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2022년 2월 석사학위 논문

Surface Morphology and Biocompatibility of

Hydroxyapatite-coated Ti-40Nb-xZr Alloys after Micro-and Nano-sized

Pore Formation

조선대학교 대학원

치의생명공학과

조 혜 리

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Surface Morphology and Biocompatibility of

Hydroxyapatite-coated Ti-40Nb-xZr Alloys after Micro-and Nano-sized

Pore Formation

Ti-40Nb-xZr 합금의 표면에 마이크로 및 나노 크기의 기공 형성 후, 수산화인회석 코팅 표면의 형상과 생체적합성

2022년 2월 25일

조선대학교 대학원

치의생명공학과

조 혜 리

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Ti-40Nb-xZr 합금의 표면에 마이크로 및 나노 크기의 기공 형성 후,

수산화인회석 코팅 표면의 형상과 생체적합성

지 도 교 수 최 한 철

이 논문을 공학석사학위신청 논문으로 제출함 2021년 10월

조선대학교 대학원

치의생명공학과

조 혜 리

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조혜리의 석사학위논문을 인준함

위 원 장 조선대학교 교 수 안 상 건 (인) 위 원 조선대학교 교 수 김 병 훈 (인) 위 원 조선대학교 교 수 최 한 철 (인)

2021년 12월

조 선 대 학 교 대 학 원

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Table 5. HA target ingredient used in RF-magnetron sputtering treatment 32 Table 6. Concentration of human plasma and SBF solution 37 Table 7. XRF results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, followed by 0℃ water quenching 42 Table 8. Elastic modulus and nano-indentation hardness value of Ti-40Nb alloy according to

phase 47

Table 9. Elastic modulus and nano-indentation hardness value of Ti-40Nb-xZr alloys 51 Table 10. Pore analysis of PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and

P ions 57

Table 11. Oxide layer thickness measurements value of Ti-40Nb-xZr alloys after PEO- treatment in a solution containing Ca and P ions 60 Table 12. The diameters of the nanotube bottom formed on the Ti-40Nb-xZr alloys:

(a) large nanotube bottom, (b) small nanotube bottom 65 Table 13. Nanotube thickness formed on the Ti-40Nb-xZr alloys according to Zr

contents 67

Table 14. Surface roughness by AFM of Ti-40Nb-xZr alloys with surface treatment

method 83

Table 15. Contact angles of Ti-40Nb-xZr alloys with surface treatment method 85

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LIST OF FIGURES

Fig. 1. Theα-β phase transformations in titanium alloys: (a) α phase crystal structure, (b) β phase crystal structure, and (c) phase transformation mechanism. 08 Fig. 2. Effect of alloying elements on phase diagrams of titanium alloys. 09 Fig. 3. Young's modulus values of α+β alloys and β alloys according to the added

elements. 11

Fig. 4. Spark discharge image of Ti-40Nb-xZr alloys according to applied voltage and alloys during PEO treatment in a solution containing Ca and P ions: (a) Ti-40Nb treated at 180V, (b) Ti-40Nb treated at 380V, (c) Ti-40Nb-15Zr treated at 180V,

(d) Ti-40Nb-15Zr treated at 380V. 16

Fig. 5. Mechanism of microcracks and micropore formation in the PEO treatment 17 Fig. 6. FESEM images of PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and P ions: (a) Ti-40Nb treated at 180V, (b) Ti-40Nb treated at 380V, (c) Ti-40Nb- 15Zr treated at 180V, and (d) Ti-40Nb-15Zr treated at 380V. 17 Fig. 7. Schematic diagrams of plasma electrolytic oxidation treatment on the titanium

surface in an electrolyte. 18

Fig. 8. Schematic diagrams of anodic oxidation treatment on the titanium surface in an

electrolyte. 21

Fig. 9. Mechanism of nanotube formation on the titanium surface. 21

Fig. 10. Mechanism of RF-magnetron sputtering system. 23

Fig. 11. The structural formula of hydroxyapatite. 24

Fig. 12. Schematic diagram of RF-magnetron sputtering system. 31 Fig. 13. Plasma images of RF-magnetron sputtering treatment using HA target: (a) pre- sputtering image, and (b) sputtering image. 32 Fig. 14. Mechanism of wettability measurement on substrate. 35 Fig. 15. XRF results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, follwed by 0℃ water quenching: (a) Ti-40Nb, (b) Ti-40Nb-3Zr,

(c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 41

Fig. 16. Optical micrographs of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, followed by 0℃ water quenching: (a) Ti-40Nb, (b) Ti-40Nb-3Zr,

Fig. 17. XRD results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar

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(c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 44 Fig. 18. Optical micrographs of Ti-40Nb alloy after nano-indentation measurement according

to phase. 46

Fig. 19. Nano-indentation test results of Ti-40Nb alloy according to phase after heat

treatment at 1050℃ for 1h in Ar atmosphere followed by 0℃ water quenching. 47 Fig. 20. Optical micrographs of Ti-40Nb-xZr alloys after nano-indentation measurement:

(a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 50 Fig. 21. Nano-indentation test results of Ti-40Nb-xZr alloys after heat treatment at 1050℃

for 1h in Ar atmosphere followed by 0℃ water quenching. 51 Fig. 22. Spark discharges images of Ti-40Nb-xZr alloys with coating time in PEO treatment in a solution containing Ca and P ions: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb

-7Zr, and (d) Ti-40Nb-15Zr. 55

Fig. 23. FESEM images of PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and P ions: (a) Ti-40Nb, (a-1) high magnification of Ti-40Nb, (b) Ti-40Nb-3Zr, (b-1) high magnification of Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, (c-1) high magnification of Ti -40Nb-7Zr, (d) Ti-40Nb-15Zr, and (d-1) high magnification of Ti-40Nb-15Zr. 56 Fig. 24. PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and Pions: (a) small and large pore size, (b) porosity and number of pore, and (c) fraction of large

pores. 57

Fig. 25. EDS result of PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and P ions: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 58 Fig. 26. FESEM images and EDS line profiling of the coating layer of PEO-treated

Ti-40Nb-xZr alloys at 280V in a solution Ca and P ions: (a) Ti-40Nb, (b) Ti-40Nb -3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 59 Fig. 27. Oxide layer thickness of the coating layer of PEO-treated Ti-40Nb-xZr alloys in a

solution containing Ca and P ions. 60

Fig. 28. XRD results for Ti-40Nb-xZr alloys after PEO-treatment in a solution containing Ca and P ions: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and

(d) Ti-40Nb-15Zr. 61

Fig. 29. FESEM images of the nanotube surface formed on the Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte by anodization for 1 h at 20V: (a) Ti-40Nb,

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Fig. 30. The diameter value of the nanotube bottom formed on the Ti-40Nb-xZr alloys:

(a) large nanotube bottom, (b) small nanotube bottom. 65 Fig. 31. FESEM images of the nanotube thickness formed on the Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte by anodization for 1 h at 20V:

(a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 66 Fig. 32. Nanotube thickness formed on the Ti-40Nb-xZr alloys according to Zr contents. 67 Fig. 33. XRD results of the nanotube surface formed on the Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte by anodization for 1 h at 20V: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 68 Fig. 34. FESEM images of HA-coated surface by RF-magnetron sputtering on Ti-40Nb-xZr alloys: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 70 Fig. 35. EDS result of HA-coated surface by RF-magnetron sputtering on Ti-40Nb-xZr alloys: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 71 Fig. 36. XRD result of HA-coated surface by RF-magnetron sputtering on Ti-40Nb-xZr alloys: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 72 Fig. 37. FESEM images of HA-coated surface by RF-magnetron sputtering on PEO-treated Ti-40Nb-xZr alloys in solution containing Ca and P: (a) Ti-40Nb, (b) Ti-40Nb-3Zr,

Fig. 38. EDS result of HA-coated surface by RF-magnetron sputtering on PEO-treated Ti-40 Nb-xZr alloys in solution containing Ca and P: (a) Ti-40Nb, (b) Ti-40Nb-3Zr,

Fig. 39. XRD result of HA-coated surface by RF-magnetron sputtering on PEO-treated Ti-40 Nb-xZr alloys in solution containing Ca and P: (a) Ti-40Nb, (b) Ti-40Nb-3Zr,

Fig. 40. FESEM images of HA-coated surface by RF-magnetron sputtering on nanotube surface formed Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte : (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 78 Fig. 41. EDS results of HA-coated surface by RF-magnetron sputtering on nanotube surface formed Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte:

(a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr. 79 Fig. 42. XRD results of HA-coated surface by RF-magnetron sputtering on nanotube surface formed Ti-40Nb-xZr alloys in 1.0M H3PO4 with 0.8 wt.% NaF electrolyte:

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Fig. 43. 3D AFM images of HA-coated surface by RF-magnetron sputtering on Ti-40Nb-xZr

alloys. 83

Fig. 44. Contact angle image of HA-coated surface by RF-magnetron sputtering on Ti-40Nb-

xZr alloys. 85

Fig. 45. FESEM images of bone growth in SBF solution for 24 hours on the surface

treated Ti-40Nb-xZr alloys. 87

Fig. 46. FESEM images of MC3TC-E1 cells cultured for 24 hours on the surface treated

Ti-40Nb-xZr alloys. 89

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국 문 초 록

Ti-40Nb-xZr 합금의 표면에 마이크로 및 나노 크기의 기공 형성 후, 수산화인회석 코팅 표면의 형상과 생체적합성

조 혜 리

지도교수 : 최 한 철, 공학/치의학 박사 치의생명공학과

조선대학교 대학원

본 연구에서는 기존 임플란트 재료로 사용되는 Ti-6Al-4V 합금을 대체하기 위해 저탄성계수를 갖는 무독성 티타늄 합금을 개발하였다. 순수한 Ti에 β형 원소인 Nb 를 40 wt.% 첨가하고 Zr 함량을 (x = 0, 3, 7, 15 wt.%) 변화시켜 새로운 3원계 합 금 (Ti-40Nb-xZr)을 제조하였다. Ti-40Nb-xZr 합금의 표면에는 플라즈마 전해 산화

(PEO)처리를 통해 마이크로 크기의 기공을 형성하였고, 양극산화 처리를 통해 나노

크기의 기공을 형성하였다. 마이크로 및 나노 크기의 기공 위에 RF-마그네트론 스 퍼터링을 이용하여 HA 코팅한 후 다양한 실험 장비를 사용하여 합금의 표면을 관 찰하였다.

새로운 3원계 합금인 Ti-40Nb-xZr (x = 0, 3, 7, 15 wt.%)은 아크 용해로를 사용하 여 Ar 분위기에서 제조하였다. 화학적 균질성을 향상시키기 위해 10회 연속으로 용 융시키고 1050℃에서 1시간 동안 균질화한 후 급랭시켰다. PEO 처리는 DC 전원 (Keysight Co. Ltd., USA)을 사용하여 수행하였으며 시편을 양극으로, 탄소 막대를 음극으로 사용하였다. Ca 및 P 이온 (0.15M calcium acetate monohydrate + 0.02M calcium glycerophosphate)을 전해질로 포함하는 용액에서 수행하였으며, 인가전압은 280V, 인가 시간은 3분으로 설정하였다. 직류 (DC) 공급장치 (E3641A, Agilent, USA)를 사용해 백금을 상대전극, 시편을 작업 전극으로 사용하는 2전극 구성으로 양극 산화 처리를 수행하였다. 합금 표면에 나노튜브는 1M H3PO4 + 0.8 wt.% NaF 전해질 속에서 1시간 동안 20V의 전압을 인가하여 형성하였다. RF-마그네트론 스 퍼터링 시스템 (A-Tech System Co., Korea)을 이용하여 Ar 가스 분위기에서 HA 박 막을 코팅하였다. 코팅 압력은 5.0 × 10⁻²Torr로 유지하였고, 50W의 RF 전력을 30 분간 인가하여 플라즈마를 형성하였다. Ti-40Nb-xZr 합금의 표면 및 기계적 특성은 광학현미경, 전계방출형 주사전자현미경, 에너지분산형 X선 분광법, X선 회절분석

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측정, HA 결정 성장 및 세포 부착 시험으로 생체 적합성을 평가하였다.

위와 같은 실험을 통하여 다음과 같은 결론을 얻었다:

1. Ti-40Nb-xZr 합금에서 Zr 함량이 증가함에 따라, 미세구조는 바늘 구조에서 등 축 구조로 바뀌었으며, β상의 피크가 강하게 검출되었다. 나노-인덴테이션 시험 결 과 Zr 함량이 증가할수록 나노-인덴테이션 경도 및 탄성계수는 감소하였다.

2. Ca 및 P 이온이 함유된 전해질에서 PEO 처리 후 Ti-40Nb-xZr 합금 표면에 전 체적으로 다공성이고 불규칙한 기공이 형성되었다. Zr 함량이 증가했을 때, 표면의 기공률 및 기공의 크기는 증가하였으며, 기공의 개수는 감소하였다. XRD 분석 결 과 PEO 층에 의해 Anatase, Rutile, 및 HA 상이 관찰되었다.

3. 1.0M H3PO4 와 0.8 wt.% NaF 전해질 속에서 Ti-40Nb-xZr 합금 표면에 형성된 나노튜브는 Zr 함량이 증가할수록 직경이 작아졌으며, 규칙적인 배열을 하였다.

XRD 분석 결과, 나노튜브층에서 Nb2O5, ZrO2, Anatase, Rutile 상이 검출되었다.

