MSc Chemistry - esc.fnwi.uva.nl · INTRODUCTION ... The experienced downside of this product is the retrieval of the polymer after 3 years ... It is learned from naturally occurring

MSc Chemistry

Literature Thesis

Determination of degradation behavior of biodegradable polymers for pharmaceutical devices

What are the physical properties that influence degradation and how is it predictable / measurable?

by

Laura Kox May 2014

Supervisor: Dr. W. Th. Kok

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Determination of degradation behavior of biodegradable polymers for pharmaceutical devices L.L.W. Kox University of Amsterdam

Abstract

Biodegradable pharmaceutical devices have as challenge that the release of the active pharmaceutical

ingredient is mostly related to the degradation rate. The degradation is mainly based on hydrolysis. During this

reaction with water the polymer chains get cleaved up to the point that the degradation products are small

enough to leave the polymer matrix. All physical properties have their own effect on the degradation.

The larger the molecular weight, the more hydrolysis reaction steps are needed to obtain soluble degradation

products. The higher the crystallinity or crosslinking, the denser the polymer which hinders the water

penetration resulting in lower degradation rates. Also the active pharmaceutical ingredient can alter the

physical state of the polymer. The porosity of the polymer results in an interesting combination of opposing

effects. The penetration of water is increased as could be expected, but on the other hand, autocatalysis is

reduced. Autocatalysis is the result of a decreasing of pH within the polymer resulting in higher hydrolysis

reaction rates. As any porosity within the polymer gives the degradation products a higher possibility to leave

the polymer, autocatalysis is reduced in more highly porous systems.

With in vitro measurements the degradation in vivo can be mimicked. To be able to predict the degradation of

the polymer in vivo by in vitro measurements a proper method should be developed in which the degradation

mechanism should be the same. Correlation of several nanoparticle formulations has successfully been

performed.

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TABLE OF CONTENTS Page

1. INTRODUCTION ...............................................................................................................4

2. BIODEGRADABLE POLYMERS .......................................................................................6 2.1 COMPOSITION OF BIODEGRADABLE POLYMERS .......................................................6 2.2 BIODEGRADATION BY HYDROLYSIS .............................................................................9

3. PHYSICAL PROPERTIES ............................................................................................... 10 3.1 MOLECULAR WEIGHT ................................................................................................... 10 3.2 HYDROPHOBICITY ........................................................................................................ 11 3.3 MORPHOLOGY .............................................................................................................. 12 3.4 CROSSLINKING ............................................................................................................. 15 3.5 POROSITY ...................................................................................................................... 16 3.6 SURFACE ....................................................................................................................... 19 3.7 API-POLYMER INTERACTION ....................................................................................... 19

4. AUTOCATALYSIS ........................................................................................................... 22 4.1 EROSION PROFILING .................................................................................................... 22 4.2 INTERNAL ACIDITY ........................................................................................................ 24 4.3 POROSITY ...................................................................................................................... 26 4.4 MODDELLING OF AUTOCATALYSIS ............................................................................. 27

5. IN VITRO / IN VIVO CORRELATION .............................................................................. 29 5.1 IN VITRO DETERMINATIONS ........................................................................................ 29 5.2 IN VIVO DETERMINATIONS ........................................................................................... 32 5.3 IN VITRO / IN VIVO CORRELATION .............................................................................. 34

6. CONCLUSION ................................................................................................................. 37

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1. INTRODUCTION

Parenteral pharmaceutical devices made from biodegradable polymers are interesting because such a device

can be metabolized by the body after implantation. Not retrieving the polymeric device after treatment spares

the patients a painful second procedure. Biodegradable polymeric medical devices are developed in several

forms: a widely used and known example is the soluble suture. At the moment mostly described

developments within this area of expertise are drug eluting nanoparticles and tissue engineering materials

which should provide temporary support to the damaged tissue while the body is healing[1-4].

In history within the pharmaceutical industry the tablet containing a highly soluble, low molecular weight active

pharmaceutical ingredient (API) which is easily absorbed trans-epithelially into the human body is the dosage

form of choice. Not all API’s have proper oral bioavailability. Parenteral dosage forms have to be used for

API’s that are not easily absorbed through the intestinal wall. The ultimate goal would be a method to extend

the API exposure with high bioavailability. Polymers are able to extend the release of an API, from the

injectable dosage form, making a daily injection unnecessary. The release from non-degradable polymeric

implants with fixed physical properties in time has been extensively studied and marketed[5]. An example is

Implanon from Organon. The experienced downside of this product is the retrieval of the polymer after 3 years

as a second painful procedure.

While developing a biodegradable polymeric device the degradation pattern will be investigated. Depending

on the demands and the final goal of the device every device will have its own set of specifications. The

degradation profiling should provide the information needed during development of the device. The release

profile of an API from an implant based on a biodegradable polymer is expected to have two stages of API

release. The first stage is diffusion of API out of the stable polymer matrix before the degradation is started. In

the second stage the release by diffusion is substantially lower than the release as a result of degradation.

This degradation profile makes it hard to predict the exposure of the patient to the API [2].

This literature thesis reviews the strategies and techniques to study degradation including the theoretical

background of the physical properties and their effects on the degradation of biodegradable polymers for

pharmaceutical purposes.

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Biodegradable polymers come in a wide variety and when developing a biodegradable polymeric device the

choice of polymer is the basis. The background of the polymers and the main degradation pathway is briefly

reviewed in chapter 2.

Every physical property of a polymer has its influence on the degradation. Because the polymer properties

can be modified to obtain the desired degradation rate it is important to consider all properties during

development of a biodegradable medical device. The physical properties and their effects on the degradation

are outlined in chapter 3.

In turn the degradation changes the physical properties of the device, increasing the complexity of the

degradation profile. A further complicating effect is autocatalysis. During the hydrolysis of the polymer acidic

degradation products are formed, these catalyze the degradation reaction. The porosity of the device, which

may increase during degradation, can have an interesting counter effect on the degradation profile as the

acidic degradation products can leave the polymer matrix. These effects and their determination methods are

described in chapter 4.

Methods for correlating the in vitro measurements with in vivo data, including the challenges involved will be

elaborated on in chapter 5.

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2. BIODEGRADABLE POLYMERS

The final goal of any pharmaceutical device is to deliver the desired dose of API without any toxicity.

For a polymeric device that will be implanted in the body, the toxicity of the polymer should be considered.

The material itself and its degradation products should not cause any toxic response or sustained

inflammation within the body. Additionally, after its period of action the complete device should get

metabolized and cleared from the body[6]. By washing away the small degradation products of the polymer

the device will dissolve and be metabolized by the body. The main degradation pathway within this area is

hydrolysis[7].

An introduction to biodegradable biocompatible polymers including their composition is given in this chapter,

followed by the basic principles of hydrolysis.

2.1 COMPOSITION OF BIODEGRADABLE POLYMERS

There is a high variety of biodegradable polymers available which can be used as biomaterial. A very

extensive review of the polymers and their properties was published by Nair and Laurencin[6]. They indexed

the polymers as hydrolytically degradable (e.g. poly(esters) and polyurethanes) or enzymatically degradable

polymers (e.g. polysaccharides, proteins and poly(amino acids)).

