Polylactide (PLA) is a thermoplastic aliphatic polyester that has been extensively studied as a type of bioplastic. Recently, PLA-made biodegradable scaffolds have attracted much attention in bone tissue engineering due to its high processability, suitable elastic modulus, and great biodegradability/bioactivity [1] [2] [3] . However, the mechanical strength and biocompatibility of pure PLA need to be improved in order to meet the harsh requirements for medical implants. Efforts have been attempted to improve the performance of PLA scaffolds by adding fillers, such as high-performing polymers [4, 5] , ceramic particles [6] [7] [8] , etc.
Specifically, polycaprolactone (PCL) has been selected as a potential polymer filler owing to its great biocompatibility and biodegradability but the mechanical strength of the PLA/PCL composites is sacrificed [9] . Ceramic fillers, especially calcium phosphate (CaP) based fillers, are also incorporated into the PLA matrix due to the great osteogenesis effect and biodegradability [6, 7] . The disadvantage of ceramic-based fillers, is the uneven distribution of ceramic powders induced by the tendency of agglomeration, resulting in poor mechanical properties especially decreased ductility [10] . Furthermore, distorted shapes or inaccurate pore dimensions can always be presented in the ceramic-filled PLA scaffolds, which should be precisely controlled [7, 8] . Given these disadvantages of polymer and ceramic fillers in reinforcing PLA scaffolds, the metallic particle becomes a promising alternative because it is more ductile and can be homogeneously distributed in the PLA matrix [11] [12] [13] .
In recent years, a variety of manufacturing methods have been utilized to fabricate the PLA-based composite scaffolds, including molding [14] , solvent casting [15, 16] , extrusion [17] , polymerization [18] , and additive manufacturing (AM) [19] [20] [21] . AM builds a three-dimensional object from a computer-aided design model by joining the materials layer by layer. Due to its
Polylactide (PLA) is a thermoplastic aliphatic polyester that has been extensively studied as a type of bioplastic. Recently, PLA-made biodegradable scaffolds have attracted much attention in bone tissue engineering due to its high processability, suitable elastic modulus, and great biodegradability/bioactivity [1] [2] [3] . However, the mechanical strength and biocompatibility of pure PLA need to be improved in order to meet the harsh requirements for medical implants. Efforts have been attempted to improve the performance of PLA scaffolds by adding fillers, such as high-performing polymers [4, 5] , ceramic particles [6] [7] [8] , etc.
Specifically, polycaprolactone (PCL) has been selected as a potential polymer filler owing to its great biocompatibility and biodegradability but the mechanical strength of the PLA/PCL composites is sacrificed [9] . Ceramic fillers, especially calcium phosphate (CaP) based fillers, are also incorporated into the PLA matrix due to the great osteogenesis effect and biodegradability [6, 7] . The disadvantage of ceramic-based fillers, is the uneven distribution of ceramic powders induced by the tendency of agglomeration, resulting in poor mechanical properties especially decreased ductility [10] . Furthermore, distorted shapes or inaccurate pore dimensions can always be presented in the ceramic-filled PLA scaffolds, which should be precisely controlled [7, 8] . Given these disadvantages of polymer and ceramic fillers in reinforcing PLA scaffolds, the metallic particle becomes a promising alternative because it is more ductile and can be homogeneously distributed in the PLA matrix [11] [12] [13] .
In recent years, a variety of manufacturing methods have been utilized to fabricate the PLA-based composite scaffolds, including molding [14] , solvent casting [15, 16] , extrusion [17] , polymerization [18] , and additive manufacturing (AM) [19] [20] [21] . AM builds a three-dimensional object from a computer-aided design model by joining the materials layer by layer. Due to its
Polylactide (PLA) is a thermoplastic aliphatic polyester that has been extensively studied as a type of bioplastic. Recently, PLA-made biodegradable scaffolds have attracted much attention in bone tissue engineering due to its high processability, suitable elastic modulus, and great biodegradability/bioactivity [1] [2] [3] . However, the mechanical strength and biocompatibility of pure PLA need to be improved in order to meet the harsh requirements for medical implants. Efforts have been attempted to improve the performance of PLA scaffolds by adding fillers, such as high-performing polymers [4, 5] , ceramic particles [6] [7] [8] , etc.
