Surface Eroding, Semicrystalline Polyanhydrides via Thiol–Ene “Click

Apr 13, 2015 - Katie L. Poetz,. † ... conditions, which may enhance healing rates, improve drug ..... since the addition of excess thiol may prematu...
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Surface Eroding, Semicrystalline Polyanhydrides via Thiol−Ene “Click” Photopolymerization Katie L. Poetz,† Halimatu S. Mohammed,† and Devon A. Shipp*,†,‡ †

Department of Chemistry and Biomolecular Science and ‡Center for Advanced Materials Processing, Clarkson University, Potsdam, New York 13699-5810, United States S Supporting Information *

ABSTRACT: Surface eroding and semicrystalline polyanhydrides, with tunable erosion times and drug delivery pharmacokinetics largely dictated by erosion, are produced easily with thiol−ene “click” polymerization. This strategy yields both linear and crosslinked network polyanhydrides that are readily and fully cured within minutes using photoinitiation, can contain up to 60% crystallinity, and have tensile moduli up to 25 MPa for the compositions studied. Since they readily undergo hydrolysis and exhibit the oft-preferred surface erosion mechanism, they may be particularly useful in drug delivery applications. The polyanhydrides were degraded under pseudophysiological conditions and cylindrical samples (10 mm diameter × 5 mm height) were completely degraded within ∼10 days, with the mass-time profile being linear for much of this time after a ∼24 h induction period. Drug release studies, using lidocaine as a model, showed pharmacokinetics that displayed a muted burst release in the early stages of erosion, but then a delayed release profile that is closely correlated to the erosion kinetics. Furthermore, cytotoxicity studies of the linear and cross-linked semicrystalline polyanhydrides, and degradation products, against fibroblast cells indicate that the materials have good cytocompatibility. Overall, cells treated with up to 2500 mg/L of the semicrystalline polyanhydrides and degradation products show >90% human dermal fibroblast adult (HDFa) cell viability indicative of good cytocompatibility.

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INTRODUCTION Many biodegradable polymers add significant value to drug delivery and tissue engineering systems because of their ability to degrade in a predictable manner under physiological conditions, which may enhance healing rates, improve drug efficacy or reduce the number of surgeries required for the patient.1 Common biodegradable polymers include polyesters, polyanhydrides, poly(ester anhydrides), and poly(ortho esters); of these, polyesters are the most widely used and commercially important.2,3 However, polyesters typically undergo bulk erosion (mass loss), which means hydrolytic degradation occurs throughout the sample because water is able to penetrate faster than hydrolysis (and hence erosion) occurs.4 This leads to significant changes to chemical and physical properties of the materials even at small mass loss and concentration spikes during drug release applications. The other idealized erosion mechanism is surface erosion, wherein water, and therefore, hydrolysis and erosion, is limited to the outer layers of the polymer due to hydrophobicity and hydrolytic instability of the labile bond in the polymer.5−8 Ideally, this leads to the conservation of thermomechanical properties, and relatively linear drug release rates that closely resemble the rate of erosion. Polyanhydrides are more hydrolytically unstable relative to polyesters and often undergo surface erosion.9−11 Poly(ortho esters) also exhibit surface erosion, but only under acidic conditions thus making them less desirable candidates for many physiological drug delivery systems. By comparison, polyanhydrides are a very promising © XXXX American Chemical Society

biodegradable polymer because they degrade over a range of physiological conditions and typically lose mass via the surface erosion mechanism, thus allowing for controlled release and conservation of thermomechanical properties throughout most of their degradation.5 Polyanhydride microparticles are another quite promising application for polyanhydrides as they have shown the potential to maintain protein and antigen stability.12−23 However, more widespread use of polyanhydrides has been severely limited in part because of their less than convenient synthesis and limited shelf life.24 Polyanhydrides are typically synthesized via condensation reactions, with the use of reagents such as phosgene or triphosgene, or under heat and high vacuum conditions to remove low molecular weight condensates.9,10,25−31 These relatively harsh conditions are not suitable to be conducted in vivo, significantly limiting their applications. One significant improvement in the synthesis of polyanhydrides was the use of methacrylated anhydride monomers; this allowed various radical initiation methods, such as photoinitiation, to be used and resulted in cross-linked polymers.32−39 These methacrylate-based polyanhydrides have glass transition temperatures above physiological temperature and mechanical properties that are comparable to cortical bone (cortical bone tensile modulus ∼ 17−20 GPa, polyReceived: March 1, 2015 Revised: April 9, 2015

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polymerization59,60 approach to polyanhydride synthesis, offers significant potential to the field of polyanhydride-based drug delivery systems by extending the degradation times and providing a wide range of tunable thermomechanical properties. Herein we show that semicrystalline linear and cross-linked polyanhydrides made quickly, efficiently, and in high yields, and their degradation may occur over weeks and gives products that are essentially nontoxic under biologically relevant concentrations.

