Read the paper and make detailed slides discussing the conclusion as well as the additional areas of research covering the rubric bellow. I will also need speaker notes in the bottom of each slide describing the slide further (Things I would be saying when I present the slide).Conclusions of Paper:What did the authors conclude? Do you agree with their conclusion? Did they provide sufficient evidence to prove their hypothesis? Additional areas of research related to this topic:What are some other areas of research/research papers that you discovered that are helpful to understand this research (this may also be covered in the background section)What other papers/additional research is being done on this topic?
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received: 19 October 2015
accepted: 26 January 2016
Published: 15 February 2016
3D Printing Surgical Implants at
the clinic: A Experimental Study
on Anterior Cruciate Ligament
An Liu1,3,*, Guang-huai Xue2,3,*, Miao Sun4,*, Hui-feng Shao2,3, Chi-yuan Ma1, Qing Gao2,3,
Zhong-ru Gou5, Shi-gui Yan1, Yan-ming Liu4 & Yong He2,3
Desktop three-dimensional (3D) printers (D3DPs) have become a popular tool for fabricating
personalized consumer products, favored for low cost, easy operation, and other advantageous
qualities. This study focused on the potential for using D3DPs to successfully, rapidly, and economically
print customized implants at medical clinics. An experiment was conducted on a D3DP-printed anterior
cruciate ligament surgical implant using a rabbit model. A well-defined, orthogonal, porous PLA screwlike scaffold was printed, then coated with hydroxyapatite (HA) to improve its osteoconductivity.
As an internal fixation as well as an ideal cell delivery system, the osteogenic scaffold loaded with
mesenchymal stem cells (MSCs) were evaluated through both in vitro and in vivo tests to observe boneligament healing via cell therapy. The MSCs suspended in Pluronic F-127 hydrogel on PLA/HA screw-like
scaffold showed the highest cell proliferation and osteogenesis in vitro. In vivo assessment of rabbit
anterior cruciate ligament models for 4 and 12 weeks showed that the PLA/HA screw-like scaffold
loaded with MSCs suspended in Pluronic F-127 hydrogel exhibited significant bone ingrowth and
bone-graft interface formation within the bone tunnel. Overall, the results of this study demonstrate
that fabricating surgical implants at the clinic (fab@clinic) with D3DPs can be feasible, effective, and
Ever since Charles Hull first proposed the three-dimensional (3D) printing process in 1986, the technology has
developed rapidly and well beyond what originally seemed possible1. Nowadays, 3D printing has been utilized
successfully in mechanical manufacturing and many areas of scientific research2. Many potential uses for 3D
printing have emerged within the medical field, not only as far as tissue and organ regeneration research3 (blood
vessels4, ears5, bones6), but also for customized medical devices such as splints and stents that can be printed in
small clinics7. There are several factors that limit the use of 3D printers in practice, however; 3D printers necessary for medical applications are specialized or industrial equipment that require unique materials, for example,
which drives up production costs and creates a high-level technical demand for skilled operators and specific
operational conditions, and the inconvenience of communicating at length between hospitals and factories during the production process delays the length of time between fabrication and application. It was reported that only
$11 million was invested in medical applications among the entire 3D printing industry which is worth around
$700 million in total8. To allow medical professionals and their patients to benefit from 3D printing technologies,
and to increase the market share value of 3D medical printing, it is crucial to develop methods that reduce production costs and increase the flexibility, maneuverability, and practicability of the process.
Department of Orthopaedic Surgery, Second Affiliated Hospital, School of Medicine, Zhejiang University, Hangzhou
310009, China. 2State Key Laboratory of Fluid Power and Mechatronic Systems, College of Mechanical Engineering,
Zhejiang University, Hangzhou 310027, China. 3Key Laboratory of 3D Printing Process and Equipment of Zhejiang
Province, College of Mechanical Engineering, Zhejiang University, Hangzhou 310027, China. 4Department of Oral and
Maxillofacial Surgery, Second Affiliated Hospital, School of Medicine, Zhejiang University, Hangzhou 310009, China.
