Elucidation of bio-inspired hydroxyapatie crystallization on oxygen-plasma modified 3D printed poly-caprolactone scaffolds

Abstract

Bioapatite formation in bones is a slow process starting with deposition of calcium phosphate and then its nucleation and crystallization into hydroxyapatite crystals. If the same process can be replicated on tissue engineered scaffolds, it will result in the formation of biomimetic bone constructs that will have comparable mechanical properties to native tissue. To mimic the same process on 3D printed polycaprolactone (PCL) scaffolds oxygen plasma treatment was performed to modify their surface chemistry. The attenuated total reflectance-fourier transform infrared (ATR-FTIR) analysis showed formation of carboxyl groups on the PCL surface with corresponding increase in roughness as analyzed by atomic force microscope (AFM) studies. A biomimetic acellular mineralization procedure was then utilized to deposit calcium minerals on these scaffolds. Though amorphous calcium phosphate was deposited on all the scaffolds with highest amount on PCL scaffolds with tricalcium phosphate (TCP), biomimetic hydroxyapatite crystals were only formed on oxygen plasma treated scaffolds, as shown by X-ray diffraction (XRD) analysis. The −COOH groups on the plasma treated scaffolds acted as nucleation sites for amorphous calcium phosphate and the crystal growth was observed in the (211) plane simulating the crystal growth in developing bones. The ATR-FTIR study demonstrated the carbonated nature of these hydroxyapatite crystals mimicking that of bioapatite. The electronegative −COOH groups mimic the negative amino acid side chains in collagen Type I present in bone tissue and the carbonated environment helps in creating bioapatite like deposits. The present study demonstrated the important role of PCL surface chemistry in mimicking a bone like mineralization process in vitro. This work details novel insights regarding improved mineralization of 3D printed PCL scaffolds useful for the development of more biomimetic bone constructs with improved mechanical properties.

1. Introduction

Bone is an atypical tissue composed of a nanocomposite between proteins and mineral crystals. The extracellular proteins mainly include collagen type I which acts as a template for the crystallization of calcium phosphate in the form of carbonated hydroxyapatite (HA) [1]. The crystals are formed between aligned collagen I fibrils that subsequently grow in the c-axis plane longitudinally along the long axis of collagen fibrils [2]. These nanocrystallite HA platelets have very high surface area [3]. Developmental studies have demonstrated that amorphous calcium phosphate (ACP) is the major mineral phase in newly formed bone tissue in fish fins, which is followed by subsequent crystallization by nucleation on the collagen type I fibrils to form nanocrystallite composite that is responsible for the mechanical properties of bone tissue [4]. Amorphous calcium phosphate can get deposited on any biomaterial surface easily by immersion in simulated body fluid (SBF) in vitro, which has been taken as a gold standard for bioactivity. This ACP deposition is not a virtue of the surface properties but due to the supersaturated ionic nature of the SBF solution, thus calcium phosphate will get deposited on any surface over time [5]. The more important aspect is the crystallization of this ACP into hydroxyapatite which can impart bone like material properties to the developing tissue construct. However, there is a significant gap in knowledge regarding the surface chemistry, roughness and its effect on the crystallization phase of the deposited ACP. Thus, current research has focused intently on developing materials that can support a similar biomineralization process, which will help in development of biomimetic bone implants.

PCL is a biodegradable aliphatic polymer which has been very widely used as a tissue engineering scaffold material. Its wide utilization is due to its excellent mechanical properties, low immunogenicity, and minimum toxicity [6]. It is a thermoplastic material with excellent inherent moldability which makes it ideal for manufacturing 3D printed scaffolds for tissue engineering while the excellent mechanical properties of these scaffolds make them suitable for bone tissue engineering specifically [7]. However, the hydrophobic nature and low surface energy of its surface makes it poor substrate for initial cell attachment and later cell spreading [8]. Another disadvantage of PCL is the absence of free functional groups or bioactive motifs that can interact with cells or extracellular material. Thus, to increase the cellular attachment potential of PCL scaffolds various surface modifications have been investigated. These include RGD decoration [9], surface hydroxylation [10] and composites with gelatin & collagen [11]. The most widely studied technique for surface modification of PCL scaffolds is oxygen plasma etching, which has been demonstrated to increase mineralization on these scaffolds when cultured with cells in vitro.

The plasma etching process has been widely adopted for surface modification of biomaterials as it changes the roughness, increasing the hydrophilicity of the biomaterials. The plasma treatment process also modifies the surface chemistry of the biomaterials without affecting the bulk physical and mechanical properties of the materials. Both the modification of surface roughness and chemistry results in enhanced cellular compatibility of biomaterials in terms of cellular adhesion and proliferation. Kim et al. demonstrated that oxygen plasma treatment of PCL scaffolds up to 4 h; which were subjected to in vitro mineralization demonstrated increased cellular attachment of pre-osteoblast cells (MC3T3-E1) [12]. The deposited calcium was quantified after cell culture, but not after SBF immersion. Thus, it can’t be predicted if the surface chemistry or the cellular behavior on this surface resulted in higher calcium deposition. Another study demonstrated that templated plasma treatment of PCL scaffolds increased the cell viability, alkaline phosphatase (ALP) activity and calcium mineralization on these scaffolds [13]. ALP activity was used to measure calcium mineralization that was significantly more in the plasma treated scaffolds, which was attributed to induction of signaling by the nanoscale roughness of templated plasma treatment. While ALP activity is an indicator of osteogenic differentiation and activity in the cells, it can’t be used for quantification of calcium mineralization. Thus, none of the studies have investigated the intrinsic potential of plasma treated PCL scaffolds as a template for mineralization, which should be an essential property for a material to be considered for bone tissue engineering.

