polycaprolactone fibrous composites

Materialia 14 (2020) 100874

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Radiopaque scaffolds based on electrospun iodixanol/polycaprolactone fibrous composites Joy Vanessa D. Perez a,b, Burapol Singhana a,c, Jossana Damasco a, Linfeng Lu a, Paul Behlau a, Raniv D. Rojo a,b, Elizabeth M. Whitley e, Francisco Heralde b, Adam Melancon f, Steven Huang a, Marites Pasuelo Melancon a,d,∗ a

Department of Interventional Radiology, The University of Texas MD Anderson Cancer Center, Houston, TX 77030, USA of Medicine, University of the Philippines Manila, Manila, National Capital Region 1000, Philippines Nanomedicine Research Unit, Chulabhorn International College of Medicine, Thammasat University, Rangsit Campus, Pathum Thani 12120, Thailand d Graduate School of Biomedical Sciences, University of Texas Health Science Center at Houston, Houston, TX 77030, USA e Department of Veterinary Medicine and Surgery, The University of Texas MD Anderson Cancer Center, Houston, TX 77030, USA f Department of Radiation Physics, The University of Texas MD Anderson Cancer Center, Houston, TX 77030, USA b College

c Innovative

a r t i c l e

i n f o

Keywords: Polycaprolactone Electrospinning Computed tomography Radiopacity Mesenchymal stem cells Polymeric scaffolds

a b s t r a c t Grafts based on biodegradable polymer scaffolds are increasingly used in tissue-engineering applications as they facilitate natural tissue regeneration. However, monitoring the position and integrity of these scaffolds over time is challenging due to radiolucency. In this study, we used an electrospinning method to fabricate biodegradable scaffolds based on polycaprolactone (PCL) and iodixanol, a clinical contrast agent. Scaffolds were implanted subcutaneously into C57BL/6 mice and monitored in vivo using longitudinal X-ray imaging and micro-computed tomography (CT). The addition of iodixanol altered the physicochemical properties of the PCL scaffold; notably, as the iodixanol concentration increased, the fiber diameter decreased. Radiopacity was achieved with corresponding signal enhancement as iodine concentration increased while exhibiting a steady time-dependent decrease of 0.96% per day in vivo. The electrospun scaffolds had similar performance with tissue culture−treated polystyrene in supporting the attachment, viability, and proliferation of human mesenchymal stem cells. Furthermore, implanted PCL-I scaffolds had more intense acute inflammatory infiltrate and thicker layers of maturing fibrous tissue. In conclusion, we developed radiopaque, biodegradable, biocompatible scaffolds whose position and integrity can be monitored noninvasively. The successful development of other imaging enhancers may further expand the use of biodegradable scaffolds in tissue engineering applications.

Statement of Significance The development of radiopaque bioresorbable scaffolds unlocks a vast array of preclinical and clinical tissue engineering applications that are presently limited by the inability to monitor scaffolds as they degrade in vivo. Previous studies demonstrated covalently incorporating iodine in polymer scaffolds but did not correlate radiopacity with degradation behavior of the materials over time. In this study, we report the feasibility of non-covalent incorporaton of iodine in polycaprolactone to generate radiopaque scaffolds and evaluate its effects on physicochemical properties and biocompatibility. We show that iodixanol incorporation confers radiopacity that enables long-term, real-time monitoring of the position and degradation of the biomaterial in vivo using computed tomography (CT), a noninvasive, clinically available imaging modality.

∗

1. Introduction Tissue engineering is an interdisciplinary field that combines principles of engineering, life sciences, and chemistry to develop strategies to repair, regenerate, or replace diseased tissue with biologically and mechanically equivalent tissue [1–3]. Scaffold-based grafts, which enable autologous cells to repopulate in vitro or in situ, provide a promising therapeutic option that overcomes the shortcomings of currently used prostheses [3]. Bioresorbable scaffolds are advantageous because the scaffolds gradually degrade as the natural tissue regenerates without leaving permanent foreign material that could initiate a chronic immune reaction [4]. Monitoring scaffold degradation is key to determining the optimal degradation rate that maintains the ability of the scaffold to bear mechanical strain while allowing formation of load-bearing tissue [5]. However, owing to challenges in noninvasively monitoring their

Corresponding author at: Department of Interventional Radiology, The University of Texas MD Anderson Cancer Center, Houston, TX 77030, USA. E-mail address: [email protected] (M.P. Melancon).

https://doi.org/10.1016/j.mtla.2020.100874 Received 16 May 2020; Accepted 17 August 2020 Available online 22 August 2020 2589-1529/© 2020 Acta Materialia Inc. Published by Elsevier B.V. All rights reserved.

J.V.D. Perez, B. Singhana and J. Damasco et al.

