【作者】 赵亚凡；

【导师】 陈传忠；

【作者基本信息】 山东大学 ， 材料学， 2007， 博士

【摘要】 生物玻璃是一种透明的生物活性材料，其化学组成与生物体自然骨骼中的硬组织相似，具有良好的生物相容性，与羟基磷灰石相比，其活性更高，成分可根据需要在较大范围内调节。然而，生物玻璃的疲劳强度与断裂韧性较低，难以在承载部位使用。若将其作为涂层材料与钛合金基体相结合，便兼有金属材料优良的机械性能与生物材料良好的生物性能，因而是理想的骨修复材料。目前生物玻璃涂层技术已得到广泛的研究。与传统的涂层制备技术，如涂覆-烧结、等离子喷涂、溶胶凝胶、电泳沉积以及离子束增强沉积等相比，脉冲激光沉积（Pulsed Laser Deposition，PLD）能够制备出与靶材保持一致的化学成分计量比的薄膜，所制备的薄膜具有连续性好、组织细小致密、结合强度高等特点。本文在钛合金表面采用脉冲激光沉积技术制备生物玻璃薄膜，通过调节PLD工艺参数及生物玻璃涂层材料的成分，使涂层与基体结合牢固，并且保持良好的生物性能。目前国际上对脉冲激光沉积生物玻璃的工艺研究才刚刚起步，关于脉冲激光工艺参数的制定还处于尝试阶段，缺乏系统的研究。本文通过调节脉冲激光工艺参数，主要包括Ar气压力、衬底温度、功率密度以及沉积时间等，研究脉冲激光工艺参数对生物玻璃薄膜的成分、表面形貌、微结构以及结合强度的影响，从而为优化工艺参数提供理论依据。Ar气压力对薄膜的形貌、微结构及沉积速率的影响表现在：（1）薄膜表面存在大量的白色颗粒，不同气压下颗粒的形态和数量略有差异。真空状态下，颗粒形状以球形为主，随气压升高，颗粒数目增多，出现许多不规则形状的小颗粒。根据固态、液态及气态三种类型颗粒的形成机制可得，白色颗粒多数为由熔融态液滴形成的球形颗粒物，另外一些不规则的大颗粒则与靶材的玻璃结构有关。真空及低背景气压下，溶蚀粒子几乎没有受到碰撞就到达基底，因而薄膜表面以球形液滴为主。随气压升高，沉积室内的气体逐渐达到过饱和状态，从而在薄膜表面形成气体凝聚的颗粒。（2）脉冲激光沉积生物玻璃所得薄膜与靶材的微观结构并不一致，表现在与靶材相比，薄膜中的活性基团Si-O-NBO与Si-O（s）的峰值之比有所降低，且随压力升高降低得越明显。可见，薄膜在形成与生长的过程中，原有生物玻璃的网络结构发生了变化，这与PLD复杂的物理机制有关。脉冲激光沉积得到的薄膜是由高能量的气体瞬间凝结而成，因而各粒子难以在极短时间内扩散到其平衡位置，导致Si-O-NBO功能团在一些区域发生偏聚，在电荷斥力的作用下，结合较薄弱的Si-O-NBO键易断开而形成牢固的Si-O-Si键。气压越高，烧蚀产物到达衬底的动能越低，扩散越不充分，越易形成Si-O-NBO功能团的偏聚，因而就有更多的Si-O-NBO功能团转化为Si-O-Si功能团，且随气压升高，这种趋势更加明显。（3）气压对沉积速度的影响表现在随气压升高，沉积速度线性降低。衬底温度对薄膜的形貌、微结构及结合力的影响表现在：（1）衬底温度对薄膜表面颗粒的形貌影响不大，但影响薄膜的粗糙度，较低的衬底温度导致薄膜表面不平整。原因是衬底温度较低时，吸附于衬底粒子的平均能量较低，粒子的迁移能力减弱，因而粗糙度增大。（2）温度对薄膜中活性基团峰值的影响表现在Si-O-NBO与Si-O（s）的峰值之比随衬底温度降低而降低，其原因同样与PLD过程中的非平衡扩散机制有关。（3）薄膜与基体的结合强度随衬底温度降低而降低。原因是衬底温度越低，温度梯度越高，产生的热应力则越高，结合强度也就越低。功率密度对薄膜成分、形貌、沉积速率以及结合力的影响表现在：（1）较低的功率密度难以保证薄膜与靶材成分的一致性，造成生物玻璃薄膜与靶材相比富Na、Ca而贫Si、P。（2）功率密度对薄膜表面形貌的影响表现在高的功率密度提高了薄膜表面的粗糙度，且使已形成的颗粒物在冲击作用下因受压而变形。（3）沉积速率随功率密度的增加而增加，但在高功率密度范围内，沉积速率的增加相对变得缓慢。（4）随功率密度增大，薄膜与基体的结合强度减小，原因与薄膜本身的质量变差有关。沉积时间通过影响薄膜的厚度而影响薄膜与基体的结合强度。沉积时间越长，薄膜的厚度越高，薄膜与基体的结合强度则越低。原因是较薄的薄膜受基体的影响较大，可以松弛一部分与其相接触的薄膜产生的应力，而较厚的薄膜则产生了应力积累，致使结合强度降低。脉冲激光沉积薄膜的物理机制非常复杂。为研究脉冲激光烧蚀靶材产生等离子体而建立的S-N模型，将靶材理想化为无限大平板，没有考虑到其实际厚度，显得较为粗糙。本文研究的是PLD技术中的第一个过程（脉冲激光对靶材的烧蚀）及第二个过程（等离子体的运动）的物理机制。在建立脉冲激光作用块状靶材的一维热流烧蚀模型的基础上，根据能量平衡原理，导出烧蚀速率的公式，结果表明，激光对靶材的烧蚀率N取决于激光和靶材的物性参数，公式中薄膜的沉积速率与功率密度的关系也与实验结果相吻合，便于指导和优化工艺参数。在建立等离子体演化的动力学模型的基础上，讨论了等离子体在等温膨胀阶段和绝热膨胀阶段的演化规律，通过将烧蚀率公式作为研究等离子体运动的一个边界条件，得出了等离子体空间的各向异性分布公式。从而将激光烧蚀靶材的过程同等离子体空间膨胀过程有机的结合起来，完整的描述了脉冲激光沉积薄膜的前两个过程。目前采用的45S5生物玻璃因热膨胀系数与Ti6Al4V基体不匹配，降低了薄膜与靶材的结合强度，从而影响了种植体的长期使用。本文设计的薄膜成分为在45S5（组分1）生物玻璃的基础上，用MgO取代10％的Na₂O（组分2），因与基体的热膨胀系数较为接近，结合强度有所提高。通过检测生物玻璃薄膜置于模拟体液过程中，表面组织形貌、物相和结构的变化，以及相应的模拟体液中离子浓度及PH值的变化，验证了生物玻璃在模拟体液中经历了如下五步生物活性反应过程：（a）碱金属及碱土金属离子的溶解；（b）硅的水解；（c）富硅胶层的形成；（d）无定形相CaO-P₂O₅的形成；（e）碳酸磷灰石（HCA）的形成；并详细阐明了每一步活性反应的具体过程。由于45S5生物玻璃的活性较高，在模拟体液中易出现磷灰石还没有来得及出现，薄膜己被溶解掉的问题。因此，PLD制备的生物玻璃薄膜需要一定的厚度（远高于羟基磷灰石），才能保证薄膜的稳定性，但是薄膜太厚不仅需要较长的沉积时间，而且由于应力的积累会降低薄膜与基体的结合力。因而存在活性与结合力之间的矛盾。体外生物活性实验结果为组分1与组分2在模拟体液中析出磷灰石的时间分别为70小时与90小时，表明组分2的生物活性低于组分1，即在45S5生物玻璃的基础上以MgO取代部分Na₂O，降低了生物活性，从而降低了薄膜具有生物活性的最小厚度。薄膜表面元素含量Ca／P的实验结果为刚开始Ca／P在增大，说明Ca上升较快，而后Ca／P开始下降，说明P上升较快，从而表明Ca先于P沉积到试样表面。未经模拟体液浸泡的组分2表面的极性基团与组分1稍有差别。组分1中的非桥氧功能团有Si-O-NBO与Si-O-2NBO两种，而组分2中仅有Si-O-NBO，并未检测到Si-O-2NBO功能团。对于生物玻璃在模拟体液中的活性机理及其结构本质，目前还缺乏深入的分析。本文从热力学与动力学的角度解释了生物玻璃的活性机理，并在此基础上阐明了45S5生物玻璃添加镁之后生物活性降低的原因：（1）降低了玻璃中离子的溶解速度（生物活性反应过程1和2）；（2）增大了磷灰石中晶格的畸变程度，延迟了晶态磷灰石形成的时间。运用双电层理论，从离子静电吸引的角度揭示了生物活性的机理，解释了Ca先于P沉积到试样表面的实验结果：由于硅凝胶粒子中含有大量带负电荷的羟基（电位低于生理环境下的PH值），其双电层Zeta电势为负，因而在库仑力的作用下，便先吸附带正电荷的Ca²⁺离子，再吸附具有负电荷的PO₄^3-离子。通过研究生物玻璃网络微结构的特征，揭示了生物玻璃活性的结构本质是非桥氧功能团的存在。碱金属及碱土金属氧化物的含量越高，非桥氧功能团的数量就越多，活性程度也就越高。对于生物玻璃中MgO及其含量对于生物活性的影响，目前说法不一，还没有定论。本文的试验结果表明，组分1结构中出现Q2功能团，而组分2没有的原因与组分2中的Mg有关。研究表明，当MgO＞8％时，一部分会成为网络形成体。组分2中MgO=10％，其中一部分MgO便与SiO₂一同成为网络形成体，相当于提高了硅含量，因而没有检测到Q2功能团。组分2的生物活性及薄膜厚度低于组分1的结构本质：组分2中含量为10％的MgO阻止了功能团Q2（Si-O-2NBO）的出现，而Q2明显促进了活性反应第二步。原因是Q2较Q3（Si-O-NBO）更有利于吸附水进入玻璃，促进硅的水解，提高活性反应速度。这样就从生物玻璃内部结构的角度解释了添加10％含量的MgO降低生物玻璃活性以及薄膜厚度的原因。更多还原