The current study therefore specifically investigated the potential of oxygen plasma treated 3D printed scaffolds as template for mineralization, as the surface chemistry plays a pivotal role in nucleation and growth of mineral crystals. We created differential surface chemistries by treating 3D printed PCL scaffolds for 10 and 30 min with oxygen plasma, while the unmodified PCL and PCL incorporated with TCP was used to compare the effect of chemistry on the mineralization process. TCP has been widely used as a bone cementing and substitute material owing to its bioactivity and bone regeneration potential [14]. The modification in surface chemistry of PCL was then analyzed and could be later correlated to the mineralization potential of these surfaces. These PCL scaffolds with differential surface chemistries were then immersed in physiologically relevant Dulbecco’s Modified Eagle Medium (DMEM) media instead of SBF for 7 days in humidified carbonated atmosphere to mimic bone development conditions in vivo. The deposited calcium minerals were then qualitatively analyzed for crystal structure and quantified through micro-computed tomography (Micro-CT). The outcomes gave an important insight into the mineralization process occurring on 3D printed PCL scaffolds for bone regeneration.

2. Methods & materials

2.1. 3D-printing of PCL scaffolds

PCL (MW ≈ 50,000, Perstorp, USA) filament was made with a custom extruder set at 53 °C and pulled with constant tension to maintain a 1.75 mm diameter. TCP-PCL filaments contained tricalcium phosphate in the ratio TCP:PCL::20:80. The filament was fed to a Flashforge Dreamer 3D printer that printed circular porous scaffolds with 6 mm diameter and 2.75 mm thickness with a nozzle temperature of 70 °C, bed temperature of 37 °C and nozzle diameter of 400 μm.

2.2. Oxygen plasma treatment

The PCL scaffolds were processed with oxygen plasma by using Technics 85 Reactive Ion Etcher at a frequency (50 kHz), power (34 W), pressure (5.3 × 10⁻¹ Torr) for 10 min and 30 min for the respective plasma itched groups (10 OP and 30 OP). The plasma chamber was first cleaned by operating an empty cycle for 10 min before the samples were subjected for plasma treatment. The scaffolds were treated on both sides for half the total time for each oxygen plasma treated group. The abbreviated names of the four scaffold groups used in this study are listed in Table 1, that will be used now onwards. Fig. 1 explains the schematic for formation of the fours test groups of 3D printed PCL scaffolds with differential surface chemistry.

Table 1.

Abbreviations for the 3D printed PCL scaffold groups.

Group	Name
Only PCL scaffolds	PCL
10 min oxygen plasma treated PCL scaffolds	10 OP
10 min oxygen plasma treated PCL scaffolds	30 OP
TCP embedded PCL scaffolds (TCP:PCL::20:80)	TCP-PCL

Fig. 1.

Schematic for the development of 3D printed PCL scaffolds with differential surface chemistry. PCL scaffolds were printed and then treated with oxygen plasma for different exposure times to make plasma etched PCL scaffolds. While the TCP-PCL scaffolds were made by making a hybrid filament of TCP-PCL in the ration 20:80 and then printing it through a 3D melt extrusion printer.

2.3. Scanning electron microscopy (SEM)

For SEM analysis, all samples were air dried and coated with gold using a gold sputter coater (Emitech SC7620, UK) at 25 mA for 1 min, to form a coating of approximately 15–20 nm thickness [15]. The PCL scaffolds with different surface modifications were imaged using SEM at an accelerating voltage of 20 kV (Zeiss EVO MA10 SEM).

2.4. Energy dispersive X-ray emission (EDX)

To confirm the elemental composition of surface deposition on the matrices (n = 3 per group), EDX analysis was performed on carbon coated samples. Imaging was performed using Zeiss EVO 50 high definition SEM and the respective Ca/P ratios were computed from the EDX spectrum.

2.5. ATR-FTIR spectroscopy

Infrared spectrum was acquired in transmittance mode in the spectral region of 4000–400 cm⁻¹ using Nicolet 6700 FTIR Spectrophotometer having a DTGS (Deuterated Triglycine Sulphate) IR detector with spectral resolution 4 cm⁻¹ and 40 scans taken for each sample [16].

2.6. Atomic force microscopy (AFM)

Surface roughness of PCL scaffolds with different surface modifications were imaged using Veeco Dimension 3100 atomic force microscope (AFM) in contact mode. Three areas were chosen randomly (Dimension: a 10 μm × 10 μm) and were imaged. The mean of their respective surface roughness (Rq) values was calculated.

2.7. X-ray diffraction (XRD)

X-ray diffraction (XRD) was used to determine the changes in crystalline content. Wide-angle XRD (X’Pert PRO PANalytical DY2022, the Netherlands) was conducted by scanning the samples with CuKα radiation (1.5405 Å) from 0° to 50° (2θ) under an accelerating voltage of 45 kV and 30 mA current [17]. For scanning the scaffolds slit width of 1.52 mm and step size 0.05 was taken, while for capturing the peaks of deposited minerals the slit width was increased to 3.03 mm and step size was decreased to 0.02.

2.8. Acellular mineralization

For achieving acellular mineralization on PCL scaffolds with differential surface chemistries, they were immersed in DMEM [low glucose, pyruvate (Gibco^™, Thermo Scientific, USA)] for 7 days at 37 °C and 5% CO₂ in a humidified atmosphere to mimic the physiological conditions including pH of ~7.4 prevailing during the formation of bioapatite in vivo. The scaffolds were taken out of the solution and air dried under overnight and then used for different analysis.

2.9. Transmission electron microscopy (TEM)

Briefly, the PCL scaffolds were ultrasonically dispersed in ethanol followed by deposition onto carbon-coated grids in dilute suspensions and scanned at 200 kV (JEOL 2000 FX-II). For morphological characterization of the crystalline deposition, high resolution phase contrast imaging of the samples was performed [18]. From the captured images, crystal dimensions were determined using ImageJ software.

