Modulating mechanical and shape-memory properties while mitigating degradation-induced inflammation of polylactides by pendant aspirin incorporation

Abstract

Synergistically modulating mechanical properties and improving shape-memory performance while mitigating degradation-induced chronic inflammation of polylactide (PLA)-based implants for biomedical applications remains elusive. We test the hypothesis that copolymerizing aspirin-functionalized glycolide with D, L-lactide could enhance the thermal processing, toughness, and shape memory efficiency of the copolymer while mitigating local inflammatory responses upon its degradation. The content of pendant aspirin was readily modulated by monomer feeds during ring opening polymerization, and the copolymers with ~10% or less aspirin pendants exhibited gigapascal-tensile moduli at body temperature and significantly improved fracture toughness and energy dissipation that positively correlated with the aspirin pendant content. The copolymers also exhibited excellent thermal-healing and shape memory efficacy, achieving >97% temporary shape fixing ratio at room temperature and facile shape recovery at 50-65 °C. These drastic improvements were attributed to the dynamic hydrophobic aggregations among aspirin pendants that strengthen glassy-state physical entanglement of PLA while readily dissociate under stress/thermal activation. When subcutaneously implanted, the copolymers mitigated degradation-induced inflammation due to concomitant hydrolytic release of aspirin without suppressing early acute inflammatory responses. The incorporation of aspirin pendants in PLA represents a straightforward and innovative strategy to enhance the toughness, shape memory performance and in vivo safety of this important class of thermoplastics for biomedical applications.

Graphical Abstract

1. INTRODUCTION

Degradable shape-memory polymers (SMPs), capable of being fixed into a temporary shape for minimally invasive surgical delivery and triggered to revert to a preprogrammed permanent shape by external stimuli,¹ are attractive for applications as self-tightening sutures, smart tissue scaffolds and drug-releasing cardiovascular stents.^2-3 Polylactides (PLAs), particularly amorphous PLAs, due to their well established and tunable in vivo safety profiles as resorbable implants and drug delivery vehicles, are attractive candidates for engineering thermal responsive degradable SMPs.^4-5 Poor temporary shape fixing and slow/incomplete shape recovery around physiological temperature, along with the brittleness of pure amorphous PLAs, however, have attracted efforts to improve their shape memory performances and toughness by covalent crosslinking with urethanes,⁶ blending with other polymers,^7-8 or block copolymerization with polyethylene glycol or polyurethanes that are commonly utilized for shape memory programming with/without inorganic additives.^9-11 However, chemically crosslinked thermoset SMPs tend to be associated with higher manufacturing cost compared to their thermoplastic counterparts whereas existing copolymerization approaches often involve significant compositional changes (>50% non-PLA components) of the polymer.

Meanwhile, applications of PLA-based degradable polymeric scaffolds for guided tissue regeneration^12-13 have revealed challenges of the excessive acidic degradation products (e.g. lactic acid) that could cause significant inflammatory cells infiltrations and even local tissue resorption (osteolysis).¹⁴ Current approaches for modulating local immune responses to biomaterial implants¹⁵ include inhibiting protein or macrophage adhesion by engineering hydrophilic anti-fouling surfaces,¹⁶ neutralizing acidic degradation products by blending PLA with basic bioceramics such as hydroxyapatite,¹⁷ and systemic administration or local delivery of non-steroid anti-inflammatory drugs (NSAIDs) such as aspirin and ibuprofen.^18-21 The first two approaches involve significant changes in material properties of the implant, while the latter approach may exert an undesired inhibitory effect on the critical early stage of tissue repair as a result of suppressing the critical inflammation/healing cascade following acute injury.^22-23

Overall, a strategy that can simultaneously enhance the mechanical and shape-memory performances of PLAs while selectively mitigating degradation-induced local inflammatory responses without perturbing the critical early-stage inflammation/tissue healing cascade remains elusive. Here, we explore the novel strategy of copolymerizing a small percentage (~10% or less) of aspirin-functionalized glycolide, (S)-(3,6-dioxo-1,4-dioxan-2-yl)methyl 2-acetoxybenzoate (AspGA), with D, L-lactide to enhance the thermal processing, toughness, shape memory efficiency of the copolymer while mitigating local inflammatory responses upon its degradation. Specifically, we hypothesize that the hydrophobic interactions among the pendant aspirin units could lead to the formation of loosely packed domains with relatively low binding energy to dynamically strengthen the glassy-state physical entanglement of the amorphous PLA chains, facilitating shape memory programming and improving the viscoelasticity and toughness of the resulting copolymer. Meanwhile, we hypothesize that the hydrolytic degradation of the copolymer would result in simultaneous hydrolysis of the aspirin pendants to release salicylic acid (active metabolite of NSAID aspirin; Scheme S1), thereby timely mitigating the inflammatory response elicited by the immunogenic acidic degradation products of PLA while avoiding premature drug release.

To test these hypotheses, we synthesized the novel monomer AspGA, and copolymerized it with D,L-lactides in readily altered feed ratios by ring opening polymerization (ROP). The resulting random copolymers and the PLA control were characterized with thermal, thermomechanical, rheological, and energy dissipative properties, and examined for thermal-healing and shape memory behaviors. By copolymerizing ~10% or less AspGA, significant improvements in glassy state tensile modulus, fracture toughness and viscoelasticity, as well as excellent thermal-healing, effective temporary shape fixing at room temperature and facile shape recovery at safe triggering temperatures were accomplished. The hydrolytic degradation kinetics of these polymers and the cytocompatibility of their degradation products were examined, revealing no significant impact on the in vitro proliferation and osteogenesis of rat bone marrow derived stromal cells (BMSCs). Using a rat subcutaneous implantation model, local immunological responses to the copolymers and PLA control prior to and during the scaffold degradation were investigated over 16 weeks. High-performance liquid chromatography and histological analysis of the explants were carried out to correlate the extent of scaffold degradation and drug release with the severity of local immune responses as a function of aspirin pendant content, supporting the mitigation of degradation-induced inflammation by the released salicylic acid.

2. EXPERIMENTAL SECTION

2.1. Materials and general instrumentations

All chemicals were purchased from Sigma-Aldrich (St. Louis, MO) or Fisher Scientific (Pittsburgh, PA) and used as received unless otherwise specified. O-Benzyl-L-serine (99.8%) was purchased from Chem-Inpex International (Wood Dale, IL). D,L-Lactide was recrystallized twice with dry toluene prior to use.

Nuclear magnetic resonance (NMR) spectra were recorded on a Varian INOVA-400 spectrometer or a Bruker 500 MHz instrument. All small molecule compounds were visualized with 0.2 wt% orcinol in 2N H₂SO4 for thin layer chromatography monitoring. Molecular weights and polydispersity of polymers were determined by gel permeation chromatography (GPC) with two 5-mm PLGel MiniMIX-D columns and a PL-ELS2100 evaporative light scattering detector and calibrated using polystyrene standards (Polymer Laboratories). THF was used as the eluent at a flow rate of 0.3 mL min⁻¹. Fourier-transform infrared (FTIR) spectra were recorded on a Nicolet IR 100 spectrometer (Thermo Electron Corporation) with 2 cm⁻¹ spectral resolution. The solid specimens were compressed into transparent discs with KBr.

