldene® Distilled water

167

Fig. 32: Plot of oedema rate against time

SlM-2a (drug) SLM-2b (zero drug) SLM-3a (drug)

SLM-3b (zero drug) SLM-1a (drug) SLM-1b (zero drug)

(Pure drug) Feldene® Distilled water

168

Fig. 33: Plot of inhibition of oedema against time

SlM-2a (drug) SLM-2b (zero drug) SLM-3a (drug)

SLM-3b (zero drug) SLM-1a (drug) SLM-1b (zero drug)

(Pure drug) Feldene® Distilled water

169

drug powder and Feldene®. However, at 4 h, Feldene

® showed superior results to the

SLMs. Generally, the drug-loaded SLMs showed better anti-inflammatory activity than

pure piroxicam powder throughout the course of the experiment. It was noted that the

Feldene® powder had much finer (smaller) particles than the pure piroxicam powder.

This difference in particle size could have been responsible for the significant difference

in terms of in vivo anti-inflammatory activity between these two positive controls.

On comparison of the drug-loaded SLMs, despite all having superior in vivo

properties to the positive controls within this time span of 1-3 h, SLM-2a had the highest

in vivo effects both in terms of the anti-inflammatory activity, the percentage

inflammation and oedema rate for the 1st two hours. SLM-3a had the greatest in vivo

effects both in terms of the anti-inflammatory activity; the percentage inflammation and

oedema rate from 3 h, while SLM-1a consistently had low or high in vivo results

throughout the course of the study.

The zero-drug formulations (SLMs-1b; 2b, and 3b) respectively, had similar

results for the 6 h post-induction of inflammation. Although, they had greater percentage

inflammation values than the negative control (distilled water), a comparison of the

inhibition of inflammation shows that they had low anti-inflammatory effects in vivo.

3.10 In vivo glibenclamide release studies from SLMs formulated with structured

Softisan® 142.

3.10.1 Induction of diabetes mellitus in the experimental rats

Diabetes was confirmed after three days and normal glucose level was 160 ± 27.2.

but 600 ± 25 in diabetic rats. Daily consumption of water and food in healthy adult rats

were 35 ± 5 ml and 11.3 g respectively. Daily urine volume in healthy adult rats was 11.1

ml but 130 ± 5 ml in diabetic rats. Daily consumption of water and food in the diabetic

170

rats were 150 ± 5 ml and 50.6 ± 4 g respectively. There was also body weight change

indicating that diabetes was accompanied by loss of weight (Fig. 34).

In addition, the changes in healthy and diabetic rats were distinctive because in

addition to loss of weight while the tails of the healthy rats were pink and had a white

velvet coat compared to the dark stained colour and white velvet to pink or grey coats. If

the environment of the rat was kept clean, this change of colour appeared from white to

pink otherwise, the change occurs from white to grey.

3.10.2 Fasting blood glucose reduction

The various glibenclamide-loaded SLMs were shown to effectively lower the

fasting blood sugar levels in the rats over a 24 h period although this trend was highest in

the SLM-3a. This suggests that the solid lipid microparticles could effectively be a carrier

for targeted and prolonged release of glibenclamide as shown in Fig 35. The relevance of

the word targeted is described below.

This is a pointer to the fact that glibenclamide could be delivered in the form of

microparticles thereby targeting the islet cells of the pancreas to further stimulate the

production of insulin from these cells in a controlled delivery rate than the conventional

tablet form. The result of this finding shows that the blood glucose levels were within

normal range before the alloxan injection and were further lowered in a gradual manner

over a period of 24 h by the prolonged release of glibenclamide from the formulated

structured lipid particles.

171

0

100

200

300

400

500

600

700

Weight

(g)

Urine

vol/day

(ml)

Blood

sugar

(mg/dl)

Water

(ml)

Food (

g)

Qu

an

titi

es

of

ass

ess

ed

pa

ra

mete

rs

Fig. 34: Physiological parameters in normal and diabetic rats

Before alloxan admin After alloxan admin

172

Legends:

DW-G = pure glibenclamide dispersed in distilled water

DW = Distilled water

Commercial sample = Daonil®

SLM-0 = SLM without drug (blank)

SLM 1-5 = SLM containing glibenclamide various concentrations of glibenclamide

(0.1, 0.2, 0.3, 0.4 and 0.5 g %)

0

20

40

60

80

100

120

140

160

180

200

0 5 10 15 20 25 30

Blo

od

glu

cose

red

uct

ion

(m

g/d

l)

Time (h)

Fig. 35: Effect of glibenclamide-loaded SLM (g) on the fasting blood

glucose of normoglycaemic rats

DW-G SLM-0 SLM-1 SLM-2 SLM-3

SLM-4 SLM-5 DW Sample

173

3.10.3 Effect of oral administration of glibenclamide SLMs formulated with

structured Softisan® 142 matrices to diabetic rats

After alloxan monohydrate injection, the blood glucose level increased and

remained high after 3 days post injection. With the maintained hyperglycaemia, the rats

showed polyurea, polydypsia, and polyphagia in addition to weight loss. Glucose levels

above 180 mg/dl were considered as diabetic especially as the animals were fasted for 12

h with access to water only.

