Polyester Dendritic Systems for Drug Delivery

Bioconjugate Chem. 2002, 13, 453−461

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Polyester Dendritic Systems for Drug Delivery Applications: In Vitro and In Vivo Evaluation Omayra L. Padilla De Jesu´s,† Henrik R. Ihre,† Lucie Gagne,‡ Jean M. J. Frećhet,*,† and Francis C. Szoka, Jr.*,‡

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Department of Chemistry, University of California, Berkeley, California 94720-1460 and School of Pharmacy, University of California, San Francisco, California 94143-0446. Received October 22, 2001

High molecular weight polymers (> 20 000 Da) have been widely used as soluble drug carriers to improve drug targeting and therapeutic efficacy. Dendritic polymers are exceptional candidates for the preparation of near monodisperse drug carriers due to their well-defined structure, multivalency, and flexibility for tailored functionalization. We evaluated various dendritic architectures composed of a polyester dendritic scaffold based on the monomer unit 2,2-bis(hydroxymethyl)propanoic acid for their suitability as drug carriers both in vitro and in vivo. These systems are both water soluble and nontoxic. In addition, the potent anticancer drug, doxorubicin, was covalently bound via a hydrazone linkage to a high molecular weight 3-arm poly(ethylene oxide)-dendrimer hybrid. Drug release was a function of pH, and the release rate was more rapid at pH < 6. The cytotoxicity of the DOX-polymer conjugate measured on multiple cancer lines in vitro was reduced but not eliminated, indicating that some active doxorubicin was released from the drug polymer conjugate under physiological conditions. Furthermore, biodistribution experiments show little accumulation of the DOX-polymer conjugate in vital organs, and the serum half-life of doxorubicin attached to an appropriate high molecular weight polymer has been significantly increased when compared to the free drug. Thus, this new macromolecular system exhibits promising characteristics for the development of new polymeric drug carriers.

INTRODUCTION

A variety of approaches have been implemented for the discovery of medicinal compounds, including the isolation of naturally occurring bioactive substances and the design and synthesis of new molecules with biological activity. Many compounds have been identified by these means, but the lack of selectivity of the majority of current highly toxic drugs used for chemotherapy has motivated the search for other means to improve their therapeutic efficacy and increase their selectivity (1). In the past few decades, the use of polymers as carriers of both covalently bound (2-5) and physically entrapped drug molecules (6-8) has been widely explored. When a drug is linked to a polymer, it loses activity because it cannot interact with the drug target either due to steric constraints or the inability of the drug to reach the target site (9). Furthermore, the drug-polymer conjugate can no longer cross the cellular membrane by simple diffusion, but must do so by endocytosis. The resulting endosomal compartment containing the polymeric conjugate fuses with the more acidic lysosomal compartment (pH 4-5), whose hydrolytic enzymes then degrade the conjugate. This mechanism allows a more selective uptake of therapeutic agents, in contrast to small drugs that can easily diffuse into and out of the cell (2, 10). In certain cells, the endocytotic uptake enables drugpolymer conjugates to bypass multidrug resistance (MDR) mediated efflux mechanisms hence to overcome drug resistance in the cancer cells (11, 12). Furthermore, the anatomical differences between normal and tumor tissue allows for preferential accumulation of a polymeric drug in neoplastic tissue. High molecular weight polymers preferentially accumulate in solid tumor tissue due to a combination of the leaky character of tumor blood vessels † ‡

University of California, Berkeley. University of California, San Francisco.

formed during neo-angiogenesis and to limited lymphatic drainage. The combination of these two factors is responsible for the enhanced permeability and retention effect (EPR) observed with tumor tissue, which leads to a passive targeting of drugs to tumors (13, 14). In addition, the larger hydrodynamic volume of polymers contributes to the increased plasma half-life of the drug-polymer conjugates, increasing the probability of accumulation of the therapeutic agent in the tumor tissue by means of the EPR effect (3, 15). Drugs have also been conjugated to polymers to improve their water solubility properties, to decrease their toxicity due to local accumulation of the drug prior to reaching the target tissue, and to protect them from possible enzymatic degradation or hydrolysis. Polymers previously evaluated for delivery of covalently attached drugs have primarily been random coil linear polymers such as poly(ethylene glycol) (PEG) (16), N-(2-hydroxypropyl)methacrylate (HPMA) (17), and styrene maleic anhydride (SMA) (5). However, the synthesis of some of these systems results in highly polydisperse materials with a wide range of molecular weights due to their mode of preparation (18). Molecules of different molecular weights found in a polydisperse sample will behave differently in the body. Therefore, either monodisperse or low-polydisperse (PDI < 1.1) polymers are required to achieve systems displaying homogeneous and reproducible pharmacokinetic properties. The intrinsic properties of dendrimers, with their well-defined structures and their flexibility for tailored functionalization, classify these molecules as good candidates for the assembly of multivalent drug carriers (19-25). The well-designed step-by-step preparation of dendrimers results in a monodisperse system (26-28). Few dendritic systems have been evaluated previously, due to their lack of availability. At present, probably the best-known dendritic system is the polyamidoamine (PAMAM) Starburst family of dendrimers (29). These are