2.10. Micro-computed tomography

MicroCT was performed using a Siemens Inveon PET/SPECT/CT scanner (Siemens Medical Solutions, Malvern, PA, USA). The cone-beam CT parameters were as follows: 360° rotation, 720 projections, 1300 ms exposure time, 1500 ms settle time, 80 kV voltage, 500 μA current, and effective pixel size 17.67 μm. Acquisitions were reconstructed using a Feldkamp algorithm with Shepp-Logan filter and beam-hardening correction, matrix size 1024 × 1472 × 721, using manufacturer-provided software. Protocol-specific Hounsfield Unit (HU) calibration factor was applied. Quality control included HU constancy using rat femur phantom standards with the same protocol.

3D scaffold segmentation was performed using ROI thresholding tools of the Siemens Research Workplace software (version 4.2) to calculate volumes of the scaffold structures and deposits. The number of all pixels which contained scaffold material were counted, ensuring voids were not included in the count by setting a threshold of HU > −850. Then a threshold of HU values > 0, was set to count pixels for deposits in the scaffold. A third threshold was applied to count extremely dense deposits with HU values > 1000. Percent volumes (# pixels in deposits/# pixels in scaffold) were then calculated for each.

2.11. Statistical analysis

Data have been presented as mean ± SD. The number of replicates has been indicated with the respective methods. Single factor paired one tailed students t-test was used for the analysis of the significance of variations. The difference between different test and control groups was calculated by p-value. The groups with p < 0.01 were considered as statistically significantly different.

3. Results

3.1. Pre-mineralization studies

3.1.1. SEM analysis

The surface morphology of the samples as observed by scanning electron microscope revealed etching of the PCL surface in oxygen plasma treated samples (Fig. 2). The high magnification images (3.5 kX and 12 kX) of 30 OP scaffolds demonstrated subjectively deeper surface etching. The subjective appearance of the depth of surface etching increased with the increase in time of plasma treatment. The control PCL scaffolds demonstrated a flat surface, while the TCP-PCL scaffolds showed a rough surface with TCP particles projecting out. This demonstrates that the oxygen plasma etching process and TCP incorporation in 3D-printed PCL scaffolds induced different surface morphologies.

Fig. 2.

Surface morphology (SEM images) of 3D printed PCL scaffolds with differential morphologies- PCL, 10 OP, 30 OP and TCP-PCL. The plasma etched groups- 10 OP and 30 OP demonstrated roughened surfaces in higher magnification images. 3.5 kX and 12 kX images of 30 OP scaffolds showed subjectively deeper retching of the PCL surface forming grooves in it.

3.1.2. AFM analysis

Atomic force microscopy was done for further analysis and quantification of the surface roughness as a result of plasma etching and TCP incorporation processes (Fig. 3). The AFM analysis demonstrated a significant increase in surface roughness of PCL scaffolds as a result of oxygen plasma treatment (p < 0.01, n = 3). The roughness increased significantly as a factor of treatment time (PCL (96.2) < 10OP (196) < 30OP (274), p < 0.001). TCP-PCL scaffolds showed significantly higher roughness values (283.67) than PCL and 10OP. These values were consistent with the observations made during SEM. The decorated TCP particles were easily visible on the surface of TCP-PCL scaffolds.

Fig. 3.

(a) Surface topography of 3D printed PCL scaffolds with differential morphologies- PCL, 10 OP, 30 OP and TCP-PCL obtained with AFM analysis- 2D surface and 3D reconstruction. The 30 OP scaffolds showed a highly roughened surface in 3D reconstruction images. TCP-PCL scaffolds demonstrated the presence of TCP particles on the surface of PCL. (b) Roughness values calculated from the AFM data showed significantly higher roughness of plasma treated and TCP-PCL scaffolds than the PCL scaffolds (n = 3, #p < 0.01 between 10 OP and PCL, *p < 0.001 between 30 OP, TCP-PCL and PCL, 10 OP).

3.1.3. Elemental analysis

Energy-dispersive X-ray spectroscopy was performed to analyze the change in elemental composition of the surface of PCL scaffolds as a result of oxygen plasma treatment and TCP incorporation (Fig. 4(a), Table 2). Oxygen plasma treatment for 10 min didn’t demonstrate any change in elemental composition of the surface. However, when PCL was treated for 30 min, the atomic percentage of oxygen on the PCL surface increased from 17.41% to 20.09% consistent with the induction of oxygen into the surface functionality of PCL scaffolds. TCP incorporation in PCL also changed the elemental composition of the surface by increasing the atomic percentage of oxygen to 21.36%, as well as introducing 1.73% of P and 2.83% of Ca.

Fig. 4.

(a) EDX data showing increase in oxygen content with oxygen plasma etching process, while the TCP-PCL scaffolds demonstrated the presence of Ca and P on the surface indicating exposure of TCP particle. (b) ATR-FTIR spectra demonstrated that 30 OP scaffold surface got significantly modified by introduction of hydroxyl and carboxyl groups as indicated by the band at 3700–3000 cm⁻¹ and hunch in the C=O peak at 1620 cm⁻¹. The 10 OP scaffold demonstrated only a slight band in the −OH region, indicating initiation of oxygen modification of the surface. (c) XRD spectra of the 3D printed scaffolds with differential surface chemistries. The spectra indicated no changes in crystallinity of the plasma treated PCL scaffolds. The TCP-PCL scaffolds demonstrated reduced crystallinity as indicated by a reduced peak at 21.3°. The small peak at 31.5° indicated the 211 plane of TCP consistent with AFM and its exposure on the scaffold surface.

Table 2.

Elemental composition of surfaces of 3D printed PCL scaffolds with differential surface chemistry.