Zhejiang-California International Nanosystem Institute, Zhejiang University, Hangzhou 310058, China. *These
authors contributed equally to this work. Correspondence and requests for materials should be addressed to S.-g.Y.
(email: email@example.com) or Y.-m.L. (email: firstname.lastname@example.org) or Y.H. (email: email@example.com)
Scientific Reports | 6:21704 | DOI: 10.1038/srep21704
Fused deposition modeling (FDM)9, when applied to the 3D printer, creates a desktop 3D printer (D3DP) that
can be used at home, in schools, and by small businesses to fabricate customized products cost-effectively. D3DPs
cost as little as $500, as opposed to the $15,00030,000 price range for 3D printers used in academic institutions.
If the D3DP can be successfully applied in the medical field, the possibility for cost-effective, personalized devices
such as implants or grafts to be fabricated in-clinic is momentous. Doctors and specialists who employ such technology would represent the pioneering edge of the medical field.
In a previous study conducted in our laboratory10, we were able to fabricate soft tissue prostheses using a
D3DP; the prostheses, which showed smooth surfaces and intricate structures, cost only about $30. The results
of this study have considerable implications as far as the future of maxillofacial repair technology. In the present
study, we focused on fabricating surgical implants and applying them in operations to demonstrate that a surgeon
can indeed customize and fabricate surgical implants his or herself using a D3DP.
Our target operation was an anterior cruciate ligament (ACL) reconstruction using a hamstring tendon graft.
This operation requires that the tendon graft within the bone tunnel heal appropriately. Tendon-to-bone tunnel
healing occurs through new bone ingrowth that initially forms between the tendon and the bone. With the help
of new bone mineralization and maturation, the grafts biomechanical properties progressively increase tendon
graft healing within the bone tunnel thus mainly depends on the osteointegration of the tendon graft within the
Bioabsorbable interference screws, made with polymers such as polylactic acid (PLA) and polyglycolic acid
(PGA), are commonly used to provide a press fit between bone, graft, and screws initially, which then degrade
mainly by hydrolysis as bone union gradually progresses12,13. According to clinical trials, PLA and PGA screws
have been shown to persist in vivo for up to 5 years and result in complete resorption at 7 to 10 years14,15. The
relatively slow degradation rate of bioabsorbable screws does not suit the speed of new bone formation, which
leads to malformation of new bone around the tendon graft, where only calcified fibrous or fatty tissue replaces
the screw in the bone tunnel15,16.
It has been reported that 3D porous structure is a key point to promote bone ingrowth by providing sufficient
growth space. Macropores (200400 µ m) enhance the migration of osteoblasts and osteoprogenitors into the scaffold and facilitate osteoid formation and mineralization17. Additionally, interconnected micropores (50100 µ m)
can increase vascularization and nutrient diffusion during bone reconstruction18. These structures cannot be
well-controlled through conventional methods19,20, but surgeons and specialists can easily and precisely manipulate them using a D3DP.
In this study, common PLA filament, the same as that used for bioabsorbable screws, was applied to D3DP
manufacturing of a 3D, porous, screw-like scaffold in-clinic. The scaffold not only could fix the tendon graft, but
also could provide adequate space for bone ingrowth around the graft. A simple surface modification was made
using hydroxyapatite (HA) on the scaffold in order to enhance osteoconductivity and cellular adhesion21, and
mesenchymal stem cells (MSCs), known as one of the most optimal cell sources for ACL regeneration due to their
high potential for proliferation and collagen production22,23, were seeded onto the scaffold as cell therapy.
We hypothesize that the 3D-printed, bioabsorbable screw-like scaffold loaded with MSCs can promote tendon graft healing within the bone tunnel by increasing bone ingrowth. We hope that the results of this study will
increase the popularity of 3D-printed surgical devices by proving that they can be customized and fabricated
feasibly, economically, and successfully in the clinic.
Fabrication and Characterization of PLA Screw-like Scaffold.
