Atmospheric pressure plasma jet: A facile method to modify the intimal surface of polymeric tubular conduits Free

Low temperature plasma (LTP) at atmospheric pressure is known to initiate chemical reactions through the production of various species, including oxygen radicals, reactive charged species, ions, and UV light. These active species can be exploited for biomedical, agricultural, and food safety applications. Several well established examples include using LTP for bacterial inactivation in water and other media, treatment of seeds for accelerated germination and growth, and the sterilization of seeds.^1–3 LTP can potentially be very toxic to cells (due in part to changes in pH) and thus can selectively kill bacteria, fungi, and viruses.^4–6 This highly reactive nature of LTP makes it an attractive tool for the modification of materials.

There has been extensive interest and work in utilizing a variety of plasmas for the modification of biomaterials.^7–9 Many techniques for surface modification of biomaterials have been reported that include physical, chemical, and biological treatments; however, chemical methods often result in undesirable byproducts or residues; thus, LTP offers a readily accessible surface modification method with innocuous byproducts and no residues.^7–9 Furthermore, LTP modification allows for facile tuning of the surface chemistry and properties while leaving the bulk properties of the material unaltered.⁸ Of particular interest in the field of tissue engineering is the fabrication of small caliber vascular grafts for use in regenerative cardiovascular medicine.¹⁰ Electrospun tubular conduits have been established as viable constructs for vascular grafts and have been proposed for neural tissue engineering as conduits for peripheral nerve grafts or for spinal cord repair for guiding neurons.^10–14 The surface properties of vascular graft materials such as surface chemistry, surface charge, wettability, and roughness play an important role in protein adsorption and cell-proliferation, all of which are important factors for tissue integration.^10–13 Similarly, these properties possess a key role in preventing the process of platelet adhesion and activation, thus preventing clotting on the blood contacting intimal surfaces in vascular grafts.¹⁴ Recent literature suggests that the current state of the art for biomaterials surface modification includes techniques such as plasma assisted chemical vapor deposition and “click” chemistry.^15,16 Both methods are offering advantages for surface modification such as ease, speed, and mild conditions. Recent work in vascular graft engineering suggests approaches based on mimicking native vessel tissues with fibrous scaffolds, using mesenchymal stem cells for enhancing hemocompatibility, and exploiting hierarchical structures in electrospun scaffolds.^17–19 

Achieving uniform surface functionalization and preventing significant deteriorative effects are two of the major challenges in 3D fibrous tissue scaffold engineering. Processing may negatively impact the mechanical properties, biodegradation, and/or biointegration of these implants and scaffolds.^10–13 The major challenges in plasma treatment of soft matter are the depth of plasma treatment, aging of the modified surface after plasma treatment, and the degradation-effect on the substrate. The depth of the treatment is easily controlled by experimental factors such as power and time.⁹ Several strategies are reported to mitigate the aging and degradation effects encountered in plasma modified polymers and include increasing the degree of crosslinking, exploiting steric hindrance, employing surface barrier coating, and postplasma processing.²⁰ 

Current blood vessel substitutes, especially small caliber substitutes (<6 mm in diameter), often suffer from low patency and can occlude within a short period of time, thus necessitating a variety of approaches that have been adopted to improve their biocompatibility.¹⁰ These approaches include transplanting a monolayer of viable endothelial cells onto the luminal surface of a graft prior to implantation and the engineering of synthetic tissues.^{10–13,17–19} Materials such as nanofibrous bioresorbable polymers like polycaprolactone (PCL) and polyglecaprone [both with and without physiological extracellular matrix (ECM) such as collagen/laminin/elastin] have been experimentally employed in cell transplantation and for regeneration of vascular and neural tissues.^10–14,21 Retention of adsorbed ECM proteins and protecting their morphology (by preventing denaturing) and retention of proteins secreted by cells on the surfaces of synthetic polymer scaffolds for the duration of the regeneration period (3–6 months typically) is a challenge.^21–23 Wet chemical methods could be employed for surface modification using coupling agents or cross linkers. However, they can have cytotoxic effects that hinder cellular growth and other bio-functions, thus necessitating additional postmodification washing and cleaning steps. Plasma treatment therefore provides a convenient route of modifying vascular tissue scaffolds for enabling increased protein adsorption. Thus, enhanced cell proliferation (perhaps utilizing stem cells or induced pluripotent stem cells, both commonly used in tissue engineering) can be achieved.