Element	Weight %
	PCL	10 OP	30 OP	TCP-PCL
C	82.59	82.66	79.91	74.07
O	17.41	17.34	20.09	21.36
P				1.73
Ca				2.83

3.1.4. ATR-FTIR analysis

FTIR analysis was performed to identify the change in functional groups present on the scaffold surface as a result of oxygen plasma treatment and TCP incorporation (Fig. 4(b)). The oxygen plasma treated groups demonstrated changes in two specific bands- 3600–3100 cm⁻¹ which corresponds to the OH stretching of the carboxyl group and a strong peak at 1720 cm⁻¹ which corresponds to C=O stretching [19]. The 30 OP group demonstrated a broad OH band which was not significant in 10 OP and TCP-PCL groups and absent in the PCL scaffolds. These findings confirmed the introduction of OH groups onto the PCL surface after oxygen plasma treatment. Simultaneously, the peak for C=O group at 1720 cm⁻¹ demonstrated widening with a hunch at 1620 cm⁻¹. This widening of the carbonyl group peak and introduction of a broad OH carboxyl band demonstrated the conversion of −CH₂−CH₂− and ester linkages present in PCL to oxygen containing hydroxy, carboxyl and carboxyl groups. The TCP-PCL scaffolds demonstrated additional peaks at 564 cm⁻¹ and 603 cm⁻¹ which represents the PO₄³⁻ group present, showing that the TCP particles incorporated in PCL during the filament production process were exposed on the surface of 3D printed TCP-PCL scaffolds, leading to an expected alteration in surface chemistry [20].

3.1.5. XRD analysis

XRD analysis was performed to study the effects of oxygen plasma treatment or TCP incorporation on the crystallinity of 3D printed PCL scaffolds (Fig. 4(c)). All the samples demonstrated peaks at 21.3° and 23.6° corresponding to the orthorhombic planes (110 and 200) of semicrystalline PCL. The peaks had a weaker intensity in TCP-PCL samples consistent with a decrease in crystallinity due to the TCP incorporation process. The TCP-PCL scaffolds also demonstrated a peak at 31.5° which represents the 211 plane of TCP, consistent with the surface exposure of TCP particles on TCP-PCL scaffolds [12].

3.2. Post-mineralization studies

3.2.1. SEM analysis

SEM analysis was performed post-mineralization to study the extent of mineral deposition on the surface of PCL scaffolds (Fig. 5). SEM micrographs demonstrated formation of subjectively larger crystalline deposits on the surface of 10 OP and 30 OP scaffolds covering their surfaces, while the PCL scaffolds had a sparse deposition of morphologically flat sheets of minerals on the surface. The crystalline networks formed on the 10 OP scaffolds had petal like branches while those on the 30 OP scaffolds had needle shaped branches, indicating a higher order of crystallinity in the networks observed on 30 OP scaffolds. The TCP-PCL scaffolds demonstrated the deposition of thick sheets of minerals which were not flat as in PCL scaffolds.

Fig. 5.

SEM images of PCL scaffolds with differential surface chemistry after acellular mineralization for 7 days in DMEM media. The oxygen plasma treated scaffolds demonstrated crystal network formation as shown in the high magnification images (5 kX and 12 kX), while the PCL scaffolds demonstrated deposition of calcium phosphate in form of sheets. TCP-PCL scaffolds showed subjectively larger calcium phosphate deposits which were not crystallized in the same fashion as on the oxygen plasma treated scaffolds (10 OP and 30 OP).

3.2.2. Elemental analysis

EDX analysis was performed to study the elemental composition of the deposits on the 3D printed scaffolds with differential surface chemistry (Fig. 6(a)). All the scaffolds showed Ca, P, Na and Cl distinct from C and O, indicating the deposition of NaCl and hydroxyapatite from the media solution onto the surfaces of these scaffolds (Table 3). The atomic percentages of the newly introduced mineral elements on the surfaces of these scaffolds, were comparable to each other indicating equal deposition of the minerals despite differential surface chemistry. The Ca/P ratio allows identification of specific calcium mineral deposits. Hydroxyapatite is reported to have a Ca/P ratio of 1.67. Notably, the mineral deposited upon the 30 OP scaffolds demonstrated a Ca/P ratio of 1.7 indicative of hydroxyapatite (1.67) [21]. However, the elemental percentages of mineral deposits on the remaining PCL scaffolds were similar to but not identical to hydroxyapatite.

Fig. 6.

(a) EDX data demonstrated that calcium phosphate was deposited on all the PCL scaffolds. (b) The FTIR spectra revealed the broad −OH band (3700–3000 cm⁻¹) in 30 OP scaffolds indicating the formation of hydroxyapatite crystals while the PCL scaffolds demonstrated no such band indicating the deposition of amorphous calcium phosphate on them. Further, the peak at 1585 cm-1indicated the carbonated nature of hydroxyapatite on 30 OP scaffolds, consistent with the formation of bioapatite like mineral deposits.

Table 3.

Elemental composition of the surfaces of PCL, 10 OP, 30 OP and TCP-PCL scaffolds post mineralization for 7 days in DMEM.

Sample	Weight %						Ca/P ratio
	C	O	P	Ca	Cl	Na
PCL	70.96	21.21	2.2	4.51	0.37	0.75	2.05
10 OP	71.32	20.05	2.36	4.85	0.47	0.95	2.05
30 OP	71.98	19.98	2.41	4.12	0.49	1.02	1.7
TCP-PCL	72.3	19.75	2.1	4.2	0.58	1.07	2

3.2.3. ATR-FTIR analysis

To further analyze the chemistry of the mineral deposits on the PCL scaffolds with differential surface chemistry, ATR-FTIR analysis was performed (Fig. 6(b)). The OH− band from 3700 to 3000 cm⁻¹ which indicates the hydroxy groups associated with deposited minerals, had a maximal intensity in 30 OP scaffolds followed by 10 OP and TCP-PCL scaffolds [19]. The PCL scaffolds didn’t show the OH-band indicating that mineral deposits on these scaffolds was not indicative of hydroxyapatite but rather tetra calcium phosphate, as indicated by the Ca/P ratio. The peak observed at 1585 cm⁻¹, which represents –CO₃^2– groups, showed highest intensity in 30 OP followed by 10 OP and TCP-PCL scaffolds, while the peak was absent in PCL scaffolds. This finding confirms the presence of the carbonate group in the hydroxyapatite deposits of oxygen plasma treated PCL scaffolds; making it more biomimetic [22]. Thus, the ATR-FTIR findings confirmed the deposition of biomimetic hydroxyapatite on oxygen plasma treated PCL scaffolds.