The PLA screw-like scaffold was
designed using Rhinoceros software (ver. 4.0, USA) according to a schematic, actual-size diagram of the implant
and tendon graft based on a rabbit ACL reconstruction model (Fig. 1). Its digital dataset was saved as a stereolithography (STL) file. Slice software Slic3r24 was used to generate G code for the D3DP (Dot Go 3D Technology
Corporation, Xiangtan, China) from the STL file. Melt PLA filament (Shenzhen Esond Technology Co., Ltd)
was extruded through a heated metal nozzle (0.4 mm in diameter, moving horizontally and vertically) at 205 °C
and deposited onto a receiving station to form the desired scaffolds. The scaffolds were then observed under a
scanning electron microscope (SEM) (S-4800, Hitachi, Japan) to measure macropore sizes. The porosity of the
scaffolds was determined using the Archimedes method, and the PLA scaffolds were weighed as dry weight (W1).
The scaffolds were then immersed in a beaker of water and held under vacuum to force the liquid into the pores
until no bubbles emerged, then re-weighed under water to determine the suspension weight (W2). The scaffolds
were then carefully taken out of the beaker and any water on the surface was removed, then they were quickly
re-weighed in air to determine the saturated wet weight (W3). The final porosity of the scaffolds was calculated
via the following equation: porosity (%) =??(W3 – W1)/(W3 – W2) ×??100%. Six specimens were measured in total.
HA Synthesis and Characterization.
HA powders were synthesized by chemical precipitation using
Ca(NO3)2·4 H2O and (NH4)2HPO4 as P and Ca precursors, respectively. Ca(NO3)2·4 H2O (Sigma-Aldrich,
Australia) was dissolved in distilled water (0.5 mol/L) and adjusted to pH 10.5 with NH3·H2O. (NH4)2HPO4
(Sigma-Aldrich, USA) was dissolved in distilled water at density of 0.3 mol/L and pH 10.5, then the Ca(NO3)2
solution was added to the (NH4)2HPO4 solution dropwise. After stirring for 12 h, the precipitate was filtered and
subsequently washed three times with distilled water followed by three washing steps with ethanol. The remaining
liquid was removed by vacuum filtration, and the precipitate was dried at 80 °C overnight. The resultant powders
were calcined at 850 °C for 3 h to obtain HA powders. The calcined HA powders were then ground and sieved
through 250 mesh sieves. The crystal morphology of the synthesized HA powder was observed using SEM, and
the phase composition of HA was characterized by X-ray diffraction (XRD, Rigaku Co., Japan).
Scientific Reports | 6:21704 | DOI: 10.1038/srep21704
Figure 1. Schematic diagrams of the implant and tendon graft within the bone tunnel in ACL
reconstruction. (A) The 3D perspective of the bone tunnel in ACL reconstruction. (B) The transverse section
view of the bone tunnel. (G: Graft; S: Screw-like scaffold; BT: Bone tunnel; M: Macropore).
Surface Modification for PLA/HA Scaffold.
Chitosan (CHI) was dissolved in 2% (v/v) acetic acid to
obtain CHI solution (1% (w/v)). Sodium alginate (SA) solution (1% (w/v)) was prepared with distilled water. HA
powders were added into the CHI and SA solutions, respectively, on a magnetic stirrer plate for 30 min to obtain
4% (w/w) HA/CHI solution and 4% (w/w) HA/SA solution. Sodium hydroxide (NaOH) solution (0.2% (w/w))
was mixed with equal volume of ethanol to prepare NaOH/ethanol solution. The PLA scaffolds were first dipped
in the NaOH/ethanol solution under vacuum for 10 min to modify the scaffolds with stable negative charge,
then washed twice with distilled water under vacuum, then freeze-dried for 30 min. Next, the scaffolds were
immersed in 4% (w/w) HA/CHI solution to force solution into the pores until no bubbles emerged from the scaffolds (10 min) followed by centrifugation (1000 r/min, 5 min). The scaffolds were dried at room temperature for
20 min, then immersed in 4% HA/SA solution under vacuum. The same procedures were repeated for all samples.
