The hysteresis behavior of the PLA/1 wt.% MWNT layers is clearer in Figure 6, which shows that after a large hysteresis in the first cycle, the track of the three successive cycles of loading/unloading is nearly identical, with a minor deviation distinguished between the second and the fourth loops. This behavior is similar to the findings of Slobodian and Saha [47], where accordingly in the MWNT network, a ratcheting strain (mean value of the maximum and minimum strain in one cycle) takes place after the first compression cycle as a consequence of the primary deformation of the porous composition and blocked the reverse mobility of nanotubes in the middle of the dense networks. Through successive cycles of loading and unloading, the nanotubes' reorder becomes stable and the MWNT network gets to a steady stress–strain hysteresis loop order. This suggests that when the carbon nanotube network is well deformed, it can be applied as a sensing component of compression stress. In Figure 6, it can also be observed that the signal is linear with a slope difference below and over 30 kPa. Figure 6b depicts a schematic representation of a 1 wt.% MWNT 3D-printed layer sandwiched between two copper electrodes. Dashed lines between the MWNT individual particles and clusters represent quantum tunneling bridges which accordingly allow charge carriers to tunnel from one cluster to another without any physical contact in composite systems.

**Figure 6.** Piezoresistive behavior of 3D-printed PLA/1 wt.% MWNT nanocomposite: (**a**) Synchronism of *Ar* with stress versus deformation. (**b**) Schematic diagram of the transduction mechanism of PLA/1 wt.% MWNT nanocomposite sandwiched between two metal electrodes towards compressive pressure. The electrical model of the FSR consists of a series of connections between the bulk (tunneling) resistance (*Rbulk*) and the contact resistance (*Rc*).

When the sample is subjected to external stress (*σ*), the inter-particle distance is reduced. According to the proposed model by Paredes-Madrid et al. [46], the total resistance across the Force Sensing Resistor (FSR) can be decomposed from Equation (3):

$$R\_{FSR} = R\_{bulk} + 2R\_c \tag{3}$$

where *Rbulk* is the resistance of the CPC caused by the quantum tunneling phenomenon and *Rc* is the contact resistance between the conductive particles and the metal electrodes. An FSR is created by the series connection between *Rbulk* and 2*Rc* as shown in Figure 6b. However, three phenomena occur when incremental stress is applied to an FSR [46]: (1) the contact resistance of the existing paths is decreased according to power laws; (2) new contact paths are constructed to a greater extent contributing to a decrease in the contact resistance; and (3) the average inter-particle distance is decreased, as a consequence decreasing the tunneling resistance, *Rbulk*. It seems that for the 1 wt.% MWNT 3D-printed layer, the contact resistance is decreased by forming new contact paths and decreasing the contact resistance of the existing paths under 30 KPa and 10% strain. However, over 30 KPa, the resistance decreases because of the diminishing of inter-particle distance and consequently decreasing the tunneling resistance. The piezoresistive coefficient, which is also called the gauge factor, *G*, can be graphically figured out from the slope of the curve in Figure 6 and calculated with Equation (4) [40]

$$G = \frac{A\_r}{\varepsilon} \tag{4}$$

where *Ar* = Δ*R R*0 (Equation (1)) is the piezoresistive response and ε= Δ*L L*0 is the deformation of the sensor. For the 3D-printed PLA/1wt.% MWNT layers, the value of *G* = 7.6 was obtained with the amplitude of the piezoresistive response of about *Ar* = −0.8 (−80%).

To find out more about the sensitivity limitations of the developed FSR such as stress rate and related strain, the piezoresistive response of the 3D-printed composite layers under cyclic compressive stress with a high speed of 18 N/min was observed. Figure 7 shows the piezoresistive responses of the PLA/1 wt.% MWNT and PLA/4 wt.% KB samples, exposed to ten cycles of compressive stress from 10 up to 100 kPa.

**Figure 7.** Comparison of 3D-printed nanocomposites piezoresistive responses in high force rate of 18 N/min (**a**) PLA/1 wt.% MWNT (**b**) PLA/4 wt.% KB.

Figure 7a shows that the piezoresistive response of the PLA/1 wt.% MWNT with an applied force rate of 18 N/min has a smaller amplitude (*Ar* = −0.60) and more noisy signals, but a higher value of *G* = 9.3 in comparison with low force rate of 1 N/min in Figure 4a. Figure 8 shows that applying a high force rate causes a smaller strain as there is insufficient time for the material to respond to stress with large-scale viscoelastic deformation or yielding [48].

In Figure 7b, the compression with a high force rate leads to responses of large sensitivity (*Ar* = −0.90) and exemption of noise for the PLA/4 wt.% KB 3D-printed layers. However, after almost seven cycles, the maximum sensitivity is not stable and starts to decrease. The high value of *G* = 47.6 in the first cycle shows the high piezoresistive properties of these layers if the cyclic functionality is not needed. The gauge factor decreases to *G* = 28 in the 10th cycle of stress. Therefore, the PLA/KB 3D-printed layers do not show stable piezoresistive behavior in a cyclic mode at low and high force rates.

**Figure 8.** Stress–strain diagrams of 3D-printed PLA/1 wt.% MWNT with different force rates in a cyclic mode.

#### **4. Conclusions**

In this paper, fused deposition modeling 3D printing is used to develop CPC layers and investigate their electrical and piezoresistive behaviors. To this aim, multi-walled carbon nanotubes (MWNT) and high-structured carbon black (Ketjenblack) (KB) were incorporated into polylactic acid and 3D-printable filaments using the melt-mixing process were developed. The 3D-printed layers were created using fused deposition modeling. The percolation threshold of the MWNT and KB 3D-printed layers are 1 wt.% and 4 wt.%, respectively, by the two-point resistance measurement method. It was shown that it was possible to 3D print piezoresistive PLA nanocomposite layers from the MWNT and KB fillers and PLA matrix. The PLA/1 wt.% MWNT 3D-printed layers with a 1 mm thickness exhibit a negative pressure coefficient (NPC) characterized by a decrease of about one decade in resistance with increasing compressive loadings up to 18 N with a maximum strain up to about 16%. In the cyclic mode with a 1 N/min force rate, the PLA/1 wt.% MWNT 3D-printed layers showed good performance with a value of *G* = 7.6 obtained with the amplitude of the piezoresistive response of about *Ar* = −0.8 (−80%). The response was linear in the range of pressure 10–100 kPa, with low noise and hysteresis that comes from the layer-by-layer architecture of the component and the tunneling effect of MWNT nanofillers in lower contents than the percolation threshold. At a high force rate of 18 N/min, the piezoresistive response of the PLA/1 wt.% MWNT has a smaller amplitude (*Ar* = −0.60) and more noisy signals but a value of *G* = 9.3. The KB composites could not show stable piezoresistive responses in a cyclic mode. However, the PLA/4 wt.% KB 3D-printed layers under high force rate compression lead to responses of large sensitivity (*Ar* = −0.90) and are exempt from noise with a high value of *G* = 47.6 in the first cycle. The results show that PLA/MWNT and PLA/KB can be considered good piezoresistive nanocomposites to be 3D printed where complex designs with functionality are needed for possible use in wearable electronics, soft robotics, and prosthetics, etc.

**Author Contributions:** Conceptualization, R.H.S., A.C., V.N. and C.C.; Data curation, R.H.S.; Formal analysis, R.H.S.; Funding acquisition, J.G., G.C., V.N. and C.C.; Investigation, R.H.S.; Methodology, R.H.S., A.C. and C.C.; Project administration, V.N. and C.C.; Resources, V.N. and C.C.; Supervision, A.C., J.G., G.C., V.N. and C.C.; Validation, R.H.S., A.C. and C.C.; Visualization, R.H.S.; Writing—original draft, R.H.S.; Writing—review and editing, R.H.S., A.C. and C.C. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research was funded by [ErasmusMundus Joint Doctorate Programme SMDTex—Sustainable Management and Design for Textile] grant number [n◦2014-0683/001-001-EMJD] and the APC was funded by [University of Borås, Sweden].

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Not applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

*Communication*

## **Characterization of 3D Printed Metal-PLA Composite Scaffolds for Biomedical Applications**

**Irene Buj-Corral 1,\*, Héctor Sanz-Fraile 2, Anna Ulldemolins 2, Aitor Tejo-Otero 1, Alejandro Domínguez-Fernández 1, Isaac Almendros 2,3 and Jorge Otero 2,3,\***

**Abstract:** Three-dimensional printing is revolutionizing the development of scaffolds due to their rapid-prototyping characteristics. One of the most used techniques is fused filament fabrication (FFF), which is fast and compatible with a wide range of polymers, such as PolyLactic Acid (PLA). Mechanical properties of the 3D printed polymeric scaffolds are often weak for certain applications. A potential solution is the development of composite materials. In the present work, metal-PLA composites have been tested as a material for 3D printing scaffolds. Three different materials were tested: copper-filled PLA, bronze-filled PLA, and steel-filled PLA. Disk-shaped samples were printed with linear infill patterns and line spacing of 0.6, 0.7, and 0.8 mm, respectively. The porosity of the samples was measured from cross-sectional images. Biocompatibility was assessed by culturing Human Bone Marrow-Derived Mesenchymal Stromal on the surface of the printed scaffolds. The results showed that, for identical line spacing value, the highest porosity corresponded to bronzefilled material and the lowest one to steel-filled material. Steel-filled PLA polymers showed good cytocompatibility without the need to coat the material with biomolecules. Moreover, human bone marrow-derived mesenchymal stromal cells differentiated towards osteoblasts when cultured on top of the developed scaffolds. Therefore, it can be concluded that steel-filled PLA bioprinted parts are valid scaffolds for bone tissue engineering.

**Keywords:** steel-filled PLA; FFF; scaffold; grid structure; cell culture

#### **1. Introduction**

Additive manufacturing (AM) is a group of techniques in which three-dimensional structures are manufactured layer-by-layer in an automated way. It offers several advantages over the traditional subtractive or forming techniques: (1) it allows manufacturing complex shapes and even porous structures, (2) cheaper parts are produced if low-cost machines are employed; (3) it implies material, waste, and energy savings. Within the AM field, there are seven different categories [1]: binder jetting (BJ) [2], directed energy deposition (DED) [3], material extrusion (includes FFF—Fused Filament Fabrication and DIW—Direct Ink Writing) [4,5], material jetting (MJ) [6], powder bed fusion (PBF) (includes SLM—Selective Laser Melting and SLS—Selective Laser Sintering) [7], sheet lamination [8], and vat photopolymerization (includes SLA—stereolithography and Digital Light Processing (DLP) printing, as well as volumetric 3D printing) [9,10].

Within the different AM techniques, FFF is one of the most widely used technologies for rapid prototyping within the biomedical field, as it presents several advantages in terms of costs and the range of materials that can be used [11]. Also known as FDM (Fused

**Citation:** Buj-Corral, I.; Sanz-Fraile, H.; Ulldemolins, A.; Tejo-Otero, A.; Domínguez-Fernández, A.; Almendros, I.; Otero, J. Characterization of 3D Printed Metal-PLA Composite Scaffolds for Biomedical Applications. *Polymers* **2022**, *14*, 2754. https://doi.org/ 10.3390/polym14132754

Academic Editors: Hamid Reza Vanaei, Sofiane Khelladi and Abbas Tcharkhtchi

Received: 31 May 2022 Accepted: 1 July 2022 Published: 5 July 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

Deposition Modelling), FFF uses a continuous filament of a thermoplastic material such as PolyLactic Acid (PLA) to build complex 3D structures in an automated way. One of the main disadvantages of FFF technology is the difficulty to ensure the correct bonding between layers [12]. The FFF 3D printing technique was patented in 1989 [13] and it bloomed up after its patent expired in 2009. After the technique became available to the general public, it has been used in different fields: automation, aeronautics, medicine, etc. Regarding the biomedical area, different applications could be highlighted, such as implants [14,15], 3D surgical planning prototypes [16,17], scaffolding [18,19], and regeneration of tissues [20].

There are numerous materials in the market for FFF 3D printing, for example acrylonitrile butadiene styrene (ABS) or Nylon, being PLA one of the most widely used, both alone and in combination with other materials such as wood, metals, or ceramics. Very little data have so far been published on systematic studies regarding the use of metal-filled filaments, since selecting compatible filler materials for the sake of improving the performance of polymeric composite materials is a difficult task [21]. In the present study, copper-, bronze-, and steel-filled PLA filaments are studied. Copper has an excellent heat and electric conductivity, it is easy to machine, bio-fouling resistant, and corrosion resistant [22]. Bronze alloy consists primarily of Cu, commonly with between 12 and 12.5% of Sn. It is a ductile alloy. Stainless steel is made of iron with typically a few tenths of carbon percentage, and with anti-corrosion elements such as Ni or Cr. It has high tensile strength, high corrosion resistance, and high biocompatibility. Therefore, it is used in a wide range of biomedical applications such as prostheses.

