Novel Bioadhesive Complexation Networks for Oral Protein Drug

Chapter 12

Novel Bioadhesive Complexation Networks for Oral Protein Drug Delivery Downloaded by UNIV OF CALIFORNIA SAN FRANCISCO on December 12, 2014 | http://pubs.acs.org Publication Date: September 24, 1998 | doi: 10.1021/bk-1998-0709.ch012

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A. M . Lowman , Nicholas A. Peppas , M . Morishita , and T. Nagai 1

Department of Chemical Engineering, Drexel University, Philadelphia, PA 19104 School of Chemical Engineering, Purdue University, West Lafayette, IN 47907-1283 Department of Pharmaceutics, Hoshi University, Tokyo, Japan 2

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Copolymer networks of poly(methacrylic acid) grafted with poly(ethylene glycol), which exhibit pH-sensitive swelling behavior due to the reversible formation/dissociation of interpolymer complexes, are investigated as potential carriers for protein drugs. The ability of these gels to serve as oral carriers for proteins was examined. The release kinetics of insulin from complexed and uncomplexed gels was studied. The rate of insulin release from the gels was strongly dependent on the pH of the environmental fluid and the copolymer composition. Because of interpolymer compelxation, the release rate of the drug in acidic fluids simulating the conditions of the stomach was an order of magnitude greater than the release rate in fluids simulating that of the upper small intestine. Additionally, these materials exhibit favorable bioadhesive characteristics for oral protein delivery. The mucoadhesive interactions were significantly stronger between the highly swollen, uncomplexed gels and the mucosa. Two problems exist in developing oral delivery systems for peptides and protein drugs, such as insulin. The major problem is the inactivation of the sensitive drugs by proteolytic enzymes in the gastrointestinal (GI) system, mainly in the stomach and the proximal regions of the small intestine (7-5). This can be overcome by designing carriers which can protect the insulin from the harsh environments of the stomach before releasing the drug into more favorable regions of the GI tract, specifically the upper small intestine or the colon (7-7). Additionally, researchers have attempted to incorporate protease inhibitors into oral formulations which serve to prevent degradation of the sensitive drug by the proteolytic enzymes (4,5,8). The other major problem is the slow transport of the macromolecular drugs across the lining of the colon into the blood stream. Researchers have attempted to bypass this hurdle with the addition of compounds known as absorption enhancers which aid the transport of macromolecules across boundaries (4,5).

156

©1998 American Chemical Society

In Tailored Polymeric Materials for Controlled Delivery Systems; McCulloch, I., et al.; ACS Symposium Series; American Chemical Society: Washington, DC, 1998.

Downloaded by UNIV OF CALIFORNIA SAN FRANCISCO on December 12, 2014 | http://pubs.acs.org Publication Date: September 24, 1998 | doi: 10.1021/bk-1998-0709.ch012

157 We have experimented with a delivery system consisting of microparticles of crossl inked copolymers of poly(methacrylic acid) which are grafted by ethylene glycol (P(MAA-g-EG)) and contain insulin. These new systems function because the structure of the copolymers exhibits pH sensitive swelling behavior due to the reversible formation of interpolymer complexes stabilized by hydrogen bonding between the protons of the carboxylic acid group and the etheric groups on the grafted chains (9). The formation of the interpolymer complexes also serves to stabilize the insulin (70,77). Additionally, the presence of the P E G grafts helps maintain the biological activity of the insulin by stabilizing the drug and preventing binding to ionizable backbone chain (72). Complex formation in the insoluble copolymers is sensitive to the nature and pH of the surrounding fluid as well as the copolymer composition and graft chain length (9). In the acidic environment of the stomach, the gels are in the complexed state. Under these conditions insulin cannot readily diffuse through the membrane because of the small pore size, ξ, and is therefore protected from the harsh environment of the stomach. As the particles pass the stomach into the intestine, the environmental pH increases above the transition pH of the gel. The complexes immediately dissociate and the network pore size rapidly increases leading to the diffusive release of insulin. Experimental Section Material Preparation. In previous work with P(MAA-g-EG) hydrogels, the gels were prepared as films (thickness = 0.7 mm) by a solution polymerization technique (9). In order to increase the response of the hydrogels, it was desireable to prepare the hydrogels in the form of micron sized polymeric microparticles. Thus, P(MAA-gEG) microspheres were prepared by a free-radical bulk, suspension polymerization. The suspending phase, 250 ml of Silicon oil (Dow 200 fluid) was added to a three-necked flask and heated to 70° C while being agitated at 250 R P M using an overhead stirrer. A reflux condenser was attached to the flask. The flask was sealed and purged with nitrogen to ensure an inert atmosphere for the reaction vessel. Vacuum distilled methacrylic acid (Sigma Chemical Co., St. Louis, MO) and poly (ethylene glycol) monomethacrylate containing P E G of molecular weight 1000 (Polysciences Inc., Warrington, PA) were mixed in appropriate molar ratios (20 g total monomers). The crosslinking agent, tetraethylene glycol dimethacrylate (Polysciences Inc., Warrington, PA), was added in the amount of 0.75 % moles of total monomers. Following complete dissolution of the monomers, nitrogen was bubbled through the well mixed solution for 30 minutes to remove dissolved oxygen, a free radical scavenger, which would act as an inhibitor. 2,2'-Azobisisobutyronitrile (AIBN) was added in the amount of 0.5% of the total monomers as the thermal reaction initiator. At elevated temperatures (T > 50° C), A I B N degrades into a stable nitrogen molecule and two free radical containing compounds at a rate rapid enough to initiate the polymerization reaction (73). Poly(dimetyl siloxane-b-ethylene oxide) (Polysciences Inc., Warrington, PA) containing 25% DMS was added in the amount of 1% weight of total monomers as a surfactant to prevent microparticle aggregation during and after the reaction. The monomer mixture was added to the oil phase, agitated at 350 rpm and allowed to react for 3 hours at 70° C. After 3 hours, the temperature was increased to 90° C and allowed to react for an additional 2 hours. Following the higher