Biodegradable polymers can be produced from either fossil resources or natural sources[8]. The

polymerization process could be a biosynthesis, using microorganisms or enzymes, or a chemical synthesis

(for example a heavy metal catalysis synthesis route)[6, 8].

Although it seems very ‘green’ to use natural polymerization processes, they have certain downsides.

Immunogenicity is sometimes seen for these polymers, purification of these polymers seems hard to control

and there may be a risk for disease transmission. These disadvantage are avoided using synthetic materials,

they are more biologically inert and show less batch-to-batch variations[6].

Poly(esters) are the main class of investigated polymers in this area. In table 1 an overview of the most

commonly used monomers and resulting linear homopolymer structures is given. Their degradation time

frame fits most biomedical applications and various synthesis routes and bacterial processes have been

developed[6].

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Table 1) Monomers (lactones) and resulting linear homopolymer structures of poly(esters); reprinted from [6]

Copolymerization is used to fine tune the physical properties of the polymer. Poly(lactide-co-glycolide) (PLGA)

is the most described and extensively researched copolymer for medical devices. It is composed of three

monomers: glycolide (GA) and the two optically active L and D forms of the chiral lactide (LA). By tuning the

ratio of these three monomers a wide variety of physico-chemical properties can be obtained[2, 4, 6]. The

polymer chains of pure poly(L-lactide) (PLLA), poly(D-lactide) PDLA and poly(glycolide) PGA are able to order

themselves resulting in a highly crystalline polymer while copolymerization results in disordered chains and

more amorphous polymers.

It is learned from naturally occurring polymers, as for example spider silk, that there are three important

dimensions of microstructure control: composition, tacticity and structural sequence[9]. This is visualized in

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figure 1. However, whereas the composition of polymers is often described extensively, the tacticity and

sequence in relation to degradation not[10]

The impact of the degree of randomness of copolymerization is shown by Li et al.[9]. They compared

randomly disordered copolymer and highly ordered sequential copolymers of lactide and glycolide and

compared their degradation behavior. Figure 2 visualizes their findings. Random copolymers mostly result in

first order degradation kinetics while sequential copolymerization leads to a stable zero order degradation.

The expected reason for this effect is the difference of degradation rate of the La-La and Ga-Ga bonds

compared to the La-Ga bonds [11].

Figure 1) Copolymerization engineering triangle;

reprinted from [9]

Figure 2) Difference in degradation rate of PLGA

prepared by random and sequential copolymerization;

reprinted from [11]

The physical properties of the polymer and therefore of the final device, are determined to a large extent by

the choice of monomers and the (co)polymerization conditions. These physical properties define the

characteristics of the degradation rate and therefore the release rate of the API[12, 13]. Within the rest of this

thesis it is chosen not to focus on a specific set of polymers but the effects of the properties are described in

general.

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2.2 BIODEGRADATION BY HYDROLYSIS

Hydrolysis is the main degradation pathway described in the literature for biodegradable polymers[2, 6, 7, 14].

There are other pathways known, for instance oxidation, but the contribution of this pathway is minimal[2, 7,

14]. Hydrolysis is a bond breaking reaction (mostly of an ester bond) with water. When backbone bonds of the

polymer are broken this results in chain cleavage. The water influx into the polymer and the resulting swelling

is therefore a crucial parameter for the reaction and can be considered as the first step of degradation.

Swelling can be monitored during in vitro measurements. When developing a medical device this effect is

usually examined under physiological conditions. Phosphate buffered saline solution at 37oC is the most used

method.

Hydrophobicity, morphology, porosity and crosslinking together influence the swelling of the polymer. By

analyzing a granule under in vitro release conditions, the swelling or moisture absorption (MA) can be

determined gravimetrically. The mass of the water within the polymer at time t is the weight of the swollen

sample (Wt) minus the initial weight of that sample (W0).

The percentage water is then calculated according to equation 1.[15, 16]

(%) = − × 100%(1)

Because the weight of sample at time t is compared to the initial weight, the decrease in weight as a result of

degradation is not considered within this equation. Determining the moisture absorption at time t (MAt) the

weight of the swollen sample (Wwet) and the weight of that same sample after drying (Wdry) gives a more

accurate result. This is the technique used by the groups of Lui and Huang [17, 18] (equation 2).

(%) = − × 100%(2)

Hydrolysis can also be catalyzed by enzymes. Proteins, poly(amino acids) and polysaccharides are defined

as enzymatically degradable polymers. The site of implantation for these polymers is therefore considered

important since the availability and concentration of enzymes can vary significantly[6]. The enzyme catalyzed

degradation will not be a part of this thesis.

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3. PHYSICAL PROPERTIES

The physical properties of the biodegradable polymer determine its degradation rate. A pharmaceutical

device, however, does not consist of only the polymer. All additives including API also influence the physical

properties. The effect of e.g. the API on the degradation is therefore also important to consider. In this chapter

the physical properties of the final device are outlined including the methods used to determine them and their

effects on the degradation rate.

3.1 MOLECULAR WEIGHT

Degradation is primarily monitored in number- or weight- average molecular weight (Mn or Mw, respectively).

Hydrolytic chain scission will result in oligomers up to the point that the dissolution of degradation products is

possible[14, 19]. The higher the molecular weight the more hydrolytic chain scissions are needed to

completely break down the polymer, therefore higher molecular weights result in low degradation rates[2].

The decrease in Mn or Mw during in vitro degradation is monitored by gel permeation chromatography (GPC),

which is a form of size exclusion chromatography (SEC). These techniques are able to separate species

based on their mass (in fact based on their hydrodynamic volume), by a HPLC system equipped with a SEC

or GPC column. A SEC column is silica based while GPC is a polymer based column. The developments

within these columns makes it possible that a wide variety of solvents is possible on both techniques. Both

columns contain a network of uniform pores. Separation of mass is mechanically based: the small molecules

are getting trapped in the pores of the column, the larger molecules move with the flow of the HPLC system.

This results in chromatograms with response of high molecular weight at low retention times and response of

low molecular weight at high retention times.[20]

Figure 3) A typical curve of the Mw in time; reprinted from [21]

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By monitoring the Mw in time, a typical curve found by most groups is first order as given in figure 3 [9, 21-23].

At the beginning of the in vitro degradation test the polymer only swells. The hydrolysis reaction starts after

enough water is present within the polymer. The point where the curve shows the first decrease in Mw is called

the erosion onset[14, 24, 25]. In figure 3 the first two points have similar Mw: the erosion onset is between the

points two and three of both curves.

As previously discussed Li et al. showed the zero order degradation of sequenced PLGA, see figure 4. During

degradation measurements they observed a very symmetrical SEC peak implying that the degradation was

more homogeneous than in the case of average random copolymerization, see figure 5 [11].