Regarding the mechanical properties of the metal-filled filaments, in some cases, increasing the metal content reduces the tensile strength and increases the thermal conductivity of the composite material studied. For example, Mohammadizadeh et al. [23] manufactured PLA filaments that contained copper, bronze, stainless steel, high carbon iron, and aluminum powders. They stated that the mechanical proper4ies of copper-filled-PLA were worse than those of PLA 3D printed parts. Additionally, they showed that the larger the layer height was, the lower the tensile strength, elastic modulus, and yield stress were. On the contrary, in different works, the mechanical properties were observed to increase when adding metals to the base polymer. Liu et al. [24] found that ceramic, copper, and aluminum-based PLA composite parts had similar or even superior mechanical properties when compared to bare PLA-made parts. Fafenrot et al. [25] developed polymer-metal materials 3D printed by FFF and concluded that the mechanical properties were similar to those of the PLA parts. On the other hand, there are other available options for the polymer matrix such as the use of ceramics. For instance, glass fiber-reinforced PLA can be employed in a wide range of applications, particularly in the biomedical, energy, and electronics industry [26]. In another example, Mahmoud et al. [27] studied the incorporation of two carbon fillers into the polypropylene: carbon nanotubes and synthetic graphite. The results showed that graphite-filled composites are more conductive than carbon nanotubesfilled composites. The flexural and tensile strength for both composites increased with the increase in the filler materials weight percentage. Later, the same authors [28] showed that flame-retardant MPP (melamine polyphosphate) had remarkable effects on the mechanical properties of the LLDPE (low-density polyethylene) composites. Five weight percentages of MPP were embedded into LLDPE, ranging from 5 to 30 wt%. It was concluded that the Young's modulus increased, and the tensile break strength and the tensile yield strength increased monotonically with the increase in MPP content.

The addition of metal components to the polymers used in FFF printing opens a new world in different fields such as bioengineering, but more knowledge needs to be obtained on the optimization of the production of these biomaterials for the fabrication of novel scaffolds. Although there are some studies about cell growth on 3D printed ceramic zirconia toughened alumina (ZTA) scaffolds [29], few and non-concluding studies have been done with metal-filled polymeric materials. Moreover, previous studies have not focused on the biological response of cells in metal-PLA 3D printed parts, as cells were cultured on metal-based scaffolds, such as titanium 3D printed bases [30]. On the other hand, several studies have evaluated the cytocompatibility of 3D printed iron-based scaffolds for bone regeneration [11,31,32] so we hypothesize that composite polymers incorporating metals will be appropriate for cell growth, since metals and alloys have been used extensively as bone substitutes [33,34]. These iron-based alloys have better mechanical properties than those based on lighter materials, such as magnesium.

Porosity is another key parameter that must be taken into consideration during the design and synthesis of a biomaterial [35]. The 3D printed porous materials should ideally fulfil conditions such as biocompatibility, noninflammatory response, tunable biodegradability, appropriate mechanic properties, defined pore structure, and, above all, promote a health improvement [36].

This work presents the characterization of three metal-reinforced PLA biomaterials for 3D printing biomedical scaffolds regarding porosity, surface roughness and cell culture. For that purpose, first the surface of the parts was analyzed, and the line spacing was measured. Surface roughness was then measured on the upper surface of the specimens. Regarding biological characterization, human-derived bone marrow mesenchymal stromal cells (hBM-MSC) were cultured on the composite scaffolds to assess their biocompatibility and the effect of the 3D printing scaffolds on determining cell fate, specifically in osteogenic differentiation.

#### **2. Materials and Methods**

#### *2.1. Materials*

The materials used in the present study were metal-filled PLA filaments of 2.85 mm diameter manufactured by ColorFabb (Belfeld, Netherlands). The specific materials used were: (1) steel-filled PLA, (2) bronze-filled PLA, and (3) copper-filled PLA. As stated by the manufacturer, these materials were developed for aesthetic purposes and need a polishing treatment after being printed if a brilliant appearance is to be required. All reagents were purchased from Sigma-Aldrich (Saint Louis, MO, USA), unless specified otherwise.

#### *2.2. 3D Printing Process*

Parts were additively manufactured by using a Sigma R19 3D printer (BCN3D Technologies, Gavà, Spain). The 3D printing parameters are presented in Table 1. Figure 1 shows a scheme of an FFF 3D printer. Cura BCN3D software was used to generate the G-code that is required to print the parts.

**Table 1.** 3D printing parameters.

**Figure 1.** FFF 3D printer scheme.

Disk-shaped samples of 6 mm in diameter and 2 mm in height were manufactured. The linear infill pattern was selected, using three different line spacing values: 0.6 mm, 0.7 mm, and 0.8 mm (Figure 2). The shell width was set to 0.4 mm and the bottom width was set to 1.2 mm. No top layer was used.

**Figure 2.** Disk-shaped 3D printed samples with line spacing of 0.7 mm of (**a**) bronze-filled PLA, (**b**) copper-filled PLA, (**c**) steel-filled PLA. Scale bars correspond to 1 mm.

#### *2.3. Pictures*

Pictures of the 3D printed scaffolds were obtained by using a Leica S8AP0 binocular magnifier (Leica Camera AG, Wetzlar, Germany) with 8× (Figure 2) and 16× (Figure 3) magnification, respectively.

#### *2.4. Porosity*

The porosity of the disk samples was quantified from the images of the cross section of the scaffolds, assuming that the length of the pores corresponds to the length of the sample and using Equation (1):

$$Pt = \frac{Vp}{Vt} \tag{1}$$

where *Pt* is the porosity, *Vp* the pore volume, and *Vt* the total volume of the scaffold. The software used for cross-sectional images quantification was ImageJ.

#### *2.5. Roughness*

Roughness was measured with a Talysurf 2 contact roughness meter from Taylor Hobson Ltd., Leicester, UK. A diamond tip was used with tip angle of 90◦ and tip radius of 2 μm. The measuring force was 0.8 mN and speed was 0.5 mm/s. A Gaussian filter was employed. A cut-off value of 0.8 mm was used according to ISO 4288 [37]. Total sampling length was 4.8 mm (6 × 0.8 mm). Roughness was measured on the upper surface of the disks, along the generatrices of the filaments in the two perpendicular directions, in order to assess if there were differences regarding their surface finish. As an example, the steel-filled samples were measured with line spacing 0.6 and 0.7 mm. Two different samples were measured for each line spacing value.

The roughness parameters that were analyzed in this present study are:

− Arithmetical mean roughness value or arithmetical mean of the absolute values of the profile deviations from the mean line of the roughness profile (Ra) (Equation (2)), which is one of the most commonly employed parameters in industry;

$$\text{Ra} = \frac{1}{L} \int\_{0}^{L} |Z(\mathbf{x})| d\mathbf{x} \tag{2}$$

$$\text{Rku} = \frac{1}{R\_q^4} \left[ \frac{1}{L} \int\_0^L Z^4(x) \, dx \right] \tag{3}$$

− Skewness (Rsk), which measures the symmetry of the profile (Equation (4)).

$$\text{Rsk} = \frac{1}{R\_q^3} \left[ \frac{1}{L} \int\_0^L Z^3(\mathbf{x}) \, d\mathbf{x} \right] \tag{4}$$

These parameters are defined in the UNE-EN-ISO 4287:1999 standard [38].

#### *2.6. Human Bone Marrow-Derived Mesenchymal Stromal Cells Culture on the Developed Scaffolds*

Primary human Bone Marrow-Derived Mesenchymal Stromal cells (hBM-MSCs, ATCC PCS-500-012, ATCC, Manassas, VA, USA) were expanded following the manufacturers' instructions. Cells from passage 3–6 were used for all the experiments presented herein.

Printed parts were coated by incubating with rat tail-derived type I collagen at a concentration of 0.1 mg/mL for 30 min at 37 ◦C. hBM-MSCs were cultured on the scaffolds at a seeding density of 4 × 104 cells/cm2 for 24 and 72 h. Cells seeded on uncoated parts were cultured in parallel. At the defined time points, cells were fixed with paraformaldehyde (PFA) 4% for further immunohistochemical imaging.

In a subsequent set of experiments, hBM-MSCs were cultured with a seeding density of 9.4 × 104 cells/cm2 for 1, 4, and 7 days, respectively, on the different collagencoated/uncoated samples and then fixed with PFA 4%.

Finally, to evaluate the impact of the developed scaffolds on the cell fate, hBM-MSCs were cultured with αMEM (12509069, Gibco, Waltham, MA, USA) 10% FBS (Gibco) on the steel-filled PLA scaffolds with 0.6, 0.7, and 0.8 mm line spacing, respectively, with a seeding density of 1.3 × 104 cells/cm2. Control cells were cultured in parallel on conventional culture plates. After 21 days, the cells were fixed with PFA 4% for further analysis.

#### *2.7. Immunohistochemical Analysis*

After PFA fixation, samples were permeabilized with Triton 0.1%, blocked with 10% FBS solution, and incubated overnight at 4 ◦C with primary antibodies (anti-hOsteocalcein 967801, R&D Systems, Minneapolis, MN, USA) and subsequently, incubated for 2 h at 37 ◦C with the secondary Alexa 488 anti-rabbit antibody for differentiation studies. For morphology analysis, nuclei were stained with NucBlue (Thermo Scientific, Waltham, MA, USA) and actin cytoskeleton with phalloidin (Thermo Scientific, Waltham, MA, USA). Images were acquired with a Nikon D-Eclipse Ci confocal microscope (Nikon, Tokyo, Japan) with 10× and 20× Plan Apo objectives (Nikon).

#### **3. Results and Discussion**

#### *3.1. 3D Printed Samples*

Figure 3 shows the pictures of bronze-filled, copper-filled, and steel-filled samples manufactured with 0.6, 0.7, and 0.8 mm line spacing, respectively. Achieving 0.6 mm line spacing was more difficult than 0.8 mm, because of smaller pores. Despite that, appropriate scaffolds were achieved for the three different materials.

Additionally, scaffolds for biomedical applications should have a porous architecture. This porosity provides the necessary environment for promoting cell migration, proliferation, etc. [39].

**Figure 3.** Surface of the samples with 16× magnification: (**A**) bronze-filled 0.6 mm, (**B**) bronze-filled 0.7 mm, (**C**) bronze-filled 0.8 mm, (**D**) copper-filled 0.6 mm. (**E**) copper-filled 0.7 mm, (**F**) copperfilled 0.7 mm, (**G**) steel-filled 0.6 mm, (**H**) steel-filled 0.7 mm, (**I**) steel-filled 0.8 mm. The scale bars correspond to 1 mm.

#### *3.2. Porosity*

As shown in Figure 4, the higher the line spacing, the higher the porosity of the 3D printed scaffolds is. Among the different scaffolds, for a certain line spacing value, most porous scaffolds are the bronze-filled ones, followed by the copper-filled ones, although they were manufactured with the same 3D printing conditions:

**Figure 4.** Porosity of the 3D printed scaffolds: bronze-filled, copper-filled and steel-filled. *N* = 3.

#### *3.3. Roughness*

Table 2 presents the roughness results for both the internal and the external surfaces of the 3D printed samples. One measurement was performed on each surface.

**Table 2.** Roughness on the external and internal layer of 3D printed steel-filled samples.

Figure 5 depicts the roughness profiles of samples with a line spacing of 0.6 and 0.7 mm, on the external (first) and internal (second) layers, respectively, starting from the top of the part. Higher Ra values were obtained on the external (first) layer (Table 2 and Figure 5a,c) than on the internal (second) layer (Table 2 and Figure 5b,d). On the external layer, slightly higher Ra values (up to 35.04 μm) were found for line spacing 0.7 mm than for line spacing 0.6 mm (up to 25.60 μm). The Rz parameter shows a similar trend than Ra.

**Figure 5.** Roughness profiles of the steel-filled 3D printed samples. (**A**) 0.6 mm external roughness (first layer). (**B**) 0.6 mm internal roughness (second layer). (**C**) 0.7 mm external roughness (first layer). (**D**) 0.7 mm internal roughness (second layer).

Rku values around 3 were found in all cases, corresponding to a normal distribution of the roughness heights in each profile. On the external layer, slightly negative Rsk values were obtained, corresponding to longer valleys than crests. On the contrary, on the second layer Rsk values are close to 0, corresponding to symmetric profiles.

The external or first layers show more regular roughness profiles (Figure 5a,c) than the internal or second layers (Figure 5b,d).

#### *3.4. Human Bone Marrow-Derived Mesenchymal Stromal Cells Cultured on the Developed Scaffolds*

hBM-MSCs showed good adhesion to both collagen-coated and untreated 3D printed steel-filled PLA samples. Cells cultured for both 24 and 72 h were well-adhered to the 3D printed PLA composites (Figure 6a,b), with no observed differences between both conditions.

70

**Figure 6.** (**a**,**b**) Adhesion test to PLA with and without collagen coating. (**c**–**e**) Viability test at day 1 of cell culture. (**f**–**h**) Viability test at day 4 of cell culture. (**i**–**k**) Viability test at day 7 of cell culture.

Steel-filled PLA samples showed a very good cytocompatibility, especially on 0.6 and 0.7 line spacing (Figure 6c,d,f,g,i,j). On the contrary, copper-filled and bronze-filled PLA presented higher cytotoxicity since there were no cells adhered to the scaffolds after 24 h of culture (data not shown).

Steel-filled scaffolds showed high biocompatibility, unlike copper-filled and bronzefilled materials. This is in concordance with Kuroda et al. [39]. Additionally, the best biological behavior was found with the lowest porosity achieved (0.6 mm line spacing). This is in accordance with data presented by Chen et al. [40], who concluded that samples with 30% porosity exhibit the best biocompatibility, which were the lowest porosity scaffolds of their research.

The different scaffolds manufactured by means of FFF showed to have different behavior. As mentioned, bronze-filled as well as copper-filled scaffolds presented high cytotoxicity since there were no cells adhered to the scaffolds after 24 h of culture.

It is interesting to highlight that there were no differences observed when a specific protein coating was used in the parts. Cells form specific adhesions to the collagen protein while they are expected to form unspecific adhesions to uncoated materials. From the experiments presented herein, it can be concluded that steel-filled PLA promotes the formation of unspecific adhesion in MSCs, while this is not happening with copper- or bronze-filled polymers.