Specifically, polycaprolactone (PCL) has been selected as a potential polymer filler owing to its great biocompatibility and biodegradability but the mechanical strength of the PLA/PCL composites is sacrificed [9] . Ceramic fillers, especially calcium phosphate (CaP) based fillers, are also incorporated into the PLA matrix due to the great osteogenesis effect and biodegradability [6, 7] . The disadvantage of ceramic-based fillers, is the uneven distribution of ceramic powders induced by the tendency of agglomeration, resulting in poor mechanical properties especially decreased ductility [10] . Furthermore, distorted shapes or inaccurate pore dimensions can always be presented in the ceramic-filled PLA scaffolds, which should be precisely controlled [7, 8] . Given these disadvantages of polymer and ceramic fillers in reinforcing PLA scaffolds, the metallic particle becomes a promising alternative because it is more ductile and can be homogeneously distributed in the PLA matrix [11] [12] [13] .
In recent years, a variety of manufacturing methods have been utilized to fabricate the PLA-based composite scaffolds, including molding [14] , solvent casting [15, 16] , extrusion [17] , polymerization [18] , and additive manufacturing (AM) [19] [20] [21] . AM builds a three-dimensional object from a computer-aided design model by joining the materials layer by layer. Due to its
Polylactide (PLA) is a thermoplastic aliphatic polyester that has been extensively studied as a type of bioplastic. Recently, PLA-made biodegradable scaffolds have attracted much attention in bone tissue engineering due to its high processability, suitable elastic modulus, and great biodegradability/bioactivity [1] [2] [3] . However, the mechanical strength and biocompatibility of pure PLA need to be improved in order to meet the harsh requirements for medical implants. Efforts have been attempted to improve the performance of PLA scaffolds by adding fillers, such as high-performing polymers [4, 5] , ceramic particles [6] [7] [8] , etc.
Specifically, polycaprolactone (PCL) has been selected as a potential polymer filler owing to its great biocompatibility and biodegradability but the mechanical strength of the PLA/PCL composites is sacrificed [9] . Ceramic fillers, especially calcium phosphate (CaP) based fillers, are also incorporated into the PLA matrix due to the great osteogenesis effect and biodegradability [6, 7] . The disadvantage of ceramic-based fillers, is the uneven distribution of ceramic powders induced by the tendency of agglomeration, resulting in poor mechanical properties especially decreased ductility [10] . Furthermore, distorted shapes or inaccurate pore dimensions can always be presented in the ceramic-filled PLA scaffolds, which should be precisely controlled [7, 8] . Given these disadvantages of polymer and ceramic fillers in reinforcing PLA scaffolds, the metallic particle becomes a promising alternative because it is more ductile and can be homogeneously distributed in the PLA matrix [11] [12] [13] .
In recent years, a variety of manufacturing methods have been utilized to fabricate the PLA-based composite scaffolds, including molding [14] , solvent casting [15, 16] , extrusion [17] , polymerization [18] , and additive manufacturing (AM) [19] [20] [21] . AM builds a three-dimensional object from a computer-aided design model by joining the materials layer by layer. Due to its 48th SME North American Manufacturing Research Conference, NAMRC 48 (Cancelled due to benefits in processing geometrically complex structures, AM is a favorable approach to fabricate scaffolds with desired pore structures for cell penetration and inhabitation in tissue engineering applications. Fused filament fabrication (FFF) advances in its ability to produce complex pore structures at higher material utilization rate and lower energy cost compared with other AM technologies such as stereolithography (SLA) and selective laser sintering (SLS) [22, 23] . During the FFF process, the thermoplastic-based feeding filament is extruded from a heated nozzle with the beads solidified on the printing bed, which is desirable for the PLA-based composite scaffolds fabrication [22] . However, very limited studies have focused on the FFF of PLA/metal composite scaffolds [12, 24] .