The blood glucose levels of the SLM-0 group were significantly high throughout

the 24 h sampling period (Fig. 36) so the rats in this group had diabetes all through the

period and some even dying due to hyperglycaemia. The release of glibenclamide from

SLMs 2-5 in the microparticles progressively controlled the diabetes and restored the

blood glucose of the rats to normal after 24 h. SLM-4 and 5 groups at some point had

somewhat burst release resulting in crashing of the blood glucose especially at 9-24 h.

This effect was however dose dependent since the SLM-5 (5 mg/ml) lowered the

hyperglycaemia to 60 mg/dl whereas SLM-4 lowered it to 80 mg/dl.

174

0

100

200

300

400

500

600

700

0 10 20 30

Blo

od

glu

cose

red

uct

ion

(m

g/d

l)

Time (h)

Fig. 36: Effect of glibenclamide-loaded SLM on hyperglycaemic rats

DW-G SLM-0 SLM-1 SLM-3 SLM-4

SLM-5 DW Sample SLM-2

175

The DW-G group had a blood glucose level of 105 ± 15.6 at 6 h. Even though this

occurred within shorter time interval, it was more or less of a burst effect. Yet the result

is not comparable to those of SLM-5 (p<0.01) that lowered the blood glucose to 80 ±

22.6 mg/dl at 24 h. Finally, the DW group remained diabetic and some even died within

the 24 h sampling period.

The SLM-1 formulation which had the lowest drug concentration had the least

blood glucose lowering of 199 mg/dl after 24 h. Even though this is not excellent the

lipid matrix was still able to release the drug systematically over a prolonged time.

3.10.4 Effect of oral administration of glibenclamide SLMs formulated with

structured goat fat -Softisan 42 matrices to diabetic rats

The results so far confirm that the lipidic matrices were composed of biocompatible

lipids as well as the surfactant having being of GRAS status. The result of oral

administration of the SLM suspensions to diabetic rats showed that the microparticles

were able to reduce blood glucose levels of the hyperglycaemic rats (Fig. 37). This shows

the suitability of the microparticles to prolong the release of the glibenclamide from the

microparticles such that there was gradual control of blood glucose.

It was interesting to observe that the structured lipid matrix containing 1:2 of goat

fat and Softisan® 142 exhibited a burst release. This confirms the result of the in vitro

release study reported earlier with 1:2 structured matrices of goat fat and Softisan® 142.

After the 3rd

h, the glucose level started increasing again showing that it had exhausted

the incorporated glibenclamide. The pure glibenclamide and the commercial sample

(Daonil®) were able to lower the blood glucose within 4 h achieving blood glucose levels

of 100 and 120 mg/dl respectively. It was clear that the drug release from these samples

was not as prolonged as was obtained with the SLMs, which gradually lowered the blood

176

0

100

200

300

400

500

600

0 2 4 6 8 10

Blo

od

glu

cose

low

erin

g (

mg/d

l)

Time (h)

Fig. 37: Comparative blood glucose lowering properties of the

drug-loaded SLMs formulated from structured

admixtures of goat fat and Softisan 142

SLM 1:1 SLM 1:2 SLM 2:1 Pure sample Daonil

SLM-1a

SLM-2a

SLM-3a

177

glucose of the treated rats over a period of 9 h. This demonstrates the suitability of the

SLM system for prolonged release [235]. The maximum glucose reduction of 120 mg/dl

was exhibited by the drug-loaded SLM containing 2:1 (I.e. SLM-3a) to the structured

lipid matrix of goat fat and Softisan® 142. The 1:1 matrix composition also prolonged the

release of glibenclamide although it was not able to restore it to the normal range of

occurrence. The profile shows a controlled release mechanism. The zero-drug SLMs

counterparts showed continual increase in blood glucose levels.

Crystalline structure, related to the chemical nature of the lipid, is a key factor in

determining whether a drug will be expelled or firmly incorporated in the long-term.

Therefore, for a controlled optimization of drug incorporation and drug loading, intensive

characterization of the physical state of the lipid particles by DSC, NMR, X-ray and other

new techniques in this area are highly essential.

3.10.5 Pathological finding

The histopathological examination, following necropsy, showed that the

pancreatic islets disappeared in the diabetic rats while the reverse was the case for normal

rats. This is probably due to destruction by the diabetogenic agent, alloxan monohydrate.

The result is shown in Fig. 38. A comparison of the images reveals irreversible damage to

the β-cells of the pancreatic langerhans.

3.11 Effect of oral administration of cimetidine SLMs formulated with structured

goat fat - tallow fat matrices to ulcerated rats

After sacrificing the rats and opening their stomachs along the greater curvature, they

were found to have developed ulcer on examination of their stomachs. A comparative

178

Fig. 38: Pancreatic biopsy: (A) normal rat (B) diabetic rat

A

B

179

result of the ulcer indices is shown in Table 16. The result of the in vitro drug release

profile does not correlate with the in vivo performance. All the three cimetidine-loaded

SLMs had better antiulcer properties than the commercial sample, CEMTAB®.

The

overall in vivo performance showed the SLMs (SLM-3a) prepared from the 2:1 goat fat –

tallow fat matrices as having some 81.20 % ulcer inhibition on the rats. This was

followed by the SLM-1a prepared from the 1:1 matrices which inhibited ulcer by 76.18

%. SLM-2a prepared from the 1:2 matrices had an ulcer inhibition of 74.62 % although

still superior to the commercial brand of cimetidine.