(dimethacrylated(1,3-bis(p-carboxy phenoxy)hexane)) tensile modulus ∼ 15 GPa),34,40 yet they suffer similar problems to nondegradable methacrylate cross-linked systems in addition to sensitivity to oxygen during polymerization and a nonuniform cross-link density. Furthermore, the degradation product of these methacrylate-based polyanhydrides is poly(methacrylic acid), which can have relatively high molecular weight and cause high localized acidity; this is a result of the anhydride functional group residing in the side chain of the polymer.41−43 Additional advancements in polyanhydride-based materials have come about through combinations of esters and anhydride linkages within the polymer backbone. In doing so, Uhrich et al. showed that salicylic acid (aspirin) and other therapeutics can be included in the repeat unit of the polymer, thus, forming a polymeric pro-drug that releases the drug as one of the byproducts during degradation.44−46 With the use of commercially available monomers, we have recently used thiol−ene “click” photopolymerizations to synthesize completely amorphous, highly cross-linked polyanhydrides with uniform cross-link densities, reaching high conversions within minutes and without the need for the use of organic solvents (see Scheme 1).47−50 Degradation also yields

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EXPERIMENTAL SECTION

Materials and Methods. The following materials were purchased from Sigma-Aldrich and were used as received after being characterized by 1H NMR spectroscopy: 4-pentenoic anhydride (98%, PNA), 1,6-hexanedithiol (≥97%, HDT), pentaerythritol tetrakis(3-mercaptopropionate) (98%, PETMP), 1-hydroxycyclohexyl phenyl ketone (99%), 4-pentenoic acid (97%), lidocaine, anisole, acetonitrile, 2-propanol, and anhydrous chloroform. Human dermal fibroblast adult (HDFa) cells were obtained from Dr. Craig Woodworth, Clarkson University. MTT (3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide), RPMI 1640 medium, trysinEDTA (0.05%), phenol red, and fetal bovine serum (FBS) were purchased from Gibco Life Technologies. Instrumentation. 1H (400 MHz) and 13C (100 MHz) NMR spectroscopy was performed on Bruker Avance 400 with a BBO probe. The UV light source for polymer curing was an Oriel Instruments, model 68811, 500 W mercury xenon arc lamp (λ ∼ 365 nm, intensity ≈ 70 mW/cm2). Gel permeation chromatography (GPC) was performed on a modular system comprised of the following: a Waters 515 high-pressure liquid chromatographic pump operating at 30 °C, with tetrahydrofuran as the eluent, a Waters 717 Autosampler, two Polymer Laboratories columns (PLgel Mixed C), and a Viscotek LR40 refractometer. Molecular weights were determined from polystyrene standards (Polymer Standards Service molecular weight range from 1270 to 1230000). High performance liquid chromatography (HPLC) was performed with a Hitachi 7100 pump, Waters 717plus Autosampler, Hitachi L-7400 UV detector, at a flow rate of 0.5 mL/ min. The mobile phase used was 88% aqueous phase (HPLC grade, contained 2% acetic acid and 0.05% ammonium acetate) and 12% acetonitrile (HPLC grade). Differential scanning calorimetry (DSC) was performed on a TA Instruments Q100 series at a ramp rate of 10 °C/min. Mechanical analysis was performed on a TA Instruments Q800 dynamic mechanical analyzer (DMA). Tensile analysis was performed isothermally at 37 °C with a ramp force rate of 0.1 N/min until failure. Powder X-ray powder diffraction (PXRD) was performed on a Bruker D2 Phaser diffractometer equipped with a Cu sealed tube (λ = 1.54178 Å). The samples for PXRD were prepared on low background discs for analyses. All absorbances of MTT samples were recorded on Agilent 8453 spectrophotometer. All images were recorded on an Olympus optical microscope. Polymerization Kinetics. For kinetic analysis of the linear polyanhydrides, 1-hydroxycyclohexyl phenyl ketone (photoinitiator, 1.2 mg, 0.006 mmol, 0.1 wt %) was weighed into a scintillation vial. PNA (0.7 mL, 3.9 mmol) was transferred by a syringe into the vial, followed by HDT (0.59 mL, 3.9 mmol), anisole was used as an internal standard to monitor conversion (0.13 mL), and 1.3 mL of anhydrous chloroform. The reaction was irradiated by ultraviolet light and polymerized for 6 min, portions were removed every minute and monomer conversion was determined by 1H NMR and molecular weight via GPC. Polymer Synthesis. Linear Polyanhydrides. 1-Hydroxycyclohexyl phenyl ketone (1.6 mg, 0.008 mmol, 0.1 wt %) was weighed into a scintillation vial. PNA (1 mL, 5.47 mmol) was transferred by syringe into the vial, followed by HDT (0.67 mL, 4.38 mmol). The equivalent volume percentage of chloroform anhydrous was also added in order to reduce viscosity and avoid crystallization during polymerization. The reaction was irradiated by ultraviolet light and polymerized for 15 min. Upon completion of polymerization, if necessary, the polymer