#### *3.5. hBM-MSCs Differentiated towards Osteoblasts When Cultured on the Developed Scaffolds*

hBM-MSCs cultured on the steel-filled 3D printed scaffolds (0.6 and 0.7 mm line spacing) without osteogenic supplements showed the presence of osteocalcin after 21 days of culture (Figure 6). Moreover, cells exhibited a broad spreading area compared with those cultured under conventional culture conditions (Figure 7).

**Figure 7.** Osteocalcin detection by immunofluorescence in hBM-MSCs (at day 21) cultured on (**a**) conventional culture conditions, steel-filled PLA scaffolds of (**b**) 0.6 mm and (**c**) 0.7 mm line spacing. Osteocalcin (green) and nucleus (blue). Scale bars correspond to 100 μm.

The images show the classical spindle-like shape of MSCs, so it seems that the steelfilled 3D printed structures are a suitable scaffold for cell culturing. The enlarged phenotype is similar to differentiated osteoblasts when cultured on rigid substrates [41–43].

#### **4. Conclusions**

In the present work, results are presented for the cell growth of stem cells on metalfilled PLA composites that were printed with a grid structure by means of the FFF technique. Three different composites were tested: bronze, copper, and stainless steel, respectively. The main conclusions are as follows:

In future work, the effect of the use of other infill patterns on both cell growth and the mechanical strength of the structures will be addressed.

**Author Contributions:** Conceptualization, I.B.-C. and J.O.; methodology, A.U., H.S.-F., A.D.-F. and A.T.-O.; software, A.D.-F. and H.S.-F.; validation, I.B.-C. and J.O.; formal analysis, A.U., H.S.-F., I.A. and A.T.-O.; investigation, A.U., H.S.-F., I.A. and A.T.-O.; resources, I.A., I.B.-C. and J.O.; data curation, A.T.-O.; writing—original draft preparation, all; writing—review and editing, all; visualization, I.B.- C., A.D.-F. and J.O.; supervision, I.B.-C. and J.O.; project administration, I.B.-C. and J.O.; funding acquisition, I.B.-C., I.A. and J.O. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research was co-funded by the European Union Regional Development Fund within the framework of the ERDF Operational Program of Catalonia 2014–2020 with a grant of 50% of total cost eligible, project BASE3D, grant number 001-P-001646. Research was also funded in part by the Spanish Ministry of Science, Innovation and Universities, grants numbers PID2019-108958RB-I00/ AEI/10.13039/501100011033 and PGC2018-097323-A-I00.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Data supporting the findings of this study are available from the corresponding authors upon reasonable request.

**Acknowledgments:** The authors thank Ramón Casado for his help with the experimental tests.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **Modeling Impact Mechanics of 3D Helicoidally Architected Polymer Composites Enabled by Additive Manufacturing for Lightweight Silicon Photovoltaics Technology**

**Arief Suriadi Budiman 1,2,3,4,\*, Rahul Sahay 3,5,\*, Komal Agarwal 3, Rayya Fajarna 4, Fergyanto E. Gunawan 4, Avinash Baji 6 and Nagarajan Raghavan 5**

	- **\*** Correspondence<EMAIL_ADDRESS>(A.S.B<EMAIL_ADDRESS>(R.S.)

**Abstract:** When silicon solar cells are used in the novel lightweight photovoltaic (PV) modules using a sandwich design with polycarbonate sheets on both the front and back sides of the cells, they are much more prone to impact loading, which may be prevalent in four-season countries during wintertime. Yet, the lightweight PV modules have recently become an increasingly important development, especially for certain segments of the renewable energy markets all over the world such as exhibition halls, factories, supermarkets, farms, etc.—including in countries with harsh hailstorms during winter. Even in the standard PV module design using glass as the front sheet, the silicon cells inside remain fragile and may be prone to impact loading. This impact loading has been widely known to lead to cracks in the silicon solar cells that over an extended period of time may significantly degrade performance (output power). In our group's previous work, a 3D helicoidally architected fiber-based polymer composite (enabled by an electrospinning-based additive manufacturing methodology) was found to exhibit excellent impact resistance—absorbing much of the energy from the impact load—such that the silicon solar cells encapsulated on both sides by this material breaks only at significantly higher impact load/energy, compared to when a standard, commercial PV encapsulant material was used. In the present study, we aim to use numerical simulation and modeling to enhance our understanding of the stress distribution and evolution during impact loading on such helicoidally arranged fiber-based composite materials, and thus the damage evolution and mechanisms. This could further aid the implementation of the lightweight PV technology for the unique market needs, especially in countries with extreme winter seasons.

**Keywords:** 3D helicoidal architecture; fiber-based polymer composite; impact resistance; lightweight photovoltaics (PV); numerical modeling

#### **1. Introduction**

Lightweight photovoltaic (PV) modules are becoming increasingly important to certain sectors of the renewable energy market. Lightweight PV technology could potentially help address global climate and sustainability issues by being deployed in countries where electricity infrastructure is often lacking in very remote, poor locations separated by oceans [1–4]. For such a region, centralized energy sources may not be the best solution. By

**Citation:** Budiman, A.S.; Sahay, R.; Agarwal, K.; Fajarna, R.; Gunawan, F.E.; Baji, A.; Raghavan, N. Modeling Impact Mechanics of 3D Helicoidally Architected Polymer Composites Enabled by Additive Manufacturing for Lightweight Silicon Photovoltaics Technology. *Polymers* **2022**, *14*, 1228. https://doi.org/10.3390/ polym14061228

Academic Editors: Hamid Reza Vanaei, Sofiane Khelladi, Abbas Tcharkhtchi and Luigi Sorrentino

Received: 24 January 2022 Accepted: 12 March 2022 Published: 18 March 2022

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**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

reducing the cost of transporting and installing solar PV systems, lightweight PV could play a critical part in building the kind of self-sufficient power infrastructure that is desperately desired in distant, rustic areas of developing countries. In order to be transported to remote locations with limited road and mobility infrastructure, PV power infrastructure must be lightweight. Lightweight PV modules are desired both for use in urban structures in advanced countries [5–8] and for easy set up in distant and impoverished locations in unusual areas of the developing world.

Furthermore, many large structures, such as exhibition halls, industrial plants, supermarkets, farms, etc., have a large footprint with few supporting pillars, resulting in a rooftop with low load-supporting capacity. Such rooftops require lightweight PV modules. If not, the cost of reinforcing such building before installing the heavy glass-based PV modules would make the renewable energy venture (building plus PV power infrastructure) inefficient and unappealing to potential business interests [5,9,10]. Lightweight PV has been extensively and comprehensively researched elsewhere as part of building-integrated PV and for predominantly urban building applications [5–8,11].

In addition, as silicon will be the predominant PV technology for the foreseeable future [5,6], we need to aid lightweight silicon-based PV modules. Structural strength, and impact resistance of lightweight PV modules, especially against heavy winds and hailstorms in four-season countries in Europe and North America [12,13], is one of the most important technological concerns in the development of lightweight PV modules. The front panels of a PV module can be made strong enough not to break under impact loads, such as hailstorms; nevertheless, the energy is transferred straight to the underlying material—first to the encapsulation (usually ethylene vinyl acetate), which would simply give way. The energy was then passed to the brittle silicon cells, which are particularly susceptible to such point impact loading, causing cracks to develop (nucleate) and/or propagate [7,13,14]. As a result, electrical performance degrades gradually or dramatically, which can lead to hotspots and potentially dangerous situations (such as fires, etc.).

Even though the idea of lightweight PV modules is tempting, nevertheless, it is not currently a practical option owing to concerns about structural stiffness and reliability [5–7,12,13]. Numerous commercial lightweight PV systems (even those that meet IEC/UL standards) have limited operational time [6,7]. Despite the fact that silicon cells remain internally delicate and extremely vulnerable to certain impact loads, recent studies with substantial material advances and ingenious design have allowed great improvements in the impact resistance of numerous polymer-based substrates used as front panels (as an alternative of glass) in conventional PV modules [13–16].

Natural structural materials, such as those found in mantis shrimp and nacre, have been demonstrated to provide superior mechanical and, particularly, impact resistant properties [17–19]. For example, the dactyl club of mantis shrimps has a 3D construction with a helicoidal shape that can disperse energy through quasi-plastic compression responses, providing a barrier to the spread of microcracks throughout recurring impacts [20–23]. Recent publications from our own research have stated higher impact resistance of such materials [24,25]. Further, the comparison in terms of mechanical properties of the 3D helicoidally aligned layered materials vs. layered materials without rotation offsets (or in other words, unidirectional layered materials) has been shown experimentally in our own previous publication by Agarwal et al. [26]. In addition, such comparisons have also been reported experimentally by other research groups using various materials (glass filament epoxy [27], carbon epoxy [27,28], and fiber sizes [27–29]). Agarwal et al. [26] reported for the smaller scale (using the custom electrospinning set up as an additive manufacturing methodology) of the other groups' studies, in terms of fiber diameters. Furthermore, comparison in terms of mechanical properties of the 3D helicoidally aligned layered materials vs. a sample of same thickness as the layer stack (or in other words, bulk sample) has also been demonstrated experimentally in our own previous publication [26], in addition to other previous reports [27–29], chiefly from Kisailus et al. [29]. Again, our group's studies—started by Agarwal et al. [26] to the more recent Sahay et al. [25] and Budiman et al. [30]—simply

further pursued this line of investigation into the smaller scale fibers (using the custom electrospinning set up as an additive manufacturing methodology) and for the potential application of the unique materials for silicon-based PV module technology. The general outcome of such studies was that the layered structure of such materials, which consists of helicoidally aligned 3D fibers compared to typical layered structure/bulk material, would effectively absorb the impact energy and transfer very slight energy to the fragile silicon solar cells. This would allow novel lightweight PV module design with improved impact resistance and structural reliability (based on polymer materials for the front and back panel) particularly against cracks in the silicon cell.

The objective of this work is to provide a numerical analysis and modeling to predict how stress changes during impact loading in 3D-architectured layered polymer systems with helicoidally oriented fibers. This model demonstrates the fundamental feasibility of the proposed concept, namely, the use of 3D-architected layered polymer assemblies of helicoidally oriented fibers to protect silicon solar cells against the nucleation and proliferation of cracks caused by impact loads (e.g., from hailstorms) in the design of lightweight PV modules. We are expanding our approaches to enable this unique material for use in lightweight PV technologies. We are building on our earlier studies on novel materials [24–26,31] as well as numerical modeling of stresses in PV module design [32–37]. Furthermore, the design of lightweight PV modules would allow the integration of PV into curved or contoured surfaces, resulting in a more appealing design for integrating PV into urban structures.

#### **2. Methodology**

#### *2.1. Material*

The material used in making the multilayered composite plate is PVDF-HFP fibers, as was used in the experimental impact testing [30]. PVDF-HFP is polyvinyl alcohol (PVA) with MW = 98,000, polyvinylidene fluoride-co-hexafluropropylene (PVDF-HFP) with MW = 400,000, acetone, and dimethylacetamide (DMAc) were obtained from Merck, Singapore, as reported in our previous study [38]. More complete information about the materials used in the experimental impact testing can be found in [24].

The multilayered composite plate was modeled with fiber alignment in the layers changed from 90◦ (i.e., grid pattern) to 45◦ and then later to 15◦ to simulate the experiments (as illustrated in Figure 1), in which the impact resistance increase was observed [30]. Thus, in the present study, three multilayered composite plates were modeled:

**Figure 1.** Schematic illustration of the multilayered composite (**a**) consisting of layers with 3Darchitected helicoidally aligned fibers; (**b**) with fiber alignments of [0◦, 45◦, 90◦, 135◦, 180◦, 225◦, 270◦, 315◦, 360◦] described in our previous report [24].

The dimensions of the composite plate in the Finite Element Model (FEM) used are 100 mm × 100 mm × 0.4 mm (length × width × thickness). Impact loading uses a steel ball with a radius of 30 mm. The mass of the ball is 0.1 kg with an impact velocity of 1 mm/ms. These parameters scale with the experimental impact testing as reported in [30]. Mechanical properties of the PVDF-HFP fiber were obtained from our previous publication [24], which reported in detail the materials used in the experimental impact testing in Ref. [30] as well as in the present numerical simulation study (See Table 1).

**Table 1.** Mechanical properties of PVDF-HFP fiber composite material (in each layer, with 0◦ and 90◦ represent longitudinal and transverse direction, respectively, of the fibers) used in simulation modeling [24].

#### *2.2. Electrospinning-Based Additive Manufacturing (Es-AM)*

Only recently has electrospinning-based additive manufacturing (Es-AM) technology enabled such intricate 3D designs [25,31]. Near-field electrospinning (NFES) has been used to fabricate helicoidally oriented fiber layers as an additive manufacturing approach [7,8]. Agarwal et al. [24] describe in detail the fabrication of helicoidally oriented fiber layers with different angular orientations. NFES typically creates helicoidally oriented fiber layers by depositing one-dimensional fibers at precise locations in a controlled manner and then stacking the fibers layer by layer with angular offsets to create a 3D helicoidally arranged synthetic structural composite (HA-SSC) (see Figure 1). Typically, in the case of 90◦ HA-SSC90, the fiber layers were deposited with 90◦ angular offsets starting from 0◦, 90◦, 180◦, 270◦, to 360◦—as mimicked in this numerical simulation study as Composite Plate A. Similarly, in HA-SSC45 with 45◦ angular offsets, the fiber layers were deposited at 0◦, 45◦, 90◦, 135◦, 180◦, 225◦, 270◦, 315◦, 360◦—as mimicked in this numerical simulation study with Composite Plate B, and schematically illustrated further in Figure 1b. In HA-SSC15 with 15◦ angular offsets, the fiber layers were deposited at 0◦, 15◦, 30◦, 45◦, 60◦, 75◦, 90◦, 105◦, 120◦, 135◦, 150◦, 165◦, 180◦, 195◦, ... , 360◦—as mimicked in the present study as Composite Plate C. The HA-SSCs samples were between 230 and 250 μm thick. In the experimental impact tests published in [30], we used only HA-SSC15 and HA-SSC45—with fiber orientation at rotation angles of 15◦ and 45◦, respectively—from the different variants of the HA-SSC samples we reported in [24]. See [30] for more information on the impact testing of HA-SSCs.