2.2. Monomer and polymer syntheses and fabrication

Bromoacetic anhydride

Bromoacetic anhydride was synthesized per a previously reported literature.²⁴ In short, to a cold solution of bromoacetic acid (5.4g, 38.6 mmol) in 10 mL of CH₂Cl₂ was added 1,3-diisopropylcarbodiimide (2.4 g, 19.3 mmol) diluted in 5 mL of CH₂Cl₂ slowly. Thereafter, the reaction mixture was stirred at room temperature for 1 h. The solution was filtered to remove the urea solid. Evaporation of the solvent from the filtrate provided bromoacetic anhydride (4.8 g, 90%) as yellow oil without further purification.

(S)-3-(benzyloxy)-2-hydroxypropanoic acid (2)

Hydroxy acid derivative 2 was synthesized from O-benzyl-L-serine (1) per literature protocols.²⁵ Briefly, compound 1 (20 g, 102 mmol) was added to an aqueous solution of trifluoroacetic acid (0.7 M, 293 mL) and stirred in an ice-bath until the white powder was totally dissolved. An aqueous solution of NaNO₂ (14.5 g, 210 mmol, in 50 mL of H₂O) was then added slowly and the resulting yellowish solution was stirred overnight before NaCl (20 g) was added. The mixture was extracted with ethyl ether three times, and the combined organic phase was washed with brine before being dried over anhydrous Na₂SO₄. After evaporating the solvent, spectroscopically pure yellowish liquid (20.10 g, 100%) was obtained. ¹H NMR spectrum for 2 agreed with literature.²⁵

(S)-3-(benzyloxy)-2-(2-bromoacetoxy)propanoic acid (3)

Bromoacetic anhydride (2.9 g, 11.2 mmol) diluted in CH₂Cl₂ (5 mL) was slowly added to a cold solution of α-hydroxy acid (2.2 g, 11.2 mmol) in CH₂Cl₂ (20 mL). The mixture was then refluxed overnight. After cooling to room temperature, the mixture was diluted with CH₂Cl₂, washed with aqueous solutions of NaHCO₃ and brine. The organic layer was collected and purified by silica gel column chromatography with an eluent of hexane/ethyl acetate (2:1, v/v) to afford the product as a yellowish oil (3.2 g, 90%). ¹H NMR (500 MHz, Chloroform-d) δ 7.39 – 7.27 (m, 5H), 5.34 (dd, J = 5.2, 2.8 Hz, 1H), 4.62 (q, J = 12.2 Hz, 2H), 4.05 – 3.90 (m, 3H), 3.86 (dd, J = 11.1, 2.8 Hz, 1H). 13C NMR (125 MHz, Chloroform-d) δ 171.76, 166.65, 137.06, 128.53, 128.50, 128.02, 127.82, 73.63, 73.08, 68.17, 25.19 ppm. ¹H NMR spectrum for 3 agreed with literature.²⁶

(S)-3-((benzyloxy)methyl)-1,4-dioxane-2,5-dione (4)

Compound 3 (1.7 g, 5.4 mmol) dissolved in acetonitrile (20 mL) was added slowly to the acetonitrile solution containing triethylamine (TEA, 1.5 mL, 10.7 mmol) under the argon protection. The reaction mixture was stirred at 65 °C for 6 h. After cooling, the solvent was removed under reduced pressure, and the obtained crude product was purified by silica gel column chromatography with an eluent of hexane/ethyl acetate (2:1, v/v) to afford the product as a yellowish oil (1.1 g, 86%). ¹H NMR (500 MHz, Chloroform-d) δ 7.41 – 7.26 (m, 5H), 5.11 – 5.01 (m, 2H), 4.81 (d, J = 16.7 Hz, 1H), 4.58 (s, 2H), 4.08 (dd, J = 10.5, 2.0 Hz, 1H), 3.96 (dd, J = 10.5, 2.6 Hz, 1H) ppm. The ¹H NMR spectrum for 4 agreed with literature.²⁷

(S)-3-(hydroxymethyl)-1,4-dioxane-2,5-dione (5)

To a solution of compound 4 (900 mg, 3.81 mmol) in EtOAc (60 mL) was added Pd(OH)₂/C (30 wt%, 270 mg), and the mixture was subjected to cycles of purging to replace air with hydrogen. After being stirred under a hydrogen balloon for 2 h, the mixture was filtered through a short pad of silica gel to yield desired compound 5 (557 mg, 92.0%) as a colorless liquid. ¹H NMR (400 MHz, d6-DMSO) δ 5.72 (t, J = 4.8 Hz, 1H), 5.23 (t, J = 3.2 Hz, 1H), 5.05 (m, 2H), 3.85 (m, 2H) ppm. ¹H NMR spectrum for 5 agreed with literature.²⁷

(S)-(3,6-dioxo-1,4-dioxan-2-yl)methyl 2-acetoxybenzoate (AspGA)

A solution of compound 5 (100 mg, 0.69 mmol) in EtOAc was stirred in an ice-bath for 10 min, before Et₃N (0.19 mL, 1.37 mmol) and DMAP (17 mg, 0.14 mmol) were added. Five minutes later, a solution of acetylsalicyloyl chloride (272 mg, 1.37 mmol) in EtOAc was added dropwise. The white suspension was stirred at room temperature for another 30 min before it was quenched with a cooled 5% NH₄Cl aqueous solution. The organic phase was washed with brine and dried over anhydrous Na₂SO₄. After evaporating the solvent, the resulting residue was purified by silica gel chromatography (n-hexane: EtOAc = 1: 1) to give a colorless liquid, which was further purified by recrystallization with Et₂O twice to afford compound AspGA as a white solid (33.3%). ¹H NMR (400 MHz, CDCl3) δ 7.91 (dd, J = 1.6, 8.0 Hz, 1H), 7.59 (m, 1H), 7.32 (m, 1H), 7.11 (dd, J = 1.2, 8.0 Hz, 1H), 5.27 (dd, J = 2.8, 4.8 Hz, 1H), 4.91 (d, J = 2.8 Hz, 2H), 4.85 (dd, J = 2.8, 12.8 Hz, 1H), 4.67 (dd, J = 4.8, 12.8 Hz, 1H), 2.33 (s, 3H) ppm; ¹³C NMR (100 MHz, CDCl3) δ 169.99, 163.84, 163.42, 163.22, 150.78, 134.83, 131.71, 126.51, 124.21, 122.54, 74.07, 65.61, 63.44, 21.19 ppm.

Poly(AspGA-co-LA) and PLA

The melt ROP was conducted with different monomer feeding ratios to prepare the random copolymers and PLA control (Table S1). In a typical procedure, the recrystallized AspGA (60 mg, 0.19 mmol) and D,L-lactide (908 mg, 6.30 mmol) were weighed into a dry Schlenk flask (10 mL), and the mixture was purged with argon three times. Benzyl alcohol (0.68 mg, 0.0063 mmol) and catalyst Sn(Oct)₂ (2.55 mg, 0.0063 mmol) in dry toluene stock solutions were then introduced via syringe injection. The reaction mixture was heated at 145 °C for 30 min. After cooling, the solidified mixture was dissolved in chloroform, precipitated in ice-cold methanol, and dried under vacuum to give the desired polymer (800 mg, 82.6%). ¹H NMR detection for residue D,L-lactide in crude product indicated a monomer conversion rate over 85%. GPC characterizations of the Mn and PDI of these polymers are summarized in Table S1. The final percentages of AspGA incorporation were calculated from ¹H NMR integrations based on the following equation: %(AspGA) = (ʃ 2.324 )/(ʃ 2.324+(ʃ 1.56)/2)

Fabrication of bulk polymer specimens

The bulk polymer specimens were prepared by hot pressing the as-synthesized polymer in a custom-made Teflon mold with a hydraulic press (CARVER® series NE, Wabash, IN). Briefly, a Teflon thin film (0.008″ thick) was folded seven times and pressed tightly before being carved or punched to create rectangular, dumbbell-shaped or cylindrical voids of specific dimensions. The as-prepared polymers, cut into small pieces, were placed into the voids, and hot-pressed under a constant force of 2000 lbs at 120 °C for 5 min to fabricate specimens for various mechanical tests or in vivo implantations. ¹H NMR and GPC analysis confirmed that there was no significant change in drug content or molecular weight of the polymers after the hot pressing. For in vivo study, the obtained circular pellets (30.25 ± 0.71 mg, 6.30 mm in diameter, 1.16 mm in thickness) were sterilized with 70% ethanol followed by UV irradiation (254 nm, 1 h each side) prior to use.