Figure 4) molecular weight against time curves of

random and sequential PLGA copolymerization;

reprinted from [11]

Figure 5) Difference in SEC peak of random (left)

and sequential (right) PLGA copolymerization;

reprinted from [11]

3.2 HYDROPHOBICITY

The attraction of water and therefore the swelling of the polymer can be directly related to the hydrophobicity

of the polymer[6]. If the hydrophobicity increases the swelling rate decreases and therefore the degradation

rate decreases[3]. In case of a copolymer both monomers together determine the hydrophobicity [12].

It is concluded by Wu and Ding [23] that increasing the content of the hydrophilic Ga in PLGA increases the

degradation rate. Comparing the degradation kinetics of the polymers PLGA 85/15 and PLGA 75/25 showed

faster degradation of PLGA 75/25.

Hydrophobicity changes also with Mn; higher molecular weight is a result of longer polymer chains, which also

results in a lower concentration of hydrophilic end groups within the polymer which decreases the

hydrophobicity of the polymer[3]. The hydrolysis reaction increases the concentration of hydrophilic end

groups, leading to a change in hydrophobicity during the degradation process. Direct determination of

hydrophobicity is complicated because other physical properties, such as morphology, change concomitantly

which also adapts the swelling behavior.

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3.3 MORPHOLOGY

Increased crystallinity results in decreased degradation rates. In the amorphous parts there is the possibility to

accommodate water, whereas in the crystalline parts the polymer is denser and there is less space for water.

Therefore the swelling is hindered by the crystallinity of the polymer[2, 14, 25].

Once the crystalline parts of the polymer do degrade, however, this will result in more amorphous parts in the

drug product. Since the amorphous parts give more space for water, the swelling will increase and therefore

the hydrolysis will be accelerated[3].

Figure 6) Crystalline and amorphous parts in a semi-crystalline polymer; reprinted from [5]

The fact that the polymer composition directly affects the morphology has been confirmed several times for

PLGA. In 1994 Vert et al. attempted to map the degradation characteristics as a result of the structure of the

several PLGA compositions [4]. Their results are visualized in figure 7

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Figure 7) schema of the morphology of PLGA; reprinted from [4]

Polymers of the two optically pure D and L forms of lactic acid are given in the bottom two corners and

poly(glycolide) in the top corner. The C areas stand for the highly crystalline compositions and A for the

completely amorphous compositions. It was also seen that during degradation of the amorphous parts

crystallization of the degradation by-products or short chains was possible making the degradation more

complex[4, 25]. Copolymers that showed crystallization of the degradation by-products or short chains within

the amorphous parts are within the B area of the triangle.

Morphology is mostly related to the glass transition temperature (Tg).

The Tg is the temperature where the polymer transits from a glassy state to a rubbery state. When crossing

the Tg the mobility of the polymer chains increases and the physical properties of the polymer undergo

significant changes[20].

The glass transition temperature is determined by differential scanning calorimetry (DCS). In this technique

the heat flow into a sample is determined while warming the sample at a constant rate of temperature

increase and subtracting the heat flow into a non-sampled reference cell. As the polymer crosses the Tg, the

heat flow changes. The heat flow is plotted in a thermogram[20]. A Tg is shown in a thermogram as a

sigmoidal shape whereas melting and crystallization are really peaks. Figure 8 is an example of a thermogram

of PLLA[25]. Herein it can be seen that the Tg decreases from 56oC to 44oC during a 110 day in vitro

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degradation measurement. The deep negative peak with values of 168oC to 153oC corresponds to the melting

temperature and the positive peaks are recrystallization peaks from the crystallization of the degradation by-

products or short chains within the amorphous parts.

Figure 8) Thermogram of PLLA during a 110 day in vitro; reprinted from[25]

A relation is expected between glass transition temperature (Tg) and morphology. Still, the relation of Tg with

degradation is not always seen[2]. Probably, this behavior can be explained by the fact that the Tg is a

property that is influenced by several other physical properties. Besides that, it is also seen that wetting

results in plasticization and reduction of Tg [10, 22]. Tg and degradation are affected by many parameters and

therefore the interpretation is generally complex.

Morphology can also be determined by spectroscopic techniques as for example nuclear magnetic resonance

(NMR), X-ray and FT-Raman. Based on the different energy responses of the randomly oriented polymer in

the amorphous parts and the highly ordered polymer in the crystalline parts, the morphology can be

determined[26-28]. The determination of Tg as an identification of the morphology is used by most literature

reviewed in this thesis. In some cases a spectroscopic techniques is used additionally.

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3.4 CROSSLINKING

By introducing covalent bonds between polymer chains, crosslinking reinforces the intermolecular structure.

Crosslinking the material increases the tensile strength while the swelling decreases[16, 29].

The tensile strength is measured by puling at the polymer up to the point of breaking. The used force on the

device, in Pascal, is plotted agents the strain. A typical stress-strain curve is given in figure 9. These are the

results of Ghanbarzadeh et al.[16]. They showed improved mechanical properties of starch (a polysaccharide)

after a crosslinking treatment with carboxymethyl cellulose (CMC). The resulting tensile properties as a

function of the degree of crosslinking (the relative amount CMC added) is given in figure 9. The fact that this

treatment also increased the water resistance of the polymer is shown in figure 10. The moisture absorption

decreased by increasing crosslinking.

Figure 9) The increase of tensile strength by the

increase of crosslinking with CMC;

reprinted from[16]

Figure 10) The decrease of swelling by the

increase of crosslinking with CMC;

reprinted from[16]

Teng et al. described the crosslinking results of silk-elastinlike copolymer (a protein polymer). After a

methanol treatment a crosslinking with glutaraldehyde was performed. The samples only treated with

methanol, non-crosslinked, were more crystalline than the crosslinked samples[30]. This crystallinity results

from the fact that the polymer is able to form inter and intra molecular hydrogen bonds resulting in beta

sheets[30].

While determining the release of API from the more crystalline samples it was observed that in these samples

the swelling was hindered due to increased density which reduced the release rate[31].

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Tensile strength is important for tissue engineering, because the polymer should be strong enough to support

the healing tissue. The relation of tensile strength to crosslinking is therefore described.

The effect of crosslinking on the degradation is not described as such. Not even by Sonam et al. who

extensively reviewed the physical properties of nanoparticles and their effect on degradation and release of

API[3]. Nor by Makadia and Siegel who reviewed PLGA for medical devices[2]. Still the examples described

here showed increased water resistance upon crosslinking indicating that hydrolysis could also be decreased.

3.5 POROSITY

The larger the pores within the polymer are, the easier water can penetrate the polymer and therefore the

higher the water influx [15]. The porosity can be managed by two mechanisms, during syntheses and as the

formation of channels during degradation[15, 17].

The porosity can be visualized using scanning electron microscopy. This is a technique in which the sample is

coated with a metal and imaged by electron scanning. A focused beam of electrons is directed at the surface

and backscatters onto an electron detector. The surface of the solid sample is scanned by the beam imaging

the surface[20].