3.2.4. XRD analysis

To study the crystalline structure of the mineral deposits on PCL scaffolds with differential surface chemistry, X-ray diffraction analysis was performed (Fig. 7(a)). All the samples demonstrated peaks at 21.3° and 23.6° corresponding to the orthorhombic planes (110 and 200) of semicrystalline PCL. In addition to those peaks, smaller peaks visible between 27°–30° were also observed. To visualize these smaller peaks, another spectra with higher intensity (by increasing receiving slit with and decreasing scanning step size) was captured in this region, which highlighted these peaks clearly.

Fig. 7.

(a) XRD spectra of deposited minerals on the 3D printed PCL scaffolds showed the formation of hydroxyapatite crystals in 30 OP scaffolds, indicated by the sharp peak for 211 plane. The PCL scaffolds showed no peak for plane 211 indicating amorphous CaP deposition while 10 OP scaffolds showed a small wide peak indicating initiation of HA crystal formation in these scaffolds. The TCP-PCL scaffolds showed a wide peak for the 211plane in addition to the one for 300, indicating disordered crystal formation. (b) The TEM micrographs demonstrated deposition of elongated crystallites in 30 OP scaffolds that were arranged in bundles (as shown in the inset image), while the minerals from 10 OP scaffolds showed minimal organization. The minerals from PCL and TCP-PCL showed no elongated crystals or bundle formation.

The PCL, 10 OP and 30 OP scaffolds showed a small and broad peak for the 210 plane, indicating the amorphous nature of these mineral deposits. IN addition to this peak, the 30 OP also demonstrated a prominent peak for 211 plane which indicated ordered hydroxyapatite crystal growth in this plane. The 10 OP scaffolds also demonstrated a smaller wide peak for the 211 plane, indicating a fraction of the crystals growing disorderly in this plane. The TCP-PCL scaffolds demonstrated a prominent wide peak for the 211 plane followed by a smaller wide peak for the 300 plane which also indicated the disordered growth of the HA crystals in both these planes, 211 being the prominent plane of growth. The XRD analysis confirmed the inferences of ATR-FTIR and SEM analysis.

3.2.5. TEM analysis

To analyze the crystal morphology of the mineral deposits on PCL scaffolds with differential surface chemistry, TEM was performed (Fig. 7(b)). The TEM images showed the deposition of nanocrystalline HA on all the PCL scaffolds, consistent with the ATR-FTIR and XRD analysis. The HA deposits on oxygen plasma treated PCL scaffolds- 10 OP and 30 OP demonstrated large aggregates of needle shaped nanocrystallites which were arranged in a parallel fashion as shown in the magnified view in the inset image. The long crystals, which were approximately 70 nm in length, were observed in the 30 OP scaffolds and confirm the sharp peak for the 211 plane in XRD data. These findings are consistent with ordered crystal growth in this plane. The PCL and TCP-PCL scaffolds demonstrated HA nanocrystallites, which were smaller and more dispersed. These HA nanocrystallites were not elongated in any plane and thus could not align themselves together as compared to those on oxygen plasma treated PCL scaffolds. This was in agreement with XRD data in which the PCL and TCP-PCL scaffolds demonstrated broad peaks for the 210 and 211 planes respectively, indicating disordered crystal formation in these planes.

3.2.6. Micro-CT analysis

Micro-CT analysis was done to quantify the weight percentage of HA deposited on the PCL scaffolds with differential surface chemistry (Fig. 8). PCL and HA were differentiated in the images depending on their densities and their relative weight percentages. The TCP-PCL scaffolds demonstrated the highest HA crystal deposition followed by 30 OP, 10 OP and PCL scaffolds. The TCP-PCL scaffolds also showed the presence of extra high-density TCP that was embedded in the PCL matrix during the filament synthesis. 30 OP scaffolds demonstrated deposition of aligned aggregates of HA crystals on the scaffold filaments consistent with results of TEM analysis and absent in all other scaffold groups.

Fig. 8.

Micro-CT analysis. (a) Top view (full projection) and (b) cross sectional view (full projection) of 3D printed PCL scaffolds with differential surface chemistry, demonstrating the mineralization occurred in the bulk of the scaffolds. The 30 OP scaffolds showed aligned deposition of hydroxyapatite crystals which was not seen in any other scaffolds. (c) The micro-CT quantification (n = 3) of mineral crystals demonstrated the highest deposition in case of TCP-PCL scaffolds followed by 30 OP, 10 OP and PCL scaffolds. The PCL scaffolds had a significantly lower mineral deposition as compared to other groups (#p < 0.01) and TCP-PCL was higher than any other groups (*p < 0.01).

4. Discussion

Several studies have been performed to date examining mineralization on PCL scaffolds [12,13], but systemic characterization of the PCL surface chemistry and subsequent characterization and quantification of the deposited minerals with respect to the changes in surface chemistry has not been previously described. This critical gap in current knowledge regarding the correlation between the surface chemistry of PCL scaffolds and the mineralization process impedes further development and translation of biomimetic scaffolds for bone tissue engineering.

As previously reviewed, PCL is an ideal material for making 3D printed scaffolds for tissue engineering purposes, owing to its minimal toxicity, excellent mechanical properties, low immunogenicity and excellent printability (moldability) [6]. Its variable mechanical properties can be tailored to withstand the physiological forces applied to load bearing tissues in vivo and thus, PCL is an ideal candidate for use in bone tissue engineering [23]. An additional prerequisite for a material to be considered for bone tissue engineering is its ability to act as a template for mineralization, as the mechanical properties of the regenerating tissue will depend on the extent of mineralization that occurs during its development.