4. 연마 후, PEO 처리 후, 나노튜브 형성 후 RF-마그네트론 스퍼터링 공정을 통 해 HA 코팅된 Ti-40Nb-xZr 합금의 표면은 HA 입자들로 덮여졌으며, 나노 크기의 기공 입구를 코팅층이 덮여진 양상을 보였다. EDS 분석 결과 균일한 코팅층에 의 한 Ca 및 P 이온이 검출되었으며, XRD 분석 결과 HA 상이 관찰되었다.

5. HA 코팅된 Ti-40Nb-xZr 합금 표면의 거칠기는 표면처리법에 따라 큰 차이를 보였다. PEO+HA에서 가장 거친 표면을 보였으며, Nano+HA와 Bulk+HA 순서로 표 면거칠기가 증가하였다. Zr 함량이 증가에 따라 표면 거칠기는 점차 감소하였으나, PEO+HA의 경우, 표면 기공의 거친 형상으로 인해 Ti-40Nb-15Zr에서 가장 거칠었 다.

6. HA 코팅된 Ti-40Nb-xZr 합금의 젖음성을 측정한 결과, Nano+HA에서 접촉각이 가장 낮았다. PEO+HA와 Bulk+HA 순서로 접촉각은 증가하였으며, 모든 표면처리법 에서 Zr 함량이 증가할수록 접촉각은 감소하였다.

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7. 표면처리방법과 관계없이, HA 코팅된 Ti-40Nb-xZr 합금 표면 위에 수산화인회 석은 잘 형성되었다. 수산화인회석의 초기 성장은 마이크로 및 나노 크기의 기공 주변에서 시작되어 표면 주위로 성장하였다.

8. HA 코팅된 Ti-40Nb-xZr 합금에 MC3T3-E1 세포를 성장시킨 결과, 나노 크기의 기공이 형성된 Nano+HA에서 세포 성장이 가장 활발하였다. 그 다음으로 마이크로 크기의 기공이 형성된 PEO+HA에서 세포 성장이 활발하였으며, 모든 표면에서 Zr 함량이 증가할수록 세포 성장 및 분화가 잘 이루어졌다.

본 연구에서는 β형원소를 첨가해 무독성이며 저탄성계수인 Ti-40Nb-xZr 합금을 개발하였다. PEO 및 양극산화 법을 이용하여 합금 표면에 마이크로 및 나노 크기 기공을 형성하였으며, Zr 함량에 따른 기공 형태 변화를 관찰하였다. 마이크로 및 나노 크기의 기공 위에 RF-마그네트론 스퍼터링을 통해 HA 코팅된 Ti-40Nb-xZr은 생체적합성이 향상되었다. 따라서 HA 코팅된 마이크로 및 나노 크기의 기공을 갖 는 Ti-40Nb-xZr은 임플란트 식립 후 조골세포의 접착을 촉진해 골융합 기간을 단축 시킬 것으로 기대된다.

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Ⅰ. INTRODUCTION

Implants are being used to replace teeth lost due to periodontal disease or accidents.

As implant materials, Ti and its alloys that meet the mechanical properties and biocompatibility required for clinical use are used [1]. Among titanium alloys, Ti-6Al-4V alloys have excellent mechanical strength, are bioinert, and are very stable under physiological conditions. In addition, it has a relatively low elastic modulus compared to other alloys and is currently most widely used as an implant material because of its excellent corrosion resistance [2,3].

Clinical reports suggest that biological and biomechanical complications lead to osseointegration failure and implant loss [4,5]. Most biological implant failure causes are traumatic overload, gradual loss of bone support around the implant, infection, or inflammation. This occurs 6 to 12 months after the implant placement. The causes of biomechanical implant failure include implant fracture and mechanical complications, and the difference in elastic modulus between the implant and the surrounding bone tissue is an important issue [4]. The elastic modulus of Ti-6Al-4V alloys currently used as an implant material is 110 GPa, and the elastic modulus of the surrounding bone is 30 GPa. This difference in elastic modulus creates a stress shielding effect, which leads to bone resorption, resulting in aseptic relaxation or fracture. Therefore, interest in the development of β-type alloys with a low elastic modulus to reduce implant failure is increasing [5].

Vanadium (V) in Ti-6Al-4V is a β-stabilizing element but toxic in elemental and oxide forms. In addition, aluminum and vanadium ions released into the body after implant placement are associated with long-term health problems such as Alzheimer's disease, neuropathy, and osteomalacia. Therefore, many studies have been conducted to replace aluminum and vanadium elements with non-toxic β-stabilizing elements such as Nb, Fe, Ta, Mo [6-8].

Ti cannot form sufficient osseointegration with natural bone tissue like most other metals due to its inert characteristic. In addition, implants made of Ti are difficult to

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the long term [9,10]. Therefore, various surface treatment studies are being conducted to improve the biocompatibility of Ti and its alloys.

Among many surface treatment studies, there is a method of coating bioactive hydroxyapatite (HA) on the alloy surface. HA-coated on the alloy surface improves bone growth and prevents the release of metal ions from the metal implant. Coating methods include sol-gel, plasma spray and pulsed laser deposition, hydrothermal technique, RF-magnetron sputtering, and plasma electrolytic oxidation (PEO) were used [11-13]. Especially, PEO of these coating methods is a relatively convenient technique for forming an oxide layer on a metal. Also, PEO effectively forms a porous or irregularly shaped TiO2 layer on a Ti substrate. In addition, it has the advantage of coating a porous and uniform layer on a metal surface with a complex shape [14]. The chemical composition of the oxide layer can be controlled by the binding of calcium and phosphate species in the electrolyte, which mainly remain as dissolved ions or amorphous phases in the PEO layer [15].

RF-magnetron sputtering coats films with high quality, high density, high adhesion, and excellent thickness at low substrate temperatures. In addition, it has the advantage of obtaining a uniformly and densely controlled film over a large area [16]. In addition, it is relatively inexpensive compared to other deposition methods and has excellent process controllability and long-term stability [17].

In addition to the coating method, an anodic oxidation method is suitable for implant surface modification. Highly ordered nanotubes increase the accumulation efficiency, increasing the surface area and decreasing the contact angle. Therefore, it not only improves the adhesion and proliferation of bone cells on the alloy surface but is also beneficial for the rate of protein adhesion and osseointegration [18]. Also, it is an effective method because the nanotube diameter and layer thickness can be controlled simply by adjusting the applied voltage, electrode spacing, response time, and other parameters [19].

In order to reduce implant failure, research is conducted to improve biocompatibility through surface treatment after manufacturing a new alloy by adding non-toxic elements or β-stabilizing elements. However, studies on the HA coating surface properties

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and bone formation are insignificant.

Therefore, this study developed non-toxic titanium alloys (Ti-40Nb-xZr) with a low elastic modulus by adding Nb and Zr elements to Ti. After that, micro-and nano-sized pores were formed using PEO and anodization methods and then HA-coated using the RF-magnetron sputtering method. The surface properties of the coated surface were analyzed by optical microscope, nano-indentation, field emission scanning electron microscope, energy dispersive x-ray spectrometer, and x-ray diffraction. Biocompatibility was investigated through atomic force microscope, wettability test, hydroxyapatite formation, and cell adhesion.