Scheme 1. Synthesis and Degradation of Thiol−Ene “Click” Polyanhydrides

relatively low molecular weight degradation products. Thiol− ene “click” chemistry is attractive for the synthesis of polyanhydrides because of the fast reaction times and orthogonality to other chemistries and also allows for the potential ease of conjugation of biomolecules, polymers, surfaces, and the like.51−58 Our early work showed these highly cross-linked polymers could be used as relatively fast delivery systems where the polymer completely erodes within a day or two. Yet, many applications demand much longer timeframes, from weeks or even months. In the present work, we address this problem by developing thiol−ene polyanhydrides that erode over longer periods (weeks). The key to this significant increase in time frame is the formation of semicrystalline polyanhydrides via thiol−ene photopolymerization. This development, coupled with the facile thiol−ene “click” B

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Biomacromolecules Scheme 2. Chemicals Used for the Synthesis of Linear and Cross-Linked Polyanhydrides Using “Click” Thiol−Ene Photopolymerization, and an Illustration of Their Subsequent Erosion in Phosphate Buffered Saline (PBS) Solution

semicrystalline polyanhydride, degradation products, and synthesized model degradation compound were investigated against HDFa cells using MTT and PI exclusion assays. To prepare samples for studies, the polymeric material and synthesized degradation products were prepared according to the method outlined elsewhere in this paper. Prior to testing, all samples were sterilized under UV lamp for about 2 h just before coincubation with the HDFa cells. To create the extracted degradation products, the sterilized semicrystalline linear polyanhydride samples were completely degraded into the RPMI growth medium resulting in a homogeneous mixture. Cells were seeded in a 12-well tissue culture plate at a density of 6 × 104 per well in 2.0 mL of RPMI growth medium in the presence or absence of specified concentrations of test material. Four types of test materials were examined. These were the semicrystalline polyanhydride (both linear and cross-linked), the degradation product hydrolyzed in growth media and subsequently extracted, and the degradation product synthesized by thiol−ene addition of pentenoic acid to HDT (details of which can be found in the Supporting Information). Test material and HDFa cells were coincubated for 24 h. After incubation, the growth medium was carefully removed and cells washed with PBS buffer at 37 °C (pH = 7.4). MTT Assay To Assess Metabolic Activity. To assess the relative viability of cells, 1 mL of RPMI medium without phenol red containing 0.5 mM MTT solution was added to each well and incubated for 4 h. After 4 h, the unreacted MTT and medium were carefully removed. To quantify cell viability, the formazan crystals were dissolved in 800 μL of acidic 2-propanol (0.04 M HCl in 2-propanol) in each well. Absorbance of the dissolved formazan in each well was recorded at 570 nm. PI Exclusion Assay. Cells were cultured in the presence of 1500 and 3000 ppm linear polyanhydride. Control cells were not treated with the polyanhydride. After 48 h of incubation, both treated and untreated cells were washed with PBS solution containing 2 mM CaCl2. A total of 1 mL of the buffer supplemented with 3 μM propidium iodide to stain cells. These samples were then incubated for about 5 min at room temperature before taking images. Cells were imaged first using a light microscope and then using a fluorescence microscope. The recorded images were analyzed using the NIH image processing software, ImageJ. These results represent the qualitative cell viability measurements.