Since the SLMs were in the dispersion form containing both the aqueous and lipid

phases, it means that the high in vivo performance was expected even though it doesn‘t

correlate with the in vitro result. If the SLMs were to be lyophilized into powders, then

the actual performance of the matrix-loaded cimetidine would have been done.

180

Table 16: In vivo cimetidine release profile in ulcerated animal models

Each value represents the mean (± SEM) of five observations.

NS= p < 0.005; ap < 0.001 Vs control;

bp < 0.01

Formulations

Cimetidine

Dose

(mg/kg)

Mean ulcer index ± SEM

%

ulcer

inhibition

SLM 1:1 100 1.65 ± 1.42 76.18±5.2

SLM 1:1 0.0 6.00 ± 1.15b 13.0±1.23

SLM 1:2 100 1.75 ±1.70b 74.62±4.6

SLM 1:2 0.0 6.15 ± 1.30a

10.9±2.1

SLM 2:1 100 1.30 ± 1.0a 81.20±5.0

SLM 2:1 0.0 5.80 ± 0.92 15.9±1.5

CEMTAB 100 1.90 ± 1.1 72.50±2.5

DW 3.0 ml/kg 6.90 ± 1.40 0.00±0.0

181

However, the possible mechanism of action of these systems aside from the drugs‘

intrinsic mechanism of action is based on their small particle sizes because such small

sized particles are easily recognized by the immune system as danger signals from where

they generally get internalized by antigen-sampling membranous (M) cells in intestinal

Peyer‘s patches (Fig. 39). These M cells (specialized epithelial cells) have a thinner

glycocalyx and less organized microvilli than enterocytes and are known to internalize

and transcytose particles to underlying lymphocytes and antigen-presenting cells [236-

238]. It is noteworthy to recall that lymphocytes arise form stem cells in bone marrow

and differentiate centrally into B-cells and T-cells (thymus) from where they move

through the bloodstream to the peripheral lymphoid tissues – the lymph nodes, spleen,

and lymphoid tissues associated with the mucosa, like the gut-associated lymphoid

tissues such as tonsils, Peyer‘s patches, and appendix, which are sites of lymphocyte

activation by antigens. Particles up to 10 μm in diameter can be internalized into Peyer‘s

patches and particles less than 5 μm can be transported to draining lymph nodes and the

spleen [239]. Lymph draining carries these particles from the tissues (extracellular fluid

as lymph) via the afferent lymphatics vessels into the thoracic duct, which returns the

lymph to the bloodstream by emptying into the left subclavian vein. Although the

organization of the spleen is similar to that of a lymph node (like Peyer‘s patches),

antigen enters the spleen from the blood (via trabecular artery into the central arteriole

from where they enter the marginal sinus and drain into a trabecular vein) rather than

from the lymph.

182

Fig. 39: General mechanism of in vivo SLM uptake

183

Another side to this anology is that exogenously administered triglycerides are

digested by the action of pancreatic lipase/colipase digestive enzymes in the small

intestine and absorbed into enterocytes. After absorption, long-chain fatty acids or lipids

are biosynthesized into triglyceride-rich lipoprotein particles (chylomicrons), which are

secreted into intestinal lymph (Fig. 39). The size of intestinal lipoproteins precludes their

absorption into the blood capillaries, and therefore they are secreted into the lymph.

Secondly, the cellular lining of the gastrointestinal tract is composed of absorptive

enterocytes interspersed with membranous epithelial (M) cells. M cells that cover

lymphoid aggregates, known as Peyer‘s patches, take up microparticles by a combination

of endocytosis or transcytosis [240, 241]. The important characteristics of microparticles

for their uptake are optimum size (10-100 nm), hydrophobicity, and surface charge [242,

243]. The uptake of fluorescent polystyrene microparticles of size ranging from 0.1 to 3.0

μm into Peyer‘s patches of rats was dependent on both the size and the nonionic nature of

the particles. Uptake of many colloidal polymeric carriers across the intestinal mucosa

[244] has been shown to occur via Peyer‘s patches or isolated lymphoid follicles after

oral administration [245]. In addition to the size of these SLMs within one week of

preparation, their hydrophobic surface, imparted by phosphatidylcholine, might have

influenced the SLM uptake by Peyer‘s patches [246, 247].

3. 12 Stability studies

The lipid microparticulate dispersions stored at 40 oC showed rapid aggregation

within 1 month of storage. Although the storage at this temperature could not be

continued due to epileptic power supply, the dispersions underwent sedimentation, and a

significant increase in particle size was observed in all samples. In contrast, the

184

dispersions stored at 4 - 6 oC remained stable, with only a slight change in particle size

upon 6 months storage. Storage of the microparticles at 4 - 6ºC did not affect the intact

spherical and smooth surfaces of the microparticles as well as the drug loading. However,

the dispersions stored at room temperature generally showed gross particle growth during

the period of storage. This is partly because of high amount of crystals due to delayed

crystallization from fractions of a supercooled amorphous melt. Yet this does not

preclude the use of the microparticulate dispersion for oral drug delivery of the actives

investigated since strict limit in particle size and particle stability may be overlooked.