In the recent work of [12] , PLA/Ti composite scaffolds with a dispersive titanium powder distribution were additively manufactured using FFF technique. The scaffolds exhibited enhanced mechanical strength and improved cytocompatibility due to a higher cell proliferation rate, indicating the considerable functions of biocompatible metallic-based fillers. However, this literature did not investigate the biodegradation behavior of PLA/metal composite scaffolds, and the effects of different metallic powder contents on their physical properties such as surface roughness and dimensional accuracy remained unknown. Compared with titanium, 316L stainless steel also possesses high strength and good cytocompatibility to allow for their wide applications in medical scaffolds [25, 26] . Because of the biodegradability of iron-based alloys, fabricating a totally biodegradable scaffold by adding 316L fillers could be of greater significance. In addition, the daily tolerance in the human body for iron is higher than that for Ti [27, 28] . However, few studies have been conducted on the FFF fabrication of PLA/316L composites [29] , wherein the effects of 316L powder contents on the performance of the PLA-based composites have not been explored.
In this work, we provide a feasibility study on the fabrication of biodegradable PLA/316L composite scaffolds with different steel particle contents using FFF process. The physical properties including pore size, strut width, and surface roughness are analyzed to evaluate the dimensional accuracy and surface quality of the printed scaffolds. The structure, thermal properties, compressive performance, and corrosion behavior are also investigated to comprehensively evaluate the in-vitro performance of the PLA/316L scaffolds. The completion of this study provides valuable insights into understanding the effects of steel particle filler contents on the product performance of the PLA-based composite scaffolds for potential practical applications.
PLA pellets (4043D, IngeoTM Biopolymer) and 316L stainless steel powder (EOS, Munich, Germany) (meeting with ASTM standard F138-12 [30] with a particle size between 20 μm and 50 μm) were mixed in a lab roll ball mill (YLK-Q-5, Yonglekang Equipment Co., Ltd, Changsha, China) at room temperature for three hours. Based on the volume ratio of 316L to PLA, samples were divided into four groups: pure PLA, PLA5, PLA10, and PLA15, as the number represented the volume percentage of 316L. After mixing, the powder mixture was fed into a single-screw extruder (EX2, FilaBot Corp., Barre, VT, USA) to fabricate filaments. The extrusion temperature was 185 ℃, and the screw rotation speed was dynamically adjusted around 30 RPM to yield a suitable flow rate. Fan airpath was also applied to provide forced convection to cool filament for a uniform filament diameter of 2.85±0.15 mm.
The FFF process was conducted on a desktop 3D printer (TAZ 6, Lulzbot Corp., Colorado, USA) with a single extruder tool head. Fig. 1 illustrates the 3D printer and designed threedimensional porous scaffold to be fabricated in this work. Specifically, the extrusion nozzle initially deposited a group of eight parallel struts with a four-layer height. Each strut owns both width and height of 0.8 mm. After that, the printing orientation was alternated by 90° followed by layering another eight struts. Thus, a total of 60 layers were deposited to create a cube with a side length of 12 mm, and an interconnected pore dimension of 0.8×0.8×0.8 mm. The main process parameters for this sample fabrication are listed in Table 1 . 
Dimensions and surface roughness of the struts were measured using a 3D structural light profilometer (VR-3000, Keyence Corp., Itasca, USA). A graphical description of the measuring methodology is presented in Fig. 2 . The pore size was determined by measuring its length and width (D1) in the side view and the strut width (D2) was measured in the top view, as shown in Figs. 2(a) and 2(b), respectively. The side view allowed for higher accuracy in measuring the pore size because struts were built in the same plane. to evaluate the areal surface roughness of the strut in the top view. Three samples were measured and a total number of 24, 120, and 12 statistical data were collected in each group for strut width, pore size, and surface roughness, respectively. Then the results were analyzed by the ANOVA methodology using the Minitab 19 software, with Tukey's methods incorporated to compare between each two groups. The surface morphologies of the prepared filament and fabricated scaffold were observed by a scanning electron microscopy (SEM) (Supra 55VP, ZEISS Corp., Oberkochen, Germany). Prior to the observation, carbon coating was conducted on the sample surface to increase conductivity. 