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commercially available, but their polycationic character must be pacified to reduce their in vivo toxicity (30-33). In addition, their polyamidoamine linkage is resistant to degradation in the body. Another well-known dendritic system is based on the polyaryl ether convergent dendrimers (34). Although promising preliminary studies were performed with the latter by coupling the drug methotrexate (MTX) (21), these dendrimers also have a backbone that is inherently resistant to degradation and thus were thought to be less suitable for drug delivery purposes than other structures incorporating hydrolyzable linkages. Thus, it is of interest to develop and evaluate novel water-soluble, biocompatible, dendritic systems. A polyester dendritic scaffold based on 2,2-bis(hydroxymethyl)propanoic acid has been investigated as a possible drug carrier. Herein, the biological evaluation of various polyester dendritic model compounds is described. The biodistribution of a doxorubicin-polymer conjugate is also reported, as well as drug release studies. The synthesis and chemical characterization of the systems evaluated are reported elsewhere (35). MATERIALS AND METHODS

Cell Culture. The murine melanoma cell line B16F10 was obtained from the UCSF Cell Culture Facility. The cells were cultured in prewarmed medium (MEM Eagle’s with EBSS medium containing 10% fetal bovine serum, 1% MEM nonessential amino acids, 110 mg/L sodium pyruvate, and 1% penicillin-streptomycin) at 37 °C in a humidified atmosphere composed of 5% CO2. MDA-MB231 and MDA-MB-435 breast cancer cell lines were cultured in a medium consisting of Dulbecco’s Modified Earl’s medium (DME) H-16, containing 10% FBS and 1% MEM nonessential amino acids. The medium for culturing the monkey kidney fibroblast CV-1 cells consisted of DME H-21 (high glucose 4.5 g/L) containing 10% fetal bovine serum, 1% MEM nonessential amino acids, 1% HEPES buffer, and 1% penicillin-streptomycin. For all of our experiments, cells were harvested from subconfluent cultures using trypsin-versene (0.05 and 0.02% respectively) and were suspended in medium. Cell viability was determined using the trypan blue exclusion method. Cell Cytotoxicity Assay. The cytotoxicity of the polymeric carriers, and of the free or conjugated doxorubicin, was determined using a target cell line and the sulforhodamine B assay (SRB) (36). The target cells used were murine B16F10 melanoma cells and the human breast cancer cell lines MDA-MD-231 and MDA-MD-435. Cells were seeded into a 96-well plate at a density of 3 × 104 cells/mL with 100 µL per well for 19 h (37 °C and 5% CO2) before the assay. The medium of each well was then replaced by 100 µL of antibiotic-free medium containing various concentrations of the polymer with no drug, drug, or conjugated doxorubicin. The tests were conducted in replicates of four to eight for each drug dose. The cells were incubated at 37 °C/5% CO2. Following the incubation period, the cells were fixed for 1 h at 4 °C by adding 25 µL of ice-cold 50% trichloroacetic acid (TCA) to the growth medium of each well. The wells were washed 5 times with cold water to remove the excess TCA and then air-dried at room temperature for several minutes. Once dry, each well received 100 µL of 0.4% SRB in 1% acetic acid followed with a 30-min incubation at room temperature. The SRB solution was then aspirated off the cells and the wells were quickly washed with 1% acetic acid. The wells were air-dried at room temperature for several minutes until moisture was no longer

Padilla De Jesu´s et al.