We proposed using an external atmospheric-pressure plasma jet (APPJ) modification as a facile method to functionalize the intimal surface of 3D tubular grafts. APPJs are unique in that they are nonequilibrium plasmas having low gas temperatures (around 300 K) but high electron temperatures (around 20 000 K) and high charged particle densities.⁹ Additionally, they do not require the use of expensive and complicated vacuum equipment. APPJs are typically used in biomedical and material processing fields for applications such as bacterial inactivation^24–27 or thin film coatings.^28–32 The low temperature of these plasmas makes them ideal for use with temperature sensitive materials without altering their bulk mechanical properties. Atmospheric pressure air/ammonia plasma techniques offer a readily accessible and efficient route for soft-matter surface modification allowing tunable functionality.^7–9 Among the options possible are the ability to retain the quality of native proteins absorbed or attached to the scaffold surface, access to antithrombotic regulatory mechanisms, and enhanced endothelialization of the intimal surfaces of the vascular graft materials. The plasma treated surfaces were characterized by Fourier transform infrared spectroscopy (FTIR), x-ray photoelectron spectroscopy (XPS), scanning electron microscopy (SEM), surface contact angle, and tensile testing.

The fabrication of our small caliber (4 mm inner diameter) tubular scaffolds was accomplished by electrospinning. The working solution was prepared as 15% w/w PCL (PCL Pellets from Absorbable Polymers, Birmingham, AL) solution in 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP). This particular concentration and our electrospinning parameters were chosen in accordance with our previously reported work, in order to create suitable 3D conduits for small caliber vascular graft or spinal cord regeneration applications.^10–13 A custom built, fully computerized electrospinning setup with a translationally mobile spinneret was used (employing a high voltage source, Model M826 by Gamma High-Voltage Research, Ormond Beach, FL and additionally equipped with a syringe pump, model PHD2000 by Harvard Apparatus, MA). The PCL-HFIP solution was fed through a 25G needle with a feed rate of 1.5 mL/h using a voltage of 15 kV. In order to achieve a tubular scaffold, a controlled rotating mandrel (4 mm in diameter and 15 cm in length) kept at a distance of 15 cm, was used as the collector target for electrospinning. The programmable stage was set to cover 15 cm range with a lateral speed of 30 mm/s. The deposited fibers were collected onto a 4 mm diameter 303 stainless steel mandrel rotating at 400 rpm. The constant translational motion of the needle assists in the uniform distribution of fibers on the scaffold. After fabrication, the scaffolds were removed from the collector after soaking in water and degassed for 48 h at room temperature.

For the plasma treatment, the electrospun PCL tubes were vertically positioned and aligned to maximize exposure of the jet-plume on the intimal surface. We utilized a previously constructed dielectric barrier discharge (DBD)-APPJ system which was custom built for surface modification at our partner facility at UAH facility.³³ The plasma source is constructed of a nested pair of quartz tubes (Fig. 1). The inner tube has a 1 mm inner diameter (ID) and a 2 mm outer diameter (OD) and is closed at one end. The outer tube has a 3 mm ID and a 6 mm OD and is open at both ends. The tubes are held in a polymer compression Tee fitting to provide an air-tight seal. The working gas, helium in this case, is fed through the third leg of the fitting and flows between the nested tubes. A steel wire is inserted inside the inner tube while a steel collar is placed around the outside of the outer tube. The wire and collar serve as the two electrodes in the system that provides the pulsed direct current (DC) power to generate the low-temperature plasma. The source design allows the plasma to flow between the quartz tubes and out the end without physically contacting the electrodes. This prevents arcing from the high voltage electrode, through the plasma, to the treated material. The power system is comprised of a 10 kV DC power supply, a high-voltage pulse generator, and a digital delay generator. For the experiments discussed in this paper, the plasma was generated at 8.5 kV, 1000 μs pulse width, and 6 kHz pulse frequency. A flow of 2 l/min of bottled helium was provided via a digital mass flow controller. These settings generated a room temperature (300 K) plasma jet of approximately 4 cm long. The tubular conduits were vertically positioned and aligned to be irradiated with jet plasma (4 mm diameter) on the intimal surface for 5 min through both ends to achieve a uniform irradiation with plasma throughout the intimal surface for 10 min total exposure time.