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Ⅱ. BACKGROUND

Ⅱ. 1. Biomaterials

Biomaterials are materials that interact with biological systems without negatively affecting their natural functions and have had a significant impact on improving the quality of life for humans and animals. The interaction between biomaterials and tissues is important, and to be used as a biomaterial, it must have good corrosion resistance, hydrophilicity, and good bonding with surrounding tissues. In addition, it must be a non-toxic substance that does not cause side effects after insertion into the body [20,21].

Medical devices to which new materials or technologies are applied must be rigorously and systematically designed and verified. In particular, for products that have been verified based on high-risk biodegradable biomaterials, relevant organizations such as Food & Drug Administration (FDA, United States of America), Notified Body (NB, European Union), National Medical Products Administration (NMPA, China) must be evaluated as a medical device [22]. In particular, all materials used or contact the body require a toxicity assessment [23]. As a result of evaluating the cytotoxicity of pure metals, Cu, Al, V, Ag, and Mn had significantly lower cell viability due to cytotoxicity. On the other hand, the cell viability of Zr and Cr was 74.1% and 60.6%, respectively, and Mo, Nb, and Ti showed excellent biocompatibility due to their high cell viability compared to other metals. Based on these results, Ti-based alloys were made, and cytotoxicity weres evaluated. As a result, the cell viability was increased compared to pure titanium, and the cell viability was the highest in the Ti-Nb alloy [24].

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Ⅱ. 2. Titanium alloys

Titanium was first discovered in England by William Gregor in 1791 but was named after the German chemist Martin Klaproth in 1795 after the Greek myth of Titan. In ancient times, titanium was considered a rare metal. However, today, due to its high strength and lightweight, it is widely used in various industries. Examples include the aerospace industry, dental implants, and materials for orthopedic surgery. Although titanium is the ninth most abundant element in the earth's crust, it is difficult to extract, smelt, and process due to its strong durability, making it a costly metal despite its abundant reserves.

Titanium has an atomic number of 22 and an atomic weight of 47.9. It belongs to the 4 th period of Mendeleev's periodic table and is one of the group IV transition elements, and has an incompletely filled d shell in its electronic structure. Due to its unstable d shell, titanium can form solid solutions with most replacement elements with size factors within 20%. The basic physical properties of titanium are shown in Table 1 [25].

Titanium, which is currently commercially available, is classified into grades 1 to 4 according to the content of impurities nitrogen, carbon, hydrogen, oxygen, and iron.

Grade 1 has low impurities, and grade 4 increases the impurity content. Since the composition of each grade is different and has different properties, it is necessary to use an appropriate grade of titanium in consideration of the characteristics of the field to be used. The detailed grade according to the impurity content is shown in Table 2.

Titanium has an HCP structure formed in an α-phase at room temperature, and when a temperature of 882.5℃. or higher is applied, it is transformed into a β-phase having a body-centered cubic structure (bcc). The crystal structure of the α and β phases and the phase change mechanism are shown in Figure 1 [26]. According to the details of the phase diagram, the β-stabilizing elements of titanium alloys are divided into β-isomorphic elements and β-eutectoid-forming elements, which are schematically shown in Figure 2. The most frequently used β isoform elements in titanium alloys

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possible to stabilize the β phase at room temperature. Cr, Fe, and Si are widely used in the β vacancy-forming elements, and the use of the remaining elements is limited [27]. In general, alloying elements are added to reconstitute the phase of titanium atoms, and the addition of alloying elements results in a change in the transformation temperature. Alloying elements with wide solubility in the β phase are called β stabilizers and lower the alpha/beta transformation temperature. Conversely, alloying elements with a wide range of solubility in the α phase are called α stabilizers [28].

The properties of the material change according to the ratio of α and β phases added to the titanium alloy, and the hardness and elastic modulus of the alloy gradually decrease from α to β phase [29].

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Table 1. Summary of physical properties of unalloyed titanium

Property Value

Atomic number 22

Atomic weight (g/mol) 47.90

Crystal structure

Alpha, hexagonal, closely packed

c (Å) 4.6832 ± 0.0004

a (Å) 2.9504 ± 0.0004

Beta, cubic, body centerd

a (Å) 3.28 ± 0.003

Density (g cm^-3) 4.54

Coefficient of thermal expansion,

α, at 20℃ (K^-1) 8.4 x 10^-6

Thermal conductivity (W/(m K)) 19.2

Melting temperature (℃) 1668

Boiling temperature (estimated) (℃) 3260

Transformation temperature (℃) 882.5

Electrical resistivity

High purity (µΩ cm) 42

Commercial purity (µΩ cm) 55

Modulus of elasticity, α, (GPa) 105

Yield strength, α, (MPa) 692

Ultimate strength, α, (MPa) 785

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Fig. 1. Theα-β phase transformations in titanium alloys: (a) α phase crystal structure, (b) β phase crystal structure, and (c) phase transformation mechanism [26].

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Fig. 2. Effect of alloying elements on phase diagrams of titanium alloys [27].

Table 2. Components of pure titanium (cp-Ti) grades 1 - 4

Grade (%)

Grade 1 Grade 2 Grade 3 Grade 4

N max. 0.030 0.030 0.050 0.050

C max. 0.080 0.080 0.080 0.080

H max. 0.015 0.015 0.015 0.015

O max. 0.180 0.250 0.350 0.400

Fe max. 0.200 0.300 0.300 0.500

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Ⅱ. 3. Titanium alloy for metallic biomaterials

Titanium and titanium alloys are widely used in biomaterials because of their low elastic modulus, excellent biocompatibility, and corrosion resistance compared to stainless steel and cobalt-based alloys used in the past. Among titanium alloys, Ti-6Al-4V alloys were originally developed for aerospace applications, but due to their high corrosion resistance and excellent biocompatibility, it has been introduced into the biomedical field and is currently the most widely used [30]. However, the aluminum (Al) element can cause neurotoxicity under certain conditions and can cause motor dysfunction. It is also associated with kidney disease and Alzheimer's disease, so aluminum accumulation is harmful to the human body. The vanadium (V) element also has a toxic effect as a carcinogen, and cases of implant failure due to the release of vanadium elemental have been reported. Therefore, many studies are being conducted to replace aluminum and vanadium elements with non-toxic and biocompatible elements [31].

Recently, the focus has shifted to the development of β-titanium alloys, also called second-generation titanium alloys. As alloying elements, niobium (Nb), zirconium (Zr), molybdenum (Mo), tantalum (Ta), iron (Fe), and the like are used. The β-type element lowers the titanium alloy's elastic modulus, solving the stress shielding effect caused by the elastic modulus mismatch between the implant and the bone. The β-titanium alloy shows a lower Young's modulus value than the existing α+β titanium alloy, and Young's modulus value according to the alloying elements is shown in Figure 3 [31].