The ability to act as a mineralization template can be assessed in an in vitro environment by mimicking the physiological conditions in which mineral crystals are formed. For this purpose, simulated body fluid (SBF) has been used previously as the gold standard [12,24]. However, there are numerous factors that makes the use of SBF quite different from physiological conditions: 1) the absence of proteins that play a crucial part in controlling the hydroxyapatite nucleation process, 2) addition of TRIS buffer to SBF and 3) lack of control over carbonates in SBF which act as buffer in physiological conditions [25]. Apart from these limitations in the composition of SBF, it is our opinion that the conditions chosen for testing mineralization potential of biomaterials also need to be reconsidered. The lack of a CO₂ rich environment during the process of mineralization does not mimic the carbonate buffered environment present during physiologic conditions. Some studies also use pretreatment of the biomaterial surfaces with calcium chloride which inhibits the direct contact of the biomaterial surface with SBF and thus the mineralization occurring on the calcified surface can’t be considered as an intrinsic property of the biomaterial to act as a template for mineralization [24]. Similarly, many prior studies were performed in an in vitro cell culture (osteoblasts or mesenchymal stem cells) environment in order to predict the in vivo mineralization potential of the biomaterial surface [12,13]. During cell culture on these biomaterial surfaces extracellular matrix is produced which directly interacts with cell culture media to initiate mineralization on the biomaterial surface. Thus, mineralization as a result of cell culture is not an optimal indicator of mineralization potential of the biomaterial surface. To avoid these experimental limitations of the mineralization studies, some new protocols for testing the potential of biomaterials to act as a template for mineralization have been suggested [25,26]. Thus, in the present study the following modifications were performed- 1) cell culture medium- Dulbecco’s modified Eagle’s medium (DMEM) was used for mineralization to mimic the physiological environment, 2) no cells were added during the mineralization studies to avoid the deposition of extracellular matrix that can enhance the mineralization on the biomaterial surface, 3) serum was not added to DMEM to avoid inhibition of mineralization by serum proteins [27]. DMEM contains inorganic salts, vitamins, amino acids, etc. which more closely mimic the physiological environment than SBF. The incubation was also performed at 5% CO₂ with humidified air at 37 °C, which better mimics physiologic conditions as it keeps DMEM buffered with carbonates.

According to crystallization theory, bioactivity or bone bonding ability and hydroxyapatite formation on biomaterial surface are two entirely different phenomenon [25]. Thermochemical calculations indicate that both SBF and serum are supersaturated ionic solutions that tend to crystallize minerals on an available surface as these solutions are in a metastable thermodynamic state and become stable by the inevitable crystallization of the mineral particles [5,28]. Thus, the potential of a biomaterial to act as a template for mineralization should not only be decided by the quantity of apatite deposited on the material surface but also on the quality of the crystals formed. The crystal structure, stoichiometry and similarity to bioapatite, rather than hydroxyapatite, of the deposited minerals should also be considered as a factor in deciding mineralization potential of biomaterials. These properties play a crucial role in determining the functionality of the bone tissue regenerating on these biomaterial scaffolds, as bone is a complex nano-composite of collagen and hydroxyapatite that is in turn responsible for its mechanical properties. Thus, in the current study emphasis was given to characterizing the quality of hydroxyapatite deposited on the surface of PCL scaffolds with differential surface chemistries.

The SEM images indicated significantly increased surface roughness in oxygen plasma treated PCL scaffolds- 10 OP and 30 OP as compared to untreated PCL scaffolds, which was confirmed and quantified by the AFM studies. The main advantage of oxygen plasma treatment process is that it doesn’t alter the bulk mechanical properties of the PCL scaffolds while also avoiding the use of any toxic chemicals (chemical etching) [29–31]. The roughness of the PCL surface increased as a function of the time of plasma treatment, similar to prior studies [31]. A prior study demonstrated that excessive treatment with oxygen plasma does not produce significant changes in surface roughness after a limit [12]. Thus, in the current study, PCL scaffolds were only treated for 10 min and 30 min, flipping them halfway to have homogenous modification of the surface chemistry of 3D printed PCL scaffolds. TCP was used as a positive control to both roughen the surface of PCL scaffolds, as well as provide biomimetic mineralization sites on the PCL surface for comparison. The TCP-PCL showed the highest roughness across all the groups and introduction of TCP particles in PCL scaffolds has also been shown to increase the contact angle and thus hydrophilicity of these scaffolds making the TCP-PCL scaffolds an excellent group for comparison as a positive control.