As fabricated, the tubular scaffolds were tested for biomechanical properties at ambient conditions. Tensile specimens were prepared by cutting the scaffolds into rectangular strips (3 × 10 mm) in accordance with ASTM standard D882 for tensile testing of thin film plastics. The sample thickness was determined by the averaging values found using a contact method on a thermal mechanical analyzer (TMA Q400, TA Instruments). A dynamic mechanical analyzer (DMA) (RSA-G2 from TA instruments) with a specific tensile fixture was used to determine the tensile properties of the samples (n = 6) at a constant force mode. The specimen was clamped to the tensile fixture of DMA machine vertically, and the gauge length of the exposed sample was measured using the electronic digital caliper. The samples were tested uniaxially using a 35 N load cell at a ramp of 0.1 N mm⁻¹. Data from each experiment were analyzed using the graphic analysis software (ta universal data analysis), and the elastic modulus, percent elongation to failure, and ultimate strength were determined from the generated stress-strain curves.

The structure and morphology of scaffolds were characterized using SEM, atomic force microscopy, and Keyence 3D-optical microscopy. Tubular scaffolds were cut, opened, and sputter-coated with Au-Pd and observed using a FE-SEM (Quanta FEG 650 from FEI, Hillsboro, OR), and images were taken at different magnifications to understand the surface morphology of APPJ treated and nontreated grafts. SEM micrographs were analyzed by an image-analyzer (Image-ProPlus, Media, and Silver Spring, MD) to obtain a statistical measurement of the fiber diameter.

FTIR and XPS were employed to elucidate the chemistry. The combined data from FTIR and XPS analyses were used to confirm the presence of oxygenated surface functional groups on the APPJ treated samples. The FTIR technique was utilized to confirm the presence of surface functionalities and absorbed/bound-water on graft. The Bruker alpha FTIR spectrometer was used with the attenuated total reflection (ATR) mode to acquire absorbance spectrum (64 scans per sample ranging from 4000 to 400 cm⁻¹). XPS characterization of APPJ treated samples (intimal surface) was obtained using a Phi 5000 Versaprobe made by Phi Electronics, Inc. (Chanhassen, WI). The x-ray source of this instrument is a monochromatic, focused, Al K-alpha source (E = 1486.6 eV) at 25 W with a 100 μm spot size. The Mg anode (λ = 1253.6 eV) was used at 300 W, and a barium oxide neutralizer eliminated the charging of the sample. The survey scans (four scans averaged per analysis) were obtained using a pass energy of 187.5 eV with a step size of 0.5 eV. The high resolution scans (eight scans average per analysis) were obtained with a pass energy of 23.5 eV and a step size of 0.1 eV.

The surface glycerol contact (GSA) angle technique was utilized to understand hydrophilicity of the scaffolds. The samples were cut into 1 × 1 cm strips (n = 3) and were then mounted onto a glass slide. Contact angles were measured using the sessile drop method²⁹ at the room temperature. The glycerol droplet size was 5 μl. The video contact angle instrument and software (VHX 600E, Keyence) were used to capture the video and snapshots to determine the contact angle.

The effect of APPJ treatment on bound water hydration was studied as a complementary method to show the increased hydrophilicity. The tubular scaffolds (treated and nontreated) were immersed in water for 12 h, then dried for 24 h, and cut-open and were evaluated by using infrared spectroscopy (ATR-FTIR) for water bound to the intimal surface due to hydrogen bonding of the water to the oxygenated functionalities. The oxygen enrichment is introduced by the APPJ treatment.

The electrospinning technique allows efficient fabrication of the tubular scaffolds. The thin nanoscopic and microscopic scale fibers produced by our electrospinning technique mimic structurally the ECM of protein fibers found in native tissues.^10–13 Figure 2 shows a representative scaffold used in this experiment for the APPJ modification of the intimal surfaces. The treatment allowed for changes in the surface chemistry, wettability, and water affinity; however, the bulk mechanical properties remained unchanged as a result of the treatment. Sections III A–III C describe in detail the results of the mechanomorphological characterization, surface chemistry characterization, and water binding affinity studies. Overall, APPJ presents an efficient method for intimal surface modification as evidenced by the data.