Cell growth and proliferation on the implant surface are important for success after implant placement. In particular, good osseointegration is required for clinical use in patients with the poor bone condition. Therefore, to improve the biological response of implants, many studies are being conducted to increase the surface area by making micro-and nano-sized pores on the implant surface [13,32,33].

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Fig. 3. Young's modulus values of α+β alloys and β alloys according to the added elements [31].

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Ⅱ. 4. Ti-40Nb-xZr alloys

Ⅱ. 4. 1. Ti-Nb alloy

Niobium (Nb) is non-toxic and lowers the elastic modulus by changing α-Ti to β -Ti when used in titanium alloys as a β-type element. As the Nb content added to the alloy increases, the alloy changes from two-phase (α”+β) to single-phase (β), and the average particle size and tensile strength of β-phase decrease. Many studies on developing Ti-Nb implants are being done because titanium alloys with Nb elements have excellent corrosion resistance [34,35].

Ⅱ. 4. 2. Ti-Zr alloy

Currently, zirconium (Zr) is being used to improve the mechanical and corrosion properties of Ti-based materials. Zr element added to improve mechanical properties lowers the elastic modulus of the titanium alloy. In addition, TiO2 and ZrO2 forming a passivation film are combined so that the Ti-Zr alloy has better corrosion resistance than pure Ti. As the Zr content increases, the corrosion resistance becomes excellent, and the corrosion properties can be improved. When used as an implant material, zirconium (Zr) is a non-toxic element and does not cause long-term health problems.

In addition, a high revelation of osteogenic genes in osteoblasts on the Ti-Zr alloy surface and excellent early osseointegration in vivo were found. Therefore, it shows better bone formation ability than pure Ti [36].

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Ⅱ. 4. 3. Ti-40Nb-xZr alloys

Recently, many alloys, including β-phase stabilizing elements such as niobium (Nb) and zirconium (Zr), have been developed. The alloy containing Nb and Zr elements has a lower elastic modulus than the existing Ti-6Al-4V. In addition, it does not contain toxic Al and V elements, so it is safe for various diseases and suitable for use as a biomaterial. In fact, as a result of the cytotoxicity test, it was found that alloys containing Nb and Zr elements did not exhibit cytotoxic effects [37]. Titanium alloys with added Nb elements have better biocompatibility than pure Ti and In alloys containing 40 wt.% Nb improved cell-material interactions were observed [38].

Therefore, Ti-40Nb-xZr is non-toxic and has low elastic modulus alloys, and its biocompatibility is also improved compared to pure Ti, which is expected to reduce implant failure.

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Ⅱ. 5. Surface treatment of titanium alloy

Ⅱ. 5. 1. Plasma electrolytic oxidation (PEO) method

The plasma electrolytic oxidation (PEO) treatment, also called micro-arc oxidation (MAO), is one of the surface treatment methods and is performed at a higher potential than the breakdown voltage (150 ~ 800V) of the primary oxide film. Figure 4 shows the spark discharge that occurs during the PEO treatment over time, and it can be observed that the spark discharge appears differently depending on alloy type and applied voltage strength. Sparks appear discontinuously at random points across the entire surface when the supply voltage exceeds the breakdown voltage, and discharge occurs due to the dielectric stability loss of the oxide layer in the low conductivity region. The main characteristics of the PEO coating are microcracks and micropores randomly scattered throughout the coating surface. The formation mechanism is shown in Figure 5. Microcracks are formed by thermal stress generated during coating, and micropores are formed by the discharge of melting products from the micro discharge channel into the solution or rapid solidification of the melting region. Figure 6 shows micropores of various shapes and sizes depending on alloys' composition and applied voltage strength. The performance of PEO coatings is directly affected by the number and size of microcracks and micropores [39]. The reaction and stoichiometry to form titanium oxide and HA product during the PEO treatment are [40]:

(1) Ti ↔ Ti⁴⁺ + 4e^-

(2) Ca²⁺ + OH⁻↔ CaO + H⁺

(3) 3Ca²⁺ + 2PO43− ↔ Ca(PO4)2

(4) 10Ca²⁺ + 6PO43-

+ 2H2O ↔ Ca10(PO4)6(OH)2 + 2H⁺

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(5) Ca²⁺+ Ti⁴⁺+ 3OH^- ↔ CaTiO3 + 3H⁺

The schematic diagram of PEO treatment on Ti alloy based on this reaction is shown in Figure 7. After setting the alloy as the working electrode and the carbon rod as the counter electrode in the electrolyte, an oxidation reaction occurs on the surface of the alloy, and a reduction reaction occurs on the carbon rod when voltage is applied. At this time, a TiO2 layer containing ions in the electrolyte is formed on the surface of the alloy.

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Fig. 4. Spark discharge image of Ti-40Nb-xZr alloys according to applied voltage and alloys during PEO treatment treatment in a solution containing Ca and P ions: (a) Ti-40Nb treated at 180V, (b) Ti-40Nb treated at 380V, (c) Ti-40Nb-15Zr treated at 180V, (d) Ti-40Nb-15Zr treated at 380V.

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Fig. 5. Mechanism of microcracks and micropore formation in the PEO treatment [39].

Fig. 6. FESEM images of PEO-treated Ti-40Nb-xZr alloys in a solution containing Ca and P ions: (a) Ti-40Nb treated at 180V, (b) Ti-40Nb treated at 380V, (c) Ti-40Nb-15Zr treated at 180V, and (d) Ti-40Nb-15Zr treated at 380V.

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Fig. 7. Schematic diagrams of plasma electrolytic oxidation treatment on the titanium surface in an electrolyte.

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Ⅱ. 5. 2. Anodizing method

Nanotubes (TiO2) are easy to form uniformly on the surface of Ti-based alloys and have a tubular structure, which can be advantageous for loading drugs or antibacterial agents, thus attracting a lot of attention from biomaterial scientists and orthopedic surgeons [41]. The formed TiO2 increases the surface-to-volume ratio due to the nanoscale topography, making the alloy surface hydrophilic and lowering the contact angle. Therefore improves osteoblast proliferation, in vivo osseointegration, and implant stability [42]. The following formula describes the oxide film (TiO2) formed on the titanium surface [43].

(1) 2Ti → 2Ti⁴⁺ + 8e^-

(2) Ti⁴⁺ + 4OH^- → Ti(OH)4

(3) Ti⁴⁺ + 2O^2- → TiO2

(4) Ti(OH)4 → TiO2 + 2H2O

(5) 8H⁺ + 8e^- → 4H2

(6) Ti + 2H2O → TiO2 + 2H2

Anodization uses a two-electrode system with platinum as the counter electrode and alloy as the working electrode, and the schematic diagram of the treatment is shown in Figure 8. The nanotubes' diameter, depth, and thickness can be adjusted according to the applied voltage, the applied time, alloying elements, and electrolyte concentrations.