【Abstract】 Among all the currently used biomaterials, bioactive glasses are of great interest in medical applications for their high bioactivity and good biocompatibility. Compared with hydroxyapatite （HA）, another important bioactive ceramic, the bioglass composition can be varied in a wide range without losing its bioactivity. Thus, by changing the composition, the desired physical, mechanical and biological properties for the bioglass can be realized. Besides, as one kind of Class A bioactive materials, bioglass reacts much faster to tissue than HA （in a few hours instead of several days） and can bond to both hard and soft tissue. However, bioactive glasses are brittle and exhibit low mechanical strength. These drawbacks limit its application as bulk material in load bearing sites. Coating metallic prostheses with bioactive materials could be a route to combine the good mechanical properties of base materials with the biological properties of biomaterials. Various coating techniques, such as plasma spraying, sputtering, sol-gel and electrophoresis deposition have been used to obtain bioactive glass films. Among them, pulsed laser deposition （PLD） is a feasible method for producing adherent and uniform films with controlled stoichiometry. In this work, bioactive glass films were prepared on titanium alloy by pulsed laser deposition. The experimental parameters of PLD and composition of bioactive glass films were adjusted in order to abtain high-quality bioactive thin films for medical applications.Up to now, how to select the experimental parameters of PLD still need attempted. Effect of experimental parameters of PLD, including Ar pressure, substrate temperature, energy density and deposition duration on the composition, morphology, bonding configuration and adherence of the films is systematically studied. The purpose is to optimize the experimental parameters of PLD.Effect of the pressure of Ar atmosphere on the morphology, bonding configuration and deposition rate of the bioglass thin films was studied.（1） micron-sized and submicron-sized particles were found on the surface of the thin films, which is the typical morphology of the PLD films. It can be observed that the particles on the surface of the film deposited in the vacuum are mainly in the ball-like shape. The percentage of particles on the films increases with Ar pressure and their shapes become irregular. A great number of droplets are first correlated with the structural feature of the targets. The phenomenon that the quantity of particles increases with the pressure can be interpreted as a consequence of condensation from vapor species under high gas pressure. （2） Compared with the target, the Si-O-NBO/Si-O-Si （s） intensity ratio decreases for the films. Besides, this tendency becomes obvious with the increase of the Ar pressure. The above results demonstrated that the bonding configuration of the target is not correctly transferred to the films. This effect is associated to the network rearrangement during the film growth, originated by the complex physical mechanisms in the PLD process. （3） As the pressure increases, the film growth decreases following an linear dependence.Effect of the temperature on the morphology, bonding configuration and adherence of the bioglass thin films was also investigated.