position and integrity, bioresorbable scaffolds are slow to gain traction in clinical applications [6–8]. Most strategies for monitoring bioresorbable scaffolds in preclinical studies focus on tracking cells and assessing function; very few focus on directly visualizing such scaffolds in vivo [6,7]. The degradation of polyester [8–10], polyurethane urea [11,12], polydioxanone [11], and collagen [13] scaffolds has been assessed using imaging modalities such as ultrasound elasticity imaging, ultrasound shear wave imaging, photoacoustic imaging, and near-infrared fluorescence imaging. However, in addition to having limited penetration depth, such imaging modalities are not routinely used nor readily available in the clinic [14,15]. Previous studies showed that clinical magnetic resonance imaging (MRI) can be used to image collagen [16] and polyvinylidene fluoride [17,18] scaffolds labeled with ultrasmall superparamagnetic iron oxide. However, these studies were unable to show whether MRI can be used to monitor the degradation of the scaffolds as collagen degradation was not observed and polyvinylidene fluoride is non-biodegradable. Computed tomography (CT), which is a routinely used, noninvasive clinical imaging modality offers improved spatial resolution and deeper tissue penetration than ultrasound-based systems [19]. CT contrast agents are based on heavy elements, such as iodine and barium, which have high atomic numbers. Other heavy metals (e.g. gold) have been used to assess scaffold degradation in vitro [20], but iodine is the most widely adapted for clinical CT. Previous studies showed that radiopaque iodine can be incorporated into non-bioresorbable polymers, such as poly(methacrylate) copolymers [21], thermoplastic polyurethane elastomers [22], and phenylalanine-based poly(ester urea) [23] for use in medical devices. In other studies, radiopaque polymers were obtained by substituting anionic groups with iodine monomers that covalently bond to biodegradable polymers such as polycaprolactone [24]. These studies focused on functionalizing iodine within the polymer chains to render the chains radiopaque. However, the effect of clinically approved CT contrast agent incorporation on bioresorbability of the scaffolds and the utility of radiopacity in monitoring polymer degradation in vivo has not been investigated. We have previously shown that although coating through the wetdipping method can be used to infuse matrices of poly-p-dioxanone sutures with radiopaque materials, it has inefficient iodine loading (< 1%) [25–28]. Direct incorporation through electrospinning may be a more effective and efficient approach to creating radiopaque scaffolds. Electrospinning is a method that applies a high voltage on a polymer solution to fabricate fibrous scaffolds with submicron fiber diameters. The fiber mesh architecture mimics the hierarchical organization of fibrous tissue such as collagen in the extracellular matrix. This facilitates enhanced cell attachment and drug loading, making it favorable for tissue engineering applications [29]. One of the polymers most often used for electrospinning is polycaprolactone (PCL), a biocompatible, bioresorbable synthetic polymer that has been gaining traction as a biomaterial and tissue-engineering scaffold owing to its low melting temperature, very good blend-compatibility, low cost, and FDA approval [30]. We hypothesize that electrospinning will allow more efficient iodine loading into PCL scaffolds, hence improved radiopacity. By imparting radiopacity to bioresorbable scaffolds, we seek to enable real time in vivo assessment that can unlock further development of clinical applications. In this study, we incorporated iodixanol, a clinical CT contrast agent, into electrospun PCL scaffolds to create radiopaque grafts that can be monitored noninvasively to assess position, structural integrity, and degradation over time. 2. Experimental 2.1. Materials Polycaprolactone (average Mn 80,000) and iodixanol (European Pharmacopoeia Reference Standard) were purchased from SigmaAldrich (St. Louis, MO), 1,1,1,3,3,3-hexafluoro-2-propanol (HFP;

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>99.5%) was obtained from Fisher Scientific (Hampton, NH). All chemicals were used without further purification unless otherwise noted. 2.2. Electrospinning Polymeric-iodixanol solutions were prepared by mixing 130 mg of PCL with 5 mg (PCL-I5), 15 mg (PCL-I15), or 40 mg (PCL-I40) of iodixanol per mL of HFP at 50 °C for 15 min to obtain homogenous mixtures. The custom-designed electrospinning apparatus consisted of a Spraybase high-voltage power supply (Avectas, Maynooth, Ireland), a Spraybase infusion pump (Avectas), a plastic syringe, a stainless steel bluntend needle (20 gauge), and a metal mandrel collector. The syringe was mounted horizontally on the infusion pump, and the sample solution was fed through the syringe at a constant rate of 3.0 mL/h to the needle tip. The distance between the needle tip and the mandrel collector was 7 cm, and the applied voltage to the needle tip was 10.5 kV. The parameters were optimized by changing the voltage, flow rate, and tipto-collector distance until a homogenous and uniform fiber morphology was produced from a single continuous jet (Taylor cone). 2.3. Morphology, fiber diameter, and pore size The morphologies of the PCL and PCL-I scaffolds were assessed by scanning electron microscopy (SEM) with a Nova NanoSEM microscope (FEI, Hillsboro, OR) using an EDAX system (Ametek, Berwyn, PA) operating at 15 kV. Images were generated at 1000× and 5000× magnification. At least five SEM images were captured from random areas to determine the average fiber diameter and pore size of each sample. From each image, at least 100 segments of fibers and pore areas were randomly selected and measured using ImageJ software (National Institutes of Health, Bethesda, MD). A histogram plotting the distribution of fiber diameters and pore sizes was generated, and the average fiber diameters and pore sizes were calculated. 2.4. Quantification of iodine loading The actual iodine loading within the PCL stents was determined by inductively coupled optical emission spectrophotometry (ICP-OES) with a Varian 720ES spectrometer (Agilent Technologies, Santa Clara, CA). Briefly, scaffolds were digested in aqua regia (concentrated nitric acid:hydrochloric acid 1:3, v/v) at 55 °C until completely dissolved. The resulting solutions were diluted in deionized water and the concentration of iodine was quantified by ICP-OES at a wavelength of 179.85 nm. The experiments were conducted in triplicates. Data were analyzed using the regression results obtained from calibration curves (R2 ≥0.998) plotted from the concentrations of the standard iodine solutions. 2.5. Porosity Porosity was determined using two different methods: calculation method [31] and liquid intrusion method [32]. The calculation method is a simple, rapid method, but significant errors may be acquired in the determination of the volume of the sample. Therefore, we also used an indirect method of determining porosity. The liquid intrusion method relies on the displacement of ethanol, a solvent that penetrates the pores easily but does not cause the material being tested to shrink or swell. For the calculation method, porosity was determined using the equations below. A microcaliper was used to measure the thickness of the fibrous scaffolds. ( ) 𝑚𝑎𝑠𝑠 (𝑔 ) 𝑔 (1) apparent density = ( ) 𝑐 𝑚3 thickness (𝑐𝑚) × 𝑎𝑟𝑒𝑎 𝑐 𝑚2 ( ) apparent density 𝑐𝑚𝑔 3 porosity = 1 − (2) ( ) bulk density ofPCL 𝑐𝑚𝑔 3

J.V.D. Perez, B. Singhana and J. Damasco et al.