2.3. In vitro characterizations

Water contact angle measurements

Water contact angles on polymer films were measured with a CAM2000 goniometer (KSV Instruments) by the dynamic sessile drop method. Polymer thin films for water contact angle measurements were prepared by evaporation of 5% (w/w) chloroform solutions (0.1 mL) applied onto a clean glass slide at room temperature. After evaporating the first application of the chloroform solution for 10 min, another 0.1 mL of the polymer solution was applied and evaporated to achieve a thicker and more uniform coating. The water contact angles on the polymer thin films were measured at 25 °C. A droplet (2 μL) of Milli-Q water was placed on the respective polymer thin films, and the contact angles of the droplet from left and right were recorded after 30 s. The reported contact angles were averaged from a total of six randomly selected regions (N=6) for each polymer.

Differential scanning calorimetry (DSC)

DSC experiments were conducted on a Q200 MDSC (TA Instruments), which was calibrated with standard indium, at a heating rate of 10 °C min⁻¹ with a polymer specimen load of around 4.0 mg.

Thermal mechanical properties

The dynamic mechanical properties of the poly(AspGA-co-LA) as a function of AspGA incorporation content and temperature were determined on a Q800 Dynamic Mechanical Analyzer (TA Instruments) equipped with a gas cooling accessory and tensile film clamps. Specimens with dimensions of 10.0 mm × 3.5 mm × 0.3 mm were used for testing. The temperature was ramped from 25 °C to 100 °C at a heating rate of 5.0 °C min⁻¹. A 0.1% strain amplitude and 1.0 Hz frequency were applied. Three specimens (N=3) were tested for each polymer composition.

Rheology

Complex viscosity and storage and loss moduli of the polymer formulations (~20 mg) as a function of temperature sweep from 25 to 90 °C were determined by oscillatory rheology using an AR2000 EX rheometer equipped with 8 mm diameter parallel plates. The temperature scan rate was set to 5 °C min⁻¹, and the frequency and strain was set to 1.0 rad s⁻¹ and 1%, respectively.

Tensile test-to-failure test

Tensile test-to-failure tests of poly(AspGA-co-LA) as a function of AspGA incorporation content were carried out at room temperature on an MTS Bionix 370 servo-hydraulic mechanical tester equipped with a 250-N load cell. The test specimens, hot-pressed into dumbbell bones shape with a dimension of 28 mm in length, 3 mm in width and 1 mm in thickness, were mounted and strained to failure at 10 mm min⁻¹. The tensile modulus was calculated as the slope in the linear region (0-5%) of the stress-strain curve recorded. The fracture toughness was calculated by the area below the stress-strain curve.

Energy dissipation properties and stress relaxation measurements

Consecutive cycles (5) of tensile loading and unloading of poly(AspGA-co-LA) as a function of AspGA incorporation content were performed on a Q800 Dynamic Mechanical Analyzer at 25 °C, with the tensile force ramped from 0 to 18 N and back to 0 at a rate of 1 N min⁻¹, respectively. The dissipated energy was calculated as the area within the hysteresis loop of each tensile loading-unloading cycle.

The stress relaxation time (τ _1/2) of poly(AspGA-co-LA) as a function of AspGA incorporation content (5.9% and 10.5%) and temperatures (40 °C, 45 °C, 50 °C and 55 °C) were determined on the Q800 Dynamic Mechanical Analyzer under tensile stress relaxation mode. All specimens were applied with a 1% constant strain and 1 Hz constant frequency while the load was recorded over time. The stress relaxation activation energy (E_a) is calculated by the Arrhenius equation τ(T) = τ₀ exp(E_a/RT), where τ(T) is the stress relaxation time at temperature T (K), τ₀ is a constant, and R is the ideal gas constant. τ(T) is deduced from the equation with reference to the respective relaxation curves per σ = σ₀ exp(−t/τ), where σ₀ is the initial stress, σ is the stress measured at time t, and the relaxation time τ is determined as the time when σ₀ = 1/e.

Thermal healing property demonstration

To demonstrate the thermal healing capability, a 10.5% poly(AspGA-co-LA) specimen (2 mm in length, 2 mm in width, and 1 mm in thickness) was cut into two separate pieces using a razor blade. Subsequently, the separate pieces were brought back into contact and allowed to thermally heal at 70 °C for 30 min. The thermal healing at the broken and realigned interface was monitored by optical microscopy on a Leica M50 stereomicroscope equipped with a Leica DFC295 digital camera at 0, 15 and 30 min.

Shape memory properties

One-way shape memory cycles were run on a Q800 Dynamic Mechanical Analyzer equipped with a gas cooling accessory and tensile film clamps. Specimens (10.0 mm × 3.5 mm × 0.3 mm) were first equilibrated at 65 °C for 10 min and then cooled down to 25 °C prior to testing. A preload force of 1 mN was applied to each specimen. Temperature was first raised to 65 °C (5 °C min⁻¹) and kept isothermal for 5 min. The specimen was deformed at 65 °C at a stress ramping rate of 0.02 MPa min⁻¹ from its “permanent” shape at the beginning of the Nth testing cycle, ε_p(N − 1), to the elongated shape under a final tensile stress of 0.16 MPa. The temperature was then cooled to 25 °C with the stress kept constant at 0.16 MPa, and the sample length was recorded as ε₁ (N) (the strained sample length at the lower temperature at the N^th cycle). After being held at 25 °C for 5 min, the applied stress was released to the 1-mN preload force, and the sample length was recorded as ε_u(N) (the sample length at the lower temperature after unloading the tensile stress at the N^th cycle). Finally, the temperature was ramped from 25 °C to 65 °C at a heating rate of 5.0 °C min⁻¹ and kept isothermal for 5 min, and the final sample length was recorded as ε_p(N) [the recovered length at the N^th cycle or permanent length at the beginning of the (N + 1)^th cycle]. The strain fixing ratio (R_f) and the strain recovery ratio (R_r) in a given cycle N were determined by using the following equations:

$${\mathrm{R}}_{\mathrm{f}}(\mathrm{N})=({\epsilon }_{\mathrm{u}}(\mathrm{N})\text{-}{\epsilon }_{\mathrm{p}}(\mathrm{N}\text{-}1))∕({\epsilon }_1(\mathrm{N})\text{-}{\epsilon }_{\mathrm{p}}(\mathrm{N}\text{-}1))$$

$${\mathrm{R}}_{\mathrm{r}}(\mathrm{N})=({\epsilon }_{\mathrm{u}}(\mathrm{N})\text{-}{\epsilon }_{\mathrm{p}}(\mathrm{N}))∕({\epsilon }_{\mathrm{u}}(\mathrm{N})\text{-}{\epsilon }_{\mathrm{p}}(\mathrm{N}\text{-}1))$$

Accelerated in vitro hydrolytic degradation of the polymers

All specimens (N = 3) were incubated in PBS (pH = 7.4, 10 mg mL⁻¹) at 70 °C and retrieved at 1.5, 5, 20, 46 and 100 h. The retrieved specimens at each timepoint were washed with MilliQ water, freeze-dried, and weighed before being returned to an equal volume of fresh PBS for continued incubation at 70 °C until the next scheduled timepoint. The averaged percentage (%) of mass residue, defined as the dry mass at a given timepoint over the original dry weight of the specimen at time 0, was plotted over time for each polymer composition.