158 temperature reaction period, the solution was cooled to 37° C and the agitation rate was decreased to 250 R P M . Once the temperature reached 37° C, 20 ml of deionized water was added to polymer suspension and mixed for an additional 2 hours. The suspension was allowed to settle and the oil was decanted. The reaction flask was filled with deionized water and the swollen particles were stirred for 24 hours at 100 R P M . After 24 hours, the particles were filtered and rinsed with deionized water. This process was continued until all of the silicon oil had been removed. Following the washing, the particles were stored in deionized water with the pH adjusted to 8 by the addition of NaOH The equilibrium swelling of the particles was determined using laser light scattering. Dilute samples of particles were placed in dimethyl glutaric acid (DMGA) buffered solutions of pH = 3.2 and 7.4 and allowed to swell to equilibrium. The ionic strength was adjusted to 0.1 M with the addition of NaCl. Assuming isotropic swelling, the ratio of the volume swelling ratios in the complexed (pH = 3.2) and uncomplexed states (pH = 7.4) was determined as the ratio of the swollen diameters cubed. Drug Loading. Drug loading was accomplished by equilibrium partitioning of the insulin into the P(MAA-g-EG) microparticles. Bovine pancreatic insulin (10 mg, 27.6 IU/mg, M W 5,743, Sigma Chemical Company, St. Louis, MO) was dissolved in 200 μΐ of 1 N HC1. The insulin solution was diluted with 20 ml of phosphate buffer solution (pH = 7.4) and normalized with 200 μΐ of 0.1 Ν NaOH. Loading was accomplished by soaking 150 mg of dry P(MAA-g-EG) microparticles, which had been dried under vacuum at 37° C, for 24 hours in the insulin solution. Additionally, some of the solutions included 5% weight aprotinin, a protease inhibitor. The concentration of insulin in the solution was monitored over time using H P L C . The particles were then filtered and washed with 100 ml of 0.1 N HC1 solution to collapse the microparticles and "squeeze out" the remaining buffer solution. The drug loaded microspheres were dried under vacuum and stored at 4° C. The degree of loading was determined from H P L C analysis of the insulin concentrations of the initial solutions and the filtrate from the washings. In Vitro Release Studies. The in vitro release of insulin from P(MAA-g-EG) microspheres was also studied. Dry, insulin loaded polymer microparticles were swollen in 200 ml fluid of pH 1.2 and ionic strength of 0.1 M (simulated gastric fluid) for 2 hours and agitated at 125 R P M . The polymer microparticles were then transferred to pH = 6.8 phosphate buffered saline solutions. Samples were taken at discrete intervals and the insulin concentration in the solutions was monitored using HPLC. Mucoadhesion Studies. Films of P(MAA-g-EG) (thickness = 0.9 mm) were prepared using a solution polymerization technique described previously (9). These copolymers contained equimolar amounts of M A A and E G and P E G grafts of molecular weight 1000. The copolymers were swollen to equilibrium in D M G A buffered saline solutions of pH = 3.2 and 7.4. Upon reaching equilibrium, the swollen gels were cut into disks with diameters of 19 cm. The polymer samples were adhered to the upper holder of a tensile tester (Instron Model 4301, 10 Ν load cell, Park Ridge, IL) at 25° C and 90% R H using cyanoacrylate medical adhesive,


159 whereas a sample of gelled bovine submaxillary mucin (Sigma, St. Louis, MO) was affixed on the lower jaws. The two jaws were brought together for 15 min to allow for equilibration and then separated at 1 mm/min. The detachment force was measured as a function of displacement.