Thermal gravimetric analysis (TGA) (Q50, TA Inc., New Castle, USA) was conducted on the filament to evaluate the thermal stability and actual composition of different samples. The temperature ranged from room temperature to 500 ℃, with a ramping rate of 10 ℃ per minute. The glass transition temperature Tg, crystalline phase transition temperature Tc, and the coefficient of thermal expansion (CTE) were determined by thermal-mechanical analysis (Q400, TA Inc., New Castle, USA) with the ramping rate of 5 ℃ per minute. The result data was processed by the software of TA Universal Analysis.
The compression test was conducted on a universal tester (Instron 3334, INSTRON Corp., Norwood, USA) in accordance with the ASTM D1621 standard [31] to evaluate the compressive properties of the as-built scaffold. Five samples for each group were compressed to at least 20% deformation with a rate of 0.1 mm per second. Compressive strength and modulus were calculated from the plotted stressstrain diagram. The surface morphology of the sample after the compression test was also observed by SEM.
Biodegradation properties were evaluated by an immersion test at 37±0.5 ℃ in phosphate-buffered saline (PBS) (ThermoFisher Scientific, Waltham, USA) following the standard of ASTM NACE/G31 [32] . The initial pH value of the PBS solution was 7.4. Three samples for each group were immersed for 7 days, and another three samples in the same group were immersed for one month. Before and after the immersion test, samples were completely dried in a 50 ℃ incubator for 12 hours, and the weights were measured accordingly. Structural and surface morphology after the immersion was also characterized by SEM. Table 2 . In Fig. 3(a) , the mean pore size within the pure PLA scaffold is 0.88 mm, which is larger than the designed value of 0.8 mm. For the PLA/316L composite scaffolds, the mean pore size significantly decreases, which is closer to the designed one. Under different 316L powder contents, the mean pore size follows the order in PLA5 > PLA15 > PLA10. On the other hand, PLA/316L scaffolds show a larger mean strut width than pure PLA in Fig. 3(b) , and PLA10 group exhibits the largest value compared with other groups. The mean pore size and mean strut width could be inversely linked because a wider strut would narrow the pore space, which is consistent with our experimental results. The measured results of surface roughness are provided in Fig. 3(c) . The average surface roughness Sa of pure PLA scaffold is 0.09 mm. After adding the steel powders at 5 vol% and 10 vol%, the surface roughness is tailored to 0.03 mm and 0.04 mm, respectively, but no statistical difference is found between these two groups. As the steel powder load further increases to 15 vol%, a significantly lower average surface roughness of 0.009 mm with a standard deviation of 0.003 mm is obtained. PLA/316L scaffolds show better dimensional accuracy and lower surface roughness than pure PLA. The pore size and strut width could be influenced by various important factors during the printing process. The results indicate that the addition of 316L powders significantly changes the physical properties of pure PLA. Generally, pore size is one of the most important characteristics in bone tissue scaffolds, which can not only determine the loading capability and modulus matching but the mass transfer and cell growth accommodation [22] . A good average pore size lies in the range of 300 to 900 μm, and it is better to fabricate the pore with less standard deviation. Homogenous distribution of pores could facilitate continuous tissue growth and prevent clogging in some regions of the scaffolds, thus accelerating the healing process. For the surface roughness, it is found in the literature that a rougher surface could assist cell adhesion and growth because a higher surface fluctuation could provide more space for tissue spreading and remodeling, thus promoting a better cytocompatibility for the scaffolds [33, 34] . However, another result shows that the cell adhesion could not be solely determined by surface roughness, but more by the surface energy, as evidenced by the optimal cell adhesion is obtained in small roughness ratios [35] . Fig. 4(a) illustrates the SEM morphology of PLA5 filament. It is notable that the texture