Cytocompatibility evaluations of the hydrolytic degradation products of the polymers

To determine the cytocompatibility of the degradation products of the respective polymers, each specimen was incubated in PBS (0.5 mg mL⁻¹) at 70 °C with frequent vortexing for 7 days, when no more bulk polymer was visible. The degradation solutions were sterile-filtered through 0.22 μm Teflon membrane before being supplemented to the expansion and osteogenic differentiation cultures of rat bone marrow derived stromal cells (BMSCs) at 12.5 μg mL⁻¹ to determine their impact on the cell proliferation and osteogenesis. Passage 1 BMSCs, isolated from the tibia and femur of skeletally mature SASCO SD rats (male, 8-12 weeks) and enriched by plastic adherent culture, were used for these experiments.

For cell proliferation, BMSCs were plated at 2300 cells cm⁻² and cultured in expansion media (αMEM without ascorbic acid, supplemented with 20% Hyclone characterized fetal bovine serum and 1% Pen-Strep). The filtered polymer degradation solutions were supplemented at a final concentration of 12.5 μg mL⁻¹ every 2 days at the time of media change. Viability of cells at 1, 3, and 5 days were determined by CCK-8 assay (Dojindo, Japan), respectively. The absorbance was read at 450 nm (with background subtraction at 620 nm) on a Multiskan FC Microplate Photometer (Thermo Scientific, Billerica, MA).

For osteogenic differentiation, BMSCs was seeded at 25000 cm⁻² and cultured in expansion media for 2 days (>70% confluency) before being switched to osteogenic media (expansion media supplemented with 10 nM dexamethasone, 20 mM β-glycerol phosphate, and 50 μM l-ascorbic acid 2-phosphate) along with of the polymer degradation products (12.5 μg mL⁻¹). Media were changed 3 times a week for 2 weeks with the supplementation of the same concentrations of degradation products. After 2 weeks of osteogenic induction, cells were fixed with 10% formalin and stained with Alizarin red S for microscopic photo-documentation (100× magnification).

2.4. In vivo studies

Subcutaneous implantation and explant retrieval

All animal surgeries were carried out according to the procedures approved by the University of Massachusetts Medical School Animal Care and Use Committee. Male SASCO-SD rats (289-300 g, Charles River Laboratories) were anesthetized with 5% isoflurane/oxygen and maintained under 2% isoflurane/oxygen throughout the surgery via a vaporizer. The rat abdomen was shaved, sprayed with 70% ethanol and wiped with Povidone-iodine. Following a small abdominal skin incision, a subcutaneous pocket was created by blunt dissection. The circular implant pellet (~30 mg each) prepared from each of the three polymers shown in Table S1 was fit into each pocket and the incision was closed with wound clips. A total of seven specimens were implanted in a randomized fashion in each rat. Kefzol (20 mgkg⁻¹) and buprenorphine (0.08 mg kg⁻¹) were given immediately post-op. The rats were euthanized at predetermined time points and the fibrous tissue-encapsulated implants were retrieved.²⁸ The explants were fixed in periodate-lysine-paraformaldehyde (PLP) fixative at 4 °C for 1 day and rinsed with PBS. In a pilot study, explants were retrieved at 2-, 8-, 12- and 16-week post-implantation for gross inspection of the degree of degradation and histological analysis, which revealed a rapid increase of inflammatory cells after 12 weeks as a result of implant degradation. For detailed histological analysis of immune cell infiltrations, a sample size of 3 was applied to the 16-week explants. In addition, one specimen from each group was retrieved at 8 weeks for SEM and at 12 weeks for HPLC detection of the drug release, respectively.

High-performance liquid chromatography (HPLC)

The in vivo degradation products of the 10.5% poly(AspGA-co-LA) within the explants were analyzed by reverse phase HPLC equipped with a Microsorb-MV 100-5 C18 250 x 4.6mm column. Briefly, the 12-week explant was extracted with methanol for 2 days. The extract solution was filtered through a 0.2-μm polytetrafluoroethylene (PTFE) filter before being injected (20 μL) for HPLC. A mobile phase of methanol/1% aq. acetic acid (35/ 65) was applied at a flow rate of 0.7 mL min⁻¹ and the signal was detected by UV at 228 nm. The obtained trace was compared with those of the aspirin and salicylic acid standard obtained under the same condition. The 12-week PLA explant was also analyzed to account for any background signals at the spectral regions of interest.

Histological analyses

Fixed explants were dehydrated in graded ethanol series and subjected to glycol methacrylate embedding and sectioning. Note that paraffin sectioning is not compatible for these polymers as they would dissolve during the deparaffinization. The 5-μm sections were stained with hematoxylin and eosin (H&E) for immune cell quantifications. The presence of macrophages, foreign body giant cells (FBGCs), mast cells, neutrophils, lymphocytes and eosinophils within the fibrous tissue capsules surrounding each implant was examined by a pathologist in a blinded fashion. Five randomly selected areas of the sections stained by H&E were used for quantification at 400× magnification.

2.5. Statistics

All quantitative data are plotted or presented in tables as mean ± standard derivation. Student’s t-test was employed for pairwise comparisons in Figure 2D inset and one-way ANOVA Tukey’s multiple comparison was carried out for Figure 4D and Figure S2. Significance level was set as p < 0.05.

Figure 2.

(A) Ultimate tensile stress/strain of 5.9% and 10.5% poly(AspGA-co-LA) determined on an MTS Bionix 370; (B)–(C) Consecutive (5) tensile loading/unloading of 5.9% and 10.5% poly(AspGA-co-LA) specimens on a Q800 DMA under force control (0 to 18 N to 0; 1 N min⁻¹); (D) Representative stress relaxation profiles of 5.9% and 10.5% poly(AspGA-co-LA) at 25 °C. Inset: stress relaxation time (τ_1/2), N=3, **P < 0.01 (Student’s t text). (E)–(F) Stress relaxation of 5.9% and 10.5 % poly(AspGA-co-LA) at different temperatures. Insets: respective activation energy determinations.

Figure 4.

In vivo degradation and drug release of polymers upon rat subcutaneous implantation. (A) Representative macroscopic images of each pellet (30 mg) before (left) and after (right) 16-week implantation. Scale bar = 3 mm. (B) HPLC chromatographs (UV detector, 228 nm) of aspirin standard, salicylic acid standard and extract of explanted 10.5% poly(AspGA-co-LA) after 12-week implantation. (C) Representative H&E staining of 16-week explants at 200× and 500× magnifications (boxed areas of the top panels), with the implant remnant positioned on the right side of each optical micrograph. Scale bar = 20 μm. BV = blood vessel; F = fibroblast; FBGC = foreign body giant cell; MΦ = macrophage. (D) Quantification (n=3) of total numbers of FBGCs in five combined random fields of view at 400× magnification. * p <0.05 vs. PLA control (one-way ANOVA Tukey’s multiple comparison).