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In the liquid intrusion method, the polymer scaffolds were weighed prior to immersion in ethanol (intruding liquid of density 𝜌EtOH = 0.789 g/mL), left overnight on a shaker table to allow diffusion of ethanol into the void volume, blotted with tissue, and reweighed. The porosity was calculated as: 𝑝𝑜𝑟𝑜𝑠𝑖𝑡𝑦 =

𝑉𝐸𝑡𝑂𝐻 𝑉𝐸𝑡𝑂𝐻 + 𝑉𝑃 𝐶𝐿

(3)

where VEtOH is the volume of the intruded ethanol (i.e., the ratio between the observed mass change after intrusion and 𝜌EtOH ) and VPCL is the volume of the PCL fibers. 2.6. Thermal property Differential scanning calorimetry was used to determine the melting temperature (Tm ) and crystallization temperature (Tc ) of PCL and PCL-I scaffolds. The thermal history was assessed using an STA PT 1000 thermogravimetric analyzer (Linseis, Selb, Germany) under temperatures ranging from 20 °C to 70 °C at a heating/cooling rate of 5 °C/min in a nitrogen atmosphere.

side and incubated at 37 °C for 30 min before the addition of culture media. After 12 h of incubation, the cell-seeded samples were imaged under an Eclipse Ti2 fluorescence microscope (Nikon, Melville, NY) and prepared for SEM after 48 h by fixing with 2.5% glutaraldehyde, dehydrating in increasing ethanol concentrations, rinsing in 1:1 tert-butyl alcohol:ethanol, and air-drying overnight in preparation for gold-sputter coating. SEM was performed as described in Section 2.3. To determine the viability and proliferation of the attached cells, 6-mm diameter circular samples were punched from the scaffolds and sterilized as previously described. GFP-MSCs were seeded at a concentration of 1 × 104 cells per sample as described, washed with PBS to remove unattached cells, and placed in a 96-well microtiter plate with culture medium overnight. Cell proliferation was measured using the alamarBlueTM assay, a non-toxic colorimetric assay that measures cell viability. Fluorescence readings were taken at an excitation wavelength of 540 nm and emission wavelength of 590 nm using a Cytation5 microplate reader (Biotek Instruments, Winooski, VT). To determine the cytotoxic potential of iodine incorporation, normalized cell viability (the fold-change in cell growth) was calculated by dividing the fluorescence of each sample by the fluorescence of the positive control (cells plated on a cell culture-treated polystyrene plate).

2.7. Mechanical property 2.11. Hemolysis The maximum stress (i.e. ultimate tensile strength) of the electrospun scaffolds was determined using an Instron 3345 single-column testing system (Instron, Norwood, MA) with a 10-N load cell under a crosshead speed of 20 mm/min at ambient conditions with the clipped length maintained at 50 mm. Three samples were tested per scaffold. All samples were prepared in the form of rectangular membranes with dimensions of 50 × 10 × 0.20 mm3 . 2.8. Micro-CT Radiopacity was determined using an eXplore Locus RS micro-CT scanner (GE Medical Systems, London, ON, Canada). This system uses a tungsten source X-ray tube operating at 80 kV and 450 mA. The X-ray source and charge-coupled device-based detector gantry revolve around the subject in roughly 1.0° increments. The “MedRes-Mimic10minMimic” protocol was used for approximately 10 min with the following calibration values: –1000 Hounsfield units (HU) for air, −300 HU for fat, and +1000 HU for bone. The radiopacity of the materials in HU was quantified using the MicroView (Parallax Innovations, Ilderton, ON, Canada) software to calculate the mean HU of at least 10 regions of interest within the scaffolds. A line profile was used to plot the HU values along a line segment transecting the scaffold and to determine the maximum HU value within it.

Heparinized pig blood was centrifuged at 2000 × g for 10 min at 4 °C. The red blood cell-rich pellet was collected and diluted with PBS to form a 1:50 solution. Scaffolds were cut into 0.5 × 0.5 cm pieces, sterilized in 70% ethanol for 30 min, and washed twice with PBS. The samples were incubated with 500 mL of 2% heparinized blood for 1 h at 37 °C under gentle, continuous shaking. These were then centrifuged at 750 × g for 15 min, and the resulting supernatant was assessed for optical density at 540 nm. Triton X was used as the positive control and heparinized blood without scaffold was used as the negative control. Experiments were carried out in triplicates and the results were normalized with the positive control designated as 100% hemolysis. 2.12. Blood clotting

An MX-20 digital specimen radiography system (Faxitron Bioptics, Tucson, AZ), a high-throughput planar system, was used for X-ray imaging. The MX-20 model has an energy range of 10-35 kV and a tube current of 300 𝜇A. The focal spot size is <20 𝜇m, and the X-ray beam angle is 40°. Samples were run at an exposure time of 19 s at 25 kV.

Clotting time assay [33] was performed using extracted pig blood anticoagulated with sodium citrate (0.3% final concentration) and centrifugated at 1100 × g for 12 min. The resulting supernatant (plateletpoor plasma; PPP) was used for the subsequent clotting experiments. Scaffolds were cut into 0.5 × 0.5 cm pieces, sterilized in 70% ethanol for 30 min, and washed with PBS twice. The samples were incubated with 200 𝜇L of citrated PPP for 30 min at 37 °C under continuous shaking. After incubation, the tubes were put in ice to inhibit the activation process. Afterwards, 200 𝜇L of 3% rabbit brain cephalin was added, and the samples were incubated for an additional 2 min at 37 °C. Then, 200 𝜇L of 30 mM calcium chloride was added to the solution and the clotting time was recorded in quadruple for each incubated sample. Citrated PPP incubated without scaffold was used as the negative control while incubation with glass was used as the positive control. Three independent batches were performed. Data were pooled by calculating the normalized clotting time per sample.