This composite fabricated and reported in [24] is not yet ready for incorporation into PV power infrastructure. An optically transparent material is required for practical PV application with similar transmission of sunlight (in terms of intensity and range of suitable wavelengths). Nevertheless, as elucidated in the Introduction and the Materials section, the emphasis of the study is to establish the viability of the idea of improved impact resistance through the 3D-architected impact-resistant encapsulant, not its complete industrial incorporation into PV module design.

#### *2.3. Finite Element Modeling of the Impact Loading on Multilayered Composite with 3D Helicoidal Architecture*

Finite element (FE) models allow us to estimate the mechanical stress that develops during impact loading of the multilayered fiber-based composite materials with the 3D helicoidal arrangement to predict the damage mechanisms associated with the stress evolution during impact loading. However, to keep computational complexities to a minimum while still gaining the fundamental deformation mechanics, we approximate the fiber-based composite layers with laminate geometry that has anisotropic mechanical properties in a certain direction in each layer. We used thin film geometry in the FE model [38]. We used commercially available general-purpose LS-DYNA (4.6.19, LSTC (Ansys, Inc.), Canonsburg, PA, USA) FE software to obtain the stresses induced in the multilaminate structures to understand the evolution of stresses during impact loading. The FE model uses shell elements of a conventional thin film. The thin film sample was modeled using regular 8-noded quadrilateral elements (CPEG8). A generalized plane strain condition was assumed.

Figure 2 shows the multilayered composite plate, which in the LS-DYNA software was modeled using 4N plat on shape meshes. The size of the mesh is 2 mm. All the edges of plates are simply supported. The plates are subjected to impact loading on the middle of a surface in the form of a ball dropping onto the plate—mimicking the ball-dropping impact test as described in our previous report on a similar multilayered composite with 3D helicoidal architecture [30]. The impact loading was restricted to the z-direction. Data on displacement in the z-direction is collected on each node element of the middle layer of the multilayered composite plate.

**Figure 2.** Schematic of the 2D plane–strain FE model (**a**) with thin film shell elements of multilaminate structures under impact loading mimicking the ball-dropping impact test (**b**) as described in our previous report [30].

The mechanics of the thin film plate during the impact loading here were modeled with the surrogate approach [39]. The strain in the film here due to the impact loading was surrogated by the scaling approach (size of the ball with respect to size of the multilayered composite geometry are equivalent between the model and the experiment) to keep the complexity to a minimum, while still obtaining important insights about stress

evolution and damage distribution mechanisms during the impact loading. Thus, the relative magnitude of the stresses with respect to time during impact loading represents the evolution of energy distribution during the ball-dropping impact testing, with actual silicon solar cells under the 3D-architected layered polymer structures consisting of helicoidally aligned fibers, as reported in [30]. Uniform deformation over the volume of each layer of the multilayered composite is assumed during the process—leaving only the asymmetric deformation due to the anisotropy of each of the laminate due to different rotational angle of fiber alignment in each layer of the 3D helicoidal architected polymer composite. Fully elastic behavior of each of the laminate was assumed, which is reasonable given the scaling approach and the actual experimental results [30]. Interfaces between laminates were modeled as surface-to-surface tie constraint [39].

#### **3. Results and Discussion**

#### *3.1. Results of the Finite Element Simulation: Impact Contact between the Steel Ball and the Plate*

The position of the steel ball (modeled as sphere) with respect to the composite plate for various time instances during the impact contact is shown in Figure 3 (from the side of the model for Composite Plate A). The steel ball impinges the plate at an initial velocity of 1 mm/ms. The simulation shows the sphere has started contact with the plate surface at the time of t = 1.4399 ms. The simulation then shows in Figure 3 that the steel ball has exerted sufficient force to deform the plate laterally at t = 1.8899 ms. Consequently, at time t = 5.5788 ms, the plate has reached its maximum deformation. Lastly, at the time instant of t = 9.9811 ms, the steel ball bounced back and lost contact with the plate. Although the above time evolution was shown for Composite Plate A, similar occurrences represent the impact contact evolution between the steel ball and the plate in the other composite plates studied in the present study (Composite Plates B and C).

**Figure 3.** FE simulation showing impact contact between the sphere and the composite plate, (**a**) t = 0.4388 ms, (**b**) t = 1.4399 ms, (**c**) t = 1.8899 ms, (**d**) t = 5.5788 ms, (**e**) t = 7.8338 ms and (**f**) t = 9.9811 ms.

#### *3.2. Results of the Finite Element Simulation: Deformation of the Composite Plate*

The distribution of the Max Principal Stress (or the first principal stress, S1) with respect to time during the impact loading is shown in Figures 4–6 for Composite Plates A, B, and C, respectively. Initially, when the sphere first impinged on the plate, Figures 4–6 all show that stress concentration at the plate center. This is expected, as it is where the impact of the sphere on the plate occurred. The stress that is initiated at the plate center spreads out quickly to the whole plate and is absorbed by the composite plate structures at different rates.

**Figure 4.** The distribution of the maximum principal stress, S1, with respect to time for Composite Plate A during impact loading, (**a**) 2.3785 ms, (**b**) 4.8788 ms, (**c**) 6.7438 ms, (**d**) 7.3798 ms, (**e**) 8.4985 ms and (**f**) 9.3785 ms.

**Figure 5.** The distribution of the maximum principal stress, S1, with respect to time for Composite Plate B during impact loading, (**a**) 2.2835 ms, (**b**) 3.4785 ms, (**c**) 4.7384 ms, (**d**) 6.3218 ms, (**e**) 7.4982 ms and (**f**) 7.8849 ms.

**Figure 6.** The distribution of the maximum principal stress, S1, with respect to time for Composite Plate C during impact loading, (**a**) 2.1589 ms, (**b**) 3.1785 ms, (**c**) 4.6478 ms (**d**) 5.5679 ms, (**e**) 6.8985 ms and (**f**) 7.3798 ms.

Figure 4 clearly shows that the stress distribution during impact loading follows the four-symmetry that is created by the grid pattern of Composite Plate A, which consists of layers with fiber alignments of [0◦, 90◦, 180◦, 270◦, 360◦]. Stress concentrated at the center of the plate and reaches the maximum principal stress (which occurred in all composite plates in the present study in the *z*-direction) of 0.095 GPa at t = 7.3798 ms. Over time, the stress reduces and finally reaches the zero-stress state after about 9.5 ms.

Figure 5 clearly shows that the stress distribution during impact loading follows the double four-fold symmetry that is created by the grid pattern of Composite Plate B, which consists of layers with fiber alignments of [0◦, 45◦, 90◦, 135◦, 180◦, 225◦, 270◦, 315◦, 360◦], as illustrated in Figure 2b. Figure 5 also shows that the stress concentrated at the center

of the plate and reaches the maximum principal stress (in the *z*-direction) of 0.035 GPa at t = 4.7384 ms. Over time, the stress reduces and finally reaches the zero-stress state after about 8 ms. These parameters clearly demonstrate a much higher dissipation rate of damage in Composite Plate B, as compared to Composite Plate A. The maximum principal stress shown here in Figure 5 (for Composite Plate B) is lower, and the rate in which the stress is distributed and subsequently reduced to zero upon the same impact loading (impact load and loading rate) is much higher. This suggests that Composite Plate B with 3D architecture consisting of layers with fiber alignments of [0◦, 45◦, 90◦, 135◦, 180◦, 225◦, 270◦, 315◦, 360◦] is much more effective and efficient in absorbing and dissipating impact energy and damage compared to Composite Plate A.

Figure 6 further suggests that Composite Plate C with 3D architecture consisting of layers with fiber alignments of [0◦, 15◦, 30◦, 45◦, 60◦, 75◦, 90◦, 105◦, 120◦, 135◦, 150◦, 165◦, 180◦, 195◦, ... , 360◦] is also much more effective and efficient in absorbing and dissipating impact energy and damage compared to Composite Plate A. It also exhibits an incremental increase in the absorption and dissipation rate of impact energy and damage compared to that of Composite Plate B. Figure 6 shows that the stress concentrated at the center of the plate and reached the maximum principal stress (in the *z*-direction) of 0.034 GPa at t = 4.6478 ms. Over time, the stress reduces and finally reaches the zero-stress state after about 7.5 ms. These parameters clearly suggest that the rotational angle of the 3D helicoidal architecture influences the impact resistance of the materials. The smaller rotational angles clearly play a key role in increasing the dissipation rate of the impact damage, although the increase could be moderated after some levels of rotational angles. This is further substantiated by the stress vs. time response data shown in Figure 7.

Figure 7 clearly indicates that the Composite Plates B and C have a much shorter time response in terms of absorbing the impact energy (and thus damage dissipation in the event of plastic deformation and fracture events). Furthermore, the much lower stress intensity in Figure 7b,c demonstrates that the smaller rotational angle steps are effective in quickly and efficiently distributing the stress concentration. Both Composite Plates B and C exhibit stress levels getting back to pre-impact level (zero stress state) within under 7.5 ms, compared to just under 10 ms for Composite Plate A. Composite Plates B and C also show maximal principal stress intensities of 0.035 GPa and 0.034 GPa, respectively. This is much lower stress level compared to that of Composite Plate A, which reaches 0.095 GPa (almost three times). It is evident through the stress evolution during the impact loading simulation that the smaller rotational angles play a key role in increasing the dissipation rate of the impact damage, although the increase could be moderated after some levels of rotational angles. The increase in the absorption rate of impact damage and dissipation rate of stress intensity with reduction of rotational angle in the 3D helicoidal architecture seems to taper off after 45◦. The FE simulation was conducted with the surrogate approach, thus while the absolute magnitudes still need to be further verified with experimental study using the same scales as in the modeling, the relative comparison between the effects of rotational angles in increasing the rate of absorption of damage and dissipation of stress intensity suggests self-consistency and excellent agreement with the experimental impact loading as reported in our earlier publication [30] and as will be further elaborated in the following section.

**Figure 7.** The time response of the distribution of maximum principal stress, S1, for (**a**) Composite Plate A, (**b**) Composite Plate B, and (**c**) Composite Plate C, during impact loading.

#### *3.3. Comparison with Impact Test of Photovoltaic (PV) Cells*

The experimental study on the impact-loading test of the 3D helicoidally architected polymer composite materials (with different rotational angles) has been reported in our recent publication [30]. In it, we use fragile silicon solar cells, which are highly susceptible to impact load underneath the polymer composite materials, such that the polymer composite materials act as a protection layer for the fragile silicon solar cells [30]. A customized impact testing setup (as illustrated in Figure 2b) was used to determine the impact resistance of such solar cells when protected by the samples [40]. More complete information about the impact test using the steel ball dropping method can be obtained in [30].

The fracture height indicates the height at which we began observing fracture of the silicon solar cells under the polymer composite materials. The nominal height of the bare Si solar cell group was found to be 25 (±5) cm, which was the real height at which all silicon cells fracture in the at least six times we repeated the tests (we performed the ball drop tests to more than six samples at this height—up to 11 samples). The uncertainty level of the experiment was ±5 cm. Consequently, the fracture data indicated the heights at which the fracture started. Heights greater than the abovementioned value would evidently fracture the silicon cells.

The fracture heights when we had typical encapsulant: ethylene vinyl acetate, on top of the silicon solar cells, was 50 ± 4 cm. The fracture heights are 69 ± 2 cm and 82 ± 4 cm for the Composite Plate B: HA-SSC15 and Composite Plate C: HA-SSC45, respectively [24,25]. Therefore, it was evident that the HA-SSC materials were the better protector of the silicon solar cells against impact loading, compared to the nominal EVA. Both HA-SSC composites (HA-SSC15 and HA-SSC45, respectively, Composite Plates B and C) enable significantly higher fracture heights and related specific potential energies (well beyond experimental uncertainties) before the underlying silicon solar cells fracture or trigger/spread catastrophic fracture events in ball drop tests. Because of their helicoidally oriented fiber-reinforced layered structures, both HA-SSC would efficiently disperse the impact energy and deflect the crack laterally by following the helicoidal progression of fiber directions, rather than immediately fracturing in a straight line through the thickness of the HA-SSCs [24,25,41]. Therefore, a smaller amount of impact energy is transferred to the underlying silicon solar cell.

The fractography images described in Ref [30] show abovementioned efficient mechanism for load dissipation as well as effective damage/energy absorption. In addition, both HA-SSC permit multiple crack lines on the silicon cells, signifying high dissipated impact energy and load transfer to the side (as evidenced by crack lines at different angles on the silicon cell surfaces). In typical monocrystalline silicon wafers, cracks seem to track the favored crystallographic directions of <110> linked to the weakest crystallographic planes of {111}, as has been extensively stated in the literature for both PV and other silicon-based devices [32–37,42,43].