3. RESULTS

3.1. Monomer synthesis, copolymer preparation and characterizations

To accomplish a delayed release of anti-inflammatory drugs from PLA-based polymeric scaffolds and introduce hydrophobic pendants to facilitate their dynamic aggregation, we chose to prepare a random copolymer bearing aromatic aspirin pendants through ROP of AspGA with D,L-lactide. To prepare the functional monomer AspGA (Scheme 1), lactone intermediate 4 was first prepared by acylation of an α-hydroxy acid (compound 2) and subsequent ring closure following the literature procedure.^29-30 The overall isolated yield of lactone 4 from O-benzyl-L-serine (compound 1) was almost 80% (20 g scale). The removal of the benzyl protection group was achieved by catalytic hydrogenation (Pd(OH)₂/C, 30% w/w in ethyl acetate) in 2 h, affording 5 in quantitative yield. The freed hydroxy group in 5 was used for subsequent conjugation with aspirin via esterification with catalytic amounts of 4-dimethylaminopyridine (DMAP) and stoichiometric acyl chloride. Functional monomer AspGA was obtained as white solid after recrystallization in a moderate isolated yield (33%). Protic solvents such as methanol were avoided throughout the synthesis as they could cause unintended ring-opening of the lactones.

Scheme 1.

Synthesis of AspGA and poly(AspGA-co-LA). TFA: trifluoroacetic acid; DCM: dichloromethane; DMF: dimethyl formaldehyde; TEA: triethyl amine; EtOAc: ethyl acetate; DMAP: 4-dimethylaminopyridine; BnOH: benzyl alcohol.

The random copolymers were prepared by melt ROP of AspGA with D,L-lactide using Sn(Oct)₂ as a catalyst and benzyl alcohol as an initiator at 145 °C with good monomer conversions (> 85%). Meanwhile, PLA was also prepared by ROP as a control. ¹H NMR spectra of the purified polymers showed that the aromatic signals associated with aspirin were detected from the copolymers (Figure S1, top), indicating that the aspirin moiety was well preserved during the melt ROP and subsequent workups. The aspirin pendant incorporation in polymers matched well with their respective monomer feeding ratios as determined by ¹H NMR integrations (Table S1). The successful incorporation of the aspirin pendants in the copolymers was further validated by FTIR (Figure S1, bottom). Characteristic aromatic C-C stretching frequencies at 1608 cm⁻¹ and 1486 cm⁻¹ as well as aromatic C-H stretching and bending frequencies around the 2900-3000 cm⁻¹ region and at 704 cm⁻¹ regions were more prominently observed in both poly(AspGA-co-LA) specimens, ascribable to the aspirin pendants,³¹ than in the PLA (which only contained a trace aromatic initiator). GPC analysis (using polystyrene as standard) showed that relatively narrow polydispersity (Ð ~ 1.4) was obtained for copolymers incorporating 5.9% and 10.5% AspGA. A relatively close range of molecular weights (MW) was obtained for the PLA control (M_n = 67800 g mol⁻¹) and the poly(AspGA-co-LA) (M_n = 67900 g mol⁻¹ for 5.9% AspGA and M_n = 46200 g mol⁻¹ for 10.5% AspGA). As excessive AspGA incorporation would result in lower MW due to reduced polymerization efficiency (likely due to steric hindrance imposed by the aspirin pendants), we did not include higher AspGA-content poly(AspGA-co-LA) compositions in this study. Water contact angle measurements of the polymer thin films formed on glass slides confirmed the expected increases in surface hydrophobicity upon aromatic aspirin pedant incorporation, with water contact angles increasing from 73° for PLA to 88° for 5.9% poly(AspGA-co-LA) and 105° for 10.5% poly(AspGA-co-LA) (Figure S2). The obtained polymers were hot-pressed (Supporting Information) into pellets for mechanical tests and for in vivo subcutaneous implantation studies.

3.2. Thermal transitions and thermomechanical properties

Differential scanning calorimetry (DSC) of the copolymers revealed a narrow glass transition and no visible melt transition, supporting their largely amorphous nature. The copolymerization of AspGA resulted in an upward shift of glass transition temperature (T_g) compared to the PLA control in an AspGA content-dependent manner, with ΔT_g = 7 °C for 5.9% poly(AspGA-co-LA) and ΔT_g = 9 °C for 10.5% poly(AspGA-co-LA), respectively (Figure 1A). The T_g’s of these poly(AspGA-co-LA), 48.5 °C and 50.3 °C (Table S2), are within a biologically safe range to trigger permanent shape recovery during shape memory programming.⁶

Figure 1.

(A) DSC thermograms (2^nd heating cycle) of PLA, 5.9% poly(AspGA-co-LA) and 10.5% poly(AspGA-co-LA) from −40 °C to 180 °C at a 10 °C min⁻¹ heating rate. (B) Thermal mechanical properties of 5.9% poly(AspGA-co-LA) vs. 10.5% poly(AspGA-co-LA) obtained on a Q800 DMA using a tensile fixture. (C) Storage/loss moduli and (D) complex viscosity as a function of rheological oscillatory temperature sweep of 10.5% poly(AspGA-co-LA), 5.9% poly(AspGA-co-LA) vs. the PLA control determined on an AR 2000EX rheometer at a scanning rate of 5 °C min⁻¹ and a constant angular frequency of 1 rad s⁻¹.

Unlike the brittle PLA control that tended to break during thermomechanical molding or mounting on the DMA (Figure S3A), the poly(AspGA-co-LA) copolymers exhibited good ductility. They could be bent and twisted without breaking (Figure S3B), which is attractive for their shape memory programming. Thermomechanical analysis of the poly(AspGA-co-LA) specimens by dynamic mechanical analysis (DMA) confirmed the upward shift of T_g’s in an AspGA incorporation content-dependent manner (Figure 1B). Both copolymers possessed high glassy-state storage moduli (>1-GPa) at body temperature (E’_{37 °C}, Table S2). This cortical bone-like mechanical strength may open the door for these copolymers as weight-bearing implants for in vivo orthopedic/dental applications. The slightly higher glassy-state tensile storage modulus of 10.5% poly(AspGA-co-LA) compared to 5.9% poly(AspGA-co-LA), despite a lower MW of the former, suggests that the hydrophobic interactions among the pendant aspirin units played an important role in strengthening the copolymer. The aggregation among the hydrophobic aspirin pendants have outweighed the relative reduction in PLA chain-chain entanglements below T_g’s in terms of the mechanical impact. In contrast, the storage modulus at elastic state declined as the AspGA incorporation content increased while PLA content decreased, suggesting that the hydrophobically aggregated aspirin pendants were dissociated at the higher temperatures and the extent of physical entanglements of PLA was the predominant factor dictating the elastic storage modulus of the copolymers.

3.3. Processing rheological properties

A key characteristic of dynamic polymer networks is their capability to flow under appropriate conditions. To better evaluate the influence of AspGA incorporation on the polymer viscoelastic characteristics and processability, oscillatory temperature sweep was performed (Figure 1C). Above a characteristic crossover temperature (T_crossover) where storage modulus G’ equals to loss modulus G’’, the material becomes fluid-like while below this temperature the material behaves like solids. Both poly(AspGA-co-LA) and the PLA control exhibited a rubbery plateau within the range of 10⁶ – 10⁸ Pa, with the rubbery moduli positively correlated with the AspGA incorporation content. T_crossover increased dramatically from 40 °C for PLA to 57 °C for 5.9% poly(AspGA-co-LA) and 60 °C for 10.5% poly(AspGA-co-LA), supporting that the presence of aspirin pendants impeded the network flow as the glassy-state physical entanglements of PLA chains were stabilized by the hydrophobic aggregation of the aspirin units. The complex viscosity, another critical indicator of polymer rheology, dynamic interaction and processability, was also determined by the temperature sweep experiments (Figure 1D). All 3 polymers examined exhibited a step change in complex viscosity around their respective flowing temperatures, with their complex viscosity holding steady below the transition while steadily declining above the transition, which is characteristic of thermoplastic networks. The complex viscosity of 10.5% poly(AspGA-co-LA) below the flowing point was significantly higher than that of the copolymer with a lower AspGA incorporation content (5.9%) or the PLA control. These rheological data are in general agreement with the observations from the DSC and thermomechanical measurements, further supporting that the incorporation of aspirin pendants significantly augmented/stabilized the glassy-state PLA entanglements.