The mechanism for forming nanotubes on the Ti alloy surface is shown in Figure 9.

As the applied voltage increases when forming the nanotubes, nanotube diameters

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In addition, the nanotube length can be adjusted according to the alloying element and electrolyte concentration [44-46].

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Fig. 8. Schematic diagrams of anodic oxidation treatment on the titanium surface in an electrolyte.

Fig. 9. Mechanism of nanotube formation on the titanium surface.

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Ⅱ. 5. 3. RF-magnetron sputtering

Magnetron sputtering was developed in the 1960s and 1970s and is currently the most widely used treatment in thin film deposition and surface engineering. Magnetron sputtering, one of the physical vapor deposition methods, has an advantage in that it can be deposited on most materials because there is no limitation on the material deposited on the substrate. Sputtering is a dynamic controlled treatment in which a source-target is bombarded with ions derived from inert gas. Ions of inert gases are widely used because they can be accelerated more quickly than neutral atoms across the cathode coating towards the cathode target. When the inert gas in a vacuum passes through a direct current, electrons are ejected from the cathode and collide with gas molecules. Most collided molecules are not ionized and are excited to return to their original stable state, but some become ionized. Ionized molecules generate plasma, and in the thin film on the substrate, the positive ions in the plasma are accelerated to the target by the force of the electric field. The accelerated cations cause the cathode material to bounce off the target, as shown in Figure 10. A film formed by sputtering is denser, has better adhesion, and smaller particle size than a film formed by thermal evaporation. Therefore, it is used to form films of overall high quality, high mass density, and low surface roughness. In order to achieve the desired level of thin-film purity, it is usually carried out under high vacuum conditions. The main source of film formation is the heavy use of solid targets. However, new material combinations and compounds can be synthesized by simultaneously using multiple cathode targets or adding reactive gases [47-49].

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Fig. 10. Mechanism of RF-magnetron sputtering system [49].

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Ⅱ. 5. 3. 1. Hydroxyapatite (HA) properties and structure

Hydroxyapatite (HA) is a chemical compound of phosphorus and hydroxyl (OH^-), chloride (Cl^-), or fluoride (F^-) ions. The chemical formula is Ca10(PO4)6(OH)2, and the molecular weight is 1004.6 g/mol. A has a hexagonal crystal structure in which the stoichiometric ratio of Ca/P is generally 1.67. Covalent bonds (P-O and O-H) and ionic bonds (Ca-O) are formed in the crystal lattice, and Figure 11 shows the structure of HA. HA, one of the bioceramics, has a mineral component and crystal structure similar to apatite in the human skeleton. Therefore, it is widely used for surface treatment of implants to increase bioactivity. Pure HA is highly hydrophilic with a contact angle of about 10°, and bone regeneration ability is improved after HA coating on the implant surface [50-52]. In addition, the lattice structure of HA enables cation and anion substitution to improve the biological performance of synthetic bone graft materials [53].

Fig. 11. The structural formula of hydroxyapatite [52].

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Ⅲ. EXPERIMENTAL MATERIALS AND METHODS

Ⅲ. 1. Preparation of Ti-40Nb-xZr alloys

Ti-40Nb-xZr alloys were designed and manufactured by varying Zr contents (x = 0, 3, 7, 15 wt.%). As alloy elements, pellet Cp-Ti (G & S Titanium, Grade 4, 99.5%

USA), and Nb (Kurt J. Lesker Company, 99.95% pure, USA), Zr (Kurt J. Lesker Company, 99.70% pure, USA)were used. The alloy was manufactured in a high-purity argon atmosphere using a vacuum arc melting furnace (Arc skull melting system, Acevacuum, Korea). The oxidation of the designed alloy was minimized by melting Cp-Ti 5 times to remove oxygen in the furnace before manufacturing the alloy. Then, the designed alloy was melted ten times in a button shape and ten times in an ingot shape using a tungsten (W) electrode to improve chemical homogeneity. The manufactured ingot-shaped alloy was cut to a thickness of 3 mm at a speed of 2000 rpm using a high-speed diamond cutting machine (Accutom-5, Struers, Denmark). Then, using a high-temperature furnace (Model MSTF-1650, MS Eng., Korea), it was homogenized at 1050℃ for 1 hour and then quenching. The homogenized alloy was polished in order of #100, #600, #800, #1000, #1200, and #2000 using SiC (silicon carbide) paper. Then, using 0.3 µm alumina powder (Al2O3), it was fine polishing until a mirror surface appeared. After polishing, the alloy was ultrasonically cleaned for 5 minutes each in the order of ethanol and distilled water and then air-dried. The composition of the prepared alloy was analyzed using an x-ray fluorescence analyzer (XRF, Analyzer Mode-Alloy, Analyzer Serial number-581331, DE-2000, Olympus, Japan).

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Ⅲ. 2. Microstructure observation of alloys

The micro-polished Ti-40Nb-xZr alloys were chemically etched using Keller's solution (190 mL of H2O + 2 mL of HF + 3 mL of HCl + 5 mL of HNO3). The etched alloys were ultrasonically washed for 5 minutes each with ethanol and distilled water, then air-dried. The alloy's microstructure was observed using an optical microscope (OM, Olympus BM60M, Japan).

Ⅲ. 3. Plasma electrolytic oxidation (PEO) treatment

Ti-40Nb-xZr alloys for PEO treatment were polished using SiC paper of #100 ~

#2000 and ultrasonically cleaned with ethanol and distilled water for 5 minutes. PEO treatment was performed using a DC power supply (KEYSIGHT Co., Ltd., USA) to use the cleaned specimen as an anode and a carbon rod as a cathode. The applied voltage was set to 280V, and the treatment time was 3 minutes. After the test was finished, the specimen was washed with ethyl alcohol and distilled water and then air-dried. The electrolyte used for PEO treatment was prepared by mixing calcium acetate monohydrate (Ca(CH3COO)2·H2O) and calcium glycerophosphate (C3H7CaO6P), and the concentration of the electrolyte is shown in Table 3.

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Table 3. The condition of plasma electrolytic oxidation

Experimental condition

Solutions

Composition of electrolyte Calcium acetate

monohydrate (CH3COO)2·H2O)

(g/L)

Calcium glycerophosphate

(C3H7CaO6P) (g/L)

Applied voltage (V)

Duration (min)

Ca/P 0.15M 0.02M 280 3

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Ⅲ. 4. Anodizing treatment

Specimens for anodization were sequentially polished from #100 to #2000 using SiC paper, followed by ultrasonically cleaned with ethanol and distilled water for 5 minutes.

Anodizing was performed using a direct current (DC) supply device (E3641A, Agilent, USA) in a two-electrode configuration using platinum as a counter electrode and a specimen as a working electrode. Nanotubes were formed in 1M H3PO4 (phosphoric acid) + 0.8 wt.% NaF (sodium fluoride) electrolyte, and detailed electrolyte concentrations are shown in Table 4. The applied voltage was applied at 20V, and the application time was set to 1 hour to experiment. After the test was finished, the specimen was washed with ethanol and distilled water and then air-dried.