The crack regularly encounters variations in the modulus of the fibers and the matrix material. The crack in HA-SSC travels along one fibrous layer, encounters a modulus deviation due to the matrix material present, penetrates further into the matrix, then reaches another fibrous layer in a dissimilar direction and plane, and finally diverges from its original route to follow a dissimilar path. The crack proliferates in multiple planes and orientations, rotates and twists inside and outside the fiber and matrix phases, and places greater stress on the fibers short of catastrophic failure. Due to the helicoidal network of the fibers in the composites, more energy is required for the propagation of the crack, so only a limited amount of impact energy/damage is transferred to the underlying silicon solar cells. According to Budiman et al. [30], the Composite Plate B: HA-SSC45 appears to be the most successful at consistently deflecting fracture/ impact damage along different angular orientations. The fracture heights show that the solar cells fracture at 82 ± 4 cm, which is the highest value among the samples investigated in this study. However, the study [35] paints a somewhat different depiction for the Composite Plate C: HA-SSC15. The breakage of the solar cells under HA-SSC15 happens at a lower height (69 ± 2 cm), resulting in a lower specific potential energy (and thus a lower impact energy/damage absorption rate).

In the present study, our FE simulation shows that the absorption rate of impact damage and dissipation rate of stress intensity increases with reduction of rotational angle in the 3D helicoidal architecture, although the effect seems to taper off after 45◦. The FE simulation here shows that Composite Plate C: HA-SSC15 exhibits faster absorption rate of impact damage and stress level reduction (back to pre-impact state), although by a much smaller margin compared to that between Composites Plate A and B (from rotational angle of 90◦ to 45◦). While this may not be in general agreement with our previous study reported in Budiman et al. [30], the FE simulation findings in the present study do agree very well with another earlier experimental report from our group, as described in Agarwal et al. [24]. The difference between these two studies is silicon solar cells were used under the HA-SSC

in the impact loading experiments in Budiman et al. [30], while glass coverslips were used in Agarwal et al. [24]. The HA-SSCs used were, however, the same [30].

We believe that the variance (in the ability to absorb impact damage between HA-SSC15 and HA-SSC45) is due to the monocrystalline silicon solar cells that were among the samples of HA-SSC, because we followed the same method of the ball-dropping test as [24,25], together with the identical size of the steel ball. Compared to the glass slip, the monocrystalline silicon samples exhibit a crystallographic dependency of the mechanical properties, such as the preferential occurrence of fractures. Consequently, the whole relationship amongst the helicoidal orientation in HA-SSC and fracture of monocrystalline silicon solar cells needs to be further explored, which is outside the scope of the current FE simulation. Nevertheless, FE simulation in the present study takes into consideration the fact that the material put under the HA-SSC was a silicon monocrystalline solar cell (with its anisotropy in the mechanical properties), and not a glass coverslip (which has mostly isotropic mechanical properties).

#### *3.4. Enabling Next-Gen Lightweight Photovoltaic (PV) Module Technology*

It is clear from the experimental results [24,30] and the FE simulation shown in the present study that the HA-SSCs are excellent at absorbing and dispersing impact energy/damage, protecting the delicate solar cells underneath from point impact loads to which silicon PV is particularly susceptible. The objective of this paper is to complement the previously published experimental data with FE simulation study on the basic viability of using HA-SSCs for PV encapsulation to shield the sensitive solar cells, particularly in the design of lightweight PV modules that are predominantly susceptible to impact loadings, such as hailstorms, as outlined in IEC 61215/61646 clause 10.17. Although many current studies have revealed that a variety of polymer-based materials [12–16] can be used to improve the fracture and impact resistance of lightweight PV modules; nevertheless, the encapsulation materials employed in these studies were all EVA—albeit with different thicknesses or slight variations, such as low curing temperature [44].

Due to the economic impact and production maturity of the entire lightweight PV module, EVA was clearly selected as the encapsulation material of choice in previous studies. Our previous experimental results, as well as the numerical simulation results reported in this manuscript, provide initial indication from the standpoint of basic technological viability for the use of other types of novel polymer films (i.e., not EVA) with unique 3D architectures to allow strong and impact resilient lightweight PV module designs.

Nevertheless, there are many potential technological challenges to be resolved that prevent this unique idea from being fully realized. First, the HA-SSCs created were not transparent.  It should be emphasized that the tests [30] and modeling of FE in the current study were performed to confirm the feasibility of integrating HA-SSC into PV modules. After our FE simulation results demonstrated the basic, complete feasibility using the electrospinning-based additive manufacturing (AM) method, we were able to identify additional polymers that we could fabricate in transparent systems, such as nylon [45] and poly(methyl methacrylate) (PMMA), or PMMA-based composites [46,47]. Furthermore, not only the interfacial adhesion with the front panel but also with the solar cell itself must be adequate [44]. The interfacial adhesion was not modeled in the current work and is the next step in the study. To maintain the 3D design throughout the lamination process, these novel materials may need further development [44,48,49]. All these may pave the way for future research of innovative polymeric composites/materials with 3D architecture to improve the use of lightweight PV modules and technologies.

#### **4. Conclusions and Future Perspectives**

In this FE simulation study, fiber-based 3D composites with helicoidally architecture were revealed to have outstanding impact energy/damage absorption and high-energy dissipation rate. Therefore, when applied to solar cells, they offer better shield against impact loads. Composite Plates C and B permit considerably higher fracture heights of 69 ± 2 and 82 ± 4, respectively, in comparison to 25 ± 5 for unprotected solar cells, and 50 ± 4 for EVA-protected solar cells, as observed during the ball-drop experiment. These helicoidally aligned synthetic composites (HA-SSCs) were fabricated using an electrospinning-based additive manufacturing (AM) technology that has only recently become possible at our group (Xtreme Materials Laboratory) in the past few years. During the ball-drop impact simulation using the Finite Element (FE) method, the HA-SSC composite materials (Composite Plates A, B, and C) showed consistent increase in damage absorption rate and stress level dissipation mechanism with reduced rotational angle of the HA-SSC materials. This is in excellent agreement with reports in the literature using isotropic materials (such as amorphous glass), although the agreement may be limited for anisotropic materials (such as monocrystalline silicon). The current FE simulation results add to the growing body of data that the innovative HA-SSC could be used for PV encapsulation to allow the design and fabrication of lightweight PV modules. The FE simulation results, as reported in the present manuscript, indicate impact protection increased with the reduction in the azimuthal angle to 45◦, after which the impact protection remained more or less constant. This is a unique insight which could have important implications for the silicon-based PV technology industry, as well as for other societally important applications of the HA-SSC materials, such as for lighter army combat vests and sports gear (helmets, etc.).

**Author Contributions:** Conceptualization, A.S.B. and R.S.; methodology, R.S., K.A. and A.B.; software, A.S.B. and F.E.G., validation, A.S.B., R.S. and K.A.; formal analysis, A.S.B.; investigation, R.S. and K.A.; resources, N.R.; data curation, R.S. and K.A.; writing—original draft preparation, R.S., R.F. and A.S.B.; writing—review and editing, A.S.B., R.F., F.E.G. and R.S.; visualization, R.S., K.A. and R.F.; supervision, A.S.B., N.R. and A.B.; project administration, N.R. and A.B.; funding acquisition, N.R. and A.B. All authors have read and agreed to the published version of the manuscript.

**Funding:** The authors would like to acknowledge the funding from the Ministry of Education (MOE) Academic Research Funds MOE2017-T2-2-175 titled "Materials with Tunable Impact Resistance via Integrated Additive Manufacturing as well as MOE2019-T2-1-197 titled "Monte Carlo Design and Optimization of Multicomponent Polymer Nano-composites". This work is supported by the La Trobe University Leadership RFA Grant, La Trobe University Start-up Grant and Collaboration and Research Engagement (CaRE) Grant offered by the School of Engineering and Mathematical Sciences (SEMS), La Trobe University. The authors also acknowledge the receipt of funding support from Temasek Labs@SUTD Singapore, through its SEED grant program for the project IGDSS1501011 and SMART (Singapore-MIT Alliance for Research and Technology) through its Ignition grant program for the project SMART ING-000067 ENG IGN. N.R. would like to acknowledge the funding from the Ministry of Education (MOE) Academic Research Fund MOE2019-T2-1-197 titled "Monte Carlo Design and Optimization of Multicomponent Polymer Nanocomposites" as well as support from EDB-IPP Surplus Funds Grant No. RGSUR08 for payment of article processing charges (APC).

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data presented in this study are available upon request from the corresponding author.

**Acknowledgments:** The authors gratefully acknowledge the critical support and infrastructure provided by the Singapore University of Technology and Design (SUTD). S. Anbazhagan, and A.S. Budiman also gratefully acknowledge the funding and support from the National Research Foundation (NRF)/Economic Development Board (EDB) of Singapore for the project under EIRP Grant "(NRF EWT-EIRP002-017)—Enabling Thin Silicon Technologies for Next Generation, Lower Cost Solar PV Systems". The authors gratefully acknowledge PT Impack Pratama Indonesia for providing materials, exploratory supports, and technical discussions related to possible implementations of the materials development for the applications in building-integrated PV (photovoltaics) technology/design. ASB gratefully acknowledges the support and infrastructure provided by the Oregon Institute of Technology (OIT), and especially the Oregon Renewable Energy Center (OREC) at OregonTech.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Article* **Clinical Application of 3D-Printed Patient-Specific Polycaprolactone/Beta Tricalcium Phosphate Scaffold for Complex Zygomatico-Maxillary Defects**

**Woo-Shik Jeong 1, Young-Chul Kim 1, Jae-Cheong Min 1, Ho-Jin Park 1, Eun-Ju Lee 2,3, Jin-Hyung Shim 2,3 and Jong-Woo Choi 1,\***

**Abstract:** (1) Background: In the present study, we evaluated the efficacy of a 3D-printed, patientspecific polycaprolactone/beta tricalcium phosphate (PCL/β-TCP) scaffold in the treatment of complex zygomatico-maxillary defects. (2) Methods: We evaluated eight patients who underwent immediate or delayed maxillary reconstruction with patient-specific PCL implants between December 2019 and June 2021. The efficacy of these techniques was assessed using the volume and density analysis of computed tomography data obtained before surgery and six months after surgery. (3) Results: Patients underwent maxillary reconstruction with the 3D-printed PCL/β-TCP scaffold based on various reconstructive techniques, including bone graft, fasciocutaneous free flaps, and fat graft. In the volume analysis, satisfactory volume conformity was achieved between the preoperative simulation and actual implant volume with a mean volume conformity of 79.71%, ranging from 70.89% to 86.31%. The ratio of de novo bone formation to total implant volume (bone volume fraction) was satisfactory with a mean bone fraction volume of 23.34%, ranging from 7.81% to 66.21%. Mean tissue density in the region of interest was 188.84 HU, ranging from 151.48 HU to 291.74 HU. (4) Conclusions: The combined use of the PCL/β-TCP scaffold with virtual surgical simulation and 3D printing techniques may replace traditional non-absorbable implants in the future owing to its accuracy and biocompatible properties.

**Keywords:** polycaprolactone; tricalcium phosphate; PCL/β-TCP; 3D printing; maxillary defect

## **1. Introduction**

The management of a maxillary defect is complicated when surgeons must replace the original 3D structure of the bone and carry out functional midfacial restoration in the periorbital and perioral region. Vascularized bone flaps have been the standard option in the field of mandibular reconstruction [1]. They provide a rigid and durable structure that allows adjuvant radiation treatment, a skin paddle for additional soft tissue defects, space for dental implant placement, and reasonable adaptation to remnant bony structures. However, no single flap can provide sufficient volume or support in larger or complex defects, especially when orbital adnexae and dental components are involved.

In complex maxillary treatments, alloplastic material has been combined with autogenous reconstruction. Titanium mesh has been widely applied because it is easy to use and biocompatible, allowing the ingrowth of connective tissue through the implant. Moreover, it can be molded into the complex maxillary structure [2]. However, it can lead to implant exposure or palpability due to the breakdown of the mucocutaneous lining. Deformative

**Citation:** Jeong, W.-S.; Kim, Y.-C.; Min, J.-C.; Park, H.-J.; Lee, E.; Shim, J.-H.; Choi, J.-W. Clinical Application of 3D-Printed Patient-Specific Polycaprolactone/Beta Tricalcium Phosphate Scaffold for Complex Zygomatico-Maxillary Defects. *Polymers* **2022**, *14*, 740. https:// doi.org/10.3390/polym14040740

Academic Editors: Hamid Reza Vanaei, Sofiane Khelladi and Abbas Tcharkhtchi

Received: 23 December 2021 Accepted: 9 February 2022 Published: 14 February 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

change can also occur during scar contracture and adjuvant radiation treatment [3,4]. To address these limitations, biodegradable or bioabsorbable materials have gained popularity; they are rigid and biocompatible, induce bone regeneration, and confer a lower chance of foreign body reaction [5–7].

Combined with computer-aided techniques, such as virtual surgical planning, various alloplastic materials have improved the accuracy of maxillofacial reconstruction [8–10]. 3D printing technology, combined with preoperative planning and modeling, enables more effective patient-specific treatment. In addition, biodegradable printing materials can now be used in a customized fashion to reconstruct complicated craniomaxillofacial defects with acceptable outcomes. Among these various biodegradable materials, PCL (polycaprolactone) has been used as guided bone regeneration (GBR) membrane owing to its favorable mechanical properties and biocompatibility with a slower degradation rate [11,12]. The beta-tricalcium phosphate (β-TCP), a bioceramic material, has been used in the field of bone tissue engineering owing to its chemical properties resembling bone minerals and excellent osteoconductivity [13–15]. The use of PCL blended with betatricalcium phosphate (β-TCP) was reported as a promising GBR membrane to promote new bone formation, with an initial stability comparable to cortical bone [11,16–18]. Traditionally, promising results were achieved in terms of osteogenic activity when a PCL scaffold was blended with 20% TCP [19–21]. To the best of our knowledge, there are few studies using a 3D-printed PCL/β-TCP scaffold in complex zygomatico-maxillary defects. The present work aims to evaluate the new bone formation and 3D conformity using a computed tomographic data and clinical outcomes in zygomatico-maxillary reconstruction with a 3D-printed PCL/β-TCP scaffold.