was diluted with additional anhydrous chloroform (1 mL) to reduce the viscosity of the polymer. The polymer was transferred to a roundbottom flask and the solvent removed by the use of rotary evaporation. The final trace amounts of solvent were removed by placing the polymer under vacuum for approximately 5 h. The polymer was stored in a freezer overnight, and the following day was heated above the crystalline melt temperature (approximately 80 °C) for 15 min and allowed to cool at room temperature. Following cooling, the polymer was characterized by DMA, DSC, and PXRD. For GPC analysis, the polymer was dissolved in chloroform and further diluted in tetrahydrofuran. Chain Extension Polymerization. For chain extension polymerization, 0.5 mL of prepolymer in chloroform was transferred to a separate vial containing photoinitiator and the appropriate amount of HDT to create a 1:1 mol ratio of enes to thiols. The polymerization was allowed to proceed for 15 min under UV irradiation. The chain extension polymer was analyzed by 1H NMR and GPC. Cross-Linked Polyanhydrides. Three different compositions of cross-linked polyanhydrides were studied. These are designated based on each components contribution to the overall functionality. For example, the composition of PNA/HDT/PETMP = 100:90:10 has 100% of the enes from PNA, 90% thiols from HDT, and 10% thiols from PETMP. In each case, there are stoichiometric amounts of ene and thiol functionality. Example of the PNA/HDT/PETMP = 100:70:30 material: 1-hydroxycyclohexyl phenyl ketone (1.4 mg, 0.0069 mmol, 0.1 wt %) was weighed into a scintillation vial. PETMP (0.22 mL, 0.57 mmol) was transferred by syringe into the vial, HDT (0.41 mL, 2.68 mmol), and PNA (0.7 mL, 3.83 mmol). The monomers and initiator were mixed to establish a homogeneous solution and transferred to a mold(s) for polymerization. Polymer samples for degradation were prepared in a cylindrical mold (10 mm diameter × 5 mm height). In some cases, lidocaine was added at the appropriate weight percentage relative to monomers was dissolved in PETMP and HDT, followed by the addition of the PNA monomer. The reaction was irradiated by ultraviolet light and polymerized for 15 min. Following curing, the polymer was allowed to sit for 30 min. Dynamic Mechanical Analysis. The polymer samples, with size for film tension was approximately 13 mm × 7 mm × 1.6 mm, were synthesized as described above and stored in the freezer for no more than 1 day until analysis was performed. Four replicates of each crosslinked polyanhydride composition were analyzed isothermally at 37 °C, at a ramp rate of 0.1 N/min until failure. Polyanhydride Degradation. Degradation of samples was commenced 30 min after synthesis by weighing the polymers and placing them into 100 mL of aqueous PBS solution at 37 °C. At select time intervals, the polymer was removed from the buffered solution, the surface quickly dried and polymer weighed (if the polymer was releasing lidocaine a portion of the buffer was saved for HPLC analysis) and new buffer solution added. Triplicates of each composition and lidocaine loading were degraded. Cytotoxicity Measurements. Cell Culture. To assess the cytotoxicity/biocompatibility of the linear (1:1 PNA/HDT) and cross-linked semicrystalline polyanhydrides (100:80:20, PNA/HDT/ PETMP), normal HDFa cells were cultured in RPMI 1640 medium supplemented with 10% FBS. The growth medium also contained 1% Pen Strep at 37 °C and cells were grown in a humidified atmosphere containing 5% CO2. In vitro cytotoxicity of the linear and cross-linked

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RESULTS AND DISCUSSION Linear Semicrystalline Polyanhydrides. In previous studies we showed that high efficiency “click” thiol−ene photopolymerizations were an excellent vehicle for the synthesis of surface eroding, cross-linked amorphous polyanhydrides.47−49 Here we utilize this facile polymerization method for the production of linear and cross-linked semicrystalline polyanhydrides through the inclusion of the aliphatic diene and dithiol monomers, as illustrated in Scheme 2, which undergo erosion in aqueous solutions and can act as a drug-release matrix. The utility of step-growth polymerization for making highly functionalized polymers is demonstrated through nonstoichiometric monomer functionality, which allows control over molecular weight and the synthesis of C

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Biomacromolecules prepolymers.61 Furthermore, the chain end functional groups can also be used for the conjugation of other materials, such as peptides, proteins, or drugs. The cross-linked semicrystalline polyanhydrides are produced through the inclusion of a tetrathiol monomer. To investigate polymerization rates and molecular weight development, particularly important parameters in terms of polymer synthesis and properties, the photoinitiated polymerization of an aliphatic dithiol monomer (HDT) and diene (PNA) was studied through monitoring monomer conversion and the evolution of molecular weight. Figure 1 shows the

of one functional group relative to another. For our prepolymer syntheses we chose to use an excess of the ene functionality since the addition of excess thiol may prematurely react with the anhydride functional group due to the nucleophilicity of thiols, as shown in our previous work where we found that the thiol functionality can add to the anhydride, creating a thioester, under ambient conditions over the course of several days.48 Here, we chose to use a 1.5:1 and a 1.25:1 stoichiometry of PNA/HDT. The synthesized prepolymers were characterized by GPC, 1H NMR, DSC, and PXRD. Data from these reactions are summarized in Table 1. In order to show that the prepolymers are capable of undergoing further polymerization, additional HDT was added to the alkene-functionalized prepolymers, enough to establish a stoichiometric equivalence of alkenes and thiols. The 1H NMR spectra show the alkene functionality is present in the two prepolymers (Supporting Information, Figure SI-1, 5.82 and 5.07 ppm), and after chain extension, the alkene end groups are consumed. Furthermore, there is no signal in the 1H NMR spectrum from excess dithiol monomer, indicating that the additional HDT was reacted with the ene end groups of the prepolymer. The GPC results (Figure 2a) show a low Mn for the initial prepolymers made with 1.5:1 and 1.25:1 PNA/HDT (Mn = 1640 and 1800, respectively), but after chain extension much higher Mn values were obtained (Mn = 6800 and 6600, respectively). These chain extension results indicate the ease and efficiency by which the alkene end groups on polyanhydrides can be further polymerized or functionalized, and forecasts their potential in various hybrid materials, such as functionalized copolymers (e.g., with cysteine-containing peptides or proteins, or thiol-containing synthetic polymers). The thermal properties of linear thiol−ene polyanhydrides with varying initial monomer ratios were studied using DSC (Figure 2b) and PXRD (Figure 2c). These were found to exhibit crystalline melt temperatures (Tm) in the range of 53− 60 °C. The crystallization temperatures (Tc) ranged between 35 and 39 °C. The 1.5:1 (ene:thiol) polyanhydride had a relatively broad Tm ranging between 53 and 58 °C and multiple melting peaks, but this is expected since it is quite a low molecular weight, and likely containing a variety of crystallites. The Tm of the “click” thiol−ene polyanhydrides are slightly higher than those reported for similar polymers (made by thiol−ene and acyclic diene metathesis chemistries),62 but lower compared to other polyanhydrides made by condensation polymerizations of diacids or their derivatives. For example, poly(sebacic acid) has a Tm of ∼75 °C.63 The glass transition temperatures (Tg) for each of these linear polymers are listed in Table 1, and each is below room temperature. The PXRD patterns of the off-stoichiometry linear polyanhydrides are comparable to the 1:1 stoichiometry polyanhydride (Figure 2c), and also exhibit similar diffraction patterns of polyanhydrides synthesized via polycondensation methods.63−65 By integrating the area underneath the crystalline regions, Ac, and the amorphous, Aa, we can