Figure 11) SEM images of PEG incorporated PLGA film during in vitro measurement; reprinted from [17]

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Huang et al. produced a poly(ethylene glycol) (PEG) incorporated PLGA film[17]. In the first stage of the in

vitro measurements the porosity of the film increased by dissolution of the PEG. Figure 11 shows the

scanning electron microscope (SEM) pictures of cross sections of the films. They managed the rate of the

channel formation by changing the molecular weight of the PEG. The hydrophilic diol character of shorter

PEG (400 Da) chains resulted in higher water influx relative to longer PEG (10k Da) chains at the same mass

concentration. The effect of the PEG additives in the samples on the swelling is given in figure 12

Figure 12) Water adsorption results of PEG incorporated PLGA film during in vitro measurement;

reprinted from [17]

Dorati et al. used sugar and salt during the preparation of PLGA scaffolds[15]. The sugar and salt crystals

resulted in different pore sizes. During preparation of the scaffolds, salt and sugar also function as

dehydrating agents resulting in reduction of the average polymer chain distance during rearrangement of the

polymer conformation. Compared to an untreated PLGA control sample the water influx was higher for their

prepared scaffolds as shown in figure 13.

Figure 13) Water uptake during in vitro measurements of scaffold 1 (sugar) and scaffold 2(salt)

compared to non-porous PLGA; reprinted from [15]

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The porosity can be determined as described by Dorati et al., by the liquid displacement method[15]. Using a

non-solvent with a known volume (V1) in a graduated cylinder, they soaked the polymer until no more air

bubbles were observed from the surface of the sample and recorded the fluid level (V2). The total volume of

polymer was V2-V1. The residue after removing the sample (V3), subtracted from the initial volume V1 (V1-

V3) is the pore volume. Even the density of the sample was determined using the initial weight of the sample

before soaked. See equations 3 and 4

ℎ = ( − ) + ( − ) = − (3) ℎ = − (4)

The effect of cracks and pores resulting from mechanical stress was investigated by Arm et al.[32]. Polymer

rods containing a protein were stressed by continuous three point bending in an in vitro system figure 14. The

release of the API was monitored for two weeks and the surface of the polymer rod investigated by SEM

analysis. They found a direct relation between rate of release and the total area of cracks and pores,

concluding that the cracks and pores increased the release of API.

Figure 14) Schematic drawing of a three point bending in an in vitro system;

reprinted from[32]

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All articles discussed in this section conclude that the degradation rate of porous samples is lower compared

to non-porous samples. The expected reason is that autocatalysis has less impact on the total degradation.

Autocatalysis and the effect that porosity has on it will be discussed in chapter 4.

3.6 SURFACE

In the literature on nanoparticles, the surface charge is often determined. The surface charge has an effect on

the stability of the polymer, the bioavailability for oral dosage forms, repulsion and attraction between particles

and interaction with tissue respectively[3]. See figure 15. The surface charge is not an intensively studied

phenomenon in the area of tissue engineering. Considering the fact that a nanoparticle has a relatively large

surface it can be expected that the surface charge has less effect on an implant. A surface property that is

important for implants, however, is its hydrophobicity. It has been shown that by adding a hydrophobic layer to

the device the swelling can be reduced[3].

Figure 15) Schematic drawing of electrostatic repulsion of nanoparticles; reprinted from[3]

3.7 API-POLYMER INTERACTION

From the polymer point of view the API is the main additive. It would therefore be expected that the presence

of the API in the polymer can influence all physical properties of the polymer[33]. Siegel et al. investigated the

influence of six different APIs on the degradation rate of PLGA[34]. Figure 16 shows the evolution of the

pellets over 21 days resulting in different shapes. Relating the water solubility of the API with the swelling the

results for the most hydrophilic API, Aspirin, and the least hydrophilic API, Haloperidol, are as expected.

However the results for Thiothixene and Ibuprofen could not be correlated to the water solubility of the API,

therefore, they concluded that more effects play a role.

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Figure 16) Change of shape of PLGA pellets during 21 day in vitro measurement;

reprinted from[34]

But the opposite is also true; the physical properties of the API can also be influenced by the polymer. It was

shown by Choi et al. that due to weakening of the intra molecular bonding the API was more amorphous in a

polymeric formulation than in its pure form[35]. This resulted in improved solubility and therefore also in

increased release rates. Since the API release is the final goal, understanding the effects of polymer-API

interactions is important in the development of any device.

Jia et al. developed a method to determine the binding constant, Kbs, of API to polymer by affinity capillary

electrophoresis (ACE)[36]. The theory is that the polymer- API charged complex will be retained on a capillary

electrophoresis (CE) column based on the polymer- API interactions. A CE system separates charged

(macro) molecules by their size/charge ratio using a buffer-filled capillary tube to which an electric field is

applied[20]. To determine the Kbs, the capillary was filled with a buffer solution containing the polymer of

interest and subsequently the API was run over the capillary. If the polymer and API interaction is strong the

affinity between them retains the API which is then measured at a later retention time at the end of the

capillary.

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They varied the polymer concentration (on the column) and buffer solution (pH 4 and 9) for several polymers

and APIs and measured the retention. The retentions where a measure for the Kbs and these results where

correlated to the octanol-water partition coefficient (log P). The final results are given in figure 17 It is

expected that this method can be used as a screening tool during early formulation development.

Figure 17) correlation of Kbs to log P; reprinted from[36]

All properties of a polymer are in relation with one another, therefore there is never a clear answer possible of

what the effects of changing one property on the final design or on the degradation will be. Sometimes

conflicting results are described in the literature, probably because of other side effects playing a role. The

theoretical description of physical properties given here should, however, provide some directions as to what

to monitor and what to expect.

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4. AUTOCATALYSIS

Hydrolysis is an acid (or base) catalyzed reaction. During an ester bond cleavage two end groups are yielded,

a carboxylic acid and a hydroxyl[37], as seen in figure 18. The carboxylic acid end group is able to catalyze

the hydrolysis reaction resulting in autocatalysis[25]. The additional complexity of this effect on the

degradation is outlined in this chapter.

Figure 18) Hydrolysis reaction of PLGA; reprinted from[37]

4.1 EROSION PROFILING

2 types of degradation are described. In bulk erosion the whole device is hydrated and hydrolysis occurs

through the whole device. Surface erosion is also possible if the water influx is hindered. As the device is then

only swollen at the surface, the surface erodes while the center stays intact. These two mechanisms are

visualized in figure 19[24].

Figure 19) Surface and bulk erosion profile; reprinted from[24]

The interplay between swelling rate and hydrolysis rate results in an erosion profile of the polymer device.

Based on this theory Burkersroda et al. showed that every polymer has a critical dimension were the erosion

profile changes from bulk erosion to surface erosion[24]. With their mathematical model they calculated the

erosion number Ɛ (which is a dimension less number).

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= ⟨ ⟩4 ln[⟨ ⟩] − / ( − 1) (5)

Herein; ⟨ ⟩is the mean travel distance of water, is a rate constant that accounts for the differences in the

reactivity of polymer functional groups, is the effective diffusion coefficient, is the number average

molecular weight, is the number of Avogadro, the degree of polymerization / number of monomers per

polymer chain and is the density of the polymer.