3.4. Ultimate tensile stress/strain, toughness and energy dissipation properties

Uniaxial tensile test-to-failure experiments showed that the incorporation of higher content of AspGA (10.5%) led to significant enhancements in tensile modulus by >55% (to 1.74 GPa), ultimate strain by 30% (to 4.3%), ultimate stress by 70% (to ~25 MPa), and tensile fracture toughness by 130% (to ~82 MJ m⁻³) compared to 5.9% poly(AspGA-co-LA) (Figure 2A; Table S3).

Energy dissipations within poly(AspGA-co-LA) were investigated by five consecutive tensile loading and unloading cycles on the DMA (Figure 2B-C, Table S3). Hysteresis loops between the loading and unloading curves, resulting from the energy dissipation within the viscoelastic polymers, were observed for both compositions. Whereas the hysteresis loops for 5.9% poly(AspGA-co-LA) after the first cycle largely overlapped, those for the 10.5% poly(AspGA-co-LA) continued to shift. The dissipated energy, quantified as the area of the hysteresis loop, steadily dropped from 7.19 ± 0.6 MJ m⁻³ (8.8% of the total input work) in the first cycle to 5.54 ± 0.7 MJ m⁻³ in the second cycle and 4.04 ± 0.5 MJ m⁻³ in the 5^th cycle for 10.5% poly(AspGA-co-LA) (Table S4) under force controlled loading/unloading (0 to 18 N to 0). In contrast, the 5.9% poly(AspGA-co-LA) dissipated 4.64 ± 0.5 MJ m⁻³ energy (13% of the total input work) in the first cycle and 3.37 ± 0.3 MJ m⁻³ in the second cycle and 3.03 ± 0.2 MJ m⁻³ in the 5^th cycle (Figure 2B-C, Table S4). Although more energy was dissipated in 10.5% poly(AspGA-co-LA) than in 5.9% poly(AspGA-co-LA), the lost energy represented a lower percentage of total input work, reflecting the much higher toughness (total energy absorbed) of the 10.5% poly(AspGA-co-LA). The steadier declines in the dissipated energy between the 1^st and 5^th cycles for the copolymer with the higher content of aspirin pendants also suggests that not all hydrophobic interactions were fully restored within the timeframe the cyclic experiments.

The viscoelastic 10.5% poly(AspGA-co-LA) was characterized with faster stress relaxation (τ_1/2) compared to its lower AspGA content counterpart (Figure 2D and inset), suggesting that the dynamic association/dissociation among the aspirin pendants directly contributed to the energy dissipation within the polymer network. Finally, stress relaxation activation energy determined by the varying temperature experiments revealed much lower activation energy for 10.5% poly(AspGA-co-LA) (E_a = 94.6 KJ mol⁻¹; Figure 2F and inset) than its lower AspGA incorporation content counterpart (E_a = 149.2 KJ mol⁻¹; Figure 2E and inset) which had higher PLA content. These trends support that the dynamic hydrophobic interactions among the aspirin pendants were more effective in energy dissipation than that of the mobility and physical entanglement/disentanglement of PLA chains.

3.5. Thermal-healing behavior

To test the hypothesis that the dynamic nature of the hydrophobic aspirin pendant interaction and PLA chain entanglements could translate into good thermal-healing, a 10.5% poly(AspGA-co-LA) specimen was bisected and examined for the ability of the re-opposed interface to heal above T_g. As shown in Figure S4, the bisected specimen exhibited impressive healing across the interface within 30 min at 70 °C, supporting that the enhanced PLA and aspirin pendants mobility at the higher temperature enabled the reestablishment of PLA physical entanglements and hydrophobic aspirin pendant interactions across the interface.

3.6. Thermal responsive shape-memory behavior

We quantitatively assessed the shape-memory performance of poly(AspGA-co-LA) via stress-controlled one-way shape-memory cycles. At both AspGA incorporation contents examined, excellent temporary shape fixing ratio (R_f : 98.7% for 5.9% AspGA, 97.7% for 10.5% AspGA) at 25 °C and shape recovery ratio (R_r: 97.3% for 5.9% AspGA, 95.4% for 10.5% AspGA) at 65 °C were achieved (Figure 3A-B). We attribute the outstanding temporary shape fixing ratios achieved by incorporating only 5-10% AspGA to the hydrophobic interaction / π-π stacking among the aspirin pendants that helped fix/stabilize the stretched PLA chains in the temporary shape. The dynamic nature of the hydrophobic interactions among aspirin pendants, on the other hand, ensured their timely dissociations upon heating to allow efficient recoiling of the stretched PLA chains to return to their elastic state. The more efficient shape recovery observed with 5.9% poly(AspGA-co-LA) was consistent with the fact that less hydrophobic interactions would have to be overcome during the shape recovery process. Finally, we demonstrated the feasibility of achieving facile shape recovery under the safe triggering temperature of 50 °C, where a bulk 10.5% poly(AspGA-co-LA) achieved good shape recovery within 30 seconds (Figure 3C). Prior studies showed that local transient temperature increases should be kept below 70 °C to avoid potential tissue necrosis including bone.³²

Figure 3.

One-way shape memory cycles of (A) 5.9% poly(AspGA-co-LA) and (B) 10.5% poly(AspGA-co-LA) along with the respective R_f’s at 25 °C and R_r’s at 65 °C determined from the second cycle. (C) Demonstration of shape recovery of 10.5% poly(AspGA-co-LA) from bent “temporary” shape upon heating at 50 °C.

3.7. In vitro hydrolytic degradation and cytocompatibility of the degradation products

With the incorporation of hydrophobic aspirin pendants, we expect slower hydrolytic degradations of the poly(AspGA-co-LA) compared to PLA due to slower water penetration of the former. Indeed, accelerated in vitro hydrolytic degradation in PBS (pH 7.4, 10 mg mL⁻¹) at 70 °C confirmed slower weight losses for both 5.9% and 10.5% poly(AspGA-co-LA) compared to PLA, resulting in 50% weight losses in 40 h, 30 h and 15 h, respectively (Figure S5). The relatively faster hydrolytic degradation observed with 10.5% poly(AspGA-co-LA) compared to its 5.9% counterpart was attributed to the slightly lower molecular weight of the former, supporting that the overall hydrolytic degradation rate can be impacted by both hydrophobicity and molecular weight of the polymers.

The in vitro cytocompatibility of the degradation products of poly(AspGA-co-LA) and PLA, obtained after full disintegration of the bulk polymer through accelerated hydrolytic degradation in PBS (pH 7.4) at 70 °C, was examined with rat BMSCs. Supplementation of 12.5 μg mL⁻¹ of the respective degradation products every other day to the BMSCs culture in expansion media did not result in statistically significant impact to their proliferation over 5 days compared to those supplemented with PBS control (Figure S6A).¹² Although the cell viability trended higher in those supplemented with the degradation products of poly(AspGA-co-LA) vs. that of PLA, the differences were statistically insignificant. Similarly, we did not observe any negative impact on the osteogenesis of rat BMSCs upon supplementation (12.5 μg mL⁻¹, 3 times a week) of the respective degradation products over the course of 2-week osteogenic induction, as evidenced by robust alizarin red staining observed in all groups (Figure S6B).