Figs. 40A – 42A show the average particles sizes obtained for the various SLM

formulations at different storage temperatures. Generally, the particle sizes were bigger

for the drug-loaded SLM samples than for the zero-drug formulations,for example no

were seen at 4 oC . In addition, the mean particle sizes increased with increase in storage

temperature especially for the drug-containing SLM. This trend can be seen in Fig. 39A

representing SLM 1a-3a formulated with structured tallow fat - Softisan® 142 matrices

(1:1, 1:2 and 2:1) containing piroxicam. The drug-loaded SLMs corresponding to 1:2

structured matrices (i. e. SLM-2a) exhibited various particle growth from 10.5 µm, to

11.7 µm and 20.5 µm for the samples stored at 4 OC, 25

OC and 40

OC respectively. This

trend more in the 2:1 matrices (i. e. SLM-3a), where crystal growth was 8.5 µm, 18.1 µm

and 27.6 µm, while for 1:1 (SLM-1a), it was 10.1 µm, 17.8 µm and 26.5 µm for the

samples stored at 4 OC, 25

OC and 40

OC respectively.

185

SLM-1a at 4

oC

SLM-1a at 25

oC

SLM-1a at 40

oC

SLM-2a at 4

oC

SLM-2a at 25

oC

SLM-2a at 40

oC

SLM-3a at 4

oC

SLM-3a 2:1 at 25

oC

SLM-3a at 40

oC

Fig. 40 A: Photomicrographs of the piroxicam-loaded SLMs formulated with

different structured tallow fat (TF) - Softisan® 142 (ST) matrices under different

storage temperatures.

186

SLM-1b at 4 oC

SLM-1b at 25

oC

SLM-1b at 40

oC

SLm-2b at 4

oC

SLM-2b at 25

oC

SLM-2b at 40

oC

SLM-3b at 4

oC

SLM-3b at 25

oC

SLM-3b at 40

oC

Fig. 40 B: Photomicrographs of the zero-piroxicam SLMs formulated with different

structured tallow fat (TF) and Softisan® 142 (ST) matrices under different storage

temperatures.

187

SLM-1a at 4

oC

SLM-1a at 25

oC

SLM-1a at 40

oC

SLM-2a at 4

oC

SLM-2a at 25

oC

SLM-2a at 40

oC

SLM-3a at 4

oC

SLM-3a at 25

oC

SLM-3a at 40

oC

Fig. 41 A: Photomicrographs of the glibenclamide-loaded SLMs formulated with

different structured goat fat (GF) and Softisan® 142 (ST) matrices under different

storage temperatures.

188

SLM-1b at 4

oC

SLM-1b at 25

oC

SLM-1b at 40

oC

SLM-2b at 4

oC

SLM-2b at 25

oC

SLM-2b at 40

oC

SLM-3b at 4

oC

SLM-3b at 25

oC

SLM-3b at 40

oC

Fig. 41 B: Photomicrographs of the zero-glibenclamide SLMs formulated with

different structured goat fat (GF) and Softisan® 142 (ST) matrices under different

storage temperatures.

189

SLM-1a at 4 oC SLM-1a at 25

oC

SLM-1a at 40 oC

SLM-2a at 4 oC

SLM-2a at 25

oC

SLM-2a at 40

oC

SLM-3a at 4 oC

SLM-3a at 25

oC

SLM-3a at 40

oC

Fig. 42 A: Photomicrographs of the cimetidine-loaded SLMs formulated with

different structured goat fat (GF) and tallow fat (TF) matrices under different

storage temperatures.

190

SLM-1b at 4 oC SLM-1b at 25

oC SLM-1b at 40

oC

SLM-2b at 4 oC SLM-2b at 25

oC SLM-2b at 40

oC

SLM-3b at 4 oC SLM-3b at 25

oC SLM-3b at 40

oC

Fig. 42 B: Photomicrographs of the zero-cimetidine SLMs formulated with different

structured goat fat (GF) and tallow fat (TF) matrices under different storage

temperatures.

191

The results show that increased temperature contributes to the instability of the

formulations evident in the growth of the particles. The piroxicam-loaded SLM samples

stored at 4 and 40 O

C had particles that occurring in clusters (Fig. 40A). In terms of

aggregation, the order follows that the SLMs prepared from the 1:2 structured matrices

mostly aggregated in clusters, while the 2:1 matrices had slight aggregation at 25 and 40

oC while the SLMs from the 1:1 matrices existed as single particles. It would therefore

seem that although SLM-2a (i.e. 1:2 matrix) had the least particle size, it was already in

the process of forming agglomerates implying greater instability when compared to

SLMs-1a and 3a containing matrix combinations of 1:1 and 2:1 respectively.

However, for the zero-piroxicam preparations (Fig. 40B) representing SLMs-1b;

2b and 3b, a different trend was observed that also generally applies to all zero-drug

preparations (Fig. 40B – 42B). For the zero drug preparations stored at 4 oC, there were

no detectable particles. There were however, trace of particles at room temperature. The

mechanism leading to the presence of these particles may be as a result of crystallization

of the cooled melts. All the zero-drug SLM formulations stored at 40 OC had detectable

particles.

Stability studies of the entire SLM formulations stored at various temperature

conditions generally increased in the following order: 4 oC > 27

oC > 40

oC. This implies

that the formulations are most stable at 4 – 6 oC.