#### **2. Materials and Methods**

We evaluated a prospective series of eight patients with complex zygomatico-maxillary defects who underwent reconstruction with 3D-printed PCL implants between December 2019 and June 2021. The inclusion criteria were as follows: (1) unilateral zygomatic maxillary defect with or without orbital floor involvement, (2) maxillary defect resulting from cancer ablation, benign tumor resection, trauma, or degenerative change of the hemiface such as Parry–Romberg syndrome, (3) requirement of immediate or delayed reconstruction due to maxillary defect, and (4) follow-up period of at least six months. The exclusion criteria were as follows: (1) bilateral defect, (2) critical infectious disease or immune deficiency, (3) current or anticipated chemotherapy or immune suppression therapy, and (4) pregnancy or possibility of pregnancy.

Demographic information regarding sex, age, underlying disease, cause of defect, onset of reconstruction (immediate or delayed), type of maxillary defect, and postoperative complications were reviewed. The maxillary defects were categorized based on the amount of vertical and horizontal maxillary defect, as suggested by Brown et al. [22]. Surgical details regarding reconstructive options, incisional approach, application of bone forming material, implant fixation method, and revisional operation were described. Each patient underwent computed tomography (CT) scans with a slice thickness of 0.6 mm at three time periods, including before surgery and six months after.

This study was conducted according to the Declaration of Helsinki and approved by the independent Ethics Committee/Institutional Review Board of the Asan Medical Center (approval number: 2021-1292), with written informed consent obtained from all patients

#### *2.1. 3D Simulation and 3D Printing of Patient-Specific Implants*

The patient-specific implants were designed using 3D modeling software (Materialise Mimics; Materialise NV, Leuven, Belgium). The anticipated maxillary defect was marked on a stereolithography model of the skull, and the contralateral normal orbit was flipped to obtain the ideal normal contours of the defect. A patient-specific implant was designed over the region of interest, fabricated and then refined, with smoothing of the contour

(Figure 1). All processes were performed under close communication between modeling experts and plastic surgeons.

**Figure 1.** Design of patient-specific PCL/β-TCP scaffold in patient #1. The implant was threedimensionally designed using 3D modeling software based on the mirror imaging of a contralateral normal zygomatico-maxillary structure.

The PCL (Evonik Industries, Essen, Germany) and β-TCP (Foster corporation, Putnam, CT, USA) were mixed in a ratio of 8:2. After the PCL was melted by heating for 15 min at 110 ◦C, powdered β-TCP was added, which was then blended for 10 min. The PCL/β-TCP mixture was 3D printed using a multi-head deposition system using computer-aided manufacturing software. It had a rectangular pore architecture with a porosity of 50% and a pore size of 500 μm, as determined by 3D modeling software (3-Matic Research 9.0, Materialise, Leuven, Belgium). The scaffolds were freeze-dried at −85 ◦C for 24 h, then sterilized under a 450W UV lamp for 4 h. All manufacture process was managed by a facility with Good Manufacturing Practice certification (T&R Biofab Co. Ltd., Seoul, Korea). The image of the 3D printed scaffold is depicted in Figure 2.

**Figure 2.** Photograph of 3D printed PCL/β-TCP scaffold with line with 500 μm, and 50% of porosity.

#### *2.2. Surgical Procedure*

Patients presented with a wide range of maxillary bone and soft tissue defects of various etiologies. The bone defect area was exposed as much as possible so that the implant could be inserted. The bone defect was covered with the 3D-printed implant, with or without osteocutaneous free flaps. The PCL implant was embedded in a betadine solution for 10 min before insertion. If necessary, it was easily molded using a No. 15 blade or scissors, depending on the actual defect. After the PCL implant was inserted into the defect, it was fixed to the adjacent bony structure using mini-plates and 6–8 mm titanium screws. Additional free flaps were indicated if the alloplastic implant necessitated soft tissue envelop to cover the defect. Immediate adverse reactions related to the implant, such as allergic reactions, were checked during surgery.

#### *2.3. Volume and Density Analysis Based on CT Data*

A CT scan was performed before surgery and six months after to evaluate volumetric and density change. The DICOM data were translated into a stereolithography model in 3D modeling software (Mimics; Materialise Software Solutions, Leuven, Belgium) to simulate a postoperative image using a volume rendering technique. The region of interest was defined before surgery along the contour of the simulated implant object, as well as six months after surgery along the outer surface of the inserted implant. Two images were superimposed based on anatomical landmarks, including the anterior nasal spine, nasion, gonion, and menton. Overlapping between the simulated implant volume and postsurgical implant volume was calculated using the Boolean operation. The volume conformity was defined as the percentage of overlapping volume between the simulated and postsurgical images (Figure 3).

**Figure 3.** Volumetric analysis between preoperative planned model and actual surgical result. (**a**) Preoperatively designed STL model. (**b**) Actual surgical result volume-rendered as STL model. (**c**) Two images were superimposed based on anatomic landmarks. Overlapping between the simulated implant volume and postsurgical implant volume was calculated using the Boolean operation.

To identify de novo bone formation, the CT images were subjected to radiodensity analysis using a 3D modeling software (Mimics, Materialise Software Solutions, Leuven, Belgium); the radiodensity was measured in Hounsfield units (HU) in the region of interest. The applied threshold to measure the bone mineral density of newly regenerated bone was 200 HU. The bone volume fraction was defined as the volume ratio of de novo bone to the total implant within the region of interest (Figure 4). In addition, the mean tissue density of the region of interest was investigated at different time periods, including before surgery and six months after surgery.

**Figure 4.** The bone volume fraction was defined as a ratio of the de novo bone volume to the total implant volume within the region of interest. Red area notes the region where the tissue density was measured over 200HU, while the purple area denotes the PCL/β-TCP scaffold.

#### *2.4. Tensile Test of the Scaffold*

Tensile testing was performed using a single column universal testing machine (Instron, Norwood, MA, USA). The dimension of the scaffold sample was standardized to 10 × 40 × 1 (mm), and porosity was 50%. The number of the sample for the test was 7. The Young's modulus was calculated by the linear curve of the stress–stain curve.

#### **3. Results**

Eight patients were included in this study, presenting a wide range of maxillary defects of various etiologies. The causes of the defects were as follows: intraosseous hemangioma in two patients, immediate reconstruction following cancer ablation in three patients, and Romberg disease, traumatic facial deformity, and fibrous dysplasia in one patient each. Five of the eight patients underwent immediate reconstruction following tumor ablation, while three underwent delayed reconstruction. There was a case of wound dehiscence caused by partial flap necrosis, which required wound coverage by a local flap. Detailed information regarding demographics are depicted in Table 1.

**Table 1.** Demographics of patients.

N.A.: Not applicable.

Regarding surgical details, in four of the eight patients, the 3D-printed implant was inserted through a perioral and conjunctival incision. The other four patients underwent concurrent free flap or free bone grafts. In patients who had undergone cancer ablation, a head and neck surgeon used lateral rhinotomy and a Weber–Ferguson incision. A boneforming substance was used in three patients: a demineralized bone matrix (DMB) in two patients and a demineralized calcium phosphate bone substitute in one patient. Revisional operation was required in four patients who underwent a secondary fat graft and one patient who underwent local wound coverage to treat partial flap necrosis (Table 2).

**Table 2.** Surgical details.

HA-PLLA: Hydroxyapatite/poly-l-lactide; DBM: Demineralized bone matrix; ALT FF: Anterolateral thigh free flap; RFFF: Radial forearm free flap.

The result of the volume analysis was as follows. The mean preoperatively planned implant volume was 11.32 mm3, ranging from 2.16 mm3 to 30.37 mm3. The mean postoperatively actual implant volume was 10.21 mm3, ranging from 1.84 mm3 to 28.22 mm3. After the superimposition of two images, the mean volume conformity was 79.71%, ranging from 70.89% to 86.31%. Postoperatively, the de novo formation of bone was calculated and the mean was 2.15 mm3, ranging from 0.22 mm3 to 7.15.mm3. The bone volume fraction was obtained as the ratio of de novo bone volume and postoperative implant volume, with a mean of 23.34%, ranging from 7.81% to 66.21%. Mean tissue density in the region of interest was 188.84 HU, ranging from 151.48 HU to 291.74 HU (Table 3).

**Table 3.** Volume and density analysis.

HU: Hounsfield unit.

In the mechanical property test, the Young's modulus of the standardized scaffold with 50% porosity was 162.7 ± 12. 8 MPa (Table 4).

**Table 4.** Experimental result of mechanical property test.

#### *3.1. Case Presentation*

Representative cases with clinical pictures are described in this section.

#### 3.1.1. Case 1

Patient #1 was 21-year-old female who underwent delayed reconstruction 24 months after ablation of intraosseous hemangioma. The maxillary bone defect was exposed using the gingivobuccal and transconjunctival approaches. A 3D-printed PCL/β-TCP scaffold was fitted into the defect, and the patient required no further resection of the bony structures. The implant was fixed using a resorbable plate and screws made of HA-PLLA (hydroxyapatite/poly-l-lactide (Figures 5 and 6).

**Figure 5.** Clinical photographs in patient #1. Contour and symmetry of left cheek region was restored. (**a**,**b**) Pre- and postoperative 6-month frontal view photographs. (**c**,**d**) Pre- and postoperative 6-month basal view photographs.

**Figure 6.** 3D CT images in patient #1. Contour and symmetry of left zygomatico-maxillary region was restored with de novo bone formation. (**a**,**b**) Pre- and postoperative 6-month frontal view CT images. (**c**,**d**) Pre- and postoperative 6-month basal view CT images.

#### 3.1.2. Case 2

Patient #5 was 21-year-old male who underwent immediate reconstruction following the en bloc resection of maxillary fibrous dysplasia, defined as a type V defect. The patient underwent reconstruction with the 3D-printed PCL/β-TCP scaffold through a conventional gingivobuccal and transconjunctival incisions. The 3D-printed implant was fixated with wire steel. There was no complication in the long-term follow-up (Figures 7 and 8).

**Figure 7.** Clinical photographs in patient #5. Contour and symmetry of left cheek region was improved. (**a,b**) Pre- and postoperative 6-month frontal view photographs. (**c,d**) Pre- and postoperative 6-month basal view photographs.

**Figure 8.** 3D CT images in patient #5. Contour and symmetry of left zygomatico-maxillary region was improved. (**a,b**) Pre- and postoperative 6-month frontal view CT images. (**c,d**) Pre- and postoperative 6-month basal view CT images.

#### **4. Discussion**

Polycaprolactone (PCL) is one of the polymers prepared by ring opening polymerization of ε-caprolactone using a variety of catalysts. It safely degrades into carbon dioxide and water over 2–3 years and provides a suitable scaffold for guided bone regeneration [23,24]. The PCL/β-TCP scaffolds used in this study had a 3D shape, moderate rigidity, and relatively high elasticity and were manufactured with a patient-specific design. This property allows surgeons to manipulate and mold the implants using a blade or scissors. In our mechanical property test, Young's modulus of the scaffold with 50% porosity was 162.7 ± 12. 8 MPa, which is a similar level to that of the human mandibular trabecular bone (6.9 to 199.5 MPa) [25]. It was strong enough to maintain a three-dimensional shape when applied to clinical practice, and also had an adequate elasticity to be carved using tools available in the operating room. However, this might be insufficient to mimic the compressive strength and modulus of cortical bone itself [26,27]. Thus, the characteristics of PCL/β-TCP should be carefully considered depending on the amount of bony defect and surrounding soft tissue. The scaffold might be insufficient to be applied alone in the reconstruction of the whole zygomatico-maxillary complex. However, it was sufficient to bear the tension and compression during biomechanics of the upper jaw as when indicated as an onlay graft onto the bony surface or interpositional graft between the bony gaps. Overall, we did not find any bony instability or occlusal complication during the follow-up period. We suggested that the loading force should be distributed to the underlying bony strut through secure fixation with titanium screws and to overlap with the surrounding bony structure.

The PCL scaffold has been widely used in craniofacial reconstruction of various forms, including mesh, membrane, plate, and 3D implants [28–31]. Several authors have used PCL mesh in rhinoplasty to replace autogenous cartilage grafts [32]. They have reported that PCL mesh with a 3D structure was a safe and effective material and that it could maintain volume without any foreign body reaction [28]. However, unlike our study, PCL implants in the previous literature have only been applied to 2D reconstruction. Recently, Han et al. used 3D PCL implants in three cases of maxillary reconstruction following cancer ablation. All patients showed favorable outcomes. No signs of infection were observed in any of the three patients, and the existing native tissue was successfully fused with filling of the pores. So far, there has been few reports on the combined use of PCL and β-TCP as a 3D scaffold in clinical cases. We applied a patient-specific PCL/β-TCP scaffold to treat various maxillary defects with a range of etiologies, including facial asymmetry due to Romberg's disease and ablation of fibrous dysplasia and hemangioma. Notably, we performed a more structured analysis in our cases, measuring volume conformity and bone density.