To mimic the native extracellular matrix, the tissue-scaffold should be fibro-porous and the fiber size should be nano- to microscale.^34,35 These biomimetic fibers can provide accelerated cell proliferation for regenerative growth. SEM images in Fig. 3 show the morphology of the intimal surface of the PCL scaffolds before and after APPJ treatment. The SEM images confirm the fiber size and bead-free fiber formation, as well as the fiber orientation. All images showed nano-/microfibers with diameters in bimodal distribution (Fig. 4) of maxima around 275 nm (for 100–600 nm) and around 1.65 μm (for 1–2.5 μm). PCL surfaces before plasma treatment exhibited a relatively stressed and ordered orientation [Fig. 3(a)] compared to the randomly oriented and relaxed fibers of APPJ treated surfaces [Figs. 3(b)–3(d)]. It is evident that the APPJ treatment did not change the fibrous morphology or diameter distribution. However, the fiber diameters are comparable to the upper range of protein fiber-sizes. For example, collagen fibers are in the range of 50–500 nm size to 3 μm.³⁵ The porosity for all three ratios calculated from apparent density measurements¹⁰ was in the range of 70%–80%.

In Fig. 5, the mechanical properties of PCL tubes before APPJ treatment are given. Over all data showed that PCL tubes exhibited tensile strength values comparable to that of human native arteries (1–3 MPa).^10–13 Specifically, small-caliber PCL scaffolds exhibited an ultimate tensile strength of 2.60 ± 0.2 MPa, with a soft-tissue-like elastic modulus of 227 ± 11 kPa and up to 150% strain under the uniaxial tension test.

APPJ treated intimal surfaces were characterized by XPS (Figs. 6 and 7) and surface contact angle. The data are summarized and given in Table I. As stated earlier in Sec. I, one of the major issues that must be addressed for nano/microfiber scaffolds intended for growing cells or for blood contact surface is their highly hydrophobic surface. For example, being hydrophobic in nature, electrospun PCL exhibited a high water contact angle values as evident from reported literature (122°–135°).²⁹ RF plasma modification under oxygen atmosphere (10 min) has shown to decrease the value to 20°. Since the uniformity of modification is important on the grafts’ surface, flipping of the grafts’ ends (after 5 min) ensured the plasma interaction uniformity. Further, XPS along the axial direction (analyses of both the middle and end regions) reveal no variation in chemistry throughout the intimal surfaces of the grafts. Another important factor in APPJ treatment is aging.²⁰ XPS analysis (Fig. 7) immediately after modification reveals the presence of nitrogen (N 1s around 400 eV), in addition to carbon and oxygen. However, after several days (1 week), the nitrogen peak disappeared and only oxygen and carbon peaks remained. This is one of the immediate effects of aging and further effects can also be experienced by changing the nature of the modified surface, including recovery of the hydrophobic character after plasma modification.²⁰ Figures 6 and 7 gave the change in the atomic ratio of carbon to oxygen on the surface of the scaffolds calculated from the XPS peak area ratios of C 1s to O 1s. Owing to the introduction of oxygenated functional groups, there was a corresponding decrease in the percentage of saturated hydrocarbon C 1s peak at 285.0 eV for APPJ treated PCL.

Not only changes in surface chemistry but also fiber roughness and alignment (orientation) can significantly influence the contact angle of liquid on the surfaces of plasma treated polymers.³¹ Oxygen has been demonstrated to have a particularly dramatic effect on the contact angle measurements of polymer surfaces treated with oxygen-containing plasmas.³⁶ Similarly, Table I shows treatment 2 (He/Air) to have the lowest contact angle of 42° ± 1.70. The other plasmas, treatments 1 and 3, have a decreased concentration of oxygen due to the other component (NH₃) and reflect contact angles of 58° ± 2.22 and 60° ± 3.69, respectively. It is also important to note the slight change of contact angle for treatment 1 freshly modified (analyzed <4 h postmodification) and treatment 1 slightly aged (analyzed after several days). This is attributed as one of the effects of short term aging, but the effect is minimal during short time frames and only becomes very noticeable after long periods.^20,37–39 Since the water contact angle for these highly hydrophilic and porous surfaces is difficult to capture with a camera due to the virtually instantaneous absorption of the droplet into the surface of the scaffold, we used glycerol to test relative hydrophilicity and to obtain quantifiable contact angles as prescribed in literature.³⁶ Owing to the fact that the LTP treatment, generally, serves as an effective method to introduce various hydrophilic functionalities (–OH, –COOH, NO, NH₂, etc.) on the surface, depending on the plasma gas mixture, the data are conclusive that the surface is indeed modified and the hydrophilicity is increased. PCL treated with He/air and He/air/NH₃ exhibited maximum surface modification and wettability as per XPS and contact angle data. Also, the O/C ratio of PCL increased from 0.30 to 0.50 with the introduction of air and NH₃ in He gas.