2.10. Cell attachment, viability, and proliferation

2.13. In vivo imaging, histology, and scaffold degradation

Primary human mesenchymal stem cells transfected with green fluorescent protein (GFP-MSC) using a standard lentiviral assay were obtained from the laboratory of Dr. Frederick Lang at MD Anderson Cancer Center. The cells were maintained in minimum essential media supplemented with 10% fetal bovine serum and 1% penicillin-streptomycin. All incubations were performed at 37 °C in a 5% CO2 atmosphere. Scaffolds were sterilized by incubation in 70% ethanol for 30 min followed by two sequential 15-min washes in phosphate-buffered saline (PBS) solution. Sterilized scaffolds were seeded with 5 × 104 GFP-MSCs per

All mouse experiments were approved by the Institutional Animal Care and Use Committee (IACUC). PCL and PCL-I15 scaffolds were cut into 1-cm pieces, sterilized with ethylene oxide gas for 48 h at room temperature, and subcutaneously implanted into 6- to 8-week old male C57BL/6 mice (Taconic Biosciences, Rensselaer, NY). The mice were anesthetized with isoflurane and once adequate analgesic depth was established by checking non-response to toe-pinch, two sagittal dorsal incisions were made with blunt dissection of the subcutaneous pocket. One PCL scaffold was implanted in one pocket and one PCL-I15 scaffold

2.9. X-ray imaging

J.V.D. Perez, B. Singhana and J. Damasco et al.

was implanted in the other pocket. Simple interrupted technique using absorbable sutures was used to close the incisions (Supplementary Fig. 1). All mice were monitored daily for adverse reactions. The mice underwent X-ray and micro-CT imaging at 1, 7, 14, 28, or 56 days after scaffold implantation. After imaging, mice were euthanized by CO2 asphyxiation followed by cervical dislocation. CT intensity was analyzed using the MicroView software. At least 10 regions-of-interest (ROIs) were selected randomly within the different samples and the mean HU was recorded for each scaffold over the different time points. HU values were plotted as a function of time to determine the changes in CT intensity as the scaffolds degraded in vivo. For histological evaluation, the scaffolds and the immediate surrounding tissue were explanted, fixed in 10% neutral buffered formalin for at least 24 h, processed routinely, and embedded in paraffin tissue blocks. The paraffin-embedded scaffolds were cut against the direction of the fiber alignment to yield 4-𝜇m sections, which were then stained with hematoxylin and eosin. The stained sections were scanned with an Aperio LV1 in vitro diagnostic medical device (Leica Biosystems, Buffalo Grove, IL) to capture images at 4X and 20X magnification and viewed with a DM2500 optical microscope paired with a DFC495 digital camera system (Leica Biosystems) to capture images at 40X magnification. Mechanical strength degradation of the scaffolds in terms of maximum stress was assessed in a separate in vivo experiment wherein sterilized PCL and PCL-I15 scaffolds measuring 3 cm in length were subcutaneously implanted into the dorsum of a C57BL/6 mouse. Scaffolds were explanted at 7, 14, 21, 28, 42, and 56 days post-implantation and degradation was measured by change in maximum stress over time, as described in Section 2.7. 2.14. Statistical analysis Where applicable, data were presented as means ± standard deviations and were compared and analyzed using one-way analysis of variance (ANOVA) and/or a two-tailed Student’s t-test, assuming the data followed a normal distribution. A p value of less than 0.05 was used to identify statistically significant differences. 3. Results and discussion 3.1. Physicochemical properties of PCL and PCL-I scaffolds The physicochemical properties of electrospun polymeric scaffolds are influenced by the electrospinning parameters and the properties of the polymer (e.g. concentration, viscosity, and the nature of the solvent used) [34–37]. In the present study, we optimized the electrospinning parameters by changing the polymer concentration, applied voltage, solution flow rate, and tip-to-collector distance. We found that a voltage of 10.5 kV, tip-to-collector distance of 7 cm, and flow rate 3.0 mL/h produced the most uniform fibers. Using these electrospinning parameters, different concentrations of iodixanol (5, 15, or 40 mg) were added to the polymer solution. The composite is created by blending the PCL and iodixanol solution in HFP prior to electrospinning. As the high voltage is applied to the needle tip, the solvent vaporizes and a solid fibrous material is collected with the iodixanol trapped in the interstices of the PCL fibers – no covalent interaction between the components is created. It was observed that as the amount of iodine added to the PCL solution increased, the loading efficiency of electrospinning decreased (PCL-I5, PCL-I15, and PCL-I40 had a loading efficiency of 96.2%, 85.0%, and 83.9% respectively). This is possibly because at higher concentrations, the hydrophilic iodixanol has decreased solubility in the presence of hydrophobic PCL in HFP. Despite this, our results show that the electrospinning method offers increased incorporation of iodine into the polymer compared with the wet-dipping method [26]. The ultra-fine fibers produced by electrospinning provide a high surface area−to−volume ratio that can enhance cell attachment, drug loading, and mass transfer properties. In the present study, SEM revealed that