3.8. In vivo polymer degradation and drug release

The extent of in vivo degradation of the polymer pellets implanted subcutaneously in rats was first grossly inspected by changes in explanted pellet dimensions over time (Figure 4A). The PLA control and the 10.5% poly(AspGA-co-LA) degraded faster, resulting in smaller explanted pellets at 16 weeks, whereas the 5.9% poly(AspGA-co-LA) exhibited a slower degradation. This observation was consistent with the in vitro hydrolytic degradation profiles of these polymers (Figure S5). The in vivo drug release within the fibrous tissue capsules was verified by HPLC. As shown in Figure 4B, HPLC traces of the methanol extract of the explanted 10.5% poly(AspGA-co-LA) pellet and its tissue capsule revealed a peak corresponding to that of salicylic acid, supporting that the conjugated drug was cleaved off the polymer scaffold and deacetylated as the bulk degradation of the polymer occurred. The peak detected around 9.4 min suggests that part of the deacetylation of aspirin, the active form of aspirin, may have occurred after its cleavage from the polymer side chain (Scheme S1).

3.9. Degradation-induced inflammatory responses

Our pilot subcutaneous implantation data first revealed a similar foreign body response to the implantation of all polymer pellets as characterized with the fibrous tissue encapsulation (Figure S7A). The initial acute inflammatory response observed at 2 weeks post-implantation subsidized over time resulting in a reduced number of immune cells (macrophages selectively quantified in the pilot study) detected within the fibrous tissue capsule and less active cellular activities by 8 weeks post-implantation (Figure S7A-B). These observations are consistent with literature reports and our prior observations with the temporal changes of foreign body responses to other biomaterial implants prior to their degradation.^33-34 The in vivo degradation of the pellets, however, triggered a secondary inflammatory response as characterized with increased cellular activities within the fibrous tissue capsules at 12 weeks post-implantation for the faster-degrading PLA and 10.5% poly(AspGA-co-LA), and at 16 weeks post-implantation for the slower-degrading 5.9% poly(AspGA-co-LA), respectively (Figure S7A-B). Notably, at 16 weeks, the secondary acute immune responses triggered by the degradation products subsidized surrounding 10.5% poly(AspGA-co-LA) while those surrounding the PLA control continued to exacerbate, suggesting a positive role of the in vivo released drug in mitigating the inflammatory response to degradation. To more thoroughly quantify the severity of the degradation-induced inflammatory responses as a function of the polymer composition, three 16-week explants of each composition were subjected to blind-evaluation by a pathologist. As shown in Figure 4A, all the explanted pellets were encapsulated with a network of fibrous tissue. Quantification of immune cells within the fibrous tissue capsules (Figure 4C) revealed less, although not statistically significant, macrophages within the fibrous tissue capsules surrounding poly(AspGA-co-LA) at 16 weeks compared to PLA (Table 1). Little difference was detected in the number of mast cells, eosinophils, neutrophils or lymphocytes across the three explanted polymer compositions. Significantly reduced number of foreign body giant cells (FBGCs, Figure 4D), however, was detected within the fibrous tissue capsules surrounding the aspirin-incorporated polymers compared to the PLA control at 16 weeks, suggesting that the salicylic acid released at this stage effectively mitigated the inflammatory responses to scaffold degradation. No statistically significant difference in the degree of immune responses, however, was detected between the 5.9% vs. 10.5% poly(AspGA-co-LA).

Table 1.

Quantification of inflammatory cells and blood vessels in 16-week explants.^a

Implant	Macrophage b	Mast cell c	Neutrophil c	Lymphocyte c	Eosinophil c	Blood vessel d
PLA	++	+	+	+	+	++
10.5%	+	+	+	+	+	++
5.9%	+	+	+	+	+	++

^aImmune cells were counted in five randomly selected 400× fields of view (FOV), and averaged.

^bFor macrophage: < 5/FOV (−), 5-10/FOV (+), 11-15/FOV (++), ≥16/FOV (+++)

^cFor mast cell, neutrophil, lymphocyte, eosinophil: < 2/FOV (+), 3-5/FOV (++), ≥5/FOV (+++).

^dFor blood vessel: < 5/FOV (+), 5-10/FOV (++), ≥11/FOV (+++).

4. DISCUSSION

Conventional amorphous PLA materials are known for their poor toughness and tendency for brittle fractures under tension and bending.³⁵ One of the main parameters in determining the glassy state mechanical properties in polymers is their molecular weights between entanglements. Highly entangled amorphous polymers often show ductile fracture due to the energetically favorable shear yielding deformation mechanism.³⁶ Interestingly, the amorphous PLAs were shown to possess low entanglement density and fracture by a crazing mechanism instead.³⁷ This limitation has made the processing and shape memory programming of conventional amorphous PLA materials very difficult, despite their faster hydrolytic degradation being better suited for certain in vivo applications than the slower-degrading crystalline PLA. Indeed, the amorphous PLA control prepared in this study was so brittle that it became nearly impossible to program temporary shapes or mount the specimens for dynamic mechanical analyses or one-way shape memory cycle experiments without breaking. In this study, we aim to address multiple key limitations of amorphous PLA-based biomaterials, including brittleness, poor shape memory properties, and degradation-induced inflammation, by strategic copolymerization of a novel aspirin-functionalized glycolide monomer AspGA with D, L-lactides. We hypothesize that the hydrophobic aggregation among the aspirin pendants could strengthen the physical entanglement of the PLA chains in the glassy state and stabilize temporarily stretched/bent configurations during shape memory programming while their dynamic dissociation provides an effective mechanism for energy dissipation under loading or heating, thereby enhancing its toughness, thermal healing and shape recovery. Meanwhile, we hypothesize that the inflammatory response elicited by the in vivo degradation of amorphous PLA scaffolds can be mitigated by the concomitant hydrolytic cleavage and release of covalently conjugated aspirin pendants.

Using ROP, random copolymers poly(AspGA-co-LA) consisting of 0, 5.9, or 10.5% aspirin-bearing pendants with comparable molecular weights and polydispersity were synthesized. Water contact angle measurements confirmed that the incorporation of the aspirin pendants resulted in increased surface hydrophobicity. These thermoplastic polymers could be readily thermal-pressed into bulk specimens with various shapes and sizes, with the aspirin-functionalized copolymers exhibiting significantly improved ductility, toughness and handling characteristics (Figure S3) desired for shape memory programming. DSC revealed an amorphous network structure free of crystalline domains for the copolymers, with an upward shift of T_g positively correlating with the AspGA content (Figure 1A). Thermomechanical analyses revealed and validated the positive correlations of glassy-state modulus and T_g with the AspGA incorporation content, respectively (Figure S2B). The gigapascal glassy-state storage moduli of poly(AspGA-co-LA), along with a T_g (48-50 °C) slightly above body temperature, make them uniquely suited for weight-bearing in vivo applications (e.g., as resorbable synthetic bone grafts). Oscillatory rheology experiments revealed a flow behavior characteristic for thermoplastic network where a high rubbery modulus plateau (Figure 1C) and higher complex viscosity (Figure 1D) were observed for both aspirin-functionalized copolymers, with an increasing AspGA incorporation content resulting in a more persistent solid-like behavior at higher temperatures. Overall, these thermal, mechanical and rheological properties collectively support our hypothesis that the hydrophobic interactions among the aspirin pendants augmented/stabilized the polymer chain entanglements, with these hydrophobic domains increasing T_g’s and enhancing the storage moduli and impeding the free-flowing of the polymer chains at the glassy state.