（l） High substrate temperature could supply more kinetic energy for the ablated atoms to diffuse sufficiently, leading to smooth surface. On the country, at low substrate temperature, low energy could be supplied for the mobility of the adatoms on the substrate surface. Thus, the film with high roughness and some structural defects was observed. （2） It can be observed that compared with the target, the Si-O-NBO （s）/Si-O-Si （s） intensity ratio decreases for the films. Besides, this tendency becomes obvious with the decrease of the temperature. （3） The bonding stress between the film and the substrate decreases with the temperature. The reason is that lower temperature of the film means higher temperature gradient, leading to the higher heat stress.Energy density also has an effect on the deposition rate, composition, morphology and adherence of the films. （1） In the low energy density, there are obvious differences of elemental composition between the films and substrate. Compared with the target, the film has higher content of Na and Ca, as well as lower content of Si and P. （2） The higher the energy density is, the higher attenuation depth is, and the more particles are produced accordingly. Besides, these particles impinge against the substrate with high kinetic energy, leading to the rough films even with the holes. （3） The deposition rate as a function of the fluence （Fig） shows a threshold of about 3J/cm², below which film growth is barely observable. Above the threshold value, the film deposition rate increases with the laser fluence. （4） As the energy density increases, bonding stress between the film and substrate decreases due to the poor quality of the film.Deposition duration decides the thickness of the film, and thus the bonding stress between the film and the substrate. The longer the deposition duration, the thicker the film, and the lower bonding stress between the film and the substrate. The reason is the critical load decreases with increasing coating thickness due to the accumulation of stresses in the coatings.The physical mechanism of PLD is very complex. According to the S-N model, the target is considered as very large plane without thickness. It is difficult to relate the experimental results with the parameters of PLD, and thus optimize the processing of PLD. In this work, the first and second processing of PLD, that is the ablation of the target and the movement of the plasma, were investigated. The equation of ablation rate was deduced on the base of the model of the massive target. Results showed the ablation rate is decided by the physical parameters of the laser and target. In the equation, the relation between the deposition rate and the energy density is consistent with the experimental results. On the base of the dynamic model, the evolutionary process of the plasma in the stage of isothermal expansion and adiabatic expansion was discussed. The equation of the anisotropic distribution of the plasma was deduced by making the equation of the ablation rate as the boundary condition of the plasma movement. Therefore, the processing of the ablation of the target and that of the expansion processing of the plasma were associated.The in vitro bioactivity process of bioglass was studied by detecting the variation of the morphology and structure of the film in the simulated body fluid （SBF）, as well as the ion concentration and PH value of the SBF. It shows that the in vitro bioactivity process of bioglass follows the five stage sequence: （1） Rapid exchange of Na⁺ or K⁺ with H⁺ or H₃O⁺ from solution; （2） Breakage of Si—O—Si bonds and formation of Si—OH （silanols） at the glass solution