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the electrospun scaffolds were homogeneous and uniform over a large area (Fig. 1A–H). The fiber diameters decreased as their iodine content increased; PCL alone had the highest average diameter (2.30 ± 0.32 𝜇m), followed by PCL-I5 (2.13 ± 0.26 𝜇m), PCL-I15 (2.02 ± 0.49 𝜇m), and PCL-I40 (1.15 ± 0.26 𝜇m) (Fig. 1I–P). Multiple studies have investigated the use of electrospinning to incorporate iodine with polyvinyl pyrrolidone, polyvinyl butyral, and polyvinyl alcohol in the synthesis of antibacterial wound dressings [38,39]. In these studies, electrospinning elemental iodine dissolved in an ethanol/water mixture resulted in increased fiber diameter as iodine concentration increased. This was attributed to an increase in the polymer solution viscosity, which increases the viscoelastic force that counteracts the stretching of the charged jet, thus resulting in larger average fiber diameters [40]. In contrast, in electrospinning studies in which iodine solutions were doped with potassium iodate to stabilize the complex, the average fiber diameters decreased up to four-fold as the iodine concentration increased. This was attributed to the increased conductivity of the polymer solution, which exerts greater elongation forces to the electrospinning jet [39]. Similarly, in the present study, the trend of decreasing average fiber diameter with increasing iodine concentration may be attributed to the increased solution conductivity brought about by the addition of iodixanol, which is highly hydrophilic. Iodixanol’s high hydrophilicity and good solubility in polar solvents are due to the hydroxyl group in the central carbon atom in the dimer linkage and the multiple hydrophilic amide side chains of the molecule [41]. Increasing the polarity of the polymer solution would increase its conductivity, allowing it to produce a surface charge density sufficient to form a Taylor cone, initiate the electrospinning process, and influence the thinning and elongation of the charged jet [42]. A high degree of porosity and adequate pore size are necessary to provide ample space for cell infiltration and migration as well as to allow adequate exchange of nutrients and waste between the scaffold and the immediate environment. Small pores would impede cell infiltration whereas large pores could decrease tensile strength and cause other complications such as blood leakage. In the present study, porosity measured by calculation and liquid intrusions methods did not differ significantly (p>0.05, 1-way ANOVA) showing 60–70% porosity for both the PCL and PCL-I scaffolds (Table 1). This range of porosity is similar to those reported previously [43-45]. Measurement of pore size by ImageJ analysis of SEM images (Fig. 1A–H) revealed that PCL (31.81 ± 19.71 𝜇m2 ), PCL-I5 (36.76 ± 17.21 𝜇m2 ), PCL-I15 (47.06 ± 27.97 𝜇m2 ), and PCL-I40 (23.21 ± 12.05 𝜇m2 ) did not differ significantly (p > 0.05, 1-way ANOVA). Densely packed stacks of two-dimensional fibrous layers are fabricated with conventional electrospinning, with the interfiber areas referred to as the pore size. Statistical models as well as experimental studies have determined that pore size is highly dependent on the fiber diameter and packing density of the material [46–49]. The decrease in the pore size for PCL-I40 is most likely due to the decrease in fiber diameter, as many previous studies have noted [37,50]. However, we noted no observable trend and no significant difference among the samples. The morphological and physicochemical properties of the scaffolds affect their rigidity and mechanical strength. The thermal properties of PCL scaffolds were unchanged despite the addition of substantial amounts of iodixanol. The published Tm of PCL scaffolds is approximately 60 °C [51,52], which was slightly higher than those of the PCL-I scaffolds but without any significant trends. Similarly, the difference in Tc of PCL and PCL-I scaffolds was insignificant. The tensile strength, termed as maximum stress, of the electrospun PCL scaffolds was 4.24 MPa, whereas those of PCL-I5, PCL-I15, and PCLI40 were 4.52, 3.63, and 5.78 MPa, respectively. No statistically significant differences among the scaffolds were observed (p > 0.05, multiple t tests), although PCL-I40, despite its poor crystallinity, had the highest numerical tensile strength. This may be attributed to its smaller fiber diameter which provided it with improved mechanical strength, as described in other studies [53,54]. Our data indicate that the PCL and

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Fig. 1. PCL morphology is generally unchanged by iodine infusion. A–D: SEM images of of PCL (A), PCL-I5 (B), PCL-I15 (C), and PCL-I40 (D) (bar = 20 𝜇m; 1000× magnification). E–H: SEM images of PCL (E), PCL-I5 (F), PCL-I15 (G), and PCL-I40 (H) (bar = 20 𝜇m; 5000× magnification). I–L: Fiber diameter distributions of PCL (I), PCL-I5 (J), PCL-I15 (K), and PCL-I40 (L); there was a marked decrease in the mean fiber diameter as iodine concentration increased. M-P: Pore size distributions of PCL (M), PCL-I5 (N), PCL-I15 (O), and PCL-I40 (P). Table 1 Physicochemical properties of PCL and PCL-I electrospun scaffolds.

PCL PCL-I5 PCL-I15 PCL-I40

% Iodine Theor Actual

Diameter (𝜇m)

0 3.70 10.34 23.53

2.30 2.13 2.02 1.15

ND 3.56 ± 0.10 8.79 ± 0.18 19.75 ± 0.15

± ± ± ±

0.32 0.26 0.49 0.26

Pore size (𝜇m2 ) 31.81± 19.71 36.76 ± 17.21 47.06 ± 27.97 23.21 ± 12.05

PCL-I scaffolds have tensile strengths that are similar to those of native human and porcine arteries [55,56]. 3.2. Radiopacity of PCL-I scaffolds Radiopaque scaffolds would provide a noninvasive modality for monitoring material placement and degradation in real time. In the present study, increasing iodine concentrations conferred greater ra-

% Porosity Calc.