Furthermore, we demonstrated that the incorporation of AspGA also resulted in significant improvements in the ultimate stress, ultimate strain, fracture toughness (Figure 2A) and energy dissipative / stress relaxation properties (Figure 2B-D) of the copolymer in an aspirin pedant content-dependent manner (Table S3). These observations support that the hydrophobic interactions among the aspirin pendants are dynamic in nature, with their effective dissociation under stress providing an important toughening mechanism for the copolymers, which is attractive for their applications as shock-absorbing materials and weight-bearing medical implants. The more effective energy dissipation (Figure 2B-C) and faster stress relaxation (Figure 2D) observed with the 10.5% poly(AspGA-co-LA) was further validated with a lower activation energy (E_a = 94.6 KJ mol⁻¹) compared to the 5.9% poly(AspGA-co-LA) counterpart (E_a = 149.2 KJ mol⁻¹; Figure 2E-F). It suggests that the dynamic disruption of the hydrophobic interaction among the aspirin pendants facilitates energy dissipation more effectively than PLA chain mobility and disentanglements. The low binding energy of aspirin pendant aggregations also translated into excellent thermal healing of the copolymer (Figure S4), where the dynamic dissociation/association of the hydrophobic pendants along with enhanced PLA chain entanglement/disentanglement above T_g enabled the healing of the bisected specimens upon heating. This property can be exploited for the thermal repair/regeneration of poly(AspGA-co-LA) as structural materials and coatings.

Finally, we demonstrated that the excellent material handling characteristics (e.g. improved ductility and toughness), suitable thermal transition window (T_g near physiological temperature), and the hydrophobic interaction among the aspirin pendants dynamically strengthening PLA chain-chain interaction translated into outstanding shape memory performances. By simply incorporating 5-10% AspGA, we achieved nearly perfect temporary shape fixing at room temperature and facile shape recovery at 50-65 °C (Figure 3), a rare accomplishment with PLA-based thermoplastics without extensive changes of the polymer compositions or formulations.

In parallel, to establish the translational biomedical application potentials of these polymers, we examined their hydrolytic degradation kinetics and the cytocompatibility of their degradation products. Consistent with the increased surface hydrophobicity as revealed by water contact angles, the incorporation of aspirin pendants led to slower hydrolytic degradation of poly(AspGA-co-LA) compared to PLA control. Meanwhile, the slightly lower molecular weight of the 10.5% poly(AspGA-co-LA) resulted in a slightly faster hydrolytic degradation compared to 5.9% poly(AspGA-co-LA) despite the higher AspGA content of the former, underscoring both molecular weight and hydrophobicity as key factors influencing the polymer hydrolytic degradation. Continuous supplementation of the polymer degradation products (12.5 μg mL⁻¹) did not negatively impact the proliferation or osteogenesis of rat BMSCs, supporting good cytocompatibility of this class of degradable polymers and their potential applications as synthetic bone graft materials.

Furthermore, we tested our hypothesis that selective inhibition of the inflammatory response elicited by the in vivo degradation of PLA scaffolds, as opposed to the early acute inflammatory response required for initiating the healing cascade,^{23, 38-39} can be accomplished by concomitant hydrolytic cleavage and release of covalently conjugated aspirin pendants from the synthetic copolymer. Gross inspection of the explanted pellet sizes of the explants (Figure 4A) revealed that 5.9% poly(AspGA-co-LA), with almost identical molecular weight as the PLA control, exhibited slower in vivo degradation, consistent with its strengthened mechanical property and increased hydrophobicity, which have likely slowed water penetration and the hydrolytic degradation. The 10.5% poly(AspGA-co-LA), with slightly lower molecular weight than the PLA control, exhibited an overall in vivo degradation rate comparable to that of the PLA control. Histological examination of the explants showed that these implants elicited normal foreign body responses upon implantation and were minimally immunogenic prior to their degradation. However, at 16 weeks, when all explants had underwent substantial in vivo degradation (accompanied by the release of salicylic acid in the case of poly(AspGA-co-LA) (Figure 4B), slightly lower numbers of macrophages and significantly less FBGCs (Figure 4C-D; Table 1) were detected within the fibrous tissue capsules surrounding poly(AspGA-co-LA). The degradation products elicited minimal hypersensitive/allergic responses (low counts of mast cells or eosinophils) and similar recruitment of neutrophils and lymphocytes (Table 1) across all formulations examined.

Macrophages are the primary immune cell type involved in both acute and chronic inflammatory responses,⁴⁰ exerting either beneficial or deteriorating roles during the dynamic tissue regeneration process.⁴¹ In advanced stages of inflammation, macrophages undergo “frustrated phagocytosis” and fuse into FBGCs, which are known for higher degradative capacities than macrophages and could inhibit neovascular infiltration and tissue regeneration.^42-43 Therefore, the detection of FBGCs within the fibrous tissue capsules surrounding biomaterial implants is often considered as a hallmark of poor biocompatibility or uncontrolled chronic inflammation.⁴² The mitigated immune responses detected within the fibrous tissue capsules surrounding poly(AspGA-co-LA) supported our hypothesis that concomitant hydrolytic cleavage of the pendant aspirin could mitigate scaffold degradation-induced inflammatory responses. Reduced number of FBGCs may also contribute to the slower immune cell-mediated degradation of the aspirin-conjugated polymers compared to PLA of similar molecular weight. For future studies, whether and how the temporally controlled local release of salicylic acid from the copolymers differentially impact macrophage polarization (e.g. M1 vs. M2 macrophages), which has been increasingly recognized for their complementary/opposing roles during various stages of inflammation and tissue repair,⁴³ will be investigated. The efficacy of poly(AspGA-co-LA) in improving their in vivo performance as tissue engineering scaffolds needs to be validated using appropriate tissue defect models.

5. CONCLUSIONS

A novel class of degradable, thermoplastic SMPs copolymerized from D,L-lactide and 5-10% aspirin-functionalized monomer AspGA was prepared by melt ROP. The dynamic hydrophobic interactions among the aspirin pendants efficiently modulated the thermal, mechanical, rheological, and energy dissipative properties of the copolymer in an aspirin incorporation content-dependent manner, resulting in gigapascal-storage moduli at body temperature, significantly enhanced toughness, ultimate tensile stress/strain, and thermal healing properties. The copolymers also exhibited outstanding shape memory performance around a physiologically relevant temperature range, achieving nearly perfect temporary shape fixing at room temperature and facile shape recovery at 50 °C. The hydrolytic degradation products of these copolymers did not negatively impact the proliferation or osteogenesis of BMSCs in vitro. Furthermore, using a rat subcutaneous implantation model, we demonstrated that the PLA degradation-induced inflammation could be mitigated by the concomitant hydrolysis of the aspirin pendants and the release of salicylic acid from the copolymer, without suppressing early acute inflammatory responses.

This straightforward yet innovative strategy provides a practical alternative to engineering novel properties of PLA-based materials for biomedical applications without significant alteration of their chemical compositions or polymer synthesis and processing. The multitude of fascinating properties introduced by the aspirin pendants and their interplay will inspire future creative designs of functional biodegradable polymeric materials. On-going and future studies of poly(AspGA-co-LA) will capitalize on their exciting properties as robust synthetic tissue scaffolds to achieve minimally invasive delivery to an area of skeletal tissue defect and safe degradation upon guiding tissue regeneration.