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Table 4. The condition of anodic oxidation

Experimental condition

Solutions

Composition of electrolyte Sodium fluoride

(NaF) (g/L)

Phosphoric acid (H3PO4)

(ml/L)

Applied voltage (V)

Duration (min)

H3PO4 + NaF 0.8 wt.% 1M 20 60

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Ⅲ. 5. RF-magnetron sputtering treatment

The HA thin film was coated using the RF-magnetron sputtering system (A-Tech System Co., Korea), and the schematic composition is shown in Figure 12. In order to obtain a high-quality HA thin film, the chamber, substrate holder, and target holder were washed with ethanol before the sputtering treatment. Sputtering was done by fixing the HA target to an RF-magnetron sputtering gun. The target area was 20.02cm², and detailed components are shown in Table 5 below. The distance between the substrate and the target was 80 mm, and the chamber pressure was initially set to less than 1.0×10^-4Torr. The coating was carried out in an Ar gas atmosphere, and 40-sccm Ar gas was constantly flowed using a mass flow controller. The chamber's pressure during coating was maintained at 5.0×10^-2 Torr, and the substrate temperature was set to 150℃ to increase ion activity. In order to clean the contaminated target before coating, pre-sputtering was performed in an Ar gas atmosphere for 30 minutes. For the HA thin-film coating, plasma was formed by applying 50W RF power for 30 minutes, and Figure 13 is an image of the sputtering treatment. HA coating was performed after polishing, PEO treatment, and nanotube treatment (corresponding sample name = Bulk+HA, PEO+HA, Nano+HA respectively).

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Fig. 12. Schematic diagram of RF-magnetron sputtering system.

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Table 5. HA target ingredient used in RF-magnetron sputtering treatment

Analysis

Target

Tests Results

(PPM)

Calcium Phosphate Hydroxide (Ca10(PO4)6(OH)2)

Al <300

As <10

Ca Matrix

Cd <5

Fe <400

Hg <5

Pb <10

C <120

H Matrix

O Matrix

P Matrix

Fig. 13. Plasma images of RF-magnetron sputtering treatment using HA target: (a) pre-sputtering image, and (b) sputtering image.

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Ⅲ. 6. Analysis of surface properties of Ti-40Nb-xZr alloys

The morphology and composition of the alloy surface were observed using a field emission scanning electron microscope (FESEM; Hitachi 4800, Japan) and an energy dispersive x-ray spectrometer (EDS; Hitachi 4800, Japan). It was performed at an operating voltage of 15 kV, and the scan rate of FESEM was set to 40 sec, and the activation time of EDS was set to 30 sec. The pore size, shape, porosity, and nanotube length of the surface-treated Ti-40Nb-xZr alloys were measured using an image analyzer (Image; Wayne Rasband, USA).

Phase identification of the alloy surface was performed using x-ray diffraction (XRD;

X'pert PRO MPD, PANalytical, The United States). Cu Kα radiation (λ = 1.5418Å) was used, and the voltage and current of the generator were 40 kV and 30 mA, respectively. After fixing the specimen, the x-ray source and detector moved together to investigate a scan range of 20 ~ 80°. The diffraction peaks on the atomic plane were analyzed compared to the Joint Committee on Powder Diffraction Standards (JCPDS) file.

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Ⅲ. 7. Characterization of surface physical properties

Ⅲ. 7. 1. Surface roughness measurement

The surface roughness of the alloy was measured using an atomic force microscope (AFM; Park XE-100, Park Systems, Korea), and then the topography and roughness of the surface were analyzed using the XE Data Acquisition program. Surface roughness was measured at a scan size of 15.00 ㎛ and a scan rate of 0.50 Hz. Surface roughness was measured in a non-contact mode, and a direction parallel to the scratch so that the scratch on the alloy did not affect the roughness value.

Ⅲ. 7. 2. Elastic modulus and hardness measurement using nano-indentation

The elastic modulus and hardness of the alloy surface were measured using a nano-indentation tester (TTX-NHT3, Anton Paar, Austria). The maximum load was set to 20 mN and the stopping time was set to 5 seconds, and the average value and standard deviation were obtained by measuring five times to obtain the minimum and maximum values of each sample.

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Ⅲ. 8. Measurement of surface wettability

The wettability of the HA-coated alloy surface was measured at room temperature using a water contact angle goniometer (KSA100, Kruss, Germany). The thickness of the needle used was 6 ㎛, and the contact angle was measured using a static droplet method and automatically falling water droplets with a video camera and a contact angle meter. The reliability of the average value was increased by measuring more than five times, and the contact angle measurement mechanism is shown in Figure 14.

Fig. 14. Mechanism of wettability measurement on substrate.

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Ⅲ. 9. Formation of hydroxyapatite in simulated body fluid (SBF) solution

To investigate the HA-coated samples' biological activity (hydroxyapatite-forming ability), samples were immersed in a similar solution (SBF, Simulated Body Fluid) for 24 hours. The SBF solution was maintained at 36.5 ± 1℃, which is the same as the human body temperature, and the detailed ion concentration of the SBF solution is shown in Table 6. The pH concentration was maintained at 7.4 ± 0.5 using tris(hydroxymethyl) aminomethane, 99.0% (C4H11NO3 = 121.14) and 1:9 HCl (hydrochloric acid, 36.46 g/mol). After the test was finished, samples were washed in distilled water and air-dried. The surface was observed with FESEM.

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Table 6. Concentration of human plasma and SBF solution

Ions Concentration (mM)

Blood plasma SBF

Na⁺ 142.0 142.0

K⁺ 5.0 5.0

Mg²⁺ 1.5 1.5

Ca²⁺ 2.5 2.5

Cl^- 103.0 103.0

HCO^3- 27.0 10.0

HPO^4- 1.0 1.0

SO42-

0.5 0.5

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Ⅲ. 10. Cell culture and observation

MC3T3-E1 was an osteoblast precursor cell line extracted from the skull of mice and appropriately distributed at a concentration of 1x10⁵ cells/well. Then, the cells were cultured in α-MEM (Alpha – minimum essential medium, without L-ascorbic acid) supplemented with 10% fetal bovin serum and 10 U/ml of penicillin/straptomycin.

The cell monolayer was washed with phosphate buffered saline (PBS) and incubated for 10 minutes in a trypsin-DTA solution (0.05% trypsin, 0.53 mM EDTA·4Na, phenol red in HBSS) at 37℃ to separate the cells. Cells were seeded on the surface of the HA-coated samples at a concentration of 1.5 x 10⁵ cells/well on a 12-well plate and grown for 24 hours. The treated specimens were washed with PBS and fixed with 10%

formaldehyde at 4℃ for 1-2 hours. The fixed samples were dehydrated with ethanol, and the morphology of the fixed cells was observed through FESEM.