Fig. 41A shows the stability of glibenclamide-loaded SLMs formulated using

structured goat fat-Softisan® 142 matrices while Fig. 41B shows the zero-glibenclamide

SLMs. The results show gross aggregation of particles at 25 and 40 oC storage

temperatures in the entire drug loaded samples of all matrices (1:1; 1:2 and 2:1) although

192

it was worst in the SLMs from the 2:1 matrices followed by 1:1 and then the 1:2. This

implies that high storage temperature causes instability in the SLM preparations. The

samples stored at 4 oC had single particles without aggregation and could be said to be

the best storage temperature for all the glibenclamide-loaded SLMs.

The cimetidine–containing SLMs (Fig. 42A) prepared from structured goat and

tallow fats showed instability in all samples stored at 25 and 40 oC. Generally, the

samples stored at 40 oC exhibited the worst instability in all the SLMs (40

oC > 25

oC >

4oC). The zero-cimetidine samples showed similar tendencies.

3.13 Syringeability studies

All formulations remained uniformly dispersed at room temperature, within 1 week

of preparation. These formulations could also be pushed through syringe with a 27-

needle (27 G). This suggests that the microparticles could be potential drug carriers for

parenteral drug targeting. Generally upon storage, the syringeability of the SLMs varied

through 25-18 G for the first 6 months and remained stable for the rest of the 12 month

study period. However, relating the storage temperature to the syriangeability, it was

found that the SLMs stored at 4 oC retained the injectability of 27 G regardless of the

duration of storage.

193

CHAPTER FOUR

SUMMARY AND CONCLUSION

The formulated SLMs were prepared from three different lipid matrices (Softisan®

142, goat and tallow fats), while employing P90G as a stabilizer heterolipid, surface

modifier and cosurfactant. Polysorbate 80 (Tween®

80) served as a mobile surfactant.

The processing parameters were optimized and the SLMs appropriately characterized.

The drugs (piroxicam, glibenclamide and cimetidine) were adequately encapsulated in

the lipid matrices and evaluated appropriately.

The results indicate that increasing the concentration of polysorbate 80 decreased the

particle size of SLMs, while increasing the concentration of the drugs increased their

particle size. With the increase of surfactant concentration from 0 to 2 % w/w, the mean

diameters of SLMs decreased from 13.4 ± 1.3 to 0.1 ± 0.01 μm across all matrices. The

SLMs had uniform sizes, smooth surfaces and monodispersity. The emulsifying time

apparently influenced the mean diameters of SLMs. From the experimental results, 5 min

was considered the best emulsifying time. The stirring speed had similar influence on

SLMs‘ morphologies as that of emulsifying time. Low stirring rate resulted in large

particle sizes and non-spherical shape while rapid stirring resulted in the aggregation of

the SLMs. The desired particle sizes and shape uniformity of the SLMs was obtained at

1.5 % w/w of surfactant concentration, 5 % w/w of lipid matrices, 5 min emulsification

time, and stirring speed of 6200 rpm.

The DSC analysis confirmed that lipid mixtures can result in increased or decreased

crystallinity depending on composition. The 1:2 mixtures showed burst release

194

mechanism typical of a crystalline matrix with perfect crystal lattice that does not allow

drug to be entrapped leading to drug expulsion which defeats the purpose of controlled

release. The SLMs prepared from the P90Gylated- 2:1 binary-structured lipid matrices

(tallow fat-Softisan® 142; goat fat- Softisan

® 142 and goat fat-tallow fat) had the best

controlled-release properties in terms of anti-inflammatory action, antidiabetic effect and

anti-ulcer properties (enhanced bioavailability), surpassing the commercial brands of

Feldene®, Daonil

® and CEMTAB

® respectively. SLMs were better formulated with low

dose hydrophobic drugs (Piroxicam and Glibenclamide) than with high dose slightly

soluble drug (cimetidine). The three drugs used showed differences with respect of in

vitro release. The SLMs enhanced the dissolution of poorly water-soluble drugs

(piroxicam and glibenclamide) than that of the slightly water-soluble drug (e.g.

cimetidine). Piroxicam and glibenclamide are practically insoluble in water and their

gastrointestinal (GI) absorption is limited by their dissolution rates. Therefore, to enhance

drug dissolution, serum concentrations and their respective controlled release anti-

inflammatory and hypoglycemic effects respectively, they can be formulated as SLMs.

The SLMs were most stable at low temperatures of 4-6 oC, suggesting that SLM

formulations are better stored in the refrigerator or freeze-dried and packaged

appropriately as suspension powders for reconstitution to avoid microbial growth. This

would reduce observed crystal growth in the particles. The SLMs could be syringed using

small-to-medium hypodermic needles within one week of preparation suggesting possible

application in parenteral drug use.

195

RECOMMENDATION

• It is recommended that since SLM-2a and SLM-3a corresponding to SLMs

formulated from 1:2 and 2:1 P90G structured tallow fat – Softisan142® matrices,

showed good in vivo activities , it is worthwhile to explore a system that will

incorporate both systems complementarily such that the SLM-2a (i.e. the 1:2)

which demonstrated a burst effect will be located externally while the SLM-3a

(i.e 2:1) which showed sustained/controlled release will be integrated into the core

of the system to maintain a steady state release devoid of burst release alone or

erratic release mechanism. A careful selection of some cationic and anionic lipids

can successfully deliver these two systems as a model multilayered control release

system.