The results for Ɛ fall into three categories; Ɛ > 1 surface erosion is expected, Ɛ < 1 bulk erosion is expected

and at Ɛ = 1 the erosion profile changes. The dimension of the polymer L can be calculated from equation 5

based on the mean travel distance of water ⟨ ⟩. Plotting the erosion number Ɛ against the dimension L and the (bond reactivity λ)/(water diffusion D) figure 20

is their result. They calculated the critical dimension (Lcritical) at which a device changes its erosion profile (from

bulk to surface erosion) for a number of polymers. Figure 22 gives the Lcritical for several polymers.

Figure 20) Graph of the mathematical model that

determines the erosion profile;

reprinted from[24]

Figure 22) Critical dimensions for change

of erosion profile; reprinted from[24]

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4.2 INTERNAL ACIDITY

During the degradation of PLGA it was seen that the inside of the polymer was degrading faster than the

surface. Samples monitored during in vitro and in vivo measurements showed that the inside first turned liquid

and later on only the empty shells were left[4]. In figure 23 photos are given of these samples. In the left panel

it can be seen that the inside has a different color than the outer walls after 2 weeks in vivo. In the right panel

an empty shell after 2 months in vivo is shown. In a later article it is concluded that autocatalysis is the reason

for this phenomenon[25].

Figure 23) Difference in degradation rate depending on the position within

the device ; reprinted from[4].

Figure 24) SEC results of the surface

(- - -) and the interior (--- · · ·); .

reprinted from[25] .

As described in 3.3 a relationship between morphology and degradation pattern was investigated by Li et al.

[25]. It was concluded that the PLGA copolymers were more amorphous and faster degrading. In a later study

of the heterogeneous degradation, these PLGA copolymers showed a difference in molecular weight of the

surface and the inside. Figure 24 are the SEC results of the surface and the interior after 3 weeks in vitro (a)

and after 7 weeks in vitro (b). The apparent bimodal size distribution of the surface after 7 weeks is probably

caused by difficulty in sampling only the surface material. The mechanism of heterogeneous degradation is

visualized in figure 24 [25].

Page 25


Figure 25) Heterogeneous degradation as a result of autocatalysis;

reprinted from[4]

Step 1 is the complete homogeneous swelling of the sample followed by homogeneous hydrolysis in step 2.

Since in the bulk the acidic degradation products are trapped but at the surface they can diffuse out, the

acidity gets heterogeneous and therefore also the degradation will be heterogeneous as seen in step 3-5 of

figure 25. The fact that smaller nanoparticles have been observed to degrade faster than larger ones is also a

result of autocatalysis[3, 21]. If the particle size decreases the diffusion path length for acidic products out of

the particles is shortened. Furthermore, the surface area increases relative to the mass of the total particle[3,

21].

Determining the inner pH on micro scale is therefore interesting for understanding the degradation pattern of

the polymer device. Liu et at. reported a determination method of the inner µpH of PLGA microspheres[18,

38]. They used a pH sensitive dye which acts as a fluorescence probe. The pH can be determined from the

intensity of fluorescence emission signal. Confocal laser scanning microscopy makes these

spectrophotometric measurements possible on micro scale within the microspheres. Figure 26 shows the pH

distribution during in vitro measurements through time of several samples (A-D are different PLGA

compositions, 1-5 are the days 1, 7, 14, 18 and 21).

Page 26


Figure 26) Results of µpH measurements within

PLGA microspheres; reprinted from[18]

Figure 27) Results of µpH measurements within

PLHMGA microspheres; reprinted from[38]

This method is not limited to PLGA microspheres. The inner µpH of poly(D,L-lactide-co-hydroxymethyl

glycolide) (PLHMGA) microspheres was also determined, results are given in figure 27. They concluded that

the inner pH was higher (more toward 6) than in the PLGA microspheres making this polymer suitable for acid

labile APIs[9].

4.3 POROSITY

Increasing porosity can alter the degradation rate. If channels are present, the acidic degradation products

can wash out of the device. The reduced acidity leads to reduced degradation rates.

Dorati et at. and Huang et at. both investigated the effect of porosity; as discussed in 3.5[15, 17]. In porous

samples the swelling is higher than in non-porous samples but the Mw is more stable through time, see figure

28 and 29 The group of Huang was interested in the increase of API release with the increase of the porosity,

they also concluded that their porous sample were more stable due to suppressed autocatalysis.

Page 27


Figure 28) Mw decrease during in vitro

measurements of (scaffold 1 and 2) compared to

non-porous PLGA; reprinted from[15]

Figure 29) water uptake decrease during in vitro

measurements of (scaffold 1 and 2) compared to

non-porous PLGA; reprinted from[15]

Furthermore Dorati et at. concluded that with increasing porosity the release of API also increased[15]. It

could be expected that this results in a change of the release profile; the release as a result of diffusion in the

first stage is higher but due to hindered autocatalysis the release as a result of degradation in the second

stage is hindered.

4.4 MODDELLING OF AUTOCATALYSIS

In an attempt to predict the change in Mn over time as a results of autocatalysis Antheunis et al. described a

autocatalytic equation for aliphatic polyesters[39].

( ) = [ ] − 11 + [ ][ ] + 1(0) (6) Herein; [ ] is the acid concentration at time point initial, [ ] is the concentration of ester bonds at time point

initial and is the density of the polymer in (g/L). For amorphous homopolyesters is ([ ] + [ ] ) with

as the reaction rate, for the crystalline regions is more complex.

The model was validated by comparing it to in vitro degradation results. The results for P4MC poly(4-

methylcaprolactone) (figure 30) and PLGA (figure 31) are given, showing a proper correlation. The difference

between the curves is due to difference in acidity at the initial time point. At the initial time point P4MC carries

relatively more acidic end groups compared to PLGA. Also the hydrolysis reaction rates differ with the ester

bond, resulting in the given time scale.

Page 28


Figure 30) compared results from P4MC

modeled and in vitro degradation;

reprinted from[39]

Figure 31) compared results PLGA

modeled and in vitro degradation;

reprinted from[39]

By combining the exact dimensions of the polymer device the mathematics become more complex. Chen et

al. and Tang et al. modeled the erosion profile of several more complex shaped polymeric devices[40, 41]. As

an example the results of a relatively simple sphere is given in figure 32[40]

Figure 32) Modeled degradation profile from two spheres with radii 7.9 µm and 55µm by

concentration profile of acidic monomers; reprinted from[40].

They concluded that the size and the architecture could play a critical role on the rate during the degradation,

in pathway.

Page 29


5. IN VITRO / IN VIVO CORRELATION

The correlation of in vitro measurements to in vivo data appears to be challenging. Mostly, the only parameter

for which the correlation is investigated is “drug release”, because that is the final purpose of the device.