Intrusion

61.9 58.7 69.1 64.4

63.7± 4.0 61.1 ± 9.7 70.2 ± 4.9 66.5 ± 2.4

± ± ± ±

4.0 8.7 3.3 6.6

Tc (°C)

Tm , (°C)

Maximum stress (MPa)

33.18 32.27 33.56 35.48

57.66 56.25 57.87 57.41

4.24± 0.859 4.52 ± 0.665 3.63 ± 0.491 5.78 ± 0.110

diopacity on X-ray and micro-CT imaging (Fig. 2). Compared with PCL alone, PCL-I had higher attenuation in a concentration-dependent manner; the greater the iodine content within the scaffold, the greater the radiopacity. The mean CT intensity of PCL was only −890.83 ± 60.97 HU, whereas PCL-I5, PCL-I15, and PCL-40 had intensities of −630.83 ± 31.77, 236.08 ± 54.14, and 284.35 ± 126.21 HU, respectively. Although the addition of 15 mg of iodixanol significantly increased the radiopacity compared with the addition of just 5 mg of iodixanol, increasing the

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Fig. 2. Greater iodine concentration increases radiopacity. Left: Photographs. Center: X-ray images. Right: micro-CT images (threshold level, ×700 HU).

iodixanol incorporation to 40 mg did not significantly increase the radiopacity further (p > 0.05, multiple t-tests). A line profile was also used to demonstrate the HU values along a line segment transecting a border of each scaffold (Supplementary Fig. 2). Although the maximum HU was higher with PCL-I40 than PCL-I15, the difference in the area under the curve did not change considerably with the additional iodine. In a previous study, we used the wet-dipping method to infuse radiopaque materials within polymer matrices by soaking poly-pdioxanone sutures in 10% dimethylsulfoxide-dichloromethane (DMSODCM) containing 15 mg of 4-iodobenzyl chloride for 24 h [26]. The iodine infusion was only 1.07% ± 0.08% but still yielded a CT intensity of 132 ± 48 HU. This technique depends on the soaking duration and the soaking solution concentration. Although the wet-dipping method adequately imbues radiopacity, the technique is limited by the amount of iodine that diffuses within the interstices of the polymer matrix. Furthermore, increasing the soaking duration and/or the solvent concentration increases the risk of changing the mechanical strength and other properties of the polymer. The present study focused on incorporating iodine within the polymer chain without changing the chemical composition of the polymer. Through the electrospinning method a higher concentration of iodine was incorporated within PCL, providing enhanced radiopacity. Compared with PCL alone, PCL with the optimum loading of 15 mg iodixanol (PCL-I15) showed no significant differences in terms of Tm , Tc , tensile strength, pore size, or porosity. Furthermore, increasing the amount of iodixanol from 15 mg to 40 mg did not significantly increase radiopacity. 3.3. MSC attachment and proliferation Interest in the use of nanostructures for tissue-engineering applications has been driven by biomimicry of the natural extracellular matrix [57–59]. Our results demonstrate that PCL and PCL-I scaffolds can support cell attachment, growth, and proliferation (Fig. 3A and B). Fluorescence microscopy imaging of GFP-MSC incubated on PCL and PCLI scaffolds for 12 h confirmed the spread-out attachment of MSCs to the surface of the scaffolds. Qualitative assessment of the fluorescence demonstrated more enhanced cell attachment with increasing iodine concentration, which may have been due to the smaller fiber diameter and increased hydrophilicity resulting from iodine incorporation. SEM revealed that MSCs interacted well with the scaffolds, with the fiber architecture guiding their growth and spread. No obvious morphological differences in cell-fiber interaction were observed across the different scaffolds, and after 48 h of incubation, cells formed a continuous layer on the surface of all scaffolds. Furthermore, the alamarBlueTM fluorescence assay revealed that cell proliferation achieved with PCL and PCL-I scaffolds is similar to that achieved with tissue culture-treated polystyrene plates. For all samples and controls, the doubling time was around 1.2 days and cell number increased as much as 7-fold after 1 week of incubation. Our results can be explained by differences in fiber diameter and surface area; the same phenomenon was observed by Chen et al. [59] Using electrospun scaffolds of different diameters, they found that cell attach-

ment was higher with smaller fiber diameters which they attributed to greater surface area for adsorption of serum proteins necessary for cell attachment. This advantage, however, was diminished 6 h after seeding, indicating that cell attachment had reached saturation. Similarly, our results suggest a trend towards enhanced cell attachment in smaller fiber diameters as seen in fluorescence microscopy in the initial 12 h of incubation. However, the alamarBlueTM assays we performed after 24 h of incubation and the SEM performed 48 h after seeding did not reveal any significant differences among the scaffolds. 3.4. In vitro blood compatibility Because one of the most common reasons for early graft failure is acute thrombogenicity of synthetic tissue scaffolds, we performed hemocompatibility and coagulation assays to determine the effect of the scaffolds on direct blood contact. The mean hemolysis percentages for the scaffolds (1.76%, 1.47%, 1.79%, and 1.58% for PCL, PCL-I5, PCL-I15, and PCL-I40, respectively), normal saline solution (2.02%), and the negative control (anticoagulated blood with no additives; 1.87%) did not differ significantly (p > 0.05, one-way ANOVA) (Fig. 4A). With a hemolysis percentage less than 5%, both PCL and PCL-I scaffolds are considered hemocompatible based on ISO 10993-4:2017 and ASTM F756:2000 [60,61]. We measured partial thromboplastin time to assess the activation of intrinsic coagulation factors by the scaffolds. Compared to the positive control (glass; normalized to 100%), clotting time of PBS, PCL, PCL-I5, PCL-I15, and PCL-I40 were 204.50%, 208.14%, 187.32%, 183.55%, and 184.93%, respectively (Fig. 4B). While clotting time tended to decrease with increasing iodine concentration, these are statistically comparable to the negative control (no material; 202.97%). Previous studies have shown that coagulation time and surface thrombogenicity are affected by the physicochemical properties of the biomaterial such as hydrophilicity and fiber morphology. In previous studies, increased hydrophilicity was observed to suppress platelet adhesion and activation [62] while smaller fiber diameter increases thrombogenic potential [63]. In our study, the smaller fiber diameters of the iodine-loaded scaffolds may have contributed to the slight decrease in coagulation time. However, this is counteracted by the increase in hydrophilicity from the addition of iodine resulting in no significant difference to the negative control. This suggests potential utility of this material as a hemocompatible scaffold. 3.5. In vivo radiopacity and histological analysis of PCL-I polymer scaffolds We found that PCL-I15 had the best PCL-to-iodine ratio for in vivo investigations as it did not significantly alter the physicochemical properties of PCL while maintaining optimal radiopacity. Subcutaneous implantation of PCL-I15 in C57BL/6 mice showed no abnormal clinical findings. The PCL-I15 scaffolds could be clearly visualized with microCT (Fig. 5A) with signal intensity decreasing steadily over time (Fig. 5B).