Figure 1. Number-average molecular weight (Mn) as a function of monomer conversion. Dispersity (Đ) values are next to Mn data points. Inset: monomer conversion as a function of reaction time.

monomer conversion of PNA, as determined by the loss of vinyl protons in 1H NMR spectroscopy, and the growth of number-average molecular weight (Mn) and dispersity (Đ) as a function of ene conversion. During the initial 120 s, approximately 20% of enes are reacted, but according to GPC there is no detectable molecular weight. At 180 s, the polymerization reaches ∼83% conversion and GPC shows polymer with Mn ≈ 1000 g/mol. is present. At 360 s, conversion of the ene monomer to polymer is ∼100% and the polymer has Mn ≈ 10600. These data indicate that the photopolymerization occurs quite fast (within minutes) and consumes all monomer, an important factor if these materials are considered for applications where in vivo curing is required. The development of molecular weight as a function of monomer conversion follows the expected trend for stepgrowth polymerizations, with low molecular weight oligomers formed at low and medium conversions, and only reaching relatively high molecular weights at high conversions. Step-growth polymers are often used as “prepolymers” that are subsequently cured in cross-linked materials that typically have significantly improved thermal and mechanical properties. Prepolymers with molecular weights normally in the few 1000 g/mol range are made through the introduction of molar excess

Table 1. GPC, DSC, and PXRD Data of Linear Prepolymers and Chain Extended Polyanhydrides

a

mole ratios (PNA/HDT)

Mna

Đa

Mnb

Đb

Tga (°C)

Tma (°C)

Tca (°C)

Xca (%)

1.5:1 1.25:1 1:1

1640 1800 N/A

1.75 1.89 N/A

6800 6400 10600

2.09 2.05 1.80

1 −2 11

53−58 61 64

35 39 36

65 64 66

Polyanhydride prepolymers. bChain extended polyanhydrides. D

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Figure 2. (a) GPC traces, (b) DSC traces, and (c) PXRD patterns of linear polyanhydrides synthesized with HDT and PNA at the HDT/PNA functional group ratios shown (1:1, 1:1.25, or 1:1.5).

determine an approximate percentage of crystallinity, Xc (eq 1). Each linear polyanhydride exhibits ∼65% Xc, which is similar to literature Xc values for polycondensation-derived polyanhydrides.63−65 Xc = Ac /(A a + Ac)

(1)

Cross-Linked Semicrystalline Polyanhydrides. Crosslinking polymers significantly alters their thermal and mechanical properties and is a primary factor in determining their suitability in many applications. Through a combination of HDT, PNA, and PETMP, the latter being a tetra-thiol and therefore a cross-linking entity, we synthesized three different compositions of cross-linked and semicrystalline polyanhydrides via thiol−ene photopolymerization. Details of the three different compositions investigated, each containing stoichiometric equivalents of alkene to thiol, can be found in Table 2. Table 2. Cross-Linked Polyanhydrides Thermal and Physical Properties mole ratios of enes and thiolsa

Tm (°C)

Tc (°C)

Tg (°C)

Xc (%)

modulusb (MPa)

100:90:10 100:80:20 100:70:30

57 52 46

21 27 −2

−24 −10 −14

66 67 65

25.6 (±2.8) 8.4 (±2.7) 3.0 (±0.4)

Figure 3. (a) DSC traces and (b) PXRD patterns of cross-linked polyanhydrides synthesized at the PNA/HDT/PETMP functional group ratios shown.