The release profile of API is generally described as starting with diffusion based release up to the erosion

onset followed by release due to erosion, if the degradation rate is higher than the diffusion rate[2, 14] which

is mostly the case. The most expected reason for failed in vitro / in vivo correlation is the presence of

enzymes in vivo, that can catalyze the reaction[2, 6, 14].

A proper development of the in vitro measurements is also important as will be seen in this chapter during the

outlining of both measurements.

5.1 IN VITRO DETERMINATIONS

To mimic the release of an API from an extended release device, the device will be submerged in a medium

and the release of API from the device into the medium will be studied. This is called an in vitro release

measurement. Based on the measured pharmaceutical device and the effects which are to be monitored,

strategies can differ.

There are several measurement systems described, Shen et al. and Amatya et al. both reviewed them[37,

42]. The most widely used method is the sample-and-separate method. Hereby the sample is introduced into

a vessel or vial containing release medium. After a certain interval of time the sample will be separated from

the medium, the medium will be sampled for analysis and replacement or refreshment of the medium is

performed, if needed. On the other hand the continuous flow cell method uses a flow cell that contains the

sample and the release medium is pumped around as given in figure 33. Inline analysis is possible with this

method. The same figure also gives several methods of fixing the sample in the flow cell, as for example

directly between glass beads. A dialysis adapter is an example that is able to fix nanoparticles. A dialysis cell

can also be used to contain micro- or nanoparticles in the sample-and-separate method (figure 34).

Page 30


Figure 34) Dialysis cell in vitro model;

reprinted from[37] .

Figure 33) Flow through cell in vitro model; reprinted from[37]

These set-ups for in vitro methods are also used to monitor the degradation of biodegradable polymeric

devices. In this case, instead of measuring the release of API the polymer itself is sampled and analyzed.

For a proper in vitro in vivo correlation (IVIVC), the optimization of in vitro conditions is important to achieve

the same release mechanisms in vitro as in vivo[37]. Parameters that influence the in vitro release are

summed up by Shen et al [37]. Although their main goal was to accelerate the measurement these effects are

also important to the ‘real time’ release.

An important parameter is the release medium. It surrounds the device and is the source of water for swelling.

Also the degradation products and API will be dissolved in the medium. It should not hinder nor accelerate the

emission of the degradation products or API from the polymer. The volume and the composition should be

optimized during method development. For studying primarily the degradation phosphate buffered saline

(PBS) is mostly used, whereas for studying the release of API sink conditions are generally required and

therefore surfactants may be used.

Temperature has a large influence on the release. The Arrhenius equation shows that managing the

temperature is important because of the natural logarithmic relation between temperature and rate constant,

see equation 7. = × / (7) Herein; k is the zero-order release rate, A a constant, Ea the activation energy, R the gas constant and T the

temperature. This equation can only be applied if the Tg is not within the temperature range. If Tg is passed

the degradation and release mechanisms are very different because the polymer is in a different physical

state.

Page 31


The hydrolysis rate is affected by the pH[24] as seen in chapter 4. To keep the erosion profile the same, the

pH and buffering capacity of the medium should be controlled.

Studying the effect of the agitation rate on the API release of PLGA microparticles Schoubben et al. showed

difference in release curves between continuous agitation and once-a-week agitation[43]. The release system

consists of a flat bottom vial containing 10 mL 0.1 M phosphate buffer solution (pH 7.4) with 0.02 % sodium

azide at 37oC. For the degradation monitoring of the polymer at each predetermined time point, the

microparticles were filtered out of the medium and if needed dried overnight under vacuum at room

temperature. Figure 35 shows the difference in Mw of two different polymers and their stirring technique.

Figure 35) Decrease of Mw of two polymer

samples at different agitation rates in vitro;

reprinted from[43]

Figure 36); 3 phases of degradation Mn average

molecular mass, E tensile strength, W weight of

the sample and D its diameter; reprinted from[22]

During in vitro measurements there are 3 phases of degradation recognized. Stage I is the quasi stable stage;

this stage is divided in two. At first the tensile strength increases by plasticization as a result of swelling [15].

After that, all measured properties seem stable except Mn. In stage II the original structure gets corroded and

this is therefore called the decrease-of-strength stage. Stage III is the loss-of-weight / disruption-of-scaffold

stage. An example graph is given in figure 36[22].

Page 32


The determination of the exact quantitative degradation effect of enzymes on the polymer is hard to determine

[2]. In some cases the effect is not significant but it should never be neglected [22]

To study the effect of enzymes, Reiche et al. reviewed the Langmuir monolayer degradation technique[19].

This is a method whereby a monolayer film of organic material, in this case polymer, can be investigated while

being in contact with a solution[19]. In this review the degrading effect of enzymes on the polymer is

described.

Figure 37) Langmuir monolayer degradation technique;

reprinted from[19]

Figure 37 is a schematic drawing of the Langmuir monolayer degradation technique. F are soluble fragments

(e.g. enzymes) within the aqueous phase to which the polymer monolayer will be exposed, SP is a sensor that

could be any kind of surface detection system. It is concluded that using this technique the hydrolytic and

enzymatic chain scission can be investigated. But also transport phenomena within the polymer can by

investigated by using computer simulations in combination with this technique[19].

5.2 IN VIVO DETERMINATIONS

For in vivo determinations the polymeric device is administered to an animal, for example a mouse, rat or

rabbit. To determine the degradation of the polymer the device should be retrieved after a certain time. To

obtain a degradation profile, multiple animals are therefore required. By determining the concentration of the

API in blood the release from the device can be monitored. A release profile can in principle be obtained from

a single animal.

Page 33


In an attempt to develop a non-invasive swelling monitoring method for in-situ forming implants Solorio et al.

used ultrasound as an imaging method[44]. The basic principle is the difference in return of the ultrasound

waves from the non-swollen part of the implant relative to the swollen part, were the incoming waves are

backscattered. With this noninvasive method the swelling front of the implant can be followed in vitro and in

vivo. The swelling front of the implants was studied using this technique; results are given in figure 38. They

confirmed a faster swelling of implants containing low molecular weight polymer.

Figure 38) Ultra sound images of the swelling front of in-situ forming implants;

reprinted from [44]

Page 34


5.3 IN VITRO / IN VIVO CORRELATION

Correlation of in vitro data with in vivo data shows whether the in vivo data can be predicted by the in vitro

data. There are several methods of correlation but the so-called level A correlation is preferred. A level A

correlation means that a point by point comparison is made between the release in vitro (in percentage of the

total) and the release in vivo (in percentage of the total) as a function of time. After statistical analysis

according to linear regression, the conclusion can be drawn whether a correlation is seen and the in vitro

results are predictive for in vivo results[45].

Yang et al. related the release of Thienorphine-loaded PLGA microspheres over a 28 day period in vitro and

after subcutaneous injection to rats[46]. In vitro API release was performed in PBS solution (0.01M, pH 7.4

and 0.02% sodium azide at 37oC 30 mL) using a dialysis cell inserted in a flask, which was shaken during the

measurement. The in vitro results are given in figure 39. In vivo results were obtained by analysis of blood

plasma by LC/MS/MS. Blood plasma results are given in figure 40. These were subsequently recalculated to

the API absorption in vivo (Fa). The level of correlation of the data is given in figure 41. It was concluded that

the shaken flask method was acceptable for IVIVC of subcutaneous administered PLGA microsphere

formulations.