detected visually. The dye bound to the cells was solubilized by adding 100 µL of 10 mM Tris base, unbuffered, per well and agitated for 5 min at room temperature. The optical densities were obtained using an OPTIMAX microplate reader (Molecular Devices, Sunnyvale, CA) at a wavelength of 564 nm. OD measured on wells containing cells that did not receive the drug were considered to represent 100% growth, and OD measured on wells containing no cells but that received drug are considered to represent 0% growth. The IC50 is evaluated by comparing the OD of wells containing cells that were exposed to at least five doses. Maximum Tolerated Polymer Intravenous Bolus Dose. The polymers were dissolved at 162 mg/mL in HBS. CD-1 female mice, 6-8 week-old, were injected with increasing volumes of polymer solution via the tail vein and observed for 24 h postinjection. All mice were then anesthetized with an intraperitoneal injection of ketamine cocktail (44 mg Ketamine; 2.5 mg Xylazine; 0.75 mg acepromazine in PBS per kg of body mass) and their liver dissected and observed under a dissecting microscope for signs of toxicity. Radiolabeling Methods. Compound I and II were iodinated using carrier-free 125I and the chloramine T reaction as previously described (37). The radiolabeled polymer was separated from noncovalently attached iodine on a Sephadex G-25 column. Compound III containing the methoxyphenol group was not readily radiolabeled when using the chloramines T reaction described in ref 37 but was readily radioiodinated using noncarrier iodine and an electrophilic protocol based upon Krummeich et al (38). In brief, 3 mg of III was dissolved in 0.25 mL of trifluoroacetic acid (TFA), and 0.01 mL of an aqueous solution of 125I was added containing 1 mCi 125I. Chloramine T (0.54 mg) dissolved in 0.05 mL of TFA was added to the mixture, and the reaction was heated at 60 °C for 5 min. The reaction was stopped by the addition of sodium metabisulfite, 1 mg in 0.01 mL of water, and the TFA was evaporated under a stream of nitrogen that was vented through a trap containing activated carbon (Dupont-New England Nuclear, MA). The residue was dissolved in 0.1 mL of 0.1 N NaOH containing 100 mM sodium metabisulfite and chromatographed on a column containing 1.8 mL bed volume of Dowex 2 anion-exchange resin. The radiolabeled polymer was eluted in the first 4 mL using water as the mobile phase. There was about a 20% incorporation of radioiodine into the polymer under these conditions. The mixture was rechromatographed two additional times, and greater than 99% of the radioactivity eluted with the polymer fraction. The radiolabeled material eluted in the void volume from a Sephadex 25 size exclusion column. A suitable quantity of radiolabeled polymer was mixed with the nonradiolabeled polymer in sterile HBS to create a 10 mg/mL polymer concentration with a specific activity of 4 µCi/mL for use in the biodistribution studies. Biodistribution Assays for Compounds I, II, and III. A known amount of radiolabeled polymer was added to a stock solution of nonradiolabeled polymer prepared at a concentration of 10 mg/mL in sterile HBS, and 100 µL was injected intravenously into 6-8 week-old CD-1 female mice. Each mouse was housed separately to collect the urine and feces. At different times postinjection, either three or four mice per group were anesthetized with an intraperitoneal injection of a ketamine cocktail and sacrificed. The feces and urine were collected. Each mouse was bled by heart puncture, and the volume of blood measured and transferred to a scintillation vial. The circulation system was perfused with PBS via the