°C, similar to the amorphous cross-linked polyanhydrides described earlier.49 The PXRD diffraction patterns of each of these cross-linked polyanhydrides (Figure 3b) are slightly different than the linear polymers described above. The cross-linked polyanhydrides show broader diffraction patterns, indicating lower periodicity. This is expected since the PETMP will act as a spacer between the crystallites and reduces mobility due to cross-linking, thus interrupting the periodicity. Each of the cross-linked polymers still exhibit approximately 65% Xc, again similar to the linear polyanhydrides reported in the literature.63−65 In summary, the PXRD and DSC data both indicate that the crystalline regions from these cross-linked semicrystalline polyanhydrides, which are less periodic than the linear polymers discussed above, are embedded in a cross-linked rubbery phase. Table 2 also provides the Young’s modulus of the crosslinked semicrystalline polymers, which was investigated using dynamic mechanical analysis (DMA). The samples were

a

Mole ratios of enes and thiols from PNA/HDT/PETMP (2 enes per PNA, 2 thiols per HDT, 4 thiols per PETMP). bStandard deviations are given in parentheses.

DSC and PXRD data are shown in Figure 3. The crystalline melt temperature (Tm) for each of these polymers is similar to the linear polyanhydrides discussed above, and are similar to previously reported linear polyanhydrides and polyanhydrides that are lightly cross-linked.63−66 However, as PETMP content increased the Tm decreased since the tetrathiol dilutes the crystalline-forming HDT, and the peak intensity in the DSC data decreased and broadness increased (Figure 3a). No crystallinity was observed at higher PETMP contents. The Tg values for each of the cross-linked polyanhydrides is below −10 E

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Figure 4. Plots of % mass remaining (closed symbols) and % lidocaine released (open symbols) for various initial lidocaine loadings (0, 3, and 5 wt %) and various cross-linked polyanhydride compositions. (a) 100:70:30 PNA/HDT/PETMP, (b) 100:80:20 PNA/HDT/PETMP, (c) 100:90:10 PNA/HDT/PETMP.

evaluated isothermally at 37 °C to mimic physiological conditions, with a ramp rate of 0.1 N/min until failure. DMA curves can be found in Figure SI-2 (Supporting Information). The polymers containing the largest loading of HDT had the highest modulus (25.6 MPa), and as the amount of HDT decreased, the Young’s modulus decreased. The polyanhydride with the lowest HDT loading studied yielded a modulus of 3.0 MPa. Degradation and Drug Release Kinetics. The erosion and release profiles of the semicrystalline, cross-linked polyanhydrides were evaluated through mass-loss studies and the release of a model drug (lidocaine) from the polyanhydrides during erosion. The lidocaine was physically mixed with the monomers and dissolved before curing. After curing, the samples were placed into 100 mL of phosphate buffered saline (PBS) solution at 37 °C. Every 12 h, the polymer was removed from the buffer solution, the surface of the polymer quickly dried, the polymer weighed and placed into fresh 100 mL of buffer. The process was repeated until the polymer could not be removed without crumbling apart (it could no longer be weighed) or had completely eroded. For lidocaine release studies, a portion of the buffer was saved and analyzed via HPLC to quantify the lidocaine in the PBS solution. Figure 4 shows the erosion profiles (as % mass remaining) for three different polymer compositions, each loaded with 0, 3, or 5 wt % lidocaine. Regardless of composition and lidocaine loading, all polymers underwent an induction period of ∼24 h prior to erosion being evident, and then showed essentially linearly erosion profiles until ∼20−30% mass remaining, after which the rate of erosion slightly slowed. The entire erosion process lasted nearly 10 days, and most samples could still be

handled (i.e., removed from the solution, dried, and weighed) without disintegrating through to less than 10 wt % of their initial mass. There are some curious observations that can be made of the early stages of erosion. The initial 3.5−4 days of degradation, some small, thin flakes broke away from the outer layer of the samples, however the sample still maintained its original shape despite the loss of this material. More interestingly, during the first 24 h or so, the samples do not appear to lose mass; in fact, they typically gain a few percent. This gain in mass is easily explained as water-uptake and swelling. However, the water uptake is not enough to explain the induction period or, indeed, the relatively sharp change in the erosion profile at the ∼24 h mark. Thus, to further investigate the amount of water taken up by the samples, we degraded a series of 100:80:20 polyanhydrides for select times (12, 24, 48, 72, and 96 h), weighed the polymer, and placed it into a vacuum oven to remove the water. Once dry, the polymer was weighed, and the amount of water absorbed could be determined. We found that at 12 h there was only 2% water content, and after drying the polymer had lost only 3.5 wt % mass. Thus, the induction time is not due to water being absorbed at a similar rate as polymer mass loss. For polymers that were degraded for 24, 48, 72, and 96 h, each contained ∼16−18% water. With the water content being relatively constant after 24 h, the rate of mass loss is due primarily (if not exclusively) to polymer erosion. A combination of other factors may be relevant to these observations, including local pH changes (due to acid generation), effective water concentration, degradation product crystallinity and solubility, and molecular mobility/diffusion. We observed a similar induction period in our previously F

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Figure 5. In vitro cytotoxicity studies to test fibroblast viability against (a) semicrystalline linear polyanhydride cells and (b) in situ degraded and synthesized degradation products compared to untreated/control cells.