Figure 39) In vitro API release (cumulative);

reprinted from[46]

Figure 40) API plasma concentrations in vivo;

reprinted from[46]

Figure 41) Linear regression plot of the correlation data;

reprinted from[46]

Page 35


That a medium of PBS is not always used is shown by D’Aurizio et al. [45]. In their investigation of a PLGA

microsphere containing the anti-Parkinson API L-dopa (LD) they used acetate buffer (0.02M) pH 4.5 as

release medium at 37oC. The results were correlated to the release in vivo after injecting the formulation

subcutaneously in rats. Figure 42 is their result on which they concluded that there was a correlation.

Figure 42) IVIVC result of LD PLGA microspheres;

reprinted from [45]

Note that Yang et al. plotted for each time point the cumulative % release in vitro and in vivo, whereas

D’Aurizio et al. plotted for each cumulative % released the required time in vitro and in vivo. The fact that there

is an IVIVC does not necessary mean that the releases are exactly equal. As seen in figure 42, e.g., it follows

that the release in vivo is 2.25 times slower than in vitro.

The injection site in the body influences the degradation rate, as confirmed by Mohammad et al. [21]. They

compared the in vivo and in vitro degradation of 200 nm and 500 nm nanoparticles in the liver and spleen of

rats and showed a first-order degradation in vitro and in vivo, according the equation 8. = ( )(8) Herein; M is the molecular weight measured as Mw, M0 is the initial molecular weight, k the rate constant of

the degradation and t is the time.

The decrease of Mw results are given in figure 43 in vitro and figure 44 and 45 in vivo. The slopes of the lines

were compared. It can be seen from figure 43 that the 500 nm nanoparticles degraded faster than the 200nm

particles, probably because of autocatalysis. The differences in vivo (figure 44 and 45) are smaller. They

reasoned that the escape of acidic degradation products could be hindered in the tissue. It was also seen that

different organs have different degradation rates. Their conclusion was that a clear IVIVC could not be made

in a global sense.

Page 36


Figure 43) In vitro degradation results of 200nm and 500nm nanoparticles;

reprinted from[21]

Figure 44) in vivo degradation results after

injection into the liver of rats;

reprinted from[21]

Figure 45) in vivo degradation results after

injection into the spleen of rats;

reprinted from[21]

The correlation examples given in this chapter are all based on nanoparticle formulations, tissue engineering

formulations are not found in the literature. Because of autocatalysis a sample with more mass could be

harder to correlate.

Page 37


6. CONCLUSION

The degradation kinetics of biodegradable pharmaceutical devices is depending on all physical properties of

the polymer including the effects of the API and other additives blended into the polymer. These effects can

be investigated in vitro using phosphate buffered saline (PBS) as medium at 37oC. The relation between the

various physical properties and the degradation rate is mostly understood, but not always straight forward.

This is probably because a certain change of unexpected effects on another physical property. An example of

a property which effect is porosity. Higher porosity increases swelling but not the hydrolysis rate. Because of

dissolution of the degradation products the autocatalysis is hindered and therefore the degradation of porous

devices is slower than of non-porous devices.

The final goal of an in vitro measurement is to predict the in vivo results. Developing an in vitro method needs

proper attention to achieve the same release mechanisms in vitro as in vivo. For a biodegradable polymer

device this is extra challenging, not only the diffusion coefficient of the API should not be influenced, also the

degradation mechanism of the polymer should stay intact. The fact that the position of the administration also

has its influence on the degradation, with the expected reason that enzymatic degradation plays a role, makes

that the development of an in vitro measurement method should be carefully considered.

In order to investigate the degradation mechanism, research groups have developed creative methods, such

as using ultrasound to determine the swelling, in vitro and in vivo. Another example is the use of a dye to

determine the inner pH on micro scale.

There is a lot of information on drug releasing biodegradable nano- and microparticles to be found in the

literature. Biodegradable tissue engineering devices are also reported although these are generally not

designed to release drug. To bring a solid, biodegradable drug-releasing implant to the marked, a lot of

research is needed because of the complexity of the release of drug as a result of the degradation rate.

Page 38


1. Gunatillake, P.A. and R. Adhikari, Biodegradable synthetic polymers for tissue engineering. European Cells and Materials, 2003. 5(1): p. 1-16.

2. Makadia, H.K. and S.J. Siegel, Poly Lactic-co-Glycolic Acid (PLGA) as Biodegradable Controlled Drug Delivery Carrier. Polymers, 2011. 3(3): p. 1377-1397.

3. Sonam, et al., Effect of Physicochemical Properties of Biodegradable Polymers on Nano Drug Delivery. Polymer Reviews, 2013. 53(4): p. 546-567.

4. Vert, M., S.M. Li, and H. Garreau, Attempts to map the structure and degradation characteristics of aliphatic polyesters derived from lactic and glycolic acids. Journal of Biomaterials Science, 1995. 6(7): p. 639-649.

5. Laarhoven, H.v., Physical-chemical aspects of a coaxial sustained release device based on poly-EVA. 2005: PhD thesis university of Utrecht.

6. Nair, L.S. and C.T. Laurencin, Biodegradable polymers as biomaterials. Progress in polymer science, 2007. 32(8): p. 762-798.

7. Karlsson, S. and A.-c. Albertsson, Biodegradable polymers and environmental interaction. Polymer Engineering & Science, 1998. 38(8): p. 1251-1253.

8. Matsumoto, K.i. and S. Taguchi, Enzyme and metabolic engineering for the production of novel biopolymers: crossover of biological and chemical processes. Current opinion in biotechnology, 2013. 24(6): p. 1054-1060.

9. Li, J., et al., The effect of monomer order on the hydrolysis of biodegradable poly (lactic-co-glycolic acid) repeating sequence copolymers. Journal of the American Chemical Society, 2012. 134(39): p. 16352-16359.

10. Pearce, R., et al., Blends of bacterial and synthetic poly (beta-hydroxybutyrate): effect of tacticity on melting behaviour. Polymer, 1992. 33(21): p. 4647-4649.

11. Li, J., R.M. Stayshich, and T.Y. Meyer, Exploiting sequence to control the hydrolysis behavior of biodegradable PLGA copolymers. Journal of the American Chemical Society, 2011. 133(18): p. 6910-6913.

12. Hong, Y., et al., Tailoring the degradation kinetics of poly(ester carbonate urethane)urea thermoplastic elastomers for tissue engineering scaffolds. Biomaterials, 2010. 31(15): p. 4249-58.

13. Shenoi, R.A., et al., Biodegradable polyglycerols with randomly distributed ketal groups as multi-functional drug delivery systems. Biomaterials, 2013. 34(25): p. 6068-81.

14. Lyu, S. and D. Untereker, Degradability of polymers for implantable biomedical devices. International journal of molecular sciences, 2009. 10(9): p. 4033-65.