Regarding the volume conformity, suboptimal results were obtained in two of the eight cases who underwent immediate reconstruction following maxillary sinus cancer ablation. Although we designed the implants with a 3D shape following the resection plan, the design did not always fit the actual resection margin. This resulted in less conformity between the preoperative simulated and postoperative actual implant volumes. However, experienced head and neck surgeons were fully capable of adjusting the shape of the implants because the material had elastic properties.

Meanwhile, our study reported a case of implant exposure in a patient who had undergone radiation treatment. We reasoned that the wound dehiscence had resulted from delayed wound healing in the irradiated field, especially in the naso-orbital region, rather than from the implant itself. It follows that the implant should be covered with a durable and thick flap, especially when patients have undergone previous radiation, and that meticulous debridement of remaining unhealthy tissue should be carried out to avoid wound complications.

Another complication of the biomaterial that should be considered is the possibility of an allergic reaction. Some rare complications have been reported with the use of biodegradable material due to the wide range of foreign body reactions [33–35]. Although there was no allergic reaction reported in our cases, the use of PCL might lead to serious foreign body reactions. Some researchers reported on long-term, late-onset inflammatory complications including granuloma formation, late allergic reaction and chronic inflammation after dermatologic application of PCL-based fillers [36,37]. This reaction seemed to result from an immune overreaction of the host tissue to the product, which is related with underlying inflammatory status of the patient. Thus, the safety of the PCL/β-TCP scaffold in our cases should be proven in the long-term study

In our previous research, we reported on the three-dimensional internal structure of a scaffold using 3D printing [26,38–40]. In the case of our 3D printed scaffold, it has an internal structure in which pores with a size of several hundred micrometers are completely interconnected by a layer-by-layer fabrication method. When implanted into the body, these perfectly connected pores are advantageous for the penetration of surrounding cells, and also help the engraftment of regenerated tissue inside the artificial scaffold as blood vessels are connected.

The effect of the material composition and porosity of a scaffold on its properties, including cell proliferation and differentiation, stiffness, and degradation, has been discussed in the literature [11,41–47]. The addition of β-TCP in PCL was shown to improve the scaffold's mechanical performance and increase osteogenic cell proliferation and differentiation [41,42]. By increasing the β-TCP concentration in the scaffolds, significantly higher mineralization was achieved compared to the pure PCL [48]. In addition, the bioceramic composition in the PCL scaffold was shown to increase water absorption and induce hydrophilic properties, which can be useful to prevent nutrient loss during bone regeneration [45]. Other considerations are the porosity, pore size, and permeability of the scaffold, which plays a significant role in biological delivery and tissue regeneration [11,46,47]. Larger pore size and porosity could be beneficial for bone tissue growth but may affect the compressive strength and modulus of the scaffold. Bruyas et al. found that both an increasing amount of β-TCP and decreasing porosity augmented the modulus of the 3D printed scaffolds, while decreasing the elasticity [43].

In our experience, when the amount ofβ-TCP in PCL is increased, viscosity also increases, and as PCL/β-TCP blend viscosity affects scaffold printing speed, 3D printer feed rate reduces, and the polymer is exposed to more thermal energy. When the weight proportion of β-TCP in PCL was more than 20% and the pore size was set to larger than 500 μm, we observed that the printing accuracy and mechanical strengths decreased. Thus, we used the PCL/β-TCP scaffold with a ratio of 80:20 and pore size of 500 μm to achieve balance between β-TCP content and printing rate.

The degradation profile of the scaffold is another factor that should be considered. The PCL has extremely slow progress of degradation, ranging from 2 to 4 years, while the TCP has an unpredictable biodegradation profile, ranging from 6 to 24 months [44,49,50]. In general, it was reported that the PCL/β-TCP composites had a faster degradation rate than that of pure PCL. Yeo et al. reported the PCL–20% TCP scaffold gradually degraded within 6 months, while maintaining its pore interconnectivity for newly mature bone to form [24]. Initial degradation of β-TCP can produce calcium ions and enhance mineralization, thereby promoting osteogenic differentiation of adipose-derived stem cells. Bruyas et al. found that higher ceramic content of over 40% TCP might lead to structural integrity of the scaffold due to the extremely high rate of degradation [43]. We agreed on their opinion in that such a manipulation of the ceramic ratio to create an ideal bioresorbable plate to match the natural healing course of bone formation. From CT findings obtained during the six-month follow-up of clinical cases, we judged that the 80:20 proportion of PCL:β-TCP and 500 μm pore size of the implant were adequate to enhance earlier bone growth and maintain durability. Other animal studies also corroborated this view, reporting neovascularization, sufficient soft tissue ingrowth, and the absence of extensive inflammation with this pore size and porosity [49].

We concluded that bone regeneration was confirmed based on CT scan results six months after surgery. In particular, it was based on the bone mineral density value from the CT image. We thought that the bone mineral density value reflected not only the purely regenerated bone but also the density of the implanted scaffold as well. However, due to the radiolucent characteristic of the biodegradable polymer, the contribution to the bone mineral density value is insignificant. Nevertheless, histological analysis from the biopsy tissue might be required for confirming the obvious bone regeneration, but it has limitation due to ethical issues. On the other hand, according to a previous study conducted by our research team, an obvious bone regeneration result was confirmed eight weeks after transplantation in an animal experiment using the same PCL/TCP scaffold applied in this study [51].

We used various materials, including a mixture of demineralized bone matrix and blood controlled thermal responsive polymer. Demineralized bone matrix has been widely used as a mixture material to enhance bone union and new bone formation [52]. Various artificial materials, including oxidized-irradiated alginate hydrogel and hydroxyapatite were combined with the 3D scaffold. Some authors have reported the combined use of bone morphogenic proteins (rhBMP-2) to treat mandibular defects [50]. However, we should be reluctant to apply this material in patients who have undergone cancer ablation as it is unclear whether rhBMP-2 promotes or inhibits tumor generation [53].

The present study had the following limitations: (1) As we assessed the density in a region of interest containing both the implant and new bone, we did not obtain the actual bone density, which might be lower than the normal bony structure outside of the implant; (2) Although a degradation period from 2 to 4 years for PCL and 6 to 24 months for TCP are known, the speed of degradation will vary depending on the transplant site due to characteristic of hydrolysis. Therefore, a long-term follow-up of more than 5 years is required for future studies; (3) The measured efficacy of PCL mesh in bone formation may have been confounded because we also applied osteoblastic agents. In the present study, we could not assess the efficacy of the combined mixture substances for bone formation, as we performed no comparative analysis. More structured investigation is necessary, with a prospective, comparative, controlled design.

#### **5. Conclusions**

The PCL/β-TCP scaffold can provide durable support and enhance bone formation in complex zygomatico-maxillary defects. The combined use of virtual surgical simulations, 3D printing techniques, and biodegradable implants may replace traditional non-absorbable implants because the method is more accurate and the materials more biocompatible.

**Author Contributions:** Conceptualization, J.-W.C.; methodology, Y.-C.K.; software, Y.-C.K.; validation, Y.-C.K.; formal analysis, Y.-C.K.; investigation, Y.-C.K., E.-J.L. and J.-H.S.; resources, J.-W.C.; data curation, Y.-C.K., J.-C.M. and H.-J.P.; writing—original draft preparation, W.-S.J.; writing—review and editing, Y.-C.K. and J.-H.S.; visualization, Y.-C.K.; supervision, J.-W.C.; project administration, J.-W.C. All authors have read and agreed to the published version of the manuscript.

**Funding:** This work was supported by the Demonstration Program of Industrial Technology for a 3D Printing Medical Device (P0008811) funded by the Korea Institute for Advancement of Technology (KIAT).

**Institutional Review Board Statement:** This study was conducted according to the Declaration of Helsinki and approved by the independent Ethics Committee/Institutional Review Board of the Asan Medical Center. (approval number: 2021-1292).

**Informed Consent Statement:** Written informed consent has been obtained from the patient(s) to publish this paper.

**Data Availability Statement:** Not applicable.

**Acknowledgments:** We specially thank to Eun-Ju Lee and Jin-Hyung Shim (Research Institute, T&R Biofab Co., Ltd., Seongnam-si, Korea, Department of mechanical engineering, Korea Polytechnic University, Sihueng-si, Korea) for adding the experimental evidence for the current study.

**Conflicts of Interest:** The funders had no role in the design of the study; in the collection, analyses, or interpretation of data; in the writing of the manuscript, or in the decision to publish the results.

#### **References**

## *Article* **Effect of Architected Structural Members on the Viscoelastic Response of 3D Printed Simple Cubic Lattice Structures**

**Ahmed Abusabir 1, Muhammad A. Khan 1,\*, Muhammad Asif 2 and Kamran A. Khan 3,4,\***

**Abstract:** Three-dimensional printed polymeric lattice structures have recently gained interests in several engineering applications owing to their excellent properties such as low-density, energy absorption, strength-to-weight ratio, and damping performance. Three-dimensional (3D) lattice structure properties are governed by the topology of the microstructure and the base material that can be tailored to meet the application requirement. In this study, the effect of architected structural member geometry and base material on the viscoelastic response of 3D printed lattice structure has been investigated. The simple cubic lattice structures based on plate-, truss-, and shell-type structural members were used to describe the topology of the cellular solid. The proposed lattice structures were fabricated with two materials, i.e., PLA and ABS using the material extrusion (MEX) process. The quasi-static compression response of lattice structures was investigated, and mechanical properties were obtained. Then, the creep, relaxation and cyclic viscoelastic response of the lattice structure were characterized. Both material and topologies were observed to affect the mechanical properties and time-dependent behavior of lattice structure. Plate-based lattices were found to possess highest stiffness, while the highest viscoelastic behavior belongs to shell-based lattices. Among the studied lattice structures, we found that the plate-lattice is the best candidate to use as a creep-resistant LS and shell-based lattice is ideal for damping applications under quasi-static loading conditions. The proposed analysis approach is a step forward toward understanding the viscoelastic tolerance design of lattice structures.

**Keywords:** 3D lattice structure; simple cubic lattice structures; plate-based lattice; shell-based lattice; truss-based lattice; ABS; PLA; 3D printing; FFF; viscoelastic behavior; relaxation; creep; cyclic loading

#### **1. Introduction**

A new generation of engineering materials, known as lattice structures (LSs), has recently found applications in biomedical [1], aerospace [2] and automotive [3]. Notable properties of LSs include their low density and high specific thermal, electrical and mechanical properties, energy absorption, and ability to reduce noise/vibration [4–6]. The overall response of LSs depend on the relative density, solid base material, and topology of the microstructure. For damping and energy absorption applications, a better understanding of the relationship between microstructure of the LS and their effective viscoelastic properties is required to obtain desired performance [7,8].

LSs consist of a solid skeleton and air pores. The architecture of microstructure influences their mechanical behaviors. Numerous architectures were proposed in the literature to describe the microstructure of LS. The architected LSs are classified into two categories: open-cell and closed-cell foams, with either a random or periodic arrangement [9]. Earlier

**Citation:** Abusabir, A.; Khan, M.A.; Asif, M.; Khan, K.A. Effect of Architected Structural Members on the Viscoelastic Response of 3D Printed Simple Cubic Lattice Structures. *Polymers* **2022**, *14*, 618. https://doi.org/10.3390/polym14030618

Academic Editors: Hamid Reza Vanaei, Sofiane Khelladi and Abbas Tcharkhtchi

Received: 30 November 2021 Accepted: 1 February 2022 Published: 5 February 2022

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**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

design of three dimensional (3D) networks of LSs are usually designed using discrete structural members such as struts or truss members. The microstructure, such as, rhombic dodecahedron [10], tetrakaidecahedron [11,12], cubic [6,13,14], Kelvin [15], Gibson-Ashby [16] and gyroids [17] have been studied. Analytical solutions for the effective response of these LSs were obtained through beam theory for elastic behavior [12,18–20] and viscoelastic behavior [21–23]. For more complicated architected LS, finite element homogenization method has been used to predict the elastic [24–29] and viscoelastic [22,30] responses.

Recently, three-dimensional network structures have been developed with interesting geometries derived from atomic crystal structures system [31]. The network of these lattice structures can be constructed with different structural members such as truss-, plate-, or shell-based (triply periodic minimal-surface (TPMS)) [32,33]. Out of these structural members, the plate-based lattice structures [34,35], offer superior stiffness which makes them excellent candidates for load-bearing applications. However, the shell based LSs such as TPMS demonstrated good energy absorption characteristics. Tancogne-Dejean et al. [36] showed that the specific energy absorption of plate-based LSs is around 45% greater than that of truss-based LSs. The elastic and viscoelastic properties of these lattice structures have been studied and investigated using the finite element method (FEM). Khan et al. used micromechanical homogenization approach to compute the apparent viscoelastic behavior such as creep, relaxation under quasi-static loading and dynamic behavior under cyclic excitation [37,38], and [39]. Previous studies highlighted the excellent viscoelastic response of the architected LSs [40]. Comprehensive studies have been conducted using theoretical and simulation approaches to investigate the properties of cellular solids; however, very limited experimental investigations have been undertaken to determine the viscoelastic response of polymeric LSs [15,40,41]. Moreover, the effect of architected structural member and base material on the viscoelastic response of 3D printed lattice structure has not been investigated

The revolution and growth in additive manufacturing have allowed the fabrication of complex and precise geometries of LSs. Additive manufacturing (AM) offers high flexibility of design and rapid prototyping. In the recent review article, it has been discussed that AM can reduce the production cost of complex components and can be implemented not only for prototyping but also production using different approaches in design [41]. Additive Manufacturing technology has enabled the porosity and architecture of cellular solids to be controlled; therefore, the density and mechanical properties can be tailored [42] for several applications [43]. Additive manufacturing includes several processes; however, the 3D printing technology using material extrusion (MEX) process [44] has been widely used to fabricate complex geometries such as cellular solids. Moreover, the base materials have significant influenced on the design of LSs. The LSs should be able to contribute to the functional purpose of structure with excellent damping performance, strengthto-weight ratio, and others. Thermoplastic polymers have been widely utilized in the fabrication of cellular solids due to their adaptability for 3D printing and their unique properties. The most utilized polymers are acrylonitrile butadiene styrene (ABS), and polylactic acid (PLA) [7,8]. The comparison of the flexural properties of ABS, PLA and a PLA–wood composite manufactured through MEX process has been presented [45]. Several authors have extensively studied the manufacturing of PLA using MEX process such an in-process monitoring of temperature evolution, multiscale damage and fatigue modeling of PLA [46–48]. The influence of process parameters has also been investigated on the mechanical properties [49], impact resistance properties [50] and interlayer adhesion on the tensile strength of 3D printed PLA [51].