The elemental analysis of scaffolds, prior to mineralization, demonstrated that percentage weight of oxygen remained approximately the same after 10 min of oxygen plasma treatment, but increased after 30 min. The time of plasma treatment plays an important role in modifying the surface chemistry of PCL surface. An earlier study demonstrated that treatment of PCL scaffolds for prolonged period with oxygen plasma resulted in decrease of oxygen content on the scaffold surface [12]. The oxygen composition of the surface for 2 h treatment group was found to be higher than that of 4 h oxygen plasma treated group. This indicated that excessive treatment of PCL scaffolds with oxygen plasma is not required and a short duration plasma treatment can produce desirable changes in PCL surface chemistry. The TCP-PCL group demonstrated the presence of Ca and P on the surface of these scaffolds, consistent with exposed TCP particles on the surface of PCL, which is very important for them to act as a nucleation site for mineralization. It has been reported earlier that the amount of TCP incorporated within PCL scaffolds is directly proportional to the roughness and hydrophilicity of these scaffolds [32, 33]. ATR-FTIR studies demonstrated that the oxygen plasma treatment process (30 OP) introduced oxygen groups into the PCL polymer chain modifying it at the ketone group to form a carboxyl group which was demonstrated by the appearance of hydroxyl band and widening of the carbonyl peak. The 10 OP and TCP-PCL groups also demonstrated a slight hydroxyl band because of partial introduction of oxygen groups and the presence of TCP respectively. Thus, the oxygen plasma treatment process introduced carboxyl functionality on the surface of 3D printed PCL scaffolds. It has been reported that oxygen plasma treatment process introduces carboxyl groups on PCL surface (up to 2 h treatment time), while it results in lowering the number of carboxyl groups on prolonged treatment (4 h) [12]. Treatment of those modified PCL scaffolds was followed by mineralization in SBF and subsequent in vitro cell culture with pre-osteoblast cells (MC3T3-E1) demonstrated enhanced cellular proliferation and mineral deposition on the modified scaffolds. As demonstrated by Kim et al., oxygen plasma treatment for 30 min raised the surface roughness and percentage composition of oxygen containing functional groups specifically −OH and −COOH significantly, over non treated samples [12]. The further increase in treatment time with oxygen plasma didn’t significantly change the surface chemistry. Thus, for the present study 30 min treatment time was chosen as it can significantly change the surface chemistry as compared to non-treated PCL scaffolds. While 10 min oxygen plasma treatment was chosen to analyze the effect of shorter exposure time, which has not been studied earlier. In a similar study, Jeon et al. demonstrated that 1 h of oxygen plasma treatment resulted in the chemical modification of PCL surface by introduction of carboxyl groups [13]. In vitro cell culture with osteoblast like cells after plasma modification of these scaffolds demonstrated enhanced cell viability and alkaline phosphatase activity. The abovementioned studies sought to define the effect of oxygen plasma treatment of PCL scaffolds on their potential to help support cellular proliferation and expression of osteoblast like phenotype. However, its effect on the deposited mineral structure and quality was not studied, which is crucial in determining the resultant mechanical properties of the developing in vitro construct. Thus, in the current study the intrinsic potential of the oxygen plasma modified PCL scaffolds for acting as a mineralization template was studied. The scaffolds were dipped in physiologically relevant conditions for 7 days in DMEM and then the deposited minerals were studied for their chemistry and crystal structure in order to avoid the influence of SBF, proteins, and cellular and extracellular matrices on the oxygen plasma treated PCL surfaces.

Strong differences in the morphology of mineral deposits on the surface of PCL scaffolds could be attributed to the different surface chemistries. The oxygen plasma treated scaffolds demonstrated large aggregated networks of mineral crystals deposited on the scaffold surfaces, while those on PCL were in the form of sheets. An earlier study also demonstrated changes in morphology of crystal deposits depending on the surface chemistry with oxygen plasma treated scaffolds showing large crystal deposits, which were more homogenously distributed on the surface with increasing time of plasma treatment [12]. Another study demonstrated the formation of plate like crystal aggregates when hydroxyapatite crystals were synthesized in the presence of glutamic acid (Glu) and phosphoserine (Ser-OPO₃); indicating a role for nucleation sites on crystal nucleation as well as determining final crystal morphology [34]. All the earlier studies were done using SBF, while the present study is the first to observe the crystal formation in more physiologically relevant DMEM media.

The elemental analysis of scaffolds after mineralization demonstrated the similar percentage compositions of calcium and phosphorous deposited on all the four types of scaffolds. However, the Ca/P ratio suggested the presence of hydroxyapatite like structure in only 30 OP scaffolds as the ration matched with the molar ratio present in hydroxyapatite (1.67), while all others had ratios similar to amorphous calcium phosphate which lies in the range of 1.2–2.2 [21]. This observation indicated that the flat sheets of mineral deposits seen on the PCL and TCP-PCL in the SEM studies, were deposits of amorphous calcium phosphate, while the crystalline structures deposited on the oxygen plasma treated scaffolds was indeed crystallized hydroxyapatite.

The ATR-FTIR results demonstrated an increase in hydroxyl groups as a function of increasing time of oxygen plasma treatment indicating hydroxyapatite formation. The 30 OP scaffolds also demonstrated a prominent peak at 1585 cm⁻¹ which indicates the presence of carbonate group. This peak had the highest intensity in 30 OP scaffolds indicating highest incorporation of carbonate groups into hydroxyapatite crystals. Hydroxyapatite forms the bone tissue but physiologically it is always associated with other minor impurities like sodium, magnesium, carbonate etc. that are essential for various physiological processes and is known as bioapatite [35]. Hydroxyapatite has several crystal sites where these interactions and substitutions can happen forming a variety of combinations. These substitutions can affect crystal structure, solubility, crystallinity and surface charge that can modify its biological properties [22]. Thus, the increased carbonate peak in the 30 OP scaffolds indicate the formation of more the biomimetic form of hydroxyapatite on that surface when compared to the remaining scaffolds. Carbonates are the major substituents of bioapatite and constitute 5–8% (by weight) of its physiological amount [36]. Carbonate ion can substitute either the hydroxyl group or the phosphate group from the hydroxyapatite structure, which are known as A-type and B-type carbonate apatite respectively. The B-type substitution of the apatite crystal causes a lot of physical changes, including reduced a-axis length, reduced c-axis length and changes in solubility, crystal size and micro-crystal strain [37]. These changes occur due to the weakness of Ca–CO₃ bonds as compared to Ca–PO₄ bonds [38]. High carbonate substituted bioapatite demonstrate a platelet morphology that results in an excellent interaction of these crystals with collagen fibers [36]. Thus, the formation of carbonated hydroxyapatite can potentially increase the interaction of deposited minerals with collagen fibrils when cultured with cells in vitro or implanted in vivo, resulting in strong nano-composite formation and thus better mechanical properties.

XRD data from the minerals deposited on the PCL scaffolds with differential surface chemistry demonstrated clear differences in the crystallinity of these deposits. The hydroxyapatite crystals formed on 30 OP scaffold prominent crystal formation elongated in towards the 211 plane. The micro-CT data suggested that with the increase of oxygen plasma treatment time the amount of deposited calcium minerals increased. The highest calcium mineral deposition was observed on TCP-PCL scaffolds. To our knowledge, this is the first report of absolute quantification of mineral deposits through Micro-CT. This suggested that phosphate groups present in TCP increased deposition of amorphous calcium phosphate, but it failed to crystallize into carbonated hydroxyapatite. Therefore, modification of the PCL surface with 30 min oxygen plasma treatment generated the necessary chemistry needed for physiologically relevant nucleation and crystal growth of carbonated hydroxyapatite, potentially useful in regenerating a more biomimetic tissue construct.