Polyester Dendritic Systems for Drug Delivery

right ventricle to reduce blood in the organs. All the organs were collected separately and transferred to preweighed scintillation vials. The carcass was sectioned into three portions and placed into preweighted vials. The radioactivity contained in each organ was counted on a Beckman Gamma 8000 Spectrometer (Irvine, CA) and expressed as counts per minute/gram of tissue. Using this procedure we were able to account for 95 ( 5% of the injected radioactivity. Blood Clearance Assay for Compounds I, II, and III. Female mice of 6-8 week-old (Charles River) were injected with 1 mg of radiolabeled polymer in 100 µL of HBS via the tail vein. At time 3, 10, 30, 60, and 90 min postinjection, each mouse was anesthetized with isoflurane, and 50 µL of blood was collected from the orbital sinus using a heparinized capillary. The level of radioactivity (cpm) found in the blood was determined using a Beckman Gamma 8000 gamma spectrometer (Irvine, CA) and the elimination rate computed using a loglinear regression of CPM/mL blood versus time after injection. Biodistribution Assay for Compound IV. A known amount of the doxorubicin conjugate was dissolved in a 5 mM HEPES and 5% glucose solution. CD-1 females mice at 6-8 weeks old were injected i. v. via tail with 200 µL of the conjugate solution. Four mice per group were anesthetized with a ketamine cocktail by i. p. injection. All mice were bled by heart puncture, and the volume of blood was measured and transferred to a scintillation vial. An injection of 1 mL of PBS to the right ventricle of the heart was used to flush the circulation system. The whole organs were dissected and their weight was recorded. The organs were put into 2 mL tubes containing 1 mL of acidified alcohol (90% 2-propanol/0.075 HCl) and zirconia beads, followed by homogenization (Bead Beater, Biospec, Bartlesville, OK) for 20 s at 5000 rpm. The blood was allowed to coagulate at 4 °C and then centrifuged for 10 min at 1000 rpm. The serum (upper phase) was collected, and its volume was recorded. An aliquot of 100 µL of serum was transferred to a 2-mL tube containing 1 mL of acidified alcohol. The homogenized organ samples and the sera were placed in the refrigerator for 24 h to extract the drug. The cell debris in the organ homogenates was removed by centrifugation at 14 000 rpm for 5 min at 4 °C. An aliquot (400 µL) of the supernatants were transferred to tubes containing 1.6 mL of acidified alcohol. For the quantification of doxorubicin present in the organs, drug-free homogenates of organs and sera from untreated mice were spiked with free doxorubicin. Various organs had different background fluorescence levels, which required separate calibration curves for each organ. Fluorescence emission from the drug was measured in fluorescence intensity units (Ex 490 nm: Em 590 nm) in a Quantech Turner fluorometer (Barnstead/Thermolyne, Dubuque, IO). Calibration curves were linear over two orders and drug levels to 10 ng/g organ could be reliably measured. Confocal Microscopy. GFP transfected and nontransfected MDA-MB-435 cells were plated on sterilized 4-well slides by the addition of 1 mL of a 1 × 105 cells/ mL solution. After incubation for 24 h, the media was gently removed and the cells were washed once with 1 mL of PBS, followed by the addition of 0.5 mL of media containing free doxorubicin and conjugated doxorubicin. The control cells contained only media. The concentration of free and conjugated doxorubicin equals the IC50 of free doxorubicin for this cell line, as determined previously in vitro experiments (0.62 mg/mL). Cells were incubated at 37 °C and 5% CO2 for 1 and 24 h. After incubation the

Bioconjugate Chem., Vol. 13, No. 3, 2002 455

Figure 1. Dendritic polyester model compound I with molecular mass of 3790 Da.

cells were fixed for 1 h at 4 °C in 1.0 mL of 4% paraformaldehyde in PBS. The cells were mounted in Vector Vectrashield (Burlingame, CA) for confocal microscopy. Confocal images were obtained using LaserSharpe software installed on a Bio-Rad (Hercules, CA) kryptonargon laser scanning confocal microscope (MRC-1024) outfitted with a Nikon Diaphot 200 microscope using an ×60 objective lens. 512 × 512 pixel sections were taken every 2 µm using all laser lines, 10% laser power, slow scan, and Kalman averaging of three scans. For both GFP transfected and untransfected MDA-MB-5 cells, doxorubicin was imaged by PMT1 using low signal amplification, an iris of 2.5, a gain of 1500, and a black level of -6. For the GFP transfected MDA-MB-5, GFP was imaged on PMT2 using an iris of 2.5, a gain of 1500, and a black level of -3. Images were analyzed using Confocal Assistant 4.02 software. Drug Release Studies as a Function of Buffer pH. A stock solution of conjugate was prepared using distilled water as a solvent. An aliquot of the stock solution was added to the respective prewarmed buffer solutions (37 °C), to initiate the reaction. The final concentration of the conjugate was of 3 mg/mL, 40 mM in buffer ion and 100 mM NaCl. Solutions at pH 2.4 (HCl), 4.5 (acetate), 5.5 (acetate), 6.5 (MES), and 7.0 (phosphate) were used to evaluate drug release as a function of pH. Quantification of drug released was performed by HPLC analysis on a model 2690 XE Alliance Separations module (Waters, USA) and a Photodiode Array detector (Waters 996). Aliquots of 10 µL were injected with builtin autoinjector at different times after the incubation started. The separation was carried out with a Waters Symmetry C18 column (2.1 × 150 mm, 3.5 µm) using a mobile phase consisting of 40% phosphate buffer pH 7.0 and 60% methanol at a flow rate of 0.3 mL/min. The data acquisition and processing were controlled by Millenium 32 software. In contrast to commonly used mobile phases for the separation of DOX conjugates from the free DOX, we used a neutral pH system, to minimize cleavage of DOX from the polymer during the analysis. A calibration curve was prepared with fresh doxorubicin and the extent of release was monitored by peak height based on the absorbance at 490 nm. RESULTS AND DISCUSSION