on the PLA5 filament surface is elongated because of the extrusion and several small cracks are observed. Those cracks may be induced due to the friction between the nozzle and filament. Another reason for cracking might be the unstable fan cooling during the extrusion process, leading to stress concentration in some areas on the surface and producing cracks. Those cracks would reduce the elongation of the filament, which ultimately affects the printing processability of the filament. Generally, the surface of PLA5 filament is smooth but a better cooling strategy should be considered. Fig. 4(b) ) could be observed on the surface of struts. As seen in Fig. 4(b) , the steel powders are fully covered by PLA and no considerable agglomeration is discovered, indicating a homogeneous distribution of PLA and 316L steel powders at 5% volume fraction. However, when the steel powder load increased to 10 vol%, voids are found in Fig. 4(d) . The generation of voids is associated with the entrapment of steel powders into the flow of the melting PLA polymer. Steel powders are thermodynamically stable at printing temperature around 205 ℃, thus impeding the polymer movement or even cutting the melting PLA matrix during the printing process. This phenomenon is not remarkable during the filament extrusion because the filament has a relatively large dimension (2.85 mm). In the printing process, however, the one-layer depositing height is only 0.2 mm which is much closer to the diameter of steel powders, resulting in a larger portion of voids when the powder loading was 10 vol% or 15 vol%. Large gaps between layers are found in PLA15, which might be caused by the contact of steel powders between layer surfaces. Those contact could weaken the surface adhesion so that the adjacent layer could be detached to generate large gaps. The existence of steel powders significantly contributes to changes in the dimensions and surface roughness of the scaffolds. According to the SEM images, gaps are generated between layers, which could be the main cause in the increase of mean strut width that narrows down the mean pore size. However, within the PLA/316L groups, PLA10 composite scaffolds exhibit the highest mean strut width. This could be explained by the varied thermal properties that will be discussed later. On the other hand, because of the entrapment effect of steel powders, the melting flow PLA matrix is inhibited and smaller surface roughness of PLA/316L struts is thus obtained.
(b) (c) (a)
Fig. 5 depicts the TGA and TMA curves of PLA/316L composites with different 316L contents. As shown in TGA curves, the actual powder composition is close to the nominal value, and a higher starting decomposition temperature is found in pure PLA sample. Table 3 lists the glass transition temperature (Tg), crystalline phase transition temperature (Tc), and the coefficient of thermal expansion (CTE) measured from TMA curves. Tg and Tc points are marked in Fig. 5 (b) and the CTE is measured from the largest tangent slope in the temperature range from Tg to Tc for each curve. It can be noted that Tg and Tc of PLA are not significantly affected by the addition of steel powders because of their thermodynamic stability within the temperature range. However, the CTE of the PLA/316L composites was decreased due to the impeding effect of the metallic powders against the expansion of the PLA matrix. Such an effect is the strongest in the PLA10 composite scaffolds and gets weakened when adding 15 vol% 316L powders. This could be caused due to the agglomeration of 316L particles at 15 vol% that results in a non-uniform thermal expansion. It should be noted that the strut width for pure PLA scaffold is less than 0.8 mm, which could stem from the highest volume shrinkage of PLA under the cooling process. With a higher CTE, the volume change of material per Celsius could be larger during heating or cooling process. On the other hand, for the FFF-built PLA/316L composite scaffolds, a larger volume of shrinkage could be obtained in PLA5 and PLA15, so the mean strut width of PLA5 or PLA15 is smaller than that of PLA 10. However, it can be considered that the detachment of PLA/316L composite layers contributes more to increasing the mean strut width to the level that the volume shrinkage can be ignored. 