Results and Discussion Equilibrium Swelling of Polymer Microparticles. The equilibrium volume swelling ratios of the polymer microspheres were determined using laser light scattering. The average diameter for swollen P(MAA-g-EG) microspheres containing a 1:1 M A A / E G ratio and graft PEG chains of molecular weight 1000 swollen in pH = 3.2 D M G A buffered solution was 10 ± 4 μηι. The diameter for the microspheres swollen in pH = 7.4 buffer solution was 30 ± 12 μπι. Assuming isotropic swelling, the ratio of the volume swelling ratios in the two states (Quncom/Qcom) was 30. The similar ratio for hydrogels prepared as films using a photo-initiated solution polymerization was 15. Thus, a greater transition existed between the uncomplexed and complexed states for the microparticles prepared by the suspension polymerization technique. In Vitro Release Results. The pulsatile release of insulin from P(MAA-g-EG) microparticles of different concentrations and P M A A microparticles is shown in Figure 1. For all of the gels studied, the rate of insulin release was significantly lower in the lower pH solution. Under these conditions, the P E G containing gels were in the collapsed state due to the formation of interpolymer complexes and the hydropholic nature of the gel in acidic media. In the collapsed state, the additional physical crosslinks were formed and the backbone chains were in the coiled conformation. The small mesh size or pore size of the gel served to hinder the diffusion of insulin through the gel. The P M A A gels were also in the collapsed state due to the hydrophobic nature of the gel in solutions of low pH. In these samples, the small network mesh size was due to the coiling of the polymer chains because of the poor compatibility between the swelling agent and the hydrophobic polymer chains. When the gels were transferred to the higher pH solution, the release rates of insulin increased dramatically. In all of the gels, the acidic groups were ionized and the gels were swelled rapidly due to the large swelling force generated due to ionization of the pendant groups and increases in the osmotic swelling force. In the PEG-containing copolymers, ionization resulted in the dissociation of the interpolymer complexes. The most rapid release of insulin occurred in the gels containing the greatest amounts of PEG. In these gels, the tethered P E G chains served to prevent binding of the peptide to the charged groups along the polymer backbone. In the other gels, the release rates slowed due to ionic interactions between the backbone polymer (PMAA) and the insulin. For the case of the gels swelling in the simulated gastric fluid, the largest amount of insulin was released from the P M A A gels, in which no complexation occured (Figure 2). However, the amount of insulin released from the gels was decreased by increasing the amount of P E G incorporated into the network structure. In gels containing PEG, interpolymer complexes formed in the gels swollen in pH = 1.2 solution. Due to the formation of complexes, the network mesh size was reduced resulting in a significant decrease in rate of release of the drug from the gels.



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Time, t (h) Figure 1. Pulsatile release of insulin in vitro from P(MAA-g-EG) microparticles containing graft PEG chains of molecular weight 1000 and M A A / E G ratios of ( ο ) 1 and (•) 4 in simulated gastric fluid (pH = 1.2) for the first two hours and phosphate buffered saline solutions (pH = 6.8) for the remaining three hours at 37° C. The release from P(MAA) microparticles is represented by Δ. 0.8 0.7 8

0.6 Έ

IS

0.5

Φ IT

0.4

Ίο c ο

0.3

1 LL

0.2 0.1

0.5

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Time, t (h) Figure 2. Controlled release of insulin in vitro from P(MAA-g-EG) microparticles containing graft PEG chains of molecular weight 1000 and M A A / E G ratios of ( ο ) 1 and (•) 4 in simulated gastric fluid (pH = 1.2) at 37° C. The release from P(MAA) microparticles is represented by A. In Tailored Polymeric Materials for Controlled Delivery Systems; McCulloch, I., et al.; ACS Symposium Series; American Chemical Society: Washington, DC, 1998.