Brief Description of Structure of Model Compounds Evaluated. Model compounds I, II, and III were evaluated as possible polymeric drug carriers (Figures 1-3). The common denominator of these systems is the polyester dendritic scaffold derived from the monomer unit 2,2-bis(hydroxymethyl)propanoic acid. These



Figure 2. Modified dendritic polyester model compound II with molecular mass of 11 500 Da.

Figure 3. Model compound III, composed of 3-arm poly(ethylene oxide) star and [G-2] polyester dendrons (23 500 Da).

three polymers differ in molecular weight and architecture, so we could evaluate these factors on toxicity and biodistribution. Compounds I and II mainly consist of two fourth generation [G-4] polyester dendrons attached to a trisphenolic core. Carrier I has a molecular mass of 3790 Da and contains a fully activated surface with 32 hydroxyl groups, which can be further modified. The surface of II was modified with methoxy terminated

tri(ethylene glycol) oligomeric units to increase the mass of the dendritic molecule by a factor of 3 in one single synthetic step, achieving a mass of 11 500 Da. An attractive feature of this polyester scaffold is that it is highly water soluble (at least 200 mg/mL) in an uncharged state, avoiding the necessity of introducing ionizable functional groups to achieve good solubility properties. This property decreases the possibilities for rapid elimination as is observed in the case for cationic systems, due to electrostatic interactions with the negatively charged cellular membranes and extracellular matrixes (30-33) or in the case of anionic systems due to their interaction with the scavenger receptor (39). Second, the sterically hindered ester bond of the monomer unit makes this backbone more stable toward both nucleophilic attack and acid-catalyzed hydrolysis avoiding its rapid and premature degradation. In contrast to I and II, compound III consists of a hybrid between a 3-arm poly(ethylene oxide) (PEO) star and three [G-2] polyester dendritic units. This system has a molecular mass of ca. 23 500 Da, and its molecular architecture has a more linear character, as compared to the compact shape of the strictly dendritic systems I and II. In this instance, the trifunctional PEO core contributes most of the mass of the system. The advantages of using a star-PEO also include its contribution to biocompatibility and water solubility (40), without jeopardizing the polydispersity of the system, since starPEO molecules with low polydispersity indices (PDI < 1.1) are commercially available. On the other hand, the [G-2] polyester dendrons of III provide the multivalency necessary for drug attachment. In contrast to linear polymers containing statistical amounts of reacting sites,



Figure 5. Cell viability results upon treatment of B16F10 cells with polymeric carriers I, II, and III after 48 h of incubation.

Figure 4. Model compound IV. Polymer drug conjugate consisting of modified III and doxorubicin using a hydrazone covalent bond as a linker.

this architecture provides a known and reproducible number of functional groups. Finally, the drug-polymer conjugate IV (Figure 4) consists of the potent anticancer drug doxorubicin (DOX) and a derivative of the polymeric carrier III. DOX was covalently bound to the polymer by means of an acidlabile hydrazone linkage (41-42). An acid-labile linkage was selected because it can remain stable under physiological conditions, but once it is internalized in the cell by endocytosis, the more acidic environment of lysosomal compartments will trigger the release of the drug. An additional advantage of a pH-sensitive linker is that the release rate of the drug attached by an acid-labile linkage is often less affected by the type of drug than in the case of enzyme cleavable linkages, which have been difficult to address (43). Finally, the environment around tumor tissue is slightly more acidic than normal tissue; therefore, an acid labile linkage might be cleaved more rapidly in the vicinity of a tumor prior being internalized by the cancer cell. Biological Evaluation of Model Compounds. In Vitro Cytotoxicity of the Carriers. In vitro studies were