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Fig. 3. Iodixanol incorporation does not impede cell attachment, viability, and proliferation. A: Fluorescence and SEM images of GFP-MSCs attached to PCL and PCL-I scaffolds (bar = 20 𝜇m; 10× magnification). Enhanced cell attachment with increasing iodine concentration was observed in fluorescence microscopy (taken after 12 h of incubation) but this was not observed in SEM (taken 48 h of incubation). B: There was no significant difference on the viability and proliferation of cells when attached to tissue culture-treated polystyrene plates (control) or to PCL and PCL-I scaffolds.

Compared with signal intensity 1 day after implantation, the signal decreased by 6.42% after 7 days, 13.71% after 14 days, 30.81% after 28 days, and 53.51% after 56 days, resulting in a linear decrease in radiopacity over time of 0.96% per day. This trend in decreasing radiopacity was compared with the degradation of the implanted PCL-I15 scaffolds using maximum strength as a surrogate marker for scaffold in-

tegrity. The scaffolds were explanted at different time points and maximum stress was measured. The decrease in radiopacity and maximum stress were graphed and the curve fitting followed first-order decay kinetics (Fig. 6A and B). The half-life calculated from CT intensity was 75.86 days (Fig. 6A) while the half-life from maximum stress was 49.04 days (Fig. 6B). The change in radiopacity is also linear at 0.96% per day,

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Fig. 4. Blood interaction effects of iodine incorporation. A and B: Hemolysis (A) and clotting (B) were not significantly different among the negative controls, NSS, PBS, PCL, and PCL-I scaffolds. Positive controls for both experiments were significantly different. (∗ ∗ ∗ ∗ p < 0.0001, multiple t tests).

Fig. 5. Iodixanol-infused PCL retains radiopacity on micro-CT for 56 days. A: X-ray and micro-CT images of subcutaneously implanted PCL (blue) and PCL-I15 (red) scaffolds over time (window level 1800/800). B: CT attenuation of PCL-I15 and PCL. (∗ p < 0.00001, Student’s t-test).

Fig. 6. Mechanical strength of PCL-I15 in vivo decreases with radiopacity. A: CT intensity (HU) of PCL-I15 scaffolds over time. B: Tensile strength (maximum stress) of explanted PCL-I15 scaffolds over time. C: Calculated % change of based on radiopacity and mechanical stress. Both trends follow first-order decay kinetics with a degradation half-life of 49.04 days based on maximum stress and 75.86 days based on CT signal intensity. Calculation of % change shows linear curves. No significant differences were observed between both data sets (p > 0.48, two-way ANOVA).

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Fig. 7. Histology of explanted scaffolds with surrounding tissue. Left column: Hematoxylin and eosin-stained sections of PCL at 4× magnification (left images; bar = 600 𝜇m) and 20× magnification (right images; bar = 200 𝜇m). Right column: Hematoxylin and eosin-stained sections of PCL-I15 at 4× magnification (left images; bar = 600 𝜇m) and 20× magnification (right images; bar = 200 𝜇m). In PCL, the inflammatory cell infiltrate consisted of a few neutrophils and macrophages (arrows). In contrast, in PCL-I15, a robust neutrophilic infiltrate progressed to a mixture of neutrophils (arrows) and increasing numbers of activated macrophages and multinucleated giant cells (arrowheads). By day 56, the PCL-I15 scaffolds were surrounded by thicker maturing fibrous tissue (wavy arrows), compared to PCL.

while that of maximum stress is 0.59% per day (Fig. 6C). The large difference in the values calculated may be attributed to the small number of animals used and inter-animal variability, as well as the use of tensile strength for measurement of polymer degradation. Other more sensitive tests for polymer chain degradation, such as thermogravimetric analysis or gel permeation chromatography could potentially be used to confirm our results [64–67]. In spite of the large numerical difference, statistical analysis revealed no significant difference between the degradation half lives calculated from radiopacity and maximum stress (p = 0.48, twoway ANOVA). This indicates that radiopacity/HU values can be used to estimate the degradation and integrity of the scaffold over time. Histological analysis showed that both the PCL and PCL-I15 scaffolds induced acute inflammation (Fig. 7 and Supplementary Fig. 3). One day

after implantation, PCL scaffold was infiltrated by only a few neutrophils and even fewer macrophages. Seven days after implantation and thereafter, the porous portions of the scaffold were expanded by clear space (consistent with edema and with degradation of the structure) and had small numbers of macrophages. By day 56, some PCL scaffolds were partially surrounded by a very thin layer of fibrous tissue. In contrast, the PCL-I15 scaffold was surrounded and infiltrated by viable and degenerating neutrophils within 1 day after implantation. The neutrophilic infiltrate subsided over several weeks. By day 7, the scaffold was surrounded and infiltrated by many activated macrophages, showing progressive development of a layer of macrophages and multinucleated (foreign body) giant cells. By day 56, most PCL-I15 scaffolds were partially surrounded by a thick layer of maturing fibrous tissue, reflecting prior mesenchymal