NAD(P)H-dependent cellular oxidoreductase enzymes. The formazan crystals can then be solubilized and absorbance of solution measured to quantitatively determine the potential toxicity levels using a simple UV/visible absorbance measurement at 570 nm. To quantify the cytocompatibility of the semicrystalline linear thiol−ene polyanhydrides and degradation products, different concentrations of the test material were evaluated by the MTT assay. For this polyanhydride, concentrations up to 20000 mg/L were measured against HDFa, a type of fibroblast cell which is one of the standard and commonly used cell types for cytotoxicity evalutions.69,70 Figure 5a shows the percentages of viable HDFa cells (viability profile) 24 h after exposure to the linear polyanhydride. Percent viability was calculated based on absorbance measurement at 570 nm. The positive control cells (untreated cells) exhibited an estimated 100.00 ± 4.5% viable cells with good culture plate surface adherence. Upon exposure to varying concentrations of the linear polyanhydride up to 2500 mg/L, HDFa cells still showed very high viabilities at 100 ± 6.0% and 98.8 ± 2.9% for 1000 and 2500 mg/L, respectively. These results are comparable to the control cells; therefore, under the tested conditions, this novel material has very good cytocompatibility, even at these high polymer concentrations. Furthermore, at polymer concentrations of 3000 and 3500 mg/L, cell viability remained above/ close to 70%, further confirming that the linear polyanhydride is nontoxic. These results suggest that the material may have good biocompatibility and is suitable for biomedical applications such as drug delivery purposes. In order for a material to be used in biomedical applications, not only does the polymer itself have to be nontoxic but the product of degradation also has to be nontoxic or have low toxicity levels. It was therefore necessary to investigate the toxicity of the degradation product of the linear polyanhydride. We investigated the cytotoxicity of the degradation product of the linear polyanhydride obtained by two methods; one was obtained by in situ degrading the polymer in RPMI growth medium (which provides the appropriate control for the polymer cytotoxicity experiment), while the second degradation product was synthesized through the thiol−ene radical addition of HDT and 4-pentenoic acid and purified separately (see Supporting Information). In comparison to the untreated or

reported amorphous polyanhydrides made by thiol−ene polymerizations, although in those materials the induction period took approximately the same length of time as the erosion period (i.e., ∼50% of the total time for erosion consisted of the induction period).47,49 Another thiol−ene polyanhydride study also showed such an induction period.66 We are currently examining this complicated system in more detail in order to fully explain the observations.67 The inclusion of 3 or 5% lidocaine in the samples increases the erosion rate slightly compared to the neat polymer samples, although there are no significant differences between the erosion rates of the 3 and 5% lidocaine samples. The increased erosion rate when lidocaine is incorporated in the polymer matrix is not unexpected, and is most likely due to a combination of decreasing crystallinity and cross-link density. Lidocaine release rates appear to generally follow the polymer erosion rates, indicative of erosion processes controlling the release of the drug, as opposed to other mechanisms that might operate such as diffusion. It is only at early stages do the release and erosion profiles differ, a commonly observed phenomena referred to as a “burst release”, although in the cases shown here this burst release is muted compared to other published data.68 In particular, some lidocaine is released (up to ∼30%) within the first ∼24 h, a period in which some water is absorbed and so the material slightly swells, as apparent from the small increase in mass during this time. It is probable that during this early period of polymer degradation the lidocaine is removed from the sample via dissolution and diffusion that is facilitated by water and not so much by erosion. This is in contrast to later in the polymer erosion process (>24 h), when the rate of mass loss is relatively constant and the release of lidocaine is likely dominated by mass loss. Total recovery of the lidocaine from each sample relative to what was loaded into the initial monomer mixture was quite high, ∼80% and above. In Vitro Cytocompatibility: Linear Polyanhydrides and Degradation Products. To be used in a drug delivery application, it is essential that the cytotoxicity levels are understood and this property is a key index in any biomedical application. We used the MTT assay to quantitatively estimate the relative cytocompatibility of the material. The MTT assay works based on the principle that, in living cells, yellow MTT is converted to an insoluble purple formazan by the action of G

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Biomacromolecules control cells, exposure of 500 mg/L for each type of the degradation product led to cell viabilities greater than 90% (Figure 5b). Overall, these results indicate that both the linear polyanhydride and its degradation products are nontoxic. One of the important parameters used in determining the in vitro toxicity levels of a material is the IC50, the concentration required to cause 50% growth inhibition (toxicity). To determine the IC50 of the semicrystalline linear polymer several more concentrations of the polymer were investigated. Figure 6