Recent studies have indicated that during the process of biomineralization, collagen and non-collagenous proteins play a very crucial role in the crystallization of the deposited minerals [39]. These proteins provide nucleation sites for hydroxyapatite crystal formation, growth and facet stability [40]. Amino acid side chains that are negatively charged in nature act as nucleation sites for formation of hydroxyapatite crystals in vivo. Similarly, specific functional groups like −COOH on biomaterial surfaces have been found to induce heterogeneous nucleation of calcium phosphate [41]. The carboxyl group is negatively charged and thus attracts Ca²⁺ ions in solution to form a chelated structure which results in heterogeneous nucleation to form hydroxyapatite crystals [42]. The crystal nucleation and growth depend on the local crystallization microenvironment. The functional groups available on the matrix surface can interact with diffusing solutes like Ca²⁺ and CO₃²⁻ which results in the modulation of the local concentration of reacting ions (supersaturation). They can also function as nucleation sites for crystal formation and can direct its orientation thereafter. Some functional groups can interact with growing crystals and the crystal facets [43]. This explains the formation of crystallized hydroxyapatite on 30 OP scaffolds, as they contain carboxyl groups developed due to the oxygen plasma treatment process.

An additional theory of the biomineralization process suggests that initially calcium phosphate gets deposited on the extracellular matrix. The interaction of the amorphous calcium phosphate (ACP) with collagen and other proteins then results into nucleation and growth of hydroxyapatite crystals [4]. This theory has also been proved in vitro where transient amorphous calcium phosphate phases lead to the formation of intrafibrillar hydroxyapatite crystals which mimic the natural bone formation [44]. This theory could help explain the different kinds of mineral deposits on PCL scaffolds with differential surface chemistries. Thus, the mechanism of in vitro acellular mineralization on PCL scaffolds can be explained as follows - The DMEM media is a supersaturated solution which initiates the deposition of ACP on all the four types of scaffolds used in this study. Now, the surface chemistry of these scaffolds guide the next step. In the case of only PCL scaffolds, there are no free functional groups present on the surface of the scaffolds and thus the deposited calcium phosphate remains in the amorphous state. In 10 OP scaffolds the scaffolds, there are some −OH and −COOH groups formed on the surface due to 10 min of oxygen plasma treatment. These negatively charged groups then act as nucleation sites for the formation of hydroxyapatite crystals, but the number of reactive functional groups on the surface is not sufficient; so only a fraction of ACP nucleates form disordered hydroxyapatite crystals. Meanwhile in the case of 30 OP scaffolds the prolonged treatment with oxygen plasma generated ample carboxyl groups on the PCL surface to act as template for hydroxyapatite crystal formation. These electronegative carboxyl groups mimic the nucleation sites present in Type I collagen. Thus, the ACP on these scaffolds in the presence of −COOH nucleation sites formed hydroxyapatite crystals. In case of TCP-PCL scaffolds the −PO₄ group of the TCP acted as a template for hydroxyapatite crystal nucleation but the crystal growth was disordered and formed crystals that were not elongated in the physiologically relevant plane (Fig. 9).

Fig. 9.

Schematic showing the mechanism of amorphous calcium phosphate deposition on PCL scaffolds and its nucleation and crystal growth depending on the surface chemistry of the corresponding scaffolds. The −COOH groups created on the PCL surface in case of 30 OP scaffolds acted as nucleation sites for crystallization of ACP into hydroxyapatite crystals and their subsequent growth. While the abundant phosphate groups on the surface TCP-PCL scaffolds from the TCP particles, resulted into HA nucleation but the crystal growth was disordered and not biomimetic.

It was also observed that the hydroxyapatite crystals grew with time on the oxygen plasma treated mineralized scaffolds (data not included) and could be seen with naked eyes. This suggested that the crystal formation process takes place after the mineral deposition and continues to form larges crystals and their network over time as it happens during bone development. This demonstrated that the 30 min oxygen plasma treatment resulted in generation of improvements in the surface chemistry of PCL scaffolds; necessary for biomimetic mineralization. Thus, this modified scaffold chemistry may prove useful in the further development of PCL scaffolds for bone tissue engineering purposes.

Another important aspect regarding surface modifications of PCL scaffolds is that it only affects the first few angstroms of the surface of the bulk material. When the surface erodes during in vitro implantation, the unmodified bulk of PCL will no longer be able to support biomimetic mineralization again. Thus, it is important to introduce this chemical modification in the bulk of PCL before printing, thus the PCL powder itself could be treated with prolonged exposure to oxygen plasma which can be then extruded into filaments and 3D printed into scaffolds. This will ensure that the biomimetic mineralization property remains available till the whole scaffold is degraded and the regenerating tissue is more like the native tissue.

5. Conclusion

The current study focused on understanding the effect of altering the surface chemistry of PCL in order to achieve biomimetic mineralization on such scaffolds to improve the in vivo performance and resultant properties of implanted scaffolds for applications in tissue engineering and regenerative medicine. In our study, 30 min of oxygen plasma treatment of PCL scaffolds introduced carboxyl groups on the surface, which acted as sites for nucleation of amorphous calcium phosphate to form hydroxyapatite crystals. It was found that the process of acellular mineralization on PCL scaffolds followed a similar biomineralization process to that which occurs during physiologic bone development. First, amorphous calcium phosphate gets deposited on the surface secondary to the supersaturated nature of the ionic solution, which then interacts with the carboxyl functionalities to form a crystalline structure. The developing hydroxyapatite crystals also incorporated carbonate groups from the environment during the development process to form bioapatite like mineral composition instead of pure hydroxyapatite. Thus, we conclude that oxygen plasma treatment induces a biomimetic surface chemistry on the PCL scaffolds that helps in a bioapatite-like crystallization cascade of calcium phosphate deposition and thus can serve as an excellent bone graft substitute for regeneration of damaged or diseased bone.