The PLA/HA scaffolds were then observed with SEM.
Cell Culture In Vitro. MSCs were obtained from bone marrow aspirates of New Zealand Rabbits25. Cells
of third passage were cultured in Dulbeccos Modified Eagle Medium (DMEM) supplemented with 10% fetal
bovine serum (FBS) (Gibco, USA) in an incubator at 37 °C with 5% CO2. Pluronic F-127 was added into complete
DMEM to prepare a 30% (w/v) solution at 4 °C. The solution was placed on a magnetic stirrer plate for 24 h to
allow complete dissolution, then the solution was filter-sterilized through a 0.22 µ m pore size bottle-top filter and
stored at 4 °C until use.
After being sterilized with ethylene oxide, the PLA scaffolds and PLA/HA scaffolds were placed into 24-well
tissue culture plates (TCPs) and immersed in DMEM with 10% FBS for 2 h, then each was seeded with 1 ×??105
MSCs. An equal number of 1 ×??105 MSCs suspended in Pluronic F-127 solution were seeded on the PLA/HA scaffolds at 4 °C to ensure the hydrogel penetrated the scaffold, then they were moved to the incubator for gelation.
Cell Morphology. After 48-hour incubation, samples were washed with phosphate buffer solution (PBS)
twice and fixed with 2.5% glutaraldehyde solution for 2 h. The fixed cells were washed with PBS three times and
treated with 1% osmium tetroxide for 2 h, then dehydrated in ascending concentrations of ethanol (30, 50, 70,
80, 90, 95, 100 (v/v)) for 5 min, respectively. The samples were then immersed in isoamyl acetate for 20 min, then
vacuum-dried at 40 °C for 4 h. The specimens were coated with gold-palladium and dried, then the MSC morphology of each was observed using SEM.
Cell Viability. MSC viabilities were analyzed with Cell Counting Kit-8 (CCK-8, Dojindo, Japan) assays at 1, 4,
and 7 days. DMEM (0.5 mL) containing 10% CCK-8 was added into each well. After 120 min, 100 µ L of the abovementioned solution was transferred to a 96-well plate. A microplate reader (Infinite F50, TECAN, Switzerland)
was used to measure solution absorbance at 450 nm, and absorbance values were corrected by subtracting the
signal of a mixture of 90 µ L DMEM and 10 µ L CCK-8. Five specimens were prepared for each sample.
Real-time Polymerase Chain Reaction (PCR) Analysis. Real-time PCR was used to detect the expression of several osteogenetic, differentiation-related marker genes (Col I, OCN, Sp7, and Runx2) at Day 7. Total
RNA was extracted using Trizol reagent (Invitrogen) according to the manufacturers instructions. NanoDrop
2000c (Thermo Fisher Scientific Inc., USA) was used to determine the total RNA concentration. First-stranded
complementary DNAs (cDNAs) were synthesized from 0.5 µ g of the isolated RNA by oligo(deoxythymidine)
(oligo (dT)) using the DyNamoTM cDNA Synthesis Kit (Fermentas) and used as templates for real-time
Scientific Reports | 6:21704 | DOI: 10.1038/srep21704
Primer sequence (5’3′)
GCG GTG GTT ACG ACT TTG GTT
AGT GAG GAG GGT CTC AAT CTG
GGC TCA GCC TTC GTG TCC AA
CCC TGC CCG TCG ATC AGT T
GGC ACG AAG AAG CCA TAC TCT GT
GGG AAA AGG CCG GGT AGT CAT
CCC AAG CAT TTC ATC CCT CAC T
CAT ACC GAG GGA CAT GCC TGA
TCA CCA TCT TCC AGG AGC GA
CAC AAT GCC GAA GTG GTC GT
Table 1. The parameters of primers utilized for detecting osteogenetic gene expression.