• Since SLMs have numerous advantages - (as against liposomes and polymeric

nano- and microparticles), feasibility of large-scale production by a high-shear

homogenization technique, in addition to relatively low raw materials and

production costs, yet novel and unexploited drug delivery system, it is

recommended that more poorly water- soluble drugs be further investigated.

• Finally, SLMs could be explored further as a carrier system for parenterally

intended actives with strict particle size control measures.

196

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APPENDICES

Appendix 1: Data for fasting blood sugar mg/dl

Time

(h) DW-G SLM-0 SLM-1

SLM-

2 SLM-3 SLM-4 SLM-5 DW Sample

0 160 165 155 160 148 175 158 160 160

1 157 162 154 159 138 173 141 158 130

3 140 160 150 159 136 152 132 158 100

6 80 162 136 140 136 139 110 156 60

9 160 129 128 120 131 96 152

12 156 120 128 120 100 80 154

24 160 96 100 110 75 60 156

Appendix 2: Data for blood glucose

lowering

Time

(h) DW-G SLM-0 SLM-1

SLM-

2 SLM-3 SLM-4 SLM-5 DW Sample

0 575 600 600 600 595 554 600 500 600

1 450 600 600 451 541 500 481 450 450

3 250 600 496 373 417 481 362 430 300

6 105 595 405 236 365 238 307 480 60

9 593 390 218 278 251 229 500

12 595 250 176 209 140 100 450

24 596 199 110 160 80 60 500

Appendix 3: Determinants of diabetes in the experimental rats before and after

alloxan administraion

Data for Fig 31:

Before alloxan

admin.

After

alloxan

admin.

Weight (g) 200 184

Urine vol/day (ml) 11.1 130

Blood sugar

(mg/dl) 160 600

Water (ml) 35 150

Food ( g) 11.3 50.6

219

Appendix 4: Data for blood glucose lowering in the experimental rats from SLMs

(containing glibenclamide) formulated with different matrix combinations of

structured goat fat and Softisan® 142

Blood Sugar Lowering (mg/dl)

Time

(h)

Batch

1:1

Batch

1:2

Batch

2:1 Daonil Pure sample

1 451 454 519 531 503

3 405 255 400 350 300

5 350 345 300 260 210

7 255 468 250 120 100

9 200 470 120

Appendix 5: Data for piroxicam release from SLMs formulated with different

matrix combinations of struectured tallow fat and Softisan® 142

Time

(h)

Cumulative Percentage

Release

1:02 2:01 1:01 pure

drug

Feldene

0.5 24.24 38.36 36.36 40.05 42.84

1 33.03 57.82 42.65 60.54 59.35

2 42.81 67.72 64.83 73.75 75.62

3 58.7 71.47 73.72 92.34 89.45

4 59.22 74.27 80.14 50.35 62.41

5 62.61 76.95 80.11 23.53 45.23

6 66.61 80.92 77.67 20.56

7 66.02 87.53 70.86

8 60.67 87.45 63.18

9 56.14 86.95 59.38

10 55.44 85.56 56.78

11 52.86 85.09 54.3

12 51.47 84.52 53.8

13 50.44 84.09 51.95

220

Appendix 6: Cimetideine release from SLMs formulated with different matrix

combinations of structured goat and tallow fats

Time

(h)

1:01

1:02 2:01

0.5 1.84

3.48

1.4

1 2.31

5.51

2.57

2 3.74 6.04 7.41

3 2.72 9.31 9

4 3.7 6.21 10.51

5 5.01 6.06 9.49

6 7.07 5.74 8.05

7 3.12 8.71 11.93

8 5.94 10 10.62

9 6.59 15.74 11.38

10 4.17 11.59 13.6

221

Appendix 7: Average paw volume displacement of the SLMs containing piroxicam.

Structured tallow

fat-Softisan® 142

Average paw volume displacement at different time (h) from

structured tallow fat-Softisan® 142 (ml)

0 hr 1 hr 2 hr 3 hr 4 hr 5 hr 6 hr

1:2 (piroxicam)

0.73 1.73 1.73 1.50 1.37 1.17 1.05

1:2

0.50 1.65 1.80 1.35 1.18 1.00 0.85

2:1 (piroxicam) 0.77 1.87 1.87 1.43 1.23 1.12 1.02

2:1

0.50 1.60 1.80 1.35 1.15 1.00 0.85

1:1 (piroxicam) 0.77 1.83 1.83 1.47 1.27 1.17 1.07

1:1

0.50 1.60 1.80 1.35 1.15 1.00 0.85

Feldene®

0.70 1.75 1.75 1.45 1.15 1.05 0.93

Piroxicam powder 0.70 1.80 1.80 1.50 1.20 1.15 1.00

Distilled water 0.70 2.10 2.20 1.70 1.50 1.30 1.10

222

Appendix 8: Oedema reduction rates of the SLMs containing piroxicam.