Figure 7. In vitro cytotoxicity of cross-linked semicrystalline polyanhydride (100:80:20 PNA/HDT/PETMP) against HDFa viability measured using MTT assay.

the morphology of cells after 24 h exposure to two concentrations of the cross-linked or linear semicrystalline polyanhydrides. Both control and treated cells appear to be well spread, elliptical with strong adherence to the culture surface. In comparison to the control cells (Figure 8a), the morphology of cells remained unchanged after the HDFa cells were exposed to 1500 and 3000 mg/L semicrystalline linear polyanhydride (Figures 8b,c). The viability of cells remained very high with very few PI staining (indicative of cell death) for both treated and untreated cells further confirming the MTT results in Figures 5 and 6. The morphology of HDFa also remained unchanged after treatment with 1500 and 3000 mg/L crosslinked semicrystalline polyanhydrides (Figure SI-5).

Figure 6. Dosage response curve of semicrystalline linear polyanhydride after 24 h exposure to HDFa cells and subsequent MTT assay.

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CONCLUSION In this work we demonstrate thiol−ene photopolymerization to be a facile method for the synthesis of semicrystalline polyanhydrides. This approach to polyanhydrides is a simple, straightforward route to linear prepolymers and cross-linked networks. These polyanhydrides possess many similar physical and thermal properties to previously studied semicrystalline polyanhydrides but have the added advantages of not needing harsh reaction conditions, long time scales for synthesis, and purification steps. The erosion profile of the cross-linked polymers has an induction period (equal to ∼10% of the total erosion time), after which the erosion rate is constant (i.e., mass loss is constant with time) for a large portion of the erosion process. We also show that drug release from the crosslinked polyanhydrides follows the erosion profile, and greater than 80% of the model drug lidocaine was recovered. The polymers exhibit low cytotoxicity, with HDFa cell viability greater than 90% at concentrations of polymer up to up 2500 mg/L. With the developments illustrated here, thiol−ene-based polyanhydrides now have a range of thermomechanical properties, tunable and predictable pharmacokinetics, and exhibit very low cytotoxicity. These highly desirable properties augur well for widespread use in drug delivery and tissue engineering applications, where they have potentially significant roles to play. Specific areas that could benefit from such polyanhydride-based biomaterials include erodible/resorbable adhesives, self-cleaning surfaces, adaptable and patternable surfaces, functionalized surfaces/particles, hierarchical structures, and scaffolds for cell growth.

shows the dosage response curve of the linear polyanhydride; the IC50 obtained for the linear polyanhydride is approximately 4700 mg/L. Similar plots were generated for the in situ degraded and synthesized degradation products (Figure SI-4 in Supporting Information). The IC50 values determined are approximately 3300 and 2900 mg/L for the in situ degraded and synthesized degradation products, respectively. Since higher IC50 values mean greater cytocompatibility (lower toxicity), our data indicate that the linear polyanhydride is slightly less toxic than the degradation products. The high IC50 values obtained for both the polymer and degradation products further confirm that, under the tested conditions, the polyanhydride and degradation product are nontoxic to HDFa cells. In Vitro Cytocompatibility: Cross-Linked Polyanhydrides. After investigating the toxicity levels of the linear semicrystalline polyanhydrides, we also studied the cytotoxicity of the cross-linked semicrystalline polyanhydrides with the 100:80:20 (PNA/HDT/PETMP) composition. Again, the MTT viability assay was used to determine the viability of HDFa cells after 24 h of coincubation with the cross-linked semicrystalline polyanhydride. As shown in Figure 7, in comparison to the control samples (untreated cells), cells treated with 250 and 2500 mg/L cross-linked semicrystalline polyanhydride (100:80:20 PNA/HDT/PETMP) showed greater than 90% viable HDFa cells, indicative of good cytocompatibility. In Vitro Cytocompatibility: Effect on HDFa Morphology. To further confirm the cytotoxic effect of the semicrystalline polyanhydrides on HDFa cells, we qualitatively investigated H

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Figure 8. In vitro cytotoxicity of linear polyanhydride against HDFa cells. Polyanhydride concentrations were (a) 0, (b) 1500, and (c) 3000 mg/L.

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ASSOCIATED CONTENT

S Supporting Information *

Experimental details, additional data, and spectra. This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Tel.: 1-315-268-2393. Fax 1315-268-6610. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We thank the Center for Advanced Materials Processing at Clarkson University, a New York State Center for Advanced Technology. We also thank Professors Mario Wriedt, Vladimir Privman, Damien Samways, and Craig Woodworth, along with Ms. Darpandeep Aulakh and Mr. Sergii Domanskyi for their assistance and helpful discussions.

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