PCR. The PCR was performed on a final volume of 25 µ L containing 1 µ L cDNA, 0.5 µ L of each primer (forward and reverse), 12.5 µ L Power SYBR Master Mix (2× ) (Applied Biosystems, Foster City, CA, USA), and
10.5 µ L dd H2O with the Bio-Rad Real-time PCR System (Bio-Rad, Hercules, CA, USA), using glyceraldehydes
-3-phosphatedehydrogenase (GADPH) as the house-keeping gene for normalization. The forward and reverse
primer sequences utilized are listed in Table 1. The conditions of real-time PCR were 95 °C for 1 min, followed by
40 cycles at 95 °C for 10 s and 64 °C for 25 s.
ACL Reconstruction. A total of 36 New Zealand male rabbits weighing 2.53.0 kg were utilized in this study
according to standard guidelines approved by the Zhejiang University Ethics Committee (ZJU2014-1-05-093).
All rabbits were randomly divided into the PLA group (PLA scaffold implantation, n =??12), PLA/HA group (PLA/
HA scaffold implantation, n =??12), or MSCs group (PLA/ HA scaffold loaded MSCs, n =??12). Next, 2 ×??105 of
MSCs suspended in Pluronic F-127 solution were loaded on the PLA/HA screw-like scaffolds at 4 °C and cultured in vitro at 37 °C with 5% CO2 over 8 h for gelation and cell adhesion before implantation. The animals were
subjected to general anesthesia with phenobarbital (30 mg/kg), followed by bilateral ACL reconstruction. The
knee joint was accessed via a medial parapatellar approach through a midline longitudinal incision. After lateral
patellar dislocation, the normal ACL was excised at femoral and tibial origins. Femoral and tibial tunnels were
created with a 3.0 mm diameter drill-bit based on the footprints of the normal ACL. The long digital extensor
tendon (2 mm in diameter and 3 cm in length) was harvested as the tendon graft. Both graft ends were braided
with Dexon 30 suture and passed through the drilling holes, then graft ends were fixed to the tunnel exits with
sutures tied over the neighboring periosteum. The PLA, PLA/HA, or PLA/HA loaded MSCs screw-like scaffolds
were then pressed into the femoral tunnel of each rabbit (Fig. 2). The rabbits were allowed free cage movement
after the operation with intramuscular injection of penicillin (800,000 U) once daily for 3 consecutive days. The
rabbits were sacrificed at 4 and 12 weeks (12 rabbits total, 6 at each time point) for magnetic resonance imagery
(MRI), micro-computed tomography (micro-CT), and histological examinations.
All specimens were examined with a 7.0 T magnetic resonance imaging (MRI) system
for small animals (Agilent VnmrJ 3.1, Agilent Technologies, USA) to observe graft and implant status in the
transverse, coronal, and sagittal sections. The scan parameters were: number of sections =??20, section thickness =??1.00 mm, TR/TE =??600 ms/8 ms, acquisition matrix =??384 ×??192, and FOV =??40 mm ×??40 mm.
Micro-CT Analysis. Micro-CT measurement was performed using a micro-CT system (vivaCT100, Scanco
Medical, Switzerland; 80 kVp, 80 mA) for quantifying mineralized tissue ingrowth inside the bone tunnel (n =??5).
Each specimen was scanned perpendicular to the long bone axis covering the entrance and exit of the femoral
tunnel. To determine the amount and quality of the newly formed mineralized tissue over time, a 3-mm circular
region of interest (ROI) inside the bone tunnel was chosen and three-dimensionally reconstructed using MicView
software (Fig. 3).
Samples were prepared for histological analysis, without decalcification, at each
respective analysis point. The samples were fixed in 4% paraformaldehyde solution for 7 days, dehydrated with
graded alcohols (70, 75, 80, 85, 90, 95, 100%), cleaned with toluene, and embedded in MMA. The embedded
specimens were then sectioned in the anterioreposterior direction and parallel to the longitudinal axis of the long
bone by saw microtome (SP1600, Leica, Germany). Finally, the sections were grinded and polished to 4050 mm
(Exakt-Micro-Grindin System, Leica, Germany) and stained with Von-Gieson to …
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