Structured tallow

fat-Softisan® 142

Oedema rate at different time (h) from structured tallow fat-

Softisan® 142

0 1 2 3 4 5 6

1:2 (piroxicam)

100.000 237.443 237.443 205.479 187.215 159.817 143.836

1:2

100.000 340.000 360.000 270.000 235.000 200.000 170.000

2:1 (piroxicam) 100.000 242.424 242.424 186.147 160.173 145.022 132.035

2:1

100.000 330.000 360.000 270.000 230.000 200.000 170.000

1:1 (piroxicam) 100.000 238.095 238.095 190.476 164.502 151.515 138.528

1:1

100.000 330.000 360.000 270.000 230.000 200.000 170.000

Feldene®

100.000 250.000 250.000 207.143 164.286 150.000 132.143

Piroxicam powder 100.000 257.143 257.143 214.286 171.429 164.286 142.857

Distilled water 100.000 300.000 314.286 242.857 214.286 185.714 157.143

223

Appendix 9: Percentage reduction of inflammation

Structured tallow fat-

Softisan® 142

Percentage Inflammation at different time (h) from structured

tallow fat-Softisan® 142

0 1 2 3 4 5 6

1:2 (piroxicam)

0.00 137.44 137.44 105.48 87.22 59.82 43.84

1:2

0.00 240.00 260.00 170.00 135.00 100.00 70.00

2:1 (piroxicam) 0.00 142.42 142.42 86.15 60.17 45.02 32.04

2:1

0.00 230.00 260.00 170.00 130.00 100.00 70.00

1:1 (piroxicam) 0.00 138.10 138.10 90.48 64.50 51.52 38.53

1:1

0.00 230.00 260.00 170.00 130.00 100.00 70.00

Feldene®

0.00 150.00 150.00 107.14 64.29 50.00 32.14

Piroxicam powder 0.00 157.14 157.14 114.29 71.43 64.29 42.86

Distilled water 0.00 200.00 214.29 142.86 114.29 85.71 57.14

224

Appendix 10: Percentage oedema inhibition of the SLMs

Structured tallow fat-

Softisan®

142

Percentage Oedema Inhibition at different time (h) from

structured tallow fat-Softisan® 142

0 1 2 3 4 5 6

1:2 (piroxicam)

- 28.57 33.33 23.33 20.83 27.78 20.83

1:2

- 14.29 13.33 15.00 15.63 16.67 12.50

2:1 (piroxicam) - 21.67 26.89 33.67 42.08 42.22 38.33

2:1

- 17.86 13.33 15.00 18.75 16.67 12.50

1:1 (piroxicam) - 24.05 29.11 30.33 37.92 33.89 25.83

1:1

- 17.86 13.33 15.00 18.75 16.67 12.50

Feldene®

- 25.00 30.00 25.00 43.75 41.67 43.75

Piroxicam powder - 21.43 26.67 20.00 37.50 25.00 25.00

Distilled water - 0.00 0.00 0.00 0.00 0.00 0.00

225

Appendix 11: Cumulative Drug Release for the Different Drug Loaded

Preparations.

Time (h) cumulative percentage release from structured tallow fat-Softisan®

142

1:2 2:1 1:1

0.5 24.24 38.36 36.36

1 33.03 57.82 42.65

2 42.81 67.72 64.83

3 58.70 71.47 73.72

4 59.22 74.27 80.14

5 62.61 76.95 70.50

6 66.61 77.92 67.33

7 60.00 87.53 65.27

8 59.85 74.05 63.18

9 56.14 70.64 59.38

10 55.44 66.74 56.78

11 52.86 65.82 54.30

12 51.47 64.45 53.80

13 50.44 61.72 51.95

226

Appendix 12: Fractional release of glibenclamide SLMs prepared from structured

binary matrices of goat fat-Softisan® 142

Time

(h)

Fraction

Released

Fraction

Released

Fraction

Released

Fraction

Released

Fraction

Released

Mt/M

2:1

Mt/M

1:1

Mt/M

1:2

Mt/M

Daonil

Mt/M

Pure

glibencla.

0.5 21.18 6.9 6.9 21 34.9

1 15.89 4.5 4.5 20.56 34.23

2 10.25 2.715 2.715 18.195 25.14

3 8.5667 2.55 2.55 16.9033 21.94

4 7.615 2.525 2.525 17.53 20.163

5 7.078 2.296 2.296 17.914 19.87

6 6.6233 2.2783 2.2783 14.833 16.5

7 6.5371 2.2714 2.2714

8 6.5287 2.1975 2.1975

9 6.68 2.1056 2.1056

10 6.537 2.02 2.02

11 5.9809 1.8618 1.8618

12 5.9982 1.7375 1.7375

13 5.4983 1.6038 1.6038

16 4.1863 1.3063 1.3063

24 2.7908 0.875 0.875

Appendix 13: Fractional release of piroxicam SLMs prepared from structured

binary matrices of tallow fat-Softisan® 142

Time (h)