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proliferation, granulation tissue formation, and extracellular matrix remodeling. These results indicate that the addition of iodixanol increases acute (neutrophils) and chronic (macrophage/giant cells) inflammatory responses and leads to the development of a more robust fibrotic capsule around the scaffold. The host inflammatory response to a biomaterial typically includes the arrival and fusion of macrophages, ultimately leading to a foreign body reaction (FBR) and often fibrotic encapsulation [68]. Historically, strategies for promoting the biocompatibility of synthetic implants focused on avoiding FBRs [69,70]. In contrast, recent strategies have been geared towards controlling biomaterial-induced immunomodulation to shift the local environment from a pro-inflammatory state to an antiinflammatory state, thereby promoting wound healing and tissue regeneration [71]. Modulating the FBR has given way to the concept of regenerating tissue in situ by modifying biomaterial surface properties to promote a suitable host regenerative response in the form of tissue remodelling [71]. The physical properties of scaffolds may influence the phenotype of recruited macrophages. Thin fibers and small pores are associated with “frustrated phagocytosis,” which lead to a pro-inflammatory response, whereas thick fibers and larger pores cause more pronounced cell infiltration, promoting an anti-inflammatory microenvironment [72,73]. Compared with the PCL scaffolds, the PCL-I15 scaffolds had more intense acute and chronic inflammatory cell infiltrate and denser encircling fibrosis. Although the iodixanol-loaded scaffolds had thinner fibers, enhanced hydrophilicity from the addition of iodixanol may have allowed a more robust host response including increased cell recruitment and infiltration that progressed to thicker layers of mature fibrous tissue formation. This suggests that the iodixanol-associated changes on the physicochemical properties of the scaffold may be able to modulate the host response from a pro-inflammatory (increased cell recruitment and infiltration) to a pro-regenerative state (thicker fibroplasia). However, further evaluation is needed to elucidate the exact mechanism for this phenomenon. Furthermore, degradation of PCL scaffolds implanted in vivo was not observed within the duration of the study. The half-life of the maximum stress of explanted PCL scaffolds was calculated to be 491,263 days using a curve for first-order decay kinetics (Supplementary Fig. 4), signifying that there was no degradation observed within the time frame. This is in accordance to degradation rates of ~2 years reported in literature for PCL [74]. In contrast, the degradation half-life calculated for PCL-I15 was 49.04 days. The changes in degradation rate may be explained by the more robust inflammatory changes observed with the addition of iodixanol. This shows that the degradation rate of the scaffolds is tunable based on the physicochemical characteristics and the host immune response, leading to a variety of different applications for scaffolds wherein length of required structural support and extent of in situ tissue regeneration are considered. 4. Conclusions We used an electrospinning method to fabricate a PCL-iodixanol nanocomposite that is radiopaque, biodegradable, and biocompatible and that supports cell adhesion and infiltration. The addition of iodixanol provided the scaffolds with advantageous properties, including smaller fiber diameter, higher surface area, and hydrophilicity without increasing toxicity to cells and tissues both in vitro and in vivo. Most importantly, our in vivo studies showed that a time-dependent decrease in radiopacity could be quantified using HU values, which may be used to monitor the integrity of the scaffold over time. In conclusion, we have developed a radiopaque biodegradable scaffold that could be monitored noninvasively at real-time allowing, more effective utilization of biodegradable devices for tissue-engineering applications. Declaration of Competing Interest The authors have no conflicts to declare.

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Acknowledgments The authors thank Joe Munch in Scientific Publications in MD Anderson’s Research Medical Library for editing the manuscript. This work was supported in part by the National Institutes of Health – National Heart, Lung, and Blood Institute (1R01HL141831; to M.P.M.), and the Department of Science and Technology, Philippine Council for Health Research and Development (to J.D.P). MD Anderson’s Small Animal Facility and Small Animal Imaging Facility are supported by the NIH through MD Anderson’s Cancer Center Support Grant (CA016672). SEM was performed under the supervision of Dr. James Gu at the Electron Microscopy Core at Methodist Research Hospital. We thank Dr. Frederick Lang and Joy Gumin (MD Anderson) for providing MSCs; Drs. Ennio Tasciotti and Jonathan Martinez (Department of Nanomedicine, Methodist Research Hospital) for assistance with the ICP-OES testing and analysis; and Dr. John Manongdo (University of Houston) for assistance with the DSC/TGA testing and analysis. Supplementary materials Supplementary material associated with this article can be found, in the online version, at doi:10.1016/j.mtla.2020.100874. References [1] M. Jafari, Z. Paknejad, M.R. Rad, S.R. Motamedian, M.J. Eghbal, N. Nadjmi, A. Khojasteh, Polymeric scaffolds in tissue engineering: a literature review, J. Biomed. Mater. Res. B 105 (2) (2017) 431–459. [2] S. Khorshidi, A. Solouk, H. Mirzadeh, S. Mazinani, J.M. Lagaron, S. Sharifi, S. Ramakrishna, A review of key challenges of electrospun scaffolds for tissue-engineering applications, J. Tissue Eng. Regen. M 10 (9) (2016) 715–738. [3] B. Dhandayuthapani, Y. Yoshida, T. Maekawa, D.S. Kumar, Polymeric scaffolds in tissue engineering application: a review, Int. J. Polym. Sci. 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