In this study, the effect of architected structural member geometry and base material on the viscoelastic response of 3D printed lattice structure has been experimentally investigated. The LSs possessing simple cubic symmetry based on plate-, truss-, and shelltype structural members were considered to describe the microstructure of the LSs. The proposed LSs were fabricated with two materials, i.e., PLA and ABS using the material extrusion (MEX) process. The quasi-static compression response of lattice structures was

investigated, and mechanical properties were obtained. Then, the creep, relaxation and cyclic viscoelastic response of the lattice structure were characterized and some interesting conclusions were presented.

#### **2. Methodology**

#### *2.1. Design of Lattice Architecture and Manufacturing*

In this study, the three lattice microstructures of simple cubic family were considered. The three designs are named as simple cubic truss-based lattice (SCTL), simple cubic platebased lattice (SCPL), and simple cubic shell-based lattice (SCSL). The SCTL, SCPL, and SCSL unit cells consist of struts, plate and shell, respectively. The arrangement of these structural members yield simple cubic LSs. Solidworks software was used to model the considered designs. The 3D designs were made with overall dimensions of 25 × 25 × 25 mm3. The investigation was conducted using two polymeric materials: Polylactic acid (PLA), and acrylonitrile butadiene styrene (ABS). Raw materials of ABS and PLA were procured in the form of filament with 1.75 mm thickness. The specifications of the utilized materials are shown below in Table 1 as provided by the manufacturing company.

Additive manufacturing based on material extrusion (MEX) process, was adopted to fabricate all specimens. In this study, we employed the Raised3D Pro2 printer, which is equipped with a 0.4 nozzle. Several attempts were made to attain the best designs in terms of lightweight, manufacturability, and flexibility. The printing parameters that were given using software Idea Maker are shown below in Table 2.

For all specimens, the faces of the infill were perpendicular to the direction of the build (out-of-plane). All samples were printed with a raft platform to ensure the flatness of the base and stability throughout the printing process. Concerning solid infill density, all candidates were designed with 27% solid infill density. Table 4 shows the unit cell CAD design, the LS with array of 5 × 5 × 5 unit cells, the design and printing parameter, and the fabricated LSs made of PLA and ABS. Throughout this study, the investigated samples will be referenced by the assigned ID codes shown in Table 3.

**Table 3.** ID codes of the 3D printed specimens.

**Table 4.** Details of considered designs.

#### *2.2. Experiments*

Four experiments were performed to understand the mechanical properties and timedependent behavior of the 3D printed polymeric LSs, i.e., quasi-static compression test, stress relaxation test, creep test, and compressive cyclic loading test, as shown in Figure 1. The experiments were conducted using an Instron universal testing machine with 5KN and 30KN load cells. The crosshead speed was 2.5 mm/min in all tests, chosen based on ASTM D1621-16 [52]. A pre-load was applied to guarantee a full initial contact between plates and specimen; all tests were conducted at room temperature. The experimental setup is shown in Figure 1. Pre-experimenting, the relative density of considered specimens were measured using a weight scale and Equation (1)

$$
\overline{\rho} = \frac{\rho\_c}{\rho} \tag{1}
$$

where *ρ*: relative density, *ρ*c: density of cellular solid, *ρ*: density of solid material.

**Figure 1.** Experiment set-up and loading program.

#### 2.2.1. Quasi-Static Compression Test

First, the quasi-static compression test was performed until fracture. The quasi-static compression test was performed according to ASTM D1621-16 "Standard Test Method for Compressive Properties of Rigid Cellular Plastics". The specimens were placed between the compression plates ensuring that the specimen centerline was aligned with the load cell centerline. Pre-loading was applied to ensure the stability of the samples and full initial contact between plates and specimens. The LSs were compressed at a constant crosshead speed of 2.5 mm/min and the effective stress–strain behavior was recorded. Many interesting characteristics of LS such as elastic modulus (*E*) and specific stiffness (*C*) were calculated using Equations (2) and (3).

$$E = \frac{\sigma}{\varepsilon} \tag{2}$$

$$\mathbf{C} = \frac{E}{\overline{\rho}}\tag{3}$$

#### 2.2.2. Stress Relaxation Test

A stress relaxation test is necessary to understand the viscoelasticity behavior (timedependent response), in which the specimen is compressed and held at a certain displacement; accordingly, the stress relaxation response is recorded as a function of time. The relaxation response can be measured by calculation stress-relaxation modulus using Equation (4).

$$E\_{\rm sr} = \frac{\sigma\_l}{\varepsilon\_0} \tag{4}$$

The samples made from the different materials were compressed to the same strain level called effective strain. The effective strain should be on or below the yield point, which was determined using the data obtained from the quasi-static compression test. It was considered to be a value below the least yield limit among the three samples. Table 5 shows the effective strain levels used during stress relaxation test. The stress relaxation test were performed according to ASTM E328 − 21: Standard Test Methods for Stress Relaxation for Materials and Structures [53]. In this study, the displacement was applied on the specimen at the strain rate of 2.5 mm/min until reaching the desired displacement. The position (displacement) was held constant for 30 min and the stress relaxation response was recorded as a function of time.

**Table 5.** Parameters of stress relaxation test.

#### 2.2.3. Creep Test

Viscoelastic behavior can also be measured by creep testing, in which constant stress is applied for a period of time and changes in strain are observed as a function of time. The viscoelastic behavior can be measured by finding creep compliance (J) using Equation (5). The creep test was performed according to ASTM D2990 − 17: Standard Test Methods for Tensile, Compressive, and Flexural Creep and Creep-Rupture of Plastics [54]. Table 6 shows the forces levels used during creep test. Here, the sample was compressed with a strain rate of 2.5 mm/min to the predetermined load limit and held constant for 30 min. While constant stress was applied, the strain will continue to increase with time and therefore recorded.

$$J\_t = \frac{\varepsilon\_t}{\sigma\_0} \tag{5}$$

**Table 6.** Parameters of creep test.

2.2.4. Compressive Cyclic Loading Test

The viscoelastic phenomenon and energy dissipation behavior of cellular materials can be observed by loading and unloading a specimen at a constant strain rate. The compressive cyclic loading test involves an appropriate repeating pattern of loading-unloading. The test may be conducted with a peak strain-controlled, or peak stress-controlled technique. In this study, the experiments were carried out with a peak stress-controlled method and the specimens were compressed with a strain rate of 2.5 mm/min to the predetermined load limit. The testing parameters are illustrated in Table 7. In total, three loading-unloading cycles were applied, and the load-displacement hysteresis loop were recorded. OriginLab software was used to calculate the area under the hysteresis curve, which represents the amount of energy absorption.

#### **3. Results and Discussion**

In this section, the data obtained from the experiments described above will be shown, analyzed, and discussed. A weight scale was used to measure the weight of the 3D printed specimens, then the relative density was calculated using Equation 1 as shown in Table 8.

The measured values show that all ABS samples having almost the same weight with a variation of ±0.06 (1.5%), similarly shown in all PLA specimens with a variation of ±0.16 (3%). The equality in weights verifies that the initial designs have the same solids infill density and the excellent accuracy of the manufacturing process. Several factors may have contributed to the slight variations, such as the uncertainty of the scaling device, the surrounding conditions in the lab, or minor uncertainties in the design or fabrication process.

#### *3.1. Quasi-Static Compression Test*

Figure 2 shows the compression stress–strain curves for the investigated LSs. The stress–strain curve provides the mechanical behavior of LSs and could help to find the Young's modulus and yield strength. The main purpose of this test is to obtain the linear stress–strain limit so that the effective load, and strain levels can be identified for creep, stress relaxation, and cyclic loading-unloading tests. It can be observed that the overall compressive behavior of LSs depends mainly on its microstructural design and relative density, and the mechanical properties of the base material. Generally, the higher the density, the higher the collapse stress. As defined early, PLA has a higher density than ABS, 1.25 and 1.03 g/cm3, respectively. Therefore, the fracture stress of the PLA samples is higher than that of the ABS specimens, as illustrated in Figure 2. With regards to the effect of the architected structural member geometry, it is evident that plate-based lattices are stiffer than others, followed by truss-based lattices then shell-based lattices made of the same material and relative density.

**Figure 2.** Compression stress–strain curves for the investigated samples.

The Young's modulus values were determined through the tangent value of the initial slope of the stress–strain curves, by using Equation 2 and the values Young's Modulus are shown in Table 9. The plate-based lattice in both materials has the highest Young's modulus values, and the least value of Young's modulus belongs to the shell-based lattice. Moreover, another interesting property that can be obtained from the stress–strain curve is the specific stiffness, whereby the stiffness-to-density ratio can be measured using Equation 3; specific stiffness values are shown in Table 9.

Another important point to be noticed that the yield limit was not clearly defined as the LSs demonstrated nonlinear stress–strain response. The method of offset point was used to compute the yield point, that indicates the limit of elastic behavior and the beginning of plastic deformation. Table 9 shows the yield stress for the considered samples, which was important to be identified for subsequent experiments.

**Table 9.** Obtained properties from quasi-static compression test.

We investigated the architected structural member geometry on the deformation mechanism. All the three structures were deformed under uni-axial compression and representative pictures were taken during the tests at different strain levels as presented in Figure 3. Noticeably, there is no physical failure in the identified yield point as shown in the 1st row in Figure 3. Moreover, it was observed that buckling occurred when compressive strain reached to some critical value and consequently led to rapid and dramatic changes of the material microstructure, as illustrated in the 2nd row in Figure 3 (in which all three structures demonstrated clear buckling). Subsequently, the middle region of structural members reached to a completely collapsed and then the deformation progressed to the neighboring cells. The plate-based lattice deformation occurred by compressing layers over each other, while truss-based lattice deformed due to buckling of its struts, whereas shell-based lattice deformed by squeezing its unit cells.

Generally, it was observed that all samples have deformed in a stretching-dominated manner; however, each specimen has its characteristics. For examples, the high stiffness in plate-based lattice is due to its plates involvement to carry load capacity and the integration or configuration of the plate-based structure. On the other hand, when a truss-based lattice experiences a compression load, and most of the load is carried by struts located in the longitudinal direction of the force, which means more stress concentration in thin struts. Therefore, vertical struts are the first to fail via buckling. Moreover, shell-based lattice has a novel geometry that doesn't contain struts or walls, the advantages of its architecture were observed during the experiment, whereby it exhibited great extension, resulting from the uniform distribution of the stresses.

#### *3.2. Stress Relaxation Test*

The stress relaxation experiment was undertaken according to the procedure explained in above methodology section. Equal effective strain was applied in each sample made of the same material, based on the outcomes of quasi-static compression test, the elastic limit of PLA samples is higher than ABS samples. Therefore, PLA samples experienced higher initial stress than ABS, as shown in Figure 4.

As can be seen in Figure 4, the stress relaxation curves can be divided into three stages. The first stage is the effective elastic stage, in which the specimens were compressed to the predetermined displacement and then held for 30 min. This initial displacement determined the starting point of stress relaxation. Then, the stress relaxation started after the first stage and can also be divided into two stages: transient stage and stable stage, representing the regions of decreasing stress relaxation rate and near-constant stress relaxation rate, respectively.

**Figure 3.** Deformation mechanism of the samples under investigation.

Figure 4 shows the plate-based lattice experienced the greatest stress to deform to the predetermined strain level, followed by the truss-based lattice. In contrast, the shell-based lattice demonstrated the least load bearing capacity. These results are due to the stiff plate-based structure, which is aligned with the conclusions drawn from the quasi-static test. As shown in Figure 4, all considered samples exhibited different stress relaxation behavior over time, which demonstrates that different viscoelastic mechanism exists in each specimen.

**Figure 4.** Stress relaxation response of all considered specimens.

For further analysis, the percentage of the normalized stress was calculated and shown in Table 10. It was found that the shell-based lattice outperformed the truss-based lattice and the plate-based lattice in terms of normalized stress over time. In addition, to determine the viscoelastic response from the stress relaxation test, the stress relaxation moduli were calculated using Equation 4. Then, the stress-relaxation moduli were converted to the relative moduli to compare based on the two considered materials as listed in Table 10. From the calculated values, it can be seen that the shell-based lattices have the greatest viscoelastic behavior, followed by the truss-based lattice, then the plate-based lattice. The outperformance of the shell-based lattice is due to its smooth geometry and curvature interconnections, by which the stress concentration is reduced, and the applied stress distributed uniformly. However, the stiffness of plate-based lattice has an adverse effect on the viscoelastic response. From the relative modulus values, it can be concluded that ABS samples have better viscoelasticity than that of PLA, resulting from the less stiffness and better elongation of ABS.

#### *3.3. Creep Test*