15. Dorati, R., et al., Effect of porogen on the physico-chemical properties and degradation performance of PLGA scaffolds. Polymer Degradation and Stability, 2009. 95(4): p. 694-701.

16. Ghanbarzadeh, B., H. Almasi, and A.A. Entezami, Physical properties of edible modified starch/carboxymethyl cellulose films. Innovative Food Science & Emerging Technologies, 2010. 11(4): p. 697-702.

17. Huang, C.L., et al., The influence of additives in modulating drug delivery and degradation of PLGA thin films. Asia Materials, 2013. 5(7): p. e54.

18. Liu, Y. and S.P. Schwendeman, Mapping microclimate pH distribution inside protein-encapsulated PLGA microspheres using confocal laser scanning microscopy. Molecular pharmaceutics, 2012. 9(5): p. 1342-50.

19. Reiche, J., et al., Current status of Langmuir monolayer degradation of polymeric biomaterials. The International journal Artificial Organs, 2011. 34(2): p. 123-8.

20. Skoog, D.A., F.J. Holler, and S.R. Crouch, Principles of instumental analysis. sixth edition ed. 2007.

21. Mohammad, A.K. and J.J. Reineke, Quantitative detection of PLGA nanoparticle degradation in tissues following intravenous administration. Molecular pharmaceutics, 2013. 10(6): p. 2183-2189.

22. Pan, Z. and J. Ding, Poly(lactide-co-glycolide) porous scaffolds for tissue engineering and regenerative medicine. Interface Focus, 2011. 2(3): p. 366-77.

23. Wu, L. and J. Ding, In vitro degradation of three-dimensional porous poly (d, l-lactide-co-glycolide) scaffolds for tissue engineering. Biomaterials, 2004. 25(27): p. 5821-5830.

24. Burkersroda, F.v., L. Schedl, and A. Gopferich, Why degradable polymers undergo surface erosion or bulk erosion. Biomaterials, 2002. 23(21): p. 4221-4231.

25. Li, S., Hydrolytic degradation characteristics of aliphatic polyesters derived from lactic and glycolic acids. Journal of biomedical materials research, 1999. 48(3): p. 342-353.

Page 39


26. Goderis, B., et al., Use of SAXS and linear correlation functions for the determination of the crystallinity and morphology of semiâ€ crystalline polymers. Application to linear polyethylene. Journal of Polymer Science, Part B: Polymer Physics, 1999. 37(14): p. 1715-1738.

27. Krimm, S. and A.V. Tobolsky, Quantitative x-ray studies of order in amorphous and crystalline polymers. Quantitative x-ray determination of crystallinity in polyethylene. Journal of Polymer Science, 1951. 7(1): p. 57-76.

28. Qin, D. and R.T. Kean, Crystallinity determination of polylactide by FT-Raman spectrometry. Applied spectroscopy, 1998. 52(4): p. 488-495.

29. Ghanbarzadeh, B. and H. Almasi, Physical properties of edible emulsified films based on carboxymethyl cellulose and oleic acid. International journal of biological Macromolecules, 2011. 48(1): p. 44-9.

30. Teng, W., J. Cappello, and X. Wu, Recombinant silk-elastinlike protein polymer displays elasticity comparable to elastin. Biomacromolecules, 2009. 10(11): p. 3028-36.

31. Teng, W., J. Cappello, and X. Wu, Physical crosslinking modulates sustained drug release from recombinant silk-elastinlike protein polymer for ophthalmic applications. Journal of Controlled Release, 2011. 156(2): p. 186-94.

32. Arm, D.M. and A.F. Tencer, Effects of cyclical mechanical stress on the controlled release of proteins from a biodegradable polymer implant. Journal of Biomedical Materials Research, 1997. 35(4): p. 433-41.

33. Jelonek, K., et al., Controlled poly(l-lactide-co-trimethylene carbonate) delivery system of cyclosporine A and rapamycine--the effect of copolymer chain microstructure on drug release rate. International journal of Pharmaceutics, 2011. 414(1-2): p. 203-9.

34. Siegel, S.J., et al., Effect of drug type on the degradation rate of PLGA matrices. European Journal of Pharmaceutics and Biopharmaceutics, 2006. 64(3): p. 287-293.

35. Choi, J.H., et al., Effect of biocompatible polymers on the physicochemical and dissolution properties of fenofibrate in nanoparticle system. Journal of Pharmaceutical Investigation, 2013. 43(6): p. 507-512.

36. Jia, Z., D.S. Choi, and H. Chokshi, Determination of drug-polymer binding constants by affinity capillary electrophoresis for aryl propionic acid derivatives and related compounds. Journal of Pharmaceutics Sciences, 2013. 102(3): p. 960-6.

37. Shen, J. and D.J. Burgess, Accelerated in vitro release testing of implantable PLGA microsphere/PVA hydrogel composite coatings. International journal of pharmaceutics, 2012. 422(1): p. 341-348.

38. Liu, Y., et al., The microclimate pH in poly (D, L-lactide-co-hydroxymethyl glycolide) microspheres during biodegradation. Biomaterials, 2012. 33(30): p. 7584-7593.

39. Antheunis, H., et al., Autocatalytic equation describing the change in molecular weight during hydrolytic degradation of aliphatic polyesters. Biomacromolecules, 2010. 11(4): p. 1118-1124.

40. Chen, Y., S. Zhou, and Q. Li, Mathematical modeling of degradation for bulk-erosive polymers: applications in tissue engineering scaffolds and drug delivery systems. Acta biomaterialia, 2010. 7(3): p. 1140-1149.

41. Tang, C.Y., et al., Damage modeling of degradable polymers under bulk erosion. Journal of Applied Polymer Science, 2012. 128(5): p. 2658-2665.

42. Amatya, S., et al., Drug release testing methods of polymeric particulate drug formulations. Journal of Pharmaceutical Investigation, 2013. 43(4): p. 259-266.

43. Schoubben, A., P. Blasi, and P.P. Deluca, Effect of agitation regimen on the in vitro release of leuprolide from poly (lactic-co-glycolic) acid microparticles. Journal of pharmaceutical sciences, 2011. 101(3): p. 1212-1220.

44. Solorio, L., et al., Noninvasive characterization of the effect of varying PLGA molecular weight blends on in situ forming implant behavior using ultrasound imaging. Theranostics, 2012. 2(11): p. 1064-77.

45. D'Aurizio, E., et al., Biodegradable microspheres loaded with an anti-Parkinson prodrug: an in vivo pharmacokinetic study. Molecular Pharmaceutics, 2011. 8(6): p. 2408-2415.

46. Yang, Y. and Y. Gao, Preparation and in vivo evaluation of thienorphine-loaded PLGA microspheres. An International Journal of Pharmaceutical Sciences, 2010. 65(10): p. 729-732.

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MSc Chemistry - esc.fnwi.uva.nl · INTRODUCTION ... The experienced downside of this product is the retrieval of the polymer after 3 years ... It is learned from naturally occurring