performed on the polymeric model compounds I, II, and III to evaluate the possible toxic effect of the polyester dendritic scaffold itself. The cell viability studies were performed on melanoma B16F10 cancer cells and cell proliferation was quantified by the sulforhodamine B assay (SRB), which measures the level of protein remaining on a cell culture dish. All of the polymers examined were well tolerated by cells in culture (Figure 5). For instance, after incubation for 48 h, cells exposed to I at a polymer concentration of 10 mg/mL retained 87% of the growth rate of the control cells. In the case of model compound II, cell proliferation was slightly suppressed to 81% of control at 5 mg/mL. Model compound III, the highest molecular weight carrier evaluated, also was very well tolerated with 87% of the growth rate of the control culture at 10 mg/mL of polymer (Figure 5). An interesting observation on this polymer is that even though there is an apparent growth inhibition at the highest concentrations, dead cells were not evident under microscopic observation even at 40 mg/mL, rather the cell density was decreased. This suggests a growth inhibition effect as a result of the very high polymer concentration. These results are very promising, considering that previously evaluated dendritic systems, such as PAMAM, polyethyleneimine and poly(ethylene oxide) grafted carbosilanes inhibit cell viability and induce hemolysis or cytolysis at concentrations much lower than 5 mg/mL (31). In Vivo Tolerability of the Dendritic Polyesters. Nonionic polymers such as the poly(ethylene glycols) are very well tolerated upon injection into animals. Given the low toxicity of the polyester dendrimers in culture, we thought they would also be well tolerated upon injection into animals. However, we did not believe it warranted conducting a complete toxicity profile of the polymers at this stage of the investigation; rather, we undertook to learn if the polymers could be administered as an intravenous bolus at levels greater than 1 g/kg body weight in mice. If the polymers could be administered at this dose, a wide variety of drug delivery applications would be possible. Compound I was injected i.v. over a 10 s period into a number of animals at increasing doses. At the dose of 1.3 g/kg body weight, 1 of the 2 animals injected survived. In the animal that died, there was no evidence for blood hemolysis. The surviving animal was sacrificed at 24 h postinjection, the organs removed and visually examined. There was no gross evidence for toxicity in any organ examined. Compound III was injected intravenous as a bolus into two mice at a dose of 1.3 g/kg body weight. Both mice


survived with no evidence for any adverse effect of the polymer. The animals were sacrificed at 24 h, the organs were removed and visually examined. There was no evidence for gross pathology in the liver, lungs, heart, kidney, or intestine. These preliminary experiments confirmed that the polymers were very well tolerated in mice when administered by i.v. bolus. Upon the basis of this result, we decided to undertake biodistribution experiments at doses of 0.1 g/kg body weight. This is comparable to what has been used to study biodistribution of dextrans (44) and HPMA, polymers that have been used as drug carriers. Biodistribution Studies. Dendrimer I was radiolabeled with 125I, and the biodistribution of the radiolabeled polymer was monitored in CD-1 mice after intravenous administration. Polymer I was renally excreted and preferential accumulation in specific organs was not observed (Figure 6a). At 10 min postinjection, only 14% of the injected material remained in the serum, and the polymer was essentially completely excreted after 4 h. The low circulation time is most likely related to the compact shape of the branched dendritic structure and its modest molecular size. Biodistribution of model compound II was evaluated and even though the mass of this system was 3 times larger than I, the polymer was also rapidly excreted in the urine. At 30 min postinjection, only 4% of the administered dose remained in serum, and after 5 h the material was essentially completely excreted from the body with no evidence for accumulation in any tissue (Figure 6b). Model compound III was labeled with 125I as well, but in contrast to I and II the radioactive label was attached to a p-methoxyphenyl acetic acid moiety that was statistically coupled to the surface of III. This compound exhibited a highly significant accumulation in the liver. At 10 min postinjection, 70% of injected dose is in the liver and 5% in serum. After 5 h, 53% of the injected dose was still found in the liver (Figure 6c). This distribution behavior was not expected. One major difference between compound III and compounds I and II is that the labeled moiety was at the periphery of the PEO chains in III, as opposed to the well protected core in compounds I and II. This could possibly account for the liver uptake since it was previously observed that iodophenol containing PEGs would expose the methoxy iodophenol to the putative liver receptors whereas the iodophenol would be sterically hindered in compounds I and II (37). Another possibility is that the iodinecontaining moiety was hydrolyzed from the polymer and incorporated into the liver rather than the model compound itself. Given the rapidity of uptake in the liver, this is unlikely. Moreover, iodinated polymer recovered from serum eluted in the high molecular weight fraction on a G-50 spin column. Thus, we attributed the biodistribution profile of compound III to the exposed iodinated phenol. Despite the uptake of III into the liver, we decided to prepare a polymer-doxorubicin conjugate with compound III for further biodistribution studies. In addition to being a highly potent anticancer agent, doxorubicin is a fluorescent compound, providing a convenient analytical tool for monitoring distribution of the conjugate IV. In contrast to the radiolabeled III, the polymer-DOX conjugate IV showed no significant accumulation in any vital organ, including the liver, heart, and lungs (Figure 7). This is a significantly different distribution pattern than is observed for the free drug, which partitions into a variety of organs such as the liver