interface; （3） Condensation and repolymerizationof an SiO₂ rich layer on the surface; （4） Migration of Ca²⁺ and PO₄³⁺ groups to the surface and growth an amorphous CaO-PaOs-rich film; （5） Crystallization of theamorphous CaO-P₂O₅ film by incorporation of OH^- and CO₃^2- anions from solutionto form a apatite layer.Generally, the thermal expansion coefficients （a） of the bioglasses in the first demonstrated system Bioglass? （4545, originally developed by Hench） are higher than that of Ti6A14V, the a mismatch between the substrate and the coatings would cause a residual tensile stress at the interface, resulting in cracks propagation and poor adherence. Therefore, tailoring the a to the substrate through regulating the composition of the bioglass was studied. Another important factor that affects the adhesion is the thickness of the coatings. It is known that the critical load decreases with increasing coating thickness due to the accumulation of stresses in the coatings. It is the precipitation of the crystal apatite layer that assures the successful bonding of the coating to host tissue, indicating the bioactivity of the bioglass. Therefore, the thickness of the glass coatings should exceed a threshold value to develop the complete bioactive process before it is dissolved in vitro. Thus, decreasing the critical thickness of the coatings is also needed to ensure the good adhesion to the substrate.In this work, a kind of bioactive glass coating is deposited on Ti6A14V by pulsed laser. The composition is designed by partial substitution of Mg for Na in Bioglass（?）. The aim is to obtain a bioactive glass coating with good adhesion to Ti6A14V substrate by adjusting its a to the substrate and decreasing the critical thickness of the coating. According to the information provided from the in vitro bioactivity, the onset time for the specimen 1 to precipitate HA is 70h, while for specimen 2, it is 90h. This means that the bioactivity index of specimen 2 is lower than that of specimen 1 and requires more time for bone bonding.Recent investigations proposed a "charged surface" theory to explain the mechanism of the bioactivity for the different bioactive materials. It is believed that the complex process of apatite formation describe above is well interpreted in terms of the electrostatic interaction of the functional groups with the ions in the fluid. The bioactivity mechanism for different biomaterials is similar in essence. The formation of bone-like apatite is induced by functional groups that have a specific arrangement. In the body environment, these functional groups assume a negative charge, and induce apatite formation via the formation of amorphous calcium compound and the subsequent formation of an amorphous calcium phosphate that finally transforms into bone mineral-like apatite.Previous work shows that the effect of MgO on the bioactivity of the bioglass is somewhat controversial. This behavior is thought of as "anomalous property of Mg" and is related to its physical property. It is showed that when the content of MgO for specimen 2 is 10mol%, part of Mg ions would become network former. This result equals to increasing the amount of SiO₂. Thus, Q2 groups can not be detected. The role of Q2 groups in the dissolution rate of the silica through the formation of Si-OH groups at the glass surface has been proved. Thus, the reason that the 10mol% content of MgO can decrease the bioactivity and the thickness of the films was revealed by discussing the characterization of the bioglass structure.更多还原