Batch

1:2

Fraction

Released

Log Log t Square

root of

t Mt/M Mt/M

0.5 24.24 48.48 1.6856 -0.301 0.707

1 33.03 33.03 1.5189 0 1

2 42.81 21.405 1.3305 0.301 1.4142

3 58.7 19.5667 1.2915 0.4771 1.372

4 59.22 14.805 1.1704 0.6021 2

5 62.61 12.522 1.0977 0.6989 2.2361

6 66.61 11.1016 1.0454 0.7781 2.4495

7 66.02 9.4314 0.9748 0.8451 2.646

8 60.67 7.5738 0.8799 0.9031 2.8284

9 56.14 6.3278 0.795 0.9542 3

227

10 55.44 5.544 0.7438 1 3.1623

11 52.86 4.8055 0.6817 1.0414 3.3166

12 51.47 4.2892 0.6324 1.0792 3.4641

13 50.44 3.88 0.5889 1.1139 3.6055

2:01

Fraction

Released

Log

Log t

Square

root of

t Mt/M Mt/M

0.5 38.36 76.72 1.8849 -0.301 0.707

1 57.82 57.82 1.7621 0 1

2 67.72 33.86 1.5297 0.301 1.4142

3 71.47 23.8332 1.377 0.4771 1.372

4 74.27 18.5675 1.2687 0.6021 2

5 76.95 15.39 1.1872 0.6989 2.2361

6 80.92 13.4867 1.1299 0.7781 2.4495

7 87.53 12.5043 1.0971 0.8451 2.646

8 87.45 10.9313 1.0387 0.9031 2.8284

9 86.95 9.6611 0.985 0.9542 3

10 85.56 8.556 0.9323 1 3.1623

11 85.09 7.7355 0.8885 1.0414 3.3166

12 84.52 7.04333 0.8478 1.0792 3.4641

13 84.09 6.4685 0.8108 1.1139 3.6055

Time (h)

1:01

Fraction

Released

Log

Log t

Square

root of

t

Mt/M Mt/M

0.5 36.36 72.72 1.88617 -0.301 0.707

1 42.65 42.65 1.6299 0 1

2 64.83 32.415 1.5107 0.301 1.4142

3 73.72 24.5733 1.3905 0.4771 1.372

4 80.14 16.028 1.2049 0.6021 2

5 80.11 16.022 1.2047 0.6989 2.2361

6 77.67 12.945 1.1121 0.7781 2.4495

7 70.86 10.1229 1.0053 0.8451 2.646

228

8 63.18 7.8975 0.8975 0.9031 2.8284

9 59.38 6.5978 0.8194 0.9542 3

10 56.78 5.678 0.752 1 3.1623

11 54.3 4.9364 0.6934 1.0414 3.3166

12 53.8 4.4833 0.6516 1.0792 3.4641

13 51.95 3.9685 0.5985 1.1139 3.6055

Time (h) pure

drug

Fraction

Released

Log Log t Square

root of

t

Mt/M Mt/M

0.5 40.05 80.1 1.903 -0.301 0.707

1 60.54 60.54 1.782 0 1

2 73.75 36.875 1.5667 0.301 1.4142

3 92.34 30.78 1.4883 0.4771 1.372

4 50.35 12.5875 1.0999 0.6021 2

5 23.53 4.706 0.6727 0.6989 2.2361

Feldene

0.5 42.84 85.68 1.9929 -0.301 0.707

1 59.35 59.35 1.7734 0 1

2 75.62 37.81 1.5776 0.301 1.4142

3 89.45 29.8167 1.4745 0.4771 1.372

4 62.41 15.6025 1.1932 0.6021 2

5 45.23 9.046 0.9565 0.6989 2.2361

6 20.56 3.4267 0.5487

229

PUBLICATIONS

ARISING SO FAR

FROM THE

WORK

230

Already published manuscripts:

1. Nnamani, P. O., Ibezim, E. C., Attama, A. A. and Adikwu, M. U. (2010).

Surface modified solid lipid microparticles based on homolipids and softisan®

142: preliminary characterization. Asian Pac. J. Trop. Med. 205-210.

2. Nnamani, P. O., Attama, A. A., Ibezim, E. C. and Adikwu, M. U.

(2010).SRMS142-based solid lipid microparticles: Application in oral

delivery of glibenclamide to diabetic rats. Eur. J. Pharm. Biopharm.

(Please cite this article in press as: P.O. Nnamani et al., SRMS142-based solid

lipid microparticles: Application in oral delivery of glibenclamide to diabetic

rats, Eur. J. Pharm. Biopharm. (2010), doi:10.1016/j.ejpb.2010.06.002).

3. Nnamani, P.O., Ibezim, E.C., Attama, A.A. and Adikwu, M.U. (2010). New

approach to solid lipid microparticles using biocompatible homolipids-

templated heterolipid microcarriers for cimetidine delivery. Nig. J. Pharm.

Res. (In Press, Accepted manuscript).

4. Nnamani, P.O., Attama, A.A., Ibezim, E.C., and Adikwu, M.U. (2010).

Piroxicam-loaded P90Gylated tallow fat-based solid lipid microparticles:

characterization and in vivo evaluation. Nig. J. Pharm. Res. (In Press,

Accepted manuscript).

Already submitted manuscript:

5. Nnamani, P. O., Attama, A. A., Ibezim, E. C. and Adikwu, M. U. (2010).

Tallowation of SRMS142-based piroxicam solid lipid microparticles:

characterization and in vitro-vivo studies. Eur. J. Pharm. Biopharm.

Manuscript under preparation:

6. Nnamani, P. O., Attama, A. A., Ibezim, E. C. and Adikwu, M. U. (2010).

Templated homolipid - Softisan® 142 conjugate as a microcarrier for intestinal

delivery of glibenclamide to diabetogenic rats. J. Control. Rel.

Paper presentation at an international conference:

7. Nnamani, P. O., Attama, A. A., Ibezim, E. C. and Adikwu, M. U. (2010).

SLMs as microcarrier for intestinal delivery of BCS classes II and III drugs

based on solidified reverse micellar solutions. Paper presented at the TWOWS

Fourth General Assembly and International Conference: Women Scientists in

a Changing World, Beijing, China, 27 – 30 June, 2010.