MSc Chemistry. Determination of degradation behavior of biodegradable polymers for pharmaceutical devices

Transcription

1 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

2 Page 2 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.

3 Page 3 TABLE OF CONTENTS Page 1. INTRODUCTION BIODEGRADABLE POLYMERS COMPOSITION OF BIODEGRADABLE POLYMERS BIODEGRADATION BY HYDROLYSIS PHYSICAL PROPERTIES MOLECULAR WEIGHT HYDROPHOBICITY MORPHOLOGY CROSSLINKING POROSITY SURFACE API-POLYMER INTERACTION AUTOCATALYSIS EROSION PROFILING INTERNAL ACIDITY POROSITY MODDELLING OF AUTOCATALYSIS IN VITRO / IN VIVO CORRELATION IN VITRO DETERMINATIONS IN VIVO DETERMINATIONS IN VITRO / IN VIVO CORRELATION CONCLUSION... 37

4 Page 4 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.

5 Page 5 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.

6 Page 6 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].

7 Page 7 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

8 Page 8 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.

9 Page 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 37 o C 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 (W t ) minus the initial weight of that sample (W 0 ). 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 (MA t ) the weight of the swollen sample (W wet ) and the weight of that same sample after drying (W dry ) 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.

10 Page 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 (M n or M w, 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 M n or M w 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 M w in time; reprinted from [21]

11 Page 11 By monitoring the M w 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 M w is called the erosion onset[14, 24, 25]. In figure 3 the first two points have similar M w : 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 M n ; 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.

12 Page 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

13 Page 13 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 (T g ). The T g is the temperature where the polymer transits from a glassy state to a rubbery state. When crossing the T g 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 T g, the heat flow changes. The heat flow is plotted in a thermogram[20]. A T g 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 T g decreases from 56 o C to 44 o C during a 110 day in vitro

14 Page 14 degradation measurement. The deep negative peak with values of 168 o C to 153 o C corresponds to the melting temperature and the positive peaks are recrystallization peaks from the crystallization of the degradation byproducts 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 (T g ) and morphology. Still, the relation of T g with degradation is not always seen[2]. Probably, this behavior can be explained by the fact that the T g 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 T g [10, 22]. T g 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 T g as an identification of the morphology is used by most literature reviewed in this thesis. In some cases a spectroscopic techniques is used additionally.

15 Page 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].

16 Page 16 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]

17 Page 17 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]

18 Page 18 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]

19 Page 19 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 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.

20 Page 20 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, K b s, 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 K b s, 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.

21 Page 21 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 K b s 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 K b s 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.

22 Page 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).

23 Page 23 = 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 (L critical ) at which a device changes its erosion profile (from bulk to surface erosion) for a number of polymers. Figure 22 gives the L critical for several polymers. Figure 20) Graph of the mathematical model that determines the erosion profile; Figure 22) Critical dimensions for change of erosion profile; reprinted from[24] reprinted from[24]

24 Page 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].

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).

26 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 M w 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.

27 Page 27 Figure 28) M w 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 M n over time as a results of autocatalysis Antheunis et al. described a autocatalytic equation for aliphatic polyesters[39]. ( ) = [ ] 1 1+ [ ] [ ] + 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.

28 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.

29 Page 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).

30 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 T g is not within the temperature range. If T g is passed the degradation and release mechanisms are very different because the polymer is in a different physical state.

31 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 37 o C. 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 M w of two different polymers and their stirring technique. Figure 35) Decrease of M w of two polymer samples at different agitation rates in vitro; reprinted from[43] Figure 36); 3 phases of degradation M n 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 M n. 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].

32 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.

33 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]

34 Page 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 37 o C 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 (F a ). 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]

35 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 37 o C. 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 M w, M 0 is the initial molecular weight, k the rate constant of the degradation and t is the time. The decrease of M w 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.