Representative compression strain-stress curves of the FFFbuilt scaffolds are shown in Fig. 6 . The compressive modulus and compressive strength are listed in Table 4 . From the strainstress curves, the initial elastic deformation stage could be determined for all samples, but the yield points for PLA and PLA5 could not be clearly identified. Following the instruction of ASTM D1621, the compressive strengths for all samples are measured in the stress at 10% strain. which is approximately six times higher than pure PLA5. This result shows that adding stainless-steel powder could increase the stiffness of PLA, and the effect is more significant when the volume percentage of steel powder exceeds 10%. But for the group of PLA15, more voids are generated in the matrix resulting in a decline of modulus. As for compressive strength, higher strength is also obtained in the group of PLA10, whereas the PLA5 exhibits a similar plastic deformation behavior with PLA. PLA15 exhibits slightly lower compression strength than PLA10, which also could be ascribed to the existence of voids and agglomeration of powders. However, the yield points for PLA10 and PLA15 are achieved at a lower strain, suggesting a smaller compressive elongation for these groups. Fig. 7 illustrates the structures of different scaffolds after compression tests. Because of low modulus, bending and fracture occur in PLA struts, as shown in Fig. 7(a) . Compared with pure PLA scaffold, the PLA/316L composites are more resilient to compressive force. The bending of struts is also found in the group PLA5, suggesting similar deformation behavior with PLA and is consistent with the results of strainstress curves. No obvious bending or cracking is observed for PLA10 and PLA15 because of higher compressive strength. It should be noted that gaps in the composite scaffolds may not influence the compressive strength because those gaps are eliminated during the compressive displacement. As shown in the high-magnification image in Fig. 7(c) , the gaps are filled by compression with several bulges observed. Those bulges could be related to the agglomeration of steel powders, which promote the generation of extra resistance against the compressive force. In the high-magnification image of PLA15, more voids are formed because of the exfoliation of steel powder. Thus, compressive strength for PLA15 is lower than that of PLA10. 
In order to evaluate the biodegradation properties of PLA/316L scaffolds, in vitro immersion tests were conducted in PBS solution at 37±0.5 ℃. After one-month immersion, the results of pH value and sample weight change are listed in Table 5 . There is not much difference that can be detected by the results of pH value and weight change between groups. PLA naturally degrades in simulated body fluid by hydrolysis where the reaction is:
Generally, the degradation of PLA causes the pH value to drop but in PBS solution, the pH value remains stable [36] . No large weight change of samples is observed after immersion for one month, which might be due to the low degradation rate of 316L in the PBS solution. Three of the four groups exhibit larger weight after immersion because the degradation products on the scaffold surface and inside the pores could be hardly eliminated, so well as the oxide product of 316L powder. Fig. 8 shows the morphologies of the PLA and PLA10 scaffolds after immersion for one week and one month. As for PLA, the delamination of struts and layers could be observed after one-week immersion as a result of the weak inter-layer connection [9] that hastens the degradation in additively manufactured scaffolds. The modulus and strength may drop rapidly because of those disconnections. After immersed in PBS for one month, swelling occurs due to the penetration of PBS solution between the layers. It also can be seen that the corrosion product is accumulated between layers rather than on the layer surfaces. On the other hand, PLA/316L composite scaffolds exhibit no fracture or disconnection in the one-week immersed sample, whereas swelling is still observed given the voids formed during the printing. For the PLA10 sample immersed in PBS for one month in Fig. 8(d) , delamination is presented, which could be generated by the gaps, while the 316L powders remain intact suggesting a high corrosion resistance. Based on the results of pH value, weight change, and corrosion morphology, no significant effects could be found on the degradation rate and the corrosion behavior after adding 316L into the PLA matrix. 
This work studied PLA/316L composite scaffolds with different metal contents fabricated by fused filament fabrication. The physical properties, meso-structure and their thermal, mechanical, and in-vitro degradation properties of the scaffolds were investigated. The detailed conclusions are drawn as follows:
• The higher dimensional accuracy based on the mean pore size and mean strut width was obtained by PLA/316L composite scaffolds, and the addition of 316L powder also decreased the surface roughness. The lowest surface roughness was obtained when the steel powder loading was at 15 vol%. • 316L powders were dispersedly distributed in the struts when the volume fraction was under 10%. For steel powder loading higher than 10 vol%, voids within a layer, gaps between two layers and powder agglomeration were observed. • The addition of 316L powder decreased the coefficient of thermal expansion of the PLA composite scaffolds, leading to the higher dimension accuracy and lower surface roughness. The adding of 316L in 10 vol% most effectively enhanced the compressive modulus and strength of the scaffolds. • Swelling and delamination of the layers were found in PLA and PLA/316L scaffolds in the immersion test for one month, but no considerable differences were discovered in the effects of 316L powder on the degradation behavior of the PLA/316L scaffolds.