Figure 6. Biodistribution of model compounds (a) I, (b) II, and (c) III. The data are expressed in % of dose injected per gram of organ tissue. The blood is expressed in % of dose injected per mL, for ca. 2 mL of total blood volume in a CD-1 mouse.

and heart (45). Most importantly, conjugate IV had circulation half-life of 72 min, which is significantly longer than both the half-life of the free drug (ca. 8 min) and the carriers I and II. This demonstrates that the polymeric form of the carrier alters the pharmacokinetics and the distribution of the drug. Drug Release Studies. One of the important features to take into consideration when designing polymeric carriers is the means of attachment of the molecule to be delivered. The nature of the attachment sites will determine whether selective release of the bioactive molecule can be attained. A hydrazone linkage was selected due to its known stability under neutral pH conditions and its ability to undergo hydrolysis under acidic conditions. The goal is to have a linkage that will be sufficiently labile at a pH range of 4-6 to provide a therapeutic level of the drug. This is the pH found in the endosomal and lysosomal environments, which are the



Figure 7. Biodistribution of IV.

Figure 8. HPLC chromatogram showing increment in free drug concentration upon release from polymer conjugate IV. The detection was observed at 490 nm.

Figure 10. Comparative cell viability results upon treatment of (a) B16F10, (b) MDA-MB-231, and (c) MDA-MB-435 cell lines with bound DOX (IV) and free DOX. Results are compared in terms of DOX equivalents.

Figure 9. pH profile for release of doxorubicin from conjugate IV.

compartments where polymeric drugs are often internalized. The linkage we used and evaluated throughout this study is a hydrazone linkage. Drug release was monitored by an HPLC method. Figure 8 shows the well-resolved peaks corresponding to the conjugate and free doxorubicin at retention times of 5.4 and 6.3 min, respectively. As shown in Figure 9, the hydrazone linkage is indeed very stable at neutral pH, while it is cleaved more rapidly under more acidic conditions. Due to the highly reversible character of the hydrolysis of hydrazone linkages, an apparent equilibrium is observed under the experimental nonsink conditions of the release assay. One can estimate the time for complete release of DOX from the conjugate based on the computed initial rates of release; a 100% release would

be achieved after 10 min, 3 h, 26.5 h, and 10 days for pH 2.5, 4.5, 5.5, and 6.5, respectively. Thus, the hydrazone linkage provides a suitable linker for a pH-dependent release that is compatible with conditions found in tumors. Cytotoxicity of Drug-Polymer Conjugate in Vitro. The cell proliferation results suggest that these systems based on polyester dendrimers do not exhibit a significant toxic effect and the drug release rates from the hydrazone linker encouraged us to further evaluate model compound IV as a polymeric drug carrier. In vitro evaluation of the conjugate was performed in various cancer cell lines to compare the cytotoxic activity of the bound drug in relation to the free drug (Figure 10). The three cell lines examined exhibited a range of sensitivity to the free doxorubicin, from 0.025 µg/mL for B16F10 to 0.62 µg/mL for the MDA-MB-435 cell lines. In all three cell lines examined, the free drug was considerably more potent than the drug-polymer conjugate (Figure 10); 6-fold in the B16F10 cells, 50-fold in the MDA-MB-231, and 9-fold in the MDA-MB-435 cells. In the experiments, the conjugated doxorubicin was chromatographed in a gel permeation column within 1 day of testing the material so that there was less than



Figure 11. Fluorescence confocal microscopy images of localized DOX and DOX-polymer conjugate IV in MDA-MD-435 cells. MDAMD-435 cells were grown to 40% confluence on sterilized coverslips and incubated for 1 and 24 h.