From the data presented, it can presumed that LTP treatment produced and bombarded free radicals on the surface which in turn reacted to form peroxides, oxoperoxide, ozonide, or other highly reactive oxygen species. These enriched the surface in oxygen functionalities as indicated by the XPS data. During the initial period of time after the modification, the surface experiences some aging effects as observed with the presence than disappearance of the nitrogen functionalities (Fig. 7).

FTIR analysis was conducted to understand the differences in vibrational spectra of the APPJ treated samples compared to control (Fig. 8). IR spectra for PCL showed prominent characteristic bands for aliphatic methylene stretching at 2941 and 2865 cm⁻¹ and a carbonyl stretching at 1727 cm⁻¹. Additional bands included C-O and C-C backbone stretching due to the amorphous phase at 1157 cm⁻¹ and the crystalline phase at 1293 cm⁻¹. The peaks at 1240 and 1170 cm⁻¹ can be assigned to C-O-C stretching, and the peak at 1190 cm⁻¹ represents the OC-O stretching. APPJ treatment did not produce any drastic change in IR vibrational peaks.^7–9 However, the plasma treated samples showed an increase in the peak intensity in the range of 3600–3300 cm⁻¹, indicating hydrogen bonded bound-water presence (Fig. 8). This is attributed to the enhanced increase in the hydrophilicity which in turn increased the absorption and retention of water to the highly hydrophilic surface for the APPJ treated PCL scaffolds. This effect is important for biomaterials engineering for the following reasons. Hydrogen bonding is one of the most prominent interactions in protein structures and contributes significantly to the stability of secondary structures (α-helices and β-sheets, etc.) and to the overall stability of adsorbed protein on the biomaterial/scaffold surface.^10–13 Water is responsible to aid in forming hydrogen-bonded bridges in ECM proteins to retain their native morphology/conformation. This would help contribute to the native, and thus favorable, structure of ECM being retained. However, further long term aging studies are required to conclude the surface functional group stability which can have a significant effect on the surface wettability, bound water content, as well as biological response. This effect on some of the challenges regarding ECM protein adsorption, platelet adsorption, and activation on modified surfaces as published elsewhere.^21–23,40

The enhancement of surface wettability for polymeric grafts and scaffolds is highly desirable. These properties are required for improved protein/surface and cell/scaffold interactions, and for enhanced cell proliferation. Specifically, it requires in PCL conduits for vascular graft applications where complete endothelialization is needed within a short duration (approximately 3–4 days). Conventional wet-chemical methods (cross-linking or copolymerization with different monomers) are harsh and can negatively affect the biomechanical, biodegradation properties, and biocompatibility of the scaffolds. APPJ produced by DBD using controlled gas flow (He/air or He/air/NH₃ gas mixtures) into ambient atmosphere, combined with 8.5 kV pulsed DC signal, proved effective to modify the intimal surface of tubular scaffolds, without jeopardizing their bulk properties. For future work combining the APPJ treatment with some of the other state of the art methods (chemical vapor deposition and click chemistry) will aid in the overall goals of improving polymeric grafting technology.

The authors acknowledge the U.S. National Science Foundation (NSF) grant (No. NSF-EPSCoR CPU2AL) under a Cooperative Agreement No. OIA-1655280 for supporting this work. Thanks are also due to Ryan Gott for assistance with the APPJ setup (UAH Propulsion Research Center) and to W. Monroe for SEM instrumental assistance (UAB Scanning Electron Microscope Facility).