161 Additionally, complexation was enhanced in gels containing equimolar amounts of M A A / E G and these gels had the smallest network mesh size due to the presence of the largest number of physical crosslinks in these materials. Accordingly, the least amount of insulin (< 10%) was released in the gastric fluid from gels containing a 1:1 ratio of M A A / E G . In designing a device for oral delivery of a protein drug, it is important to protect the drug in the stomach and release the drug into the more favorable regions of the GI tract. Therefore, in an effective carrier the release rates and diffusion coefficients will be significantly greater in neutral or basic conditions than acidic conditions. One significant parameter in evaluating the viability of a particular hydrogel for oral delivery of proteins in peptides is the ratio of the diffusion coefficients of the drug in the carrier in the stomach (acid environment) and the intestine (neutral environment). This parameter can be calculated from the Fickian expression at short times where the fractional drug released is proportional to the square root of the diffusion coefficient. Therefore, the ratio of diffusion coefficients was calculated as: D .8 [(Mt/MJigAt^l The ratio of diffusion coefficients for insulin in the gels is shown in Table I. The greatest change in diffusion coefficients between the two pH solutions was for the gels containing a 1:1 ratio of M A A / E G due to the formation of complexes in the acidic media and the reduction of polymer/drug interactions in the high pH solution due to the presence of the PEG grafts. In these gels, the diffusion coefficients varied by greater than three orders of magnitude corresponding to greater than an order of magnitude difference in the release rate between the two solutions. In gels containing lower amounts of PEG, fewer complexes formed and increased drug/polymer binding occurred resulting in a significantly lower variation in release rates between the two solutions. 6

Table I· Ratio of diffusion coefficients of insulin in the complexed and uncomplexed hydrogels as compared to P M A A .

Composition

Di.2/D .8

1:1 M A A / E G

1.79 χ ΙΟ"

4:1 M A A / E G

0.152

PMAA

0.359

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Mucoadhesive Behavior of P(MAA-g-EG) Hydrogels. The primary goal of bioadhesive controlled drug delivery is to localize a delivery device within the body to enhance the drug absorption process in a site-specific manner (14-17). Hydrogels of P(MAA-g-EG) exhibit mucoadhesive characteristics due to the presence of the graft PEG chains which serve as adhesion promoters. Adhesion promoters, such as polymer grafts or even linear polymers, function by penetrating the gel/mucosa or gel/gel interface and forming temporary anchors (18,19).



162 The mucoadhesive characteristics of P(MAA-g-EG) hydrogels were strongly dependent on the pH of the environmental fluid (Figure 3). The work of mucoadhesion was calculated as the area under the curves for each pH value (Table II). The work of adhesion was significantly higher at the pH value of 7.4. However, to truly compare the mucoadhesive characeteristics of the gels, the work of adhesion was normalized to account for the polymer gel fraction. The normalized work of adhesion was two-orders of magnitude greater for gels in the uncomplexed state (Table II). Accordingly, the gels would adhere to the mucosa of the intestine to a much greater extent than the stomach. Therefore, the residence of time of the gels would be much greater in regions where the insulin could be absorbed. Table IL Work of adhesion for P(MAA-g-EG) gels containing a 1:1 MAA/EG ratio and graft PEG chains of molecular weight 1000.

pH

Normalized Work of Adhesion, W/(V 2, Î^*l(fi(J)

Work of Adhesion, W*1&(J)

Polymer Volume Fraction, ν χ,

3.2

5.38

0.693

62.1

7.4

9.34

0.049

6720

s

The differences in the adhesive characteristics of the gels at different pH values was due to differences in the mobility of the PEG chains in each material (Figure 4). In the highly swollen, uncomplexed state, the graft PEG chains were free and readily penetrated the mucosa to serve as anchors for adhesion. In the complexed state, the graft P E G chains in the P(MAA-g-EG) formed complexes with the backbone chains and were unable to penetrate the gel/mucosa interface and form temporary anchors. Conclusions The rate of insulin permeation through P(MAA-g-EG) gels was strongly dependent on the pH of the environmental fluid and the copolymer composition. Because of interpolymer compelxation, the release rate of the drug in acidic fluids simulating the conditions of the stomach was an order of magnitude greater than the release rate in fluids simulating that of the upper small intestine (pH = 6.8). Addtionally, the largest change in the release rates between the two fluids was greatest for gels containing equimolar amounts of M A A / E G in which the largest number of complexes formed in acidic media. Additionally, the interactions between the mucosa and the gel were significantly stronger when the gel was in the highly swollen, uncomplexed state. Interpolymer complexation in acidic prevented penetration of the PEG grafts through gel/mucosa interface. Under complex breaking conditions, the tethered P E G chains were able to penetrate the gel/mucosa interface and serve as anchors to promote adhesion.


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0.006

Displacement (cm) Figure 3. Adhesive behavior of P(MAA-g-EG) gels containing a 1:1 M A A / E G ratio and graft PEG chains of molecular weight 1000 at pH values of ( — ) 3.2 and ( )7.4 in contact with bovine submaxillary gland mucin.

Figure 4. Proposed mechanism of adhesion between P(MAA-g-EG) hydrogels in the complexed and uncomplexed states and the mucosa.


164 Acknowledgments This work was supported by grants from the National Institutes of Health and the Nagai Foundation, Tokyo.


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Novel Bioadhesive Complexation Networks for Oral Protein Drug

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