1% free doxorubicin in the conjugate preparation. Moreover, the amount of polymer at the highest dose tested (100 µg/mL) was substantially below the concentration where antiproliferation effects of the polymer were observed. Therefore, the antiproliferation effects of the conjugate were most likely due to the drug that had been released from the polymer in the course of the incubation. These results are encouraging since the doxorubicin conjugate is less toxic than the free drug yet more toxic than the polymer alone. The question arises as to where the doxorubicin is being released from the polymer. At neutral pH in buffer, drug release is very slow. However, at low pH drug release is substantially faster. Therefore, we examined if the drug-polymer conjugate was internalized by cells. MDA-MB-431 cells were treated for 1 h and 24 h with the drug-polymer conjugate (IV) or the free doxorubicin. The fluorescence of doxorubicin provides a means for monitoring cell uptake of the drug using fluorescence confocal microscopy. Both the cytoplasm and the nuclei were highly fluorescent both at 1 h and 24 h after exposition to free doxorubicin. When cells were exposed to polymeric drug, fluorescence is observed in the cytosol, confirming the cell uptake of the conjugate by endocytosis (Figure 11). However, as also observed with the HPMA derivative of doxorubicin (46), little fluorescence was observed in the cell nucleus with the drug-polymer conjugate. This may be due to the slow release of the drug from the carrier at the time of evaluation. CONCLUSIONS

Various polyester dendritic carriers were evaluated as possible drug carriers. Each of the three drug carriers (I, II, III) are highly water soluble and nontoxic. Model compound III exhibited the longest circulatory half-life (72 min) as compared to the lower molecular weight systems I and II and was selected for further investigation. A doxorubicin-polymer conjugate was synthesized using III as the polymeric carrier (IV), and the drug was attached by means of an acid-labile hydrazone linkage. The cytotoxicity of the drug was significantly reduced (80-98%) after attaching it to the polymer. However, doxorubicin could be rapidly released from the polymer in an active form at pH values found in the lysosome. Uptake of IV by cells in culture was also observed by

fluorescence confocal microscopy. In biodistribution experiments, none of these systems accumulated in any organ examined, including the liver, heart, and lungs. The results suggest that this polyester dendritic backbone is a highly water soluble, nontoxic, and biocompatible polymer. However, a higher molecular weight system must be prepared to further increase the circulation halflife to effectively exploit the EPR phenomenon. This new system is a promising polymeric backbone for use as scaffolds in the development of well-defined macromolecular drug carriers. ACKNOWLEDGMENT

We thank Dr. Nikolay Vladimirov and Dr. Miroslav Janco (UCB) for their significant contribution in the development of the HPLC method for the performance of the drug release studies. We also thank John Andrew MacKay (UCSF) for his valuable assistance in the experiments involving confocal microscopy. Financial support of this research by UC Breast Cancer Research and Development Program (6JB-0137), the National Science Foundation (DMR-9816166), the Department of Energy (Basic Energy Sciences), and the National Institutes of Health (Grant GM 65361) is acknowledged with thanks. LITERATURE CITED (1) Langer, R. (1990) New Methods of Drug Delivery. Science 249, 1527-1533. (2) Duncan, R., Dimitrijevic, S., and Evagorou, E. G. (1996) The Role of Polymer Conjugates in the Diagnosis and Treatment of Cancer. S. T. P. Pharma Sciences 6, 237-263. (3) Maeda, H., Seymour, L. W., and Miyamoto, Y. (1992) Conjugates of Anticancer Agents and Polymers: Advantages of Macromolecular Therapeutics in Vivo. Bioconjugate Chem. 3, 351-362. (4) Mehvar, Reza (2000) Dextrans for Targeted and Sustained Delivery of Therapeutic and Imaging Agents. J. Control. Release 69, 1-25. (5) Matsumura, R. B., and Maeda, H. (1986) A New Concept for Macromolecular Therapeutics in Cancer Chemotherapy: Mechanism of Tumoritropic Accumulation of Proteins and the Antitumor Agent SMANCS. Cancer Res. 46, 6387-6392. (6) Ellens, H., Bentz, J., and Szoka, F. C., Jr. (1985) H+-Induced and Ca-2+-Induced Fusion and Destabilization of Liposomes. Biochemistry 24, 3099-3106. (7) Watkins, D. M., Sayed-Sweet, Y., Klimash, J. W., Turro, N. J., and Tomalia, D. A. (1997) Dendrimers with hydrophobic

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Polyester Dendritic Systems for Drug Delivery

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