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Ⅳ. RESULTS AND DISCUSSION

Ⅳ. 1. Microstructure and phase analysis of Ti-40Nb-xZr alloys

Figure 15 shows the composition of Ti-40Nb-xZr (x = 0, 3, 7, 15 wt.%) alloys analyzed using an XRF. Figure 15 (a) is Ti-40Nb, (b) is Ti-40Nb-3Zr, (c) is Ti-40Nb-7Zr, and (d) is Ti-40Nb-15Zr alloy. In Figure 15 (a), the Zr peak was not observed because Zr was not added. In Figure 15 (b ~ d), the Zr peak was detected due to the addition of Zr, and as the Zr content increased, the Zr peak was strongly detected. To increase chemical homogeneity, a homogeneous alloy was obtained as a result of continuous melting during alloy manufacturing and then heat treatment at 105 0℃. The detailed alloy's composition is shown in Table 7, and it was confirmed that the alloy was well manufactured, similar to the first designed composition.

Figure 16 shows images of the microstructure of Ti-40Nb-xZr according to the Zr content observed on an optical microscope. Figure 16 (a) is a microstructure images of Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr alloy. Figure 16 (a) shows the microstructure composed of needle-like structure (α”) and equiaxed structure (β), and Figure 16 (d) shows the microstructure composed of only equiaxed structure (β). As the Zr content increased, β-phase could be observed, which is a change caused by a decrease in the proportion of Ti in the alloy as the Zr element increases [54].

Figure 17 (a ~ d) shows the x-ray diffraction diagram of the manufactured alloys, showing Ti-40Nb, Ti-40Nb-3Zr, Ti-40Nb-7Zr, and Ti-40Nb-15Zr. At 2θ = 29.23°, the α”-phase peak was detected in Ti-40Nb. However, it was not detected in the Zr-added alloy. For Ti-40Nb, α”-and β-phase peaks were observed, and for Ti-40Nb-15Zr, only β-phase peaks were observed. This shows that alloys change from the α”+β phase to the β phase by adding Zr. The result is consistent with Figure 16, when the microstructure was observed with an optical microscope. The β-phase peaks at 2θ =

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the Zr content increased Ti-40Nb-xZr alloys. As the Zr content of the Ti-40Nb-xZr alloys increased, the FWHM (full width at half maximum) of x-ray peaks became narrower, which means that the grain size of the alloy increased and the hardness decreased [55]. In addition, as the Zr content increased, the peak shifted to the left, which is a phenomenon that occurs when Zr added, which has a larger atomic radius than Ti and Nb elements, is inserted into a solid solution, increasing lattice parameters [56].

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Fig. 15. XRF results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, follwed by 0℃ water quenching: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr.

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Table 7. XRF results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, followed by 0℃ water quenching

Elements (%) Specimen

Ti Nb Zr

Ti-40Nb 60.26 ± 0.52 39.74 ± 0.51 -

Ti-40Nb-3Zr 55.96 ± 0.66 40.94 ± 0.61 3.10 ± 0.06 Ti-40Nb-7Zr 53.12 ± 0.66 39.77 ± 0.54 7.11 ± 0.12 Ti-40Nb-15Zr 45.15 ± 0.76 39.68 ± 0.53 15.17 ± 0.23

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Fig. 16. Optical micrographs of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, followed by 0℃ water quenching: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr.

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Fig. 17. XRD results of Ti-40Nb-xZr alloys after heat treatment at 1050℃ for 1h in Ar atmosphere, followed by 0℃ water quenching: (a) Ti-40Nb, (b) Ti-40Nb-3Zr, (c) Ti-40Nb-7Zr, and (d) Ti-40Nb-15Zr.

저작자표시

2022년 2월 석사학위 논문

Surface Morphology and Biocompatibility of

Hydroxyapatite-coated Ti-40Nb-xZr Alloys after Micro-and Nano-sized

Pore Formation

조선대학교 대학원

치의생명공학과

조 혜 리

Surface Morphology and Biocompatibility of

Hydroxyapatite-coated Ti-40Nb-xZr Alloys after Micro-and Nano-sized

Pore Formation

Ti-40Nb-xZr 합금의 표면에 마이크로 및 나노 크기의 기공 형성 후, 수산화인회석 코팅 표면의 형상과 생체적합성

2022년 2월 25일

조선대학교 대학원

치의생명공학과

조 혜 리

Ti-40Nb-xZr 합금의 표면에 마이크로 및 나노 크기의 기공 형성 후,

수산화인회석 코팅 표면의 형상과 생체적합성

지 도 교 수 최 한 철

이 논문을 공학석사학위신청 논문으로 제출함 2021년 10월

조선대학교 대학원

치의생명공학과

조 혜 리

조혜리의 석사학위논문을 인준함

위 원 장 조선대학교 교 수 안 상 건 (인) 위 원 조선대학교 교 수 김 병 훈 (인) 위 원 조선대학교 교 수 최 한 철 (인)

2021년 12월

조 선 대 학 교 대 학 원

CONTENTS

국 문 초 록

Ⅰ. INTRODUCTION

Ⅱ. BACKGROUND

Ⅱ. 1. Biomaterials

Ⅱ. 2. Titanium alloys

Ⅱ. 3. Titanium alloy for metallic biomaterials

Ⅱ. 4. Ti-40Nb-xZr alloys

Ⅱ. 4. 1. Ti-Nb alloy

Ⅱ. 4. 2. Ti-Zr alloy

Ⅱ. 4. 3. Ti-40Nb-xZr alloys

Ⅱ. 5. Surface treatment of titanium alloy

Ⅱ. 5. 1. Plasma electrolytic oxidation (PEO) method

Ⅱ. 5. 2. Anodizing method

Ⅱ. 5. 3. RF-magnetron sputtering

Ⅱ. 5. 3. 1. Hydroxyapatite (HA) properties and structure

Ⅲ. EXPERIMENTAL MATERIALS AND METHODS

Ⅲ. 1. Preparation of Ti-40Nb-xZr alloys

Ⅲ. 2. Microstructure observation of alloys

Ⅲ. 3. Plasma electrolytic oxidation (PEO) treatment

Ⅲ. 4. Anodizing treatment

Ⅲ. 5. RF-magnetron sputtering treatment

Ⅲ. 6. Analysis of surface properties of Ti-40Nb-xZr alloys

Ⅲ. 7. Characterization of surface physical properties

Ⅲ. 7. 1. Surface roughness measurement

Ⅲ. 7. 2. Elastic modulus and hardness measurement using nano-indentation

Ⅲ. 8. Measurement of surface wettability

Ⅲ. 9. Formation of hydroxyapatite in simulated body fluid (SBF) solution

Ⅲ. 10. Cell culture and observation

Ⅳ. RESULTS AND DISCUSSION

Ⅳ. 1. Microstructure and phase analysis